Stimulation and Excitation of Cardiac Tissues

3 

Stimulation and Excitation of Cardiac Tissues G. NEAL KAY  |  RICHARD B. SHEPARD

T

his chapter reviews the fundamental concepts of artificial electrical cardiac stimulation, including the cellular aspects of myocardial stimulation, the influence of external current on cardiac tissue, waveform, and electrode considerations, clinical applications and considerations, and ongoing research regarding cardiac stimulation. The cardiac myocardium is characterized by the capacity for spontaneous initiation of rhythmic beating, excitability with the opening and closing of membrane ion channels resulting in cardiac action potentials, and the propagation of those action potentials from cell to cell and from chamber to chamber. Tissues with the property of excitability are characterized by having a resting transmembrane potential that undergoes rapid, transient reversal in polarity upon activation.1,2 These electrical properties are precisely coupled to mechanical contraction and relaxation of the myocardium to generate the synchronized pumping action of the heart.

ANATOMIC AND PHYSIOLOGIC CONSIDERATIONS The structural and physiologic properties of the myocardium are essential to excitability and are relevant for understanding artificial electrical stimulation. Three structural and functional aspects of cardiac myocytes will be highlighted3: the lateral sarcolemmal membrane, the intercalated disc (ID) (Fig. 3-1), and the t-tubule system as it interacts with the cisternae of the sarcoplasmic reticulum (SR) and sarcomeres. THE LATERAL MEMBRANE Cardiac myocytes are elongated rod-shaped cells that are electrically coupled to adjacent myocytes, predominantly along their long axis. The lateral membrane of the cardiac myocyte is composed of a phospholipid bilayer, in which the polar, phosphate heads of the phospholipids are oriented toward the periphery of the membrane (both inside and outside of the cell), whereas the hydrophobic aliphatic lipid chains are oriented toward each other in the center of the membrane.3 The polar regions of the bilipid membrane interface with the aqueous environment inside and outside of the cell (Fig. 3-2). The lateral membrane is therefore relatively impermeable to the passage of charged ions. Cholesterol and membrane-spanning ion channel proteins (see below) are also incorporated into the lateral membrane. There are two major structures that can be identified in the lateral membrane, the costamere4,5 and the t-tubule system.6,7 As cardiac myocytes undergo rhythmic contraction with each heartbeat, the structural integrity of the myocardium depends on linking of myocytes to an extracellular matrix. This linking to the extracellular matrix is achieved by the costamere (Fig. 3-3). Costameres are structures oriented transverse to the long axis of the lateral membrane. They are anchored to the Z lines (see Fig. 3-3) of the sarcomere to form physical links between the Z lines, the sarcolemma, and the extracellular matrix. Costameres both sense and transmit external mechanical forces.8-12 Two distinct protein complexes are important for joining the costamere to the extracellular matrix: integrins13 and the dystrophinglycoprotein complex.14,15 Integrins (Fig. 3-4) are heterodimers with both alpha and beta chains. The alpha and beta chains include a long

extracellular domain that binds to laminin in the extracellular space.3 Each integrin chain spans the lateral membrane with a short cytoplasmic tail that links to the actin cytoskeleton. Integrins function both as strong adhesion proteins but also transduce mechanical forces to the nucleus. Dystrophic-glycoprotein complexes (DGC) (see Fig. 3-3) are large proteins that also have extracellular, transmembrane, and intracellular components. The extracellular portion of the DGC binds to laminin, whereas the intracytoplasmic portion binds to actin.3 Thus both integrins and DGC provide structural integrity at the level of the costamere during physical stresses. t-TUBULES AND SARCOPLASMIC RETICULUM t-Tubules are narrow, tubular invaginations of the lateral membrane that interface tightly with the terminal cisternae of the SR to form a dyadic cleft (Fig. 3-5).6,7,16 It is at the level of the dyadic cleft between the t-tubule and the SR that the calcium communication that drives electrical-mechanical coupling occurs.17,18 Although the membrane of the t-tubule and the closely apposed SR membrane are not physically in contact, calcium released in this cleft is responsible for the coupling of electrical excitation with mechanical contraction (EC coupling). EC coupling begins with the opening of L-type calcium channels located within the t-tubules in response to membrane depolarization. The calcium released by the L-type calcium channels diffuses into the dyadic cleft and activates the release of calcium from ryanodine receptors (RyRs) on the SR membrane. The regulation of calcium release from RyRs is tightly regulated by the proteins calsequestrin, junctin, and triadin.19,20 Atrial myocytes have fewer t-tubules than ventricular myocytes whereas sinoatrial (SA) nodal, atrioventricular (AV) nodal, and His-Purkinje cells have almost none.21 INTERCALATED DISC The ability of an electrical stimulus to propagate in a syncytium-like fashion through the myocardium is due to extensive interconnection between myocytes primarily at the level of the ID.3,22 Although action potentials propagate along the lateral membrane to the IDs at the ends of the cell, it is through gap junctions (GJs) in the IDs where propagation primarily occurs from one cell to another. Electrical coupling is far less supported along their transverse axis. Thus propagation through a three-dimensional sheet of myocardium is approximately three times faster along the long axis of fiber orientation than in the transverse axis, a property known as anisotropic conduction. In addition, the safety factor for conduction along the long axis of fibers is higher than transverse to the direction that the fibers are oriented. Three types of adhesion structures are present in IDs.3 These are the fascia adherens junctions, desmosomes, and GJs. The fascia adherens junctions and desmosomes are composed of adhesion molecules that cross the sarcolemma of the ID to bind proteins from adjacent cells in the extracellular space.23 At the cytoplasmic side of the ID, these proteins are bound to the cytoskeleton. Thus fascia adherens junctions and desmosomes anchor the ID of one myocyte to the ID of another myocyte and function to sense mechanical stresses along the longitudinal axis of the myocardium. The transmembrane protein N-cadherin is connected to intracytoplasmic cytoskeletal protein actin by linker proteins such as vinculin and β-catenin in fascia adherens junctions.

61

62

SECTION 1  Basic Principles

Lateral membrane

A

ID Nucleus

B Plakoglobin Desmoplakin Desmocollin Plakophilin Desmosome

Desmin

Desmoglein

ulin

Vinc N-Cadherin

α-Catenin β-Catenin/Plakoglobin

p120

Actin

α-actinin

Fascia adherens junction

Myofibril Z-band Microtubules Connexon

p150

EB1 ZO-1 Protein 4.3

Gap junction

Figure 3-1  The Specialized Substructures of the Intercalated Discs of Cardiac Myocytes. A, Scheme of the syncytial organization of the myocardium. Cardiomyocytes are rod-shaped cells that are electrically and mechanically coupled in the longitudinal axis through a digitated specialized structure, the intercalated disc (ID). B, Diagram focusing on one ID and its three specialized substructures: desmosome, fascias adherens junction, and gap junction. The thick blue line corresponds to the ID membranes and the double green line represents lateral sarcolemma. (From Balse E, Steele DF, Abriel H, et al: Dynamic of ion channel expression at the plasma membrane of cardiomyocytes. Physiol Rev 92(3): 1317-1358, 2012.) PM

Scaffolding protein

Cholesterol Ion channel

Figure 3-2  Sarcolemmal membrane composed of a bilayer of phospholipids with the lipophilic tails oriented toward the interior of the membrane and hydrophilic heads oriented to the periphery. Ion channels composed of macromolecular, complex folded proteins span the membrane and are anchored by scaffolding proteins. Cholesterol is located in the lipophilic interior of the membrane. PM, Plasma membrane. (From Balse E, Steele DF, Abriel H, et al: Dynamic of ion channel expression at the plasma membrane of cardiomyocytes. Physiol Rev 92(3):1317-1358, 2012.)

In desmosomes, the cadherin proteins desmoglein and desmocollin from adjacent cells attach in the intercellular space.24-26 At the cytoplasmic side of the desmosome, the proteins plakoglobin, plakophilin, and desmoplakin connect to the cadherin proteins and anchor these structures to the actin cytoskeleton via the linker protein desmin.27,28 The third adhesion structure of the ID, and the one that is most important for cell-to-cell transmission of action potentials, is the GJ.29-31 GJs allow the diffusion of molecules up to approximately 1000 Daltons between cells, including ions, metabolites, and second messengers.32,33 GJs are highly plastic and undergo remodeling in response to hypertrophy, pulsatile mechanical stress, dilation, and chemical mediators.34,35 GJs are formed by two connexons, each serving as half of a channel (hemichannel) that connects one cell to its neighbor.31,36 Each connexon is formed by six connexins that span the lipid bilayer with a pore in the center.31 Each of the connexin proteins has four membrane-spanning domains with both intracellular N- and C-termini. There are 21 different human genes that encode connexins with each isoform having different conduction properties.37 In addition, connexons can be formed by six of the same connexin proteins (homomeric) or by combinations of different connexins (heterotypic).



CHAPTER 3  Stimulation and Excitation of Cardiac Tissues

63

Lateral membrane 2.

Costamere Extracellular matrix

t-Tubules

Laminin

1.

Integrin complex

Dystrophin-glycoprotein complex

3.

Actin α-actinin

Nucleus

SR

Intercalated disk

Myofibril Z line

Nucleus

SARCOLEMMA Figure 3-3  Specialized Sarcolemmal Domains for Channel Expression in Cardiac Myocytes. Schematic representation of the three well-identified specialized domains for channel expression in cardiac myocyte: (1) the “calcium synapse” composed of t-tubule and terminal cisternae of the sarcoplasmic reticulum (SR), (2) the costamere at the lateral membrane, and (3) the intercalated disc. (From Balse E, Steele DF, Abriel H, et al: Dynamic of ion channel expression at the plasma membrane of cardiomyocytes. Physiol Rev 92(3):1317-1358, 2012.)

The human heart expresses Cx40, Cx43, and Cx45 as the dominant isoforms, with Cx43 being the main connexin in the working ventricular myocardium and a combination of all three connexins being expressed in the atrium.31,38 Cx45 is expressed in the SA and AV nodes whereas Cx40 and Cx45 are important in the His-Purkinje system.39 Connexins Cx40 and Cx43 are highly selective for cations with permeability for both Na+ and K+. The major intracellular ion carrying intercellular current via GJs is K+.40,41 GJs are in mainly the open state though lower levels of conductivity occur with depolarization of the cell membrane. GJs demonstrate voltage gating by both the transjunctional voltage (Vj) or transmembrane voltage (Vm) gradients. Cx45, found in Purkinje fibers, demonstrates significant voltage gating, whereas Cx43, found in ventricular myocardium, shows less voltage gating.31,42 GJs also demonstrate chemical gating with decreased conductance at lower pH and higher intracellular Ca2+ concentrations.31 GJ conductance is also modulated by phosphorylation. Protein kinases tend to enhance GJ conductance, whereas phosphatases tend to reduce conductance.31 In addition, several forms of cardiac disease have been reported to reduce GJ density or cause heterogeneous expression of GJs resulting in slow conduction and increasing the likelihood of unidirectional conduction block. Pharmacologic agents that increase GJ conductance have not become commercially available. The peptides rotigaptide and GAP-134 are among the compounds that have been shown to increase GJ conductance.31

THE CARDIAC ACTION POTENTIAL The action potential (Fig. 3-6) reflects the constantly changing transmembrane voltage gradient across the cardiac sarcolemmal membrane.43,44 The membrane voltage is controlled by increases and decreases in inward and outward currents that flow through specialized membrane ion channels and ion pumps that maintain the resting potential (Fig. 3-7).45 Each of the ion channels has a characteristic time course for opening and closing of the channel, as well as different transmembrane voltage gradients that regulate channel conductance (Table 3-1).45 As the expression of each channel is different in the various regions of the heart, the action potential varies accordingly in the SA node, atria, AV node, His-Purkinje network, and ventricular

myocardium.43,44 There are also differences in channels between the endocardium and epicardium, leading to differences in the action potential across the myocardium. The action potential is characterized by five phases (Fig. 3-8). The normal ventricular action potential is characterized by a steep upstroke followed by a sustained plateau that gradually decays to the resting state, giving it a spike and dome shape. In the resting phase of diastole (phase 4), the concentration of Na+ is maintained higher outside than inside the cell by the Na+-K+ pump (which moves three Na+ ions from the inside to the outside of the cell in exchange for two K+ ions).44 Because the sarcolemmal membrane is otherwise impermeable to Na+ and Ca2+ ions during phase 4, but there is a steady, controlled flow of K+ ions from the inside to the outside of the cell via potassium channels (IK1), the membrane is polarized to approximately −90 mV with a net negative charge inside the cell. Similar to IK1, the acetyl choline regulated K+ current (IKAch) is operative during periods where the membrane potential is in its resting state.46 Depolarization of the membrane to a threshold potential of approximately −70 mV results in the rapid opening of voltage-gated Na+ channels (INa) (Fig. 3-9) such that Na+ ions flood into the cell resulting in depolarization of the membrane to a net positive potential (phase 0).47,48 The Na+ channels go from the open to the closed state within 1 msec as the membrane depolarizes. Depolarization of the membrane opens both L-type Ca2+ channels (inward current)49,50 and a series of outward conducting K+ currents (IKUR, IKto, IKr, and IKs).44,46 The overshoot of the membrane potential from negative to positive, which occurs with the influx of Na+ ions, initiates opening of a rapidly inactivating, transient outward K+ current (Ito) during phase 1. Following phase 1, the cell enters phase 2 or the plateau phase of the action potential, which is characterized by a period of 200 to 300 msec during which the net current carried by inward Ca2+ and outward K+ ions is relatively balanced. During phase 2 the membrane is depolarized and the cell is unexcitable for stimulation regardless of stimulus intensity (absolute refractory period). The depolarized membrane opens L-type Ca2+ channels (ICa) allowing Ca2+ release from t-tubules into the dyadic cleft that results in Ca2+-induced Ca2+ release from RyRs in the SR (Fig. 3-10).6,18 The majority of calcium is stored within the Text continued on p. 68

64

SECTION 1  Basic Principles

ECM

A

α7

B

α7

β1D

Nucleus

Mitochondria

Intercalated disc

Intercalated disc

Laminin

α7

β1D

β1D

Costamere F-actin Sarcomere Z-line

Z-line

Z-line

ECM α7

C

α7

β1D

Mechanical stress

β1D

α7

D

Laminin

β1D

α7

Mechanical stress

β1D

Laminin

Nck2 Paxillin ILK PINCH

Tln

Kindlin2 Wech

Vcl/MtVcl β-parvin

α-actinin α-actinin

PINCH RSU1 PP1α Akt

F-actin

F-actin

Melusin Paxillin ILK β-Parvin

JNK

FAK Paxillin Hsp90 IQGAP1 FAK Pyk2 Src ERK, Akt

ERK, Akt

Tln1

ERK, p38, Akt

NFKB Growth, Hypertrophy, Survival

Figure 3-4  Role of Integrins in the Myocyte. A, Integrins contact the extracellular matrix (ECM) and transverse the sarcolemmal membrane at the costamere, where they connect to Z lines. Integrins bridge the ECM to the sarcomere and may transmit information to the nucleus. Integrins are also located in the intercalated disc of the cardiac myocyte (CM). B, Enlarged portion from A showing that integrin α7β1D, a predominant laminin receptor in the mature CM, can bridge the ECM to the sarcomere across the costamere, where the integrin forms a complex of structural and signaling proteins. C, Some important structural connectors that bind to the tails of the integrin receptor in the CM. Integrins connect and aggregate a range of adapter and signaling proteins, such as integrinlinked kinase (ILK), focal adhesion kinase (FAK), vinculin (Vcl), talin (Tln), kindlin, cysteine-histidine–rich protein (PINCH), parvin, actinin, and even actin. This allows both bridging of ECM to the cytoskeleton and perhaps sarcomere and also allows propagation of signals bidirectionally across the cell membrane. D, Illustration of some integrin signaling pathways in the CM. When integrins bind ECM ligands (e.g., laminin), they assemble proteins on their intracellular cytoplasmic domains and ultimately orchestrate signals down a variety of pathways. Of note is that mechanical events occurring outside the CM can stimulate the ECM-integrin interaction and also lead to intracellular biochemical changes, a process termed mechanotransduction. See text for details. MtVcl, Metavinculin; PP1α, protein phosphatase 1 α; Pyk2, proline-rich tyrosine kinase-2; RSU1, Ras-suppressor-1. (From Israeli-Rosenberg S, Manso AM, Okada H, Ross RS: Integrins and integrinassociated proteins in the cardiac myocyte. Circ Res 114(3):572-586, 2014.)



CHAPTER 3  Stimulation and Excitation of Cardiac Tissues

A

B

Surface membrane Cytosol

t-Tubule membrane

65

Longitudinal section t-Tubule NKA

jSR

Actin

jSR

Network SR

NCX

RyR (feet)

Myosin

RyR

Z line I-band

LTCC

Z line A-band

Na+-channel NCX

I-band

C

NKA

High Ca2+

Low Ca2+

Figure 3-5  Structural Basis for EC-Coupling in Ventricular Cardiomyocytes. A, t-Tubules are invaginations of the cell membrane, and these invaginations allow close interaction between the sarcoplasmatic reticulum (SR) and the sarcolemma. The SR is divided into junctional SR (jSR) close to the t-tubules and network SR between two t-tubules. The myofilaments with actin and myosin are regularly organized in sarcomeres between two Z lines, surrounded by the SR. B, The dyadic clefts are close connections between the junctional SR and the t-tubule membrane. L-type Ca2+-channels (LTCCs) in the t-tubular membrane are closely apposed to groups of ryanodine receptors (RyRs), creating the structural basis for Ca2+ sparks during membrane depolarization. A subset of such LTCC–RyR-couplons has also been shown to colocalize with NCX, supporting that NCX can promote CICR in cardiomyocytes. However, the localization of voltage-gated Na+ channels, NCX and Na,K-pump (NKA) is not known in detail, and the positions of these membrane proteins relative to the dyadic cleft are of crucial importance for cardiac EC-coupling. Experimental data indicate functional roles for both voltage-gated Na+-channels and NKA α2-isoforms in close proximity to the LTCC–RyR-couplon. C, The concentration gradient of Ca2+ in the cytosol is a major regulator of cardiac EC-coupling. CICR rapidly increases the local Ca2+ concentration in the dyad (bright color), leading to Ca2+ diffusion out of the dyadic cleft to increase the global Ca2+ levels in the cell (dark color), which triggers contraction. CICR, Calcium-induced calcium release; NCX, sodium calcium exchanger. (From Aronsen JM, Swift F, Sejersted OM: Cardiac sodium transport and excitation–contraction coupling. J Mol Cell Cardiol 61:11-19, 2013.)

SECTION 1  Basic Principles

Na+

Ca2+

50

Na+

Phase 1 Phase 2

–50

Phase 0

0 Threshold

Transmembrane potential (mV)

66

K+

Phase 3

Phase 4 –100 100

200

300

400

500

Time (msec) Figure 3-6  Microelectrode Recording of Transmembrane Potential (Vm) From Left Ventricular Endocardium of Human Heart. In phase 0 (depolarization), sodium ions (Na+) rapidly enter the cell through fast channels. In phase 1, the initial repolarization is primarily the result of activation of a transient outward potassium ion (K+) current and inactivation of the fast Na+ current. In phase 2 (plateau), the net current is very small, although the individual Na+, Ca2+, and K+ currents are about an order of magnitude larger. Phase 3 (final repolarization) completes the cycle, with the Na+-K+ pump bringing the membrane potential to a stable point at which inward and outward currents are again in balance. During phase 4, the cell is polarized and gradually undergoes slow depolarization. (From Stokes K, Bornzin G: The electrode-biointerface [stimulation]. In Barold SS, editor: Modern cardiac pacing, Mt. Kisco, NY, 1985, Futura, 1985, pp 33-78.)

Ca2+

K+

Figure 3-7  Ion Channels Underlie Cardiac Excitability. Key ion channels (and electrogenic transporter) in cardiac cells. K+ channels (green) mediate K+ efflux from the cell; Na+ channels (purple) and Ca2+ channels (orange) mediate Na+ and Ca2+ influx, respectively. The Na+/ Ca2+ exchanger (red) is electrogenic, because it transports three Na+ ions for each Ca2+ ion across the surface membrane. (From Marban E: Cardiac channelopathies. Nature 415(6868):213-218, 2002. © Nature Publishing Group, http://www.nature.com).

TABLE Major Ionic Currents Contributing to the Cardiac Action Potential* 3-1  Current INa

Description Sodium current

Phase 0

Activation Mechanism Voltage, depolarization

Clone Nav1.5

Gene SCN5A

ICa,L ICa,T Ito,f Ito,S

Calcium current, L-type Calcium current, T-type Transient outward current, fast Transient outward current, slow

1 2 1 1

Voltage, depolarization Voltage, depolarization Voltage, depolarization Voltage, depolarization

CaV1.2 Cav3.1/3.2 KV 4.2/4.3 KV 1.4/1.7/3.4

IKUR

Delayed rectifier, ultrarapid

1

Voltage, depolarization

KV1.5/3.1

IKr IKS IK1 IKATP IKAch IKP

Delayed rectifier, rapid Delayed rectifier, slow Inward rectifier ADP activated K+ current Muscarinic-gated K+ current Background current

3 3 3&4 1&2 4 All

Voltage, depolarization Voltage, depolarization Voltage, depolarization [ADP]/[ATP] increase Acetylcholine Metabolism, stretch

IF

Pacemaker current

4

Voltage, hyperpolarization

HERG KVLQT1 Kir 2.1/2.2 Kir 6.2 Kir 3.1/3.4 TWK-1/2 TASK-1 TRAAK HCN 2/4

CACNA1C CACNA1G KCND2/3 KCNA4 KCNA7 KCNC4 KCNA5 KCNC1 KCNH2 KCNQ1 KCNJ2/12 KCNJ11 KCNJ3/5 KCN1/6 KCNK3 KCNK4 HCN 2/4

From Grant AO: Cardiac ion channels. Circ Arrhythm Electrophysiol 2(2):185-194, 2009. *The name of the current, the ion responsible for the current, the phase of the action potential during which the current is active, the mechanism of activation of the channel proteins, the clone name, and the gene name are provided.



CHAPTER 3  Stimulation and Excitation of Cardiac Tissues

Early repolarization (1)

Plateau (2) ICaL ↓ IKs ↑

0

IKr ↑

mV

INa ↓

Upstroke (0)

Final repolarization (3)

IKACh ↑

If ↓

IK1↑ IKACh ↑

Pacemaker function

IK1 ↑ –80

A

Figure 3-8  The Cardiac Action Potential Is Depicted for a Typical Atrial or Ventricular Cell. Panel A illustrates the transmembrane potential as well as the inward currents (downward arrows) and outward currents (upward arrows). Panel B represents the onset and offset of each major current during the action potential. The action potential is the electrical potential differences across cardiac cell membranes as a function of time and is controlled by the flow of ions through ion channels (gray boxes). Inward currents are depicted with arrow down and outward currents with arrow up. In the resting state (phase 4), ion pumps maintain a higher Na+ and Ca2+ concentration outside the cell and a higher K+ concentration inside the cell. The cell is impermeable to Na+ and Ca2+ but is permeable to K+ through a channel designated IK1. As a result, positively charged K+ ions flow slowly out of the cell, keeping the interior electrically negative. When the membrane potential reaches a crucial threshold voltage, there is a rush of Na+ ions into the atrial and ventricular cells. The primary inward current for sinoatrial and atrioventricular nodal cells is carried by Ca2+. As positively charged ions enter the cell, the transmembrane potential is rapidly depolarized (phase 0) with an overshoot to a somewhat positive potential (phase 1). Following depolarization of the membrane, Ca2+ channels open, allowing Ca2+ to enter the cell through the L-type Ca2+-channel (ICa,L). This permeability to Ca2+ maintains the cell membrane in a depolarized state (phase 2) and allows Ca2+ to interact with the contractile proteins to induce mechanical contraction of the myocardium. During phase 2, the inward Ca2+ and outward K+ currents are approximately equal and the membrane potential remains flat for 200 to 300 msec. Return of the membrane to its repolarized state is produced by an outward flow of K+ as a series of fast and slowly inactivating K+ currents are active. Rapidly activating K+ currents such as Ito and IKur are responsible for very rapid initial repolarization (phase 1), whereas slower components (IKr and IKs) produce the final repolarization (phase 3). The time from AP upstroke and the return to the resting potential is referred to as AP duration (APD). During the depolarized phase of the AP, myocytes cannot be induced to trigger another AP upstroke, a period of relative inexcitability called the refractory period. Cells in the SA and AV nodes have a “pacemaker current,” If, that induces a gradual depolarization during phase 4 with the triggering of a new AP when the threshold potential is reached. The Na+ and Ca2+ that entered the cell during the AP are extruded, whereas K+ is brought back inside the cell by exchanging intracellular Na+ for extracellular K+ through the Na+,K+-pump). Na+ is exchanged for Ca2+ by the Na+,Ca2+exchanger (NCX). (From Nattel S, Carlsson L: Innovative approaches to anti-arrhythmic drug therapy. Nat Rev Drug Discov 5(12):1034-1049, 2006.)

Ito ↑

IKur ↑

APD

Resting (4)

NCX ↓

0.2 seconds

IK1 INa Ito IKur ICaL IKr IKs NCX If IKACH

B

Domain II

Domain I

67

Domain III

Domain IV

NH2

β

S1 S2 S3 S4 S5

S6

S1 S2 S3 S4 S5

S6

S1 S2 S3 S4 S5

* COOH

S6

IF

S1 S2 S3 S4 S5

M

T

Inactivation particle

NH2

*

*

S6

*

COOH Docking site

Figure 3-9  Schematic Representation of the a- and b-Subunits of the Voltage-Gated Sodium Channel. The four homologous domains (I-IV) of the a-subunit are represented; S5 and S6 are the pore-forming segments, and S4 is the core of the voltage sensor. In the cytoplasmic linker between domains III and IV, the IFMT (isoleucine, phenylalanine, methionine, and threonine) region is indicated. This is a critical part of the “inactivation particle” (inactivation gate), and substitution of amino acids in this region can disrupt the inactivation process of the channel. The “docking site” consists of multiple regions that include the cytoplasmic linker between S4-S5 in domains III and IV and the cytoplasmic end of the S6 segment in domain IV (*). Depending on the subtype of b-subunit, they could interact (covalently or noncovalently) with the a-subunit. (From Savio-Galimberti E, Gollob MH, Darbar D: Voltage-gated sodium channels: biophysics, pharmacology, and related channelopathies.  Front Pharmacol 3:124-143, 2012.)

68

SECTION 1  Basic Principles

Extracellular

ICaL

Ca2+

Ca2+

NCX

Na+

PKA CamKII

Calstabin 2

PP1

RyR2

SERCA

P –PLB

P

P

PLB– P

Ca2+

ATRIAL VERSUS VENTRICULAR ACTION POTENTIALS

Ca2+

P –PLB Cytosol

SR

PP1

P

membrane back to its polarized state, K+ currents with delayed rectification are opened in response to depolarization of the membrane, which inactivate relatively rapidly or slowly (IKr and IKs) (phase 3 or repolarization phase).44,46 Toward the end of phase 3 the cell gradually becomes excitable for stimulation, although a higher stimulus voltage is required to initiate firing of another action potential (relative refractory period). As the transmembrane potential becomes progressively more negatively repolarized, the IK1 current is reinitiated and the membrane potential approaches its resting value and normal excitability is reestablished.46 The time from the onset of phase 0 until reestablishment of the resting membrane potential and normal excitability is the action potential duration (APD). The membrane potential during diastole is maintained by the Na+/K+ pump (NKA), which removes three Na+ ions from the inside of the cell in exchange for two K+ ions, whereas the Na+/Ca2+ (INCX) exchange pump which extrudes Ca2+ from the cell is an important regulator of intracellular calcium during diastole.

P PLB– P SERCA

Figure 3-10  During depolarization of the cell membrane, calcium enters the cell via activated L-type Ca2+-channels. The resulting Ca2+ influx promotes intracellular Ca2+ release from subcellular stores in the sarcoplasmic reticulum (SR) by Ca2+-induced Ca2+ release, which greatly amplifies the initial signal. Ca2+-induced Ca2+ release occurs via specialized Ca2+-release channels, also called ryanodine receptors (RyRs). The RyR2 constitutes part of a large macromolecular complex of key accessory proteins including calmodulin, calstabin 2, protein kinase A (PKA), Ca2+/calmodulin-dependent protein kinase (CaMKII) and protein phosphatases 1 (PP1) and 2A. Calmodulin is a key Ca2+-binding protein that regulates the action of key intracellular enzymes such as the kinase CaMKII and the protein phosphatase calcineurin. Protein kinase A and CaMKII are both stimulated by adrenergic activation and phosphorylate key intracellular regulatory proteins. Calstabin 2 binds to and stabilizes the open and closed states of the RyR. SR Ca2+ stores are determined by the rate of SR Ca2+ uptake and the rate of Ca2+ release. Ca2+ uptake occurs via the SR Ca2+-ATPase, SERCA (cardiac form is SERCA2A). SERCA function is negatively regulated by phospholamban (PLB), but PLB phosphorylation removes this inhibitory influence. Abnormalities in intracellular Ca2+ handling occurs in conditions such as congestive heart failure, ischemic heart disease, myocardial hypertrophy and atrial fibrillation, as well as in inherited arrhythmogenic diseases such as catecholaminergic polymorphic ventricular tachycardia (CPVT) in which obvious structural heart disease is absent. Spontaneous diastolic Ca2+ release triggers Na+,Ca2+ exchange NCX), which mitigates Ca2+ loading by extruding Ca2+ in exchange for extracellular Na+. NCX carries three Na+ ions in for each single Ca2+ ion extruded, and therefore causes movement of one extra positive ion into the cell for each functional cycle. (From Nattel S, Carlsson L: Innovative approaches to anti-arrhythmic drug therapy. Nat Rev Drug Discov 5(12):1034-1049, 2006.)

SR and its release into the cytosol triggers the initiation of myocyte contraction. In order for the myocyte to relax from its contracted state, the SR must then remove and sequester calcium from the cytosol via SR Ca2+ ATPase (SERCA).19 Intracellular calcium is buffered by the Ca2+ binding proteins calmodulin (CaM) and troponin.6,16,18 To bring the

Compared with the typical spike and dome shape of the ventricular action potential, atrial action potentials typically have a more triangular shape (Fig. 3-11).43,51-53 The resting membrane potential (phase 4) in the atrium is typically −65 to −80 mV compared with −80 to −90 mV in the ventricle. The lower degree of resting membrane polarization in the atrium is related to significantly lower density of the background inward rectifier, IK1. In addition, the upstroke velocity of phase 0 is lower in the atrium (150-300 V/S) than the ventricle (300-400 V/S). The lower upstroke velocity in the atrium is likely related to the relatively depolarized resting potential as compared with the ventricle which reduces INa. In contrast, inward Ca2+ current is increased in the atria as compared with the ventricles.53 The L-type Ca2+ current and Ca2+-induced Ca2+ release are responsible for the plateau phase of the atrial action potential (AP), whereas the T-type Ca2+ current is not present in atrial myocytes. Phase 1 of the atrial AP is caused by increased fast component of the transient outward K+ current Ito,f. The slow component of the transient outward current (Ito,S) is absent in the atrium.54 A major difference between the atria and ventricles is the outward ultrarapid K+ current (IKUR), which is important in the atria but absent in the ventricle. In addition, the delayed rectifier currents IKr and IKs are significantly reduced in the atria, likely related to the triangular, more negative plateau phase. The main route of Na+ efflux from atrial myocytes is the Na+/K+ exchange pump, whereas the Na+/ Ca2+ exchange pump is responsible for Ca2+ extrusion.55 SA NODAL IMPULSE GENERATION In the human heart, pacemaking normally occurs in specialized cells within the SA node that have limited contractile function. Secondary pacemaker cells with slower rates are found in the AV node and HisPurkinje system. The SA nodal cells are located within the crista terminalis with a higher mixture of atrial cells near the periphery of the node and higher density of specialized pacemaker cells in the interior of the node. The pacemaker cells have the highest frequency, due to the most rapid rise in membrane potential to threshold during diastolic depolarization in phase 4 (Fig. 3-12). The amount of collagen in the SA node is higher than for normal atria and increases with age. Coupling within the SA node is relatively weak and electrotonic inhibition of SA automaticity by the atrial myocardium is observed. The rate of firing in the SA node is closely controlled by a variety of factors in response to demands such as neurotransmitters, circulating hormones, and stretch.56,57 The underlying mechanisms controlling automaticity include specialized ion channels in the membrane,58 as well as a Ca2+ “clock” that determines Ca2+ cycling (Fig. 3-13).59,60 It has been demonstrated that the site of pacemaking activity may shift markedly from lower in the crista terminalis at slow rates to higher at faster rates.61



CHAPTER 3  Stimulation and Excitation of Cardiac Tissues

69

Sinus node Atrial muscle AV node Common bundle Bundle branches Purkinje fibers Ventricular muscle R T

P

U

Q Time (msec) 0

100

200

S 300

400

500

600

700

Figure 3-11  Action Potentials From Different Cell Types in the Heart. Specialized pacemaker and conduction cell types have distinct action potential characteristics. Activation times for a normal heartbeat demonstrate the spread of activation throughout the whole heart. AV, Atrioventricular. (From Malmivuo J, Plonsey R: Bioelectromagnetism: principles and applications of bioelectric and biomagnetic fields, New York, 1995, Oxford University Press.)

The predominant pacemaker currents in the SA node flow through hyperpolarization-activated cyclic nucleotide-gated channels (HCN), L-type and T-type Ca2+ channels (ICa,L and ICa,T), and acetyl cholineactivated channels.50,57 The SA node has reduced inward rectifier (IK1) current density and is devoid of connexin Cx40 but has more expression of the low conductance Cx45. Cells at the periphery of the SA node express Cx43 allowing for conduction to the atrium. The gene HCN4 encodes membrane channels that carry the hyperpolarizingactivated, or “funny,” current (If ).62 The If current is a mixed cationic current carried by both Na+ and K+ ions. It is activated by membrane hyperpolarization with a threshold between −50 and −65 mV. These channels have a binding site for cAMP, resulting in acceleration of the rate of spontaneous diastolic membrane depolarization in SA nodal cells. In contrast, muscarinic agonists reduce the funny current thereby decreasing the heart rate. The upstroke of the SA node action poten­tial (phase 0) is predominantly determined by ICa2+ with a much lower upstroke velocity than in the ventricles.57 In addition, there is very little Ito within the SA node with repolarization determined by the delayed rectifiers (IKr and IKS). The plateau phase is markedly reduced with a characteristic AP morphology shown in Figure 3-11. The voltage-dependent calcium currents ICa2+ ,L and ICa2+ ,T carry the depolarizing current in the SA and AV nodes during phase 0. However, spontaneous diastolic release of cytosolic Ca2+ in phase 4, which acts as a Ca2+ clock, is essential for pacemaker automaticity.60 This spontaneous cytosolic Ca2+ increase of the calcium clock peaks during late diastolic depolarization leads to opening of the membrane calcium currents. This spontaneous release of Ca2+ is from the SR and depends on RyR activation.63 The NCX exchange pump is also important in diastolic depolarization because it is a net inward current.55 Although

the relative importance of the transmembrane If current and spontaneous oscillations of the calcium clock has been debated, the available evidence suggests both are important for pacemaking activity.60 Both mechanisms are highly responsive to changes in beta adrenergic stimulation via G-protein (Gs) activation and cAMP/phosphokinase A phosphorylation. ATRIOVENTRICULAR NODE PROPAGATION The AV node is composed of three types of cells: spindle-shaped cells within the compact AV node near its junction with the His bundle; transitional cells having features that are intermediate between atrial myocytes and spindle-shaped cells typical of the compact AV node; and left and right inferior extensions of the AV node that approach the compact AV node from the left and right atria (Fig. 3-14).64-66 The unique aspect of AV nodal conduction is that there is significant delay in propagation of impulses from the atria to the His bundle, allowing time for the mechanical transport of blood through the AV valves. Transitional cells approach the compact AV node (AVN) from the limbus of the fossa ovalis, spreading across the tendon of Todaro. The inferior extension of the AVN approaches the compact AV node inferiorly from the left atrium and from fibers located parallel to the septal leaflet of the tricuspid valve.64 These inferior extensions have spontaneous depolarizations and pacemaker activity expressing If current, features not found in transitional cells. As compared with atrial myocytes, transitional cells of the AV node express connexin Cx43, whereas the compact AV nodal cells have minimal Cx43 expression. The action potential of transitional cells has high upstroke velocity typical of INa, whereas the compact AV nodal cells and inferior extension cells have a lower upstroke velocity related to ICa2+ .

70

SECTION 1  Basic Principles

SAN

CL APD

HIS-PURKINJE SYSTEM

E1h LDD

EDD

diastolic membrane potential with faster pacing rates leading to slowed conduction velocity as the relative refractory period is progressively encroached.73,74 Dual AV nodal pathways are common and have been demonstrated in both experimental animal models65,75 and in humans where AV nodal reentrant tachycardia is a common form of supraventricular tachycardia.76

20 mV 50 ms AVN

20 mV 50 ms PFN

20 mV 250 ms Figure 3-12  Recordings of automaticity and action potential waveforms of isolated mouse SAN, AVN, and PFN cells. APD, Action potential duration; AVN, atrioventricular node; Cl, cycle length; EDD, exponential part of the diastolic depolarization; Eth, action potential threshold; LDD, linear part of the diastolic depolarization; PF, Purkinje fibers; SAN, sinoatrial node. (From Mangoni ME, Nargeot JL: Genesis and regulation of the heart automaticity. Physiol Rev 88(3):919-982, 2008.)

The compact AVN expresses low conductance Cx30.2 and Cx45, whereas Cx43 is much less expressed in the AVN.67 As compared with the His bundle (which has large amounts of INa), there is much less INa in the AVN. These differences in connexins and sodium channels explain the rapid conduction typical of the His bundle and slow conduction typical of the compact AVN. Based on characteristics of the action potential, AV nodal cells can be distinguished as atrionodal (AN), nodal (N), and nodo-His (NH) cells.68-70 Compared with atrial and AN cells, N cells in the compact AVN have a less negative resting potential, slower upstroke velocity, smaller amplitude, and pacemaker activity. NH cells demonstrate a more negative resting potential, higher upstroke velocity, and a more prolonged plateau phase of the action potential.71 The pacemaker activity of the AV junction occurs in the NH cells and His bundle where If and the calcium clock are more highly expressed (Figs. 3-15 and 3-16).65 Decremental conduction, whereby conduction delay is greater at faster pacing rates, is characteristic of AV nodal conduction.71,72 Decremental conduction can be explained by a gradual increase in the

The insulated His bundle, right and left bundle branches, and the branching Purkinje network are designed to ensure an effective sequence of ventricular contraction that proceeds rapidly and evenly from apex to base (Fig. 3-17).77 In the human heart there are very few Purkinje fibers that penetrate from the endocardium to the epicardium.78 The Purkinje network includes free-running fibers known as false tendons and branching subendocardial fibers. The action potentials of the Purkinje network are insulated over large distances and activate the ventricular myocardium at distinct Purkinje-muscular junctions.79 Purkinje fibers are enriched by the high conductance connexin Cx40 and the INa channel, Nav1.5. Purkinje conduction is minimally responsive to autonomic stimulation. The AP duration of Purkinje cells is longer than for the ventricular myocardium, with a more rapid rate of rise and longer plateau.79 The plateau potential of Purkinje fibers is lower than for the ventricular myocardium with no significant difference in resting potential. The Purkinje-ventricular muscle junctions are formed by specialized cells with high resistance, which allows the localized stimulation of the ventricular myocardium while limiting retrograde conduction from the muscle to the Purkinje fibers. The Purkinje-ventricular muscle junctions demonstrate discontinuous conduction properties with significant conduction delay.80 Purkinje cells demonstrate automaticity at a lower frequency than in the SA node and demonstrate diastolic If current.78 Similar to the SA node, Purkinje cells also demonstrate spontaneous intracellular Ca2+ oscillations that may be the primary cause of pacemaker activity. Purkinje cells lack t-tubules but have junctional and nonjunctional types of SR.79 L-type calcium channels open in response to depolarization resulting in Ca2+-induced Ca2+ release as calcium binds to RyRs, triggering release from the SR. Thus as compared with ventricular myocytes, Ca2+ release in Purkinje cells relies less on voltage-gated Ca2+ channels and more on Ca2+-induced Ca2+ release from the SR.78,79

CONCEPTS RELATED TO ELECTRICAL STIMULATION ELECTRIC FIELDS AND CHARGE Physical objects acquire an electric charge when they have a net excess or deficit of electrons relative to the number of protons. When the number of electrons exceeds the number of protons, the object is said to have a negative charge, whereas a net deficit of electrons results in the object acquiring a positive charge. An electrically charged object is surrounded by an electric field such that charged objects act at a distance on other objects having similar or opposite charge. The strength of the electric field is related to the magnitude of its charge. A gravitational field also acts at a distance, but an electric field has an important difference, polarity. Thus electric fields have directionality with the convention that field lines are drawn away from positively charged and toward negatively charged objects. The electric fields surrounding charged objects interact with each other such that the presence of two electrically charged objects results in a force that either attracts the two objects (if they have opposite charges) or repels the two objects (if they have the same qualitative charge). This attractive or repulsive force acts at a distance, without requiring that the charged objects be in contact. The magnitude of the attractive or repulsive force (F) is given by Coulomb’s law: F=k

Q1Q2 r2



CHAPTER 3  Stimulation and Excitation of Cardiac Tissues

71

NE β1-AR

AC

+

+

Gβγ Gαs

cAMP

Ist

+

? +

PKA

+

+ P

+

P

SERCA

ERG1

P

PLB Ca2+

Cav1.3 Cav3.1

HCN

Ca2+

RyR

Ca2+

NCX

SR

?

Figure 3-13  Summary of the Ionic Mechanisms Contributing to the Diastolic Depolarization in a SAN Pacemaker Cell. Voltage-dependent ion channels and the proposed “Ca2+ clock” are represented together. Possible interactions between these mechanisms are indicated. In pacemaker cells, high basal cAMP-mediated protein kinase A (PKA)-dependent phosphorylation stimulates a perpetual “free running” Ca2+ cycling by pumping Ca2+ into the sarcoplasmic reticulum (SR) via SR Ca2+ ATPase 2 (SERCA2) and L-type calcium induced calcium release (LCICR) via ryanodine receptors (RyRs). PKA also stimulates Ca2+ entry through Cav1.3-mediated ICa,L. The thick red line indicates the persistent spontaneous Ca2+ cycling. The possibility that Cav3.1-mediated ICa,T can contribute to replenishment of SR Ca2+ stores is suggested. Spontaneous LCICR from SR is linked to the diastolic depolarization via Ca2+ activation of inward INCX current. Direct cAMP-dependent activation of hyperpolarizationactivated cyclic nucleotide-gated (HCN) channels or cAMP-mediated, PKA-dependent phosphorylation of Cav1.3 channels and Ist (dashed lines) strongly stimulates the pacemaker cycle driven by the “membrane ion channels clock” (MCC). It is conceivable that the MCC can entrain the intracellular Ca2+ clock because of the dependency of SR Ca2+ content from voltage dependent calcium currents (VDCCs). On the other hand, the Ca2+ clock may trigger oscillations of the membrane voltage and initiate normal pacemaking via the MCC. It is thus likely that under physiologic conditions the MCC and the Ca2+ clock mutually entrain one another. (Adapted from Vinogradova TM, Lyashkov AE, Zhu W, et al: High basal protein kinase A-dependent phosphorylation drives rhythmic internal Ca2 store oscillations and spontaneous beating of cardiac pacemaker cells. Circ Res 98(4):505-514, 2006. In Mangoni ME, Nargeot JL: Genesis and regulation of the heart automaticity, Physiol Rev 88(3):919-982, 2008.)

where Q1 and Q2 are the magnitude of the charges (measured in coulombs), r is the distance separating the charged objects, and k is a constant whose magnitude depends on the medium between the two charged objects. According to Coulomb’s law, the force attracting or repelling charged objects increases with the magnitude of their charges and decreases with the square of the distance that separates them. POTENTIAL DIFFERENCE Electrical potential energy is possessed by a charged particle by virtue of the magnitude of the charge and the position of the particle in relation to other charged particles. That is, a charged object in space has potential energy because of the forces that other charges exert on that object. When two charged objects repel one another, energy is required to move them closer to each other, thereby increasing the potential energy. However, if the charged object is attracted to another charged object, potential energy will be converted into kinetic energy as the charged objects move closer together. Thus the force generated by the interaction of electrical charge is electromotive, tending to move charged objects either toward one another if they have opposite polarity or away from one another if they have the same polarity. Voltage is the potential energy per unit charge for a charged object in an electric field. Thus in order to move a charge from point A to point B, the work required (measured in joules) is calculated by multiplying the charge, Q, by the voltage difference between points A and B, VAB. Because voltage is potential energy per unit of charge, it is measured in terms of joules/coulomb, as follows:

1 volt = 1 joule coulomb When one refers to “voltage,” the reference is to a difference in potential between two points in space. ELECTRIC CURRENT An electric current, I, is present when there is a movement of electric charge. Although electric charge may move through several mechanisms, for clinical purposes, electric charge is usually carried by the flow of electrons through a wire (e.g., pacemaker lead) or by the movement of ions in blood, interstitial fluid, across cell membranes, or within the cytoplasm of cells. By historical convention, current is considered to flow in the direction that positive charges would move. In reality, however, an electric current in a wire is carried by the movement of electrons that are negatively charged. For clarity in this chapter, current is stated in terms of electron or ion motion. Because electric current is the movement of charge, it is measured in terms of coulombs per second (It = dQ/dt), with 1 ampere of current equal to the movement of 1 coulomb/sec. The electromotive force for cardiac pacemakers or implantable cardioverter-defibrillators (ICDs) is determined by the chemistry of its battery. For lithium-iodine batteries at beginning of life, the chemical reaction generates approximately 2.8 volts of electromotive force. The total amount of charge that is available in the battery is measured in terms of the amount of current that can be provided multiplied by the duration that the current can be sustained. For pacemaker or ICD batteries the amount of charge that can be stored in the battery is

72

SECTION 1  Basic Principles

Transitional tissue

A

INE (end) - upper nodal tissue

E INE (end) - lower nodal tissue

Densely packed atrial muscle

F

B Ventricular muscle

C

Penetrating bundle (start) - compact node

G INE (middle)

D

Penetrating bundle (start) - lower nodal tissue

H 100 µm

Figure 3-14  High-Magnification Images of Different Myocyte Types at the Atrioventricular Conduction Axis. A, Transitional tissue. B, Atrial muscle. C, Ventricular muscle. D, Middle of INE. E, Upper nodal tissue of end of INE. F, Lower nodal tissue of end of the INE. G, Compact node at the start of the penetrating bundle. H, Lower nodal tissue at the start of the penetrating bundle. Sections stained with Masson’s trichrome are shown on the left, and adjacent Cx43-labeled sections are shown on the right. (From Li J, Greener ID, Inada S, et al: Computer three-dimensional reconstruction of the atrioventricular node, Circ Res 102(8):975-985, 2008.)



CHAPTER 3  Stimulation and Excitation of Cardiac Tissues

Model Atrial muscle

40

30 20

mV

15

10

5

25

–40

mV 30 2 3

4

20

50

45 50

45

Dorsal

55

40

60

Cranial

Caudal

0

4

0 –40 5

0 –40 –80

Ventral

Ventricular muscle

3

–40

–80 40 mV

50

55

65 70

–40

–80 40

75 mV

5

40 15

mV

1

2

0

–80 40

30

1

0 –80 40

10

De Calvalho and De Almeida

35

Stimulating electrodes

73

50 ms

2 mm Figure 3-15  Structure-Function Relationships of the Atrioventricular Node (AVN). Left, anterograde conduction as calculated using a monodomain model. The preparation (shown on the left) was stimulated at crista terminalis. The activation sequence is shown as isochrones at 5-msec intervals. Arrows highlight conduction pathway. Calculated intracellular action potentials recorded at sites 1 to 5 (see preparation) and comparable intracellular action potentials recorded experimentally are shown on the right. (From Li J, Greener ID, Inada S, et al: Computer threedimensional reconstruction of the atrioventricular node. Circ Res 102(8):975-985, 2008.)

22 Atrial muscle

ms Dorsal Caudal

Cranial

Leading pacemaker site Tricuspid valve Ventricular muscle

Ventral

2 mm

0

Figure 3-16  Activation Sequence of AVN During Pacemaker Activity. The activation map is shown as a color contour map and is superimposed on a grayscale image of the model, which in turn, is superimposed on a photograph of the preparation. Note the earliest  activation in blue with rapid spread of activation in the cranial direction toward the bundle of His. (From Li J, Greener ID, Inada S, et al:  Computer three-dimensional reconstruction of the atrioventricular node. Circ Res 102(8):975-985, 2008.)

74

A

SECTION 1  Basic Principles

B

C Figure 3-17  Purkinje Fibers in the Sheep Heart. A, Actinin staining shows that both the subendocardially located Purkinje fibers and the working cardiomyocytes have well organized contractile apparatus. B, Hematoxylin-eosin staining shows clearly that they are larger than the working ventricular myocytes and their spatial and fibrous isolation. C, Staining for connexin43 shows very high density of these gap junction proteins on the entire cell membrane (compare with neighboring working myocardium, where it is localized mostly at cell ends). Wheat germ agglutinin (WGA) staining highlights cell boundaries and fibrous tissue. Scale bars 25 µm. (From Sedmera D, Gourdie RG: Why do we have Purkinje fibers deep in our heart? Physiol Res 63(Suppl 1):S9-S18, 2014.)



CHAPTER 3  Stimulation and Excitation of Cardiac Tissues

measured in terms of ampere-hours, with 1 A-hr equal to 3600 coulombs of charge.

ELECTRIC CIRCUITS An electric circuit is an electric charge–conducting pathway that ends at its beginning. For electric circuits involved in myocardial stimulation by pacemakers or ICDs, the potential difference (measured in volts) in the circuit is provided by a battery. The difference in voltage between the anode and cathode in the battery results in a flow of charge (current) from the pulse generator through the conductors in the leads, electrodes, extracellular electrolytes, cell membranes, and intracellular ions and charged molecules. The principle of conservation of electric charge (Kirchoff ’s current law) implies that at any node in an electrical circuit, the sum of currents flowing into that node is equal to the sum of currents flowing out of that node. As current flows through the complete electric circuit, the net voltage must decrease to zero (Kirchoff ’s voltage law). In other words, the voltage drop across each element in the circuit must sum to the total potential difference across the entire circuit. Thus the electromotive force (voltage) generated by the battery (an increase in potential energy) must be completely dissipated (a decrease in potential energy) as current flows through all the elements of the circuit to end at the battery. Electric circuits in the clinical practice of pacing have multiple elements, including the pulse generator battery, the lead conductor(s), the electrodes, the myocardium, and blood within the great veins and cardiac chambers. All these elements introduce opposition to the flow of current. SERIES CIRCUITS In a series circuit, or circuit module, the elements are connected so that current must flow sequentially through each element in the circuit. Thus the current flowing through all elements in series is the same, with the voltage difference decreasing sequentially as it passes through each element in the circuit. The total resistance (RT) of a circuit with two resistive elements is series (R1 and R2) is additive, such that (RT = R1 + R2).

75

cathode-tissue interface and return to the anode, which may be located on a pacing lead or the pulse generator casing. The terminology used for electrode polarity may seem confusing as applied to lead electrodes and the electrodes of a battery. The electrode in a battery at which oxidation occurs (e.g., oxidation of lithium to yield Li+ plus an electron, e−) is the battery anode. The battery anode, by continuing oxidation becomes positively charged (Li+) while furnishing electrons to the circuit external to it. Therefore the terminal of the battery where electrons are provided to the circuit is the battery anode. From the battery anode, electrons are conducted through circuitry in the pulse generator and eventually enter the pacemaker lead, where they are conducted through the conductor to an electrode that is in contact with the myocardium. This electrode, receiving electrons from the pulse generator and furnishing electrons to the tissue, is the lead cathode. The return electrode located more proximally on a lead within the heart or on the pulse generator casing is the lead anode. It collects electrons from the tissue and returns them through the pulse generator circuitry to the positive electrode of the battery, the battery cathode, where reduction occurs (e.g., I2 + 2e− yields 2I−). The consistency in the terminology is that, when oxidation occurs, it occurs at an anode, and in the circuitry, an anode connects to a cathode that subsequently connects to another anode, and so on.

FACTORS THAT OPPOSE THE FLOW OF ELECTRIC CURRENT IMPEDANCE VERSUS RESISTANCE In cardiac pacing and defibrillation, there are multiple elements in the electric circuit that oppose the flow of electric current. As current flows through some of these elements, such as conductor wires in the leads, the opposition to current flow results in energy being lost as heat. These elements are known as ohmic, and the opposition to current flow is called resistance (R). The instantaneous voltage developed across a perfect resistor is linearly proportional to the instantaneous current flow through the resistor. If a steady voltage across the resistor is represented by V, the current by I, and an unchanging resistance by R, the relationship is expressed as follows in Ohm’s law: V = IR

PARALLEL CIRCUITS In a parallel circuit, two or more elements are joined at each end to a common conductor (node in the circuit). Therefore the potential difference is the same before each element and after each element, and current will flow from one of the common conductors to the other through any or all of the elements. The quantity of current flow in each element is inversely related to the factors that oppose the flow of electric charge in that element. At any node (junction) in an electric circuit, the sum of currents flowing into that node is equal to the sum of currents flowing out of that node (Kirchoff ’s current law). Most biologic circuits are made of various combinations of series and parallel modules or subcircuits. For example, because of electrochemical effects, an electrode placed in the heart may act as a capacitor in parallel with a resistor, both in series with the lead joining the pulse generator to the electrode. The total resistance in a parallel circuit (RT) composed of two possible current pathways (R1 and R2) is less than the resistance in any of the individual elements: 1 1 1 = ( + ). RT R1 R 2 ELECTRODE POLARITY IN BATTERIES AND LEADS All defibrillator and pacemaker electric circuits have both a positively charged electrode and a negatively charged electrode. The negatively charged electrode of a pacing or ICD lead (the cathode) is typically the tip electrode. Electrons from the pulse generator flow through the

In reality, when all the elements of a cardiac pacing circuit are considered, these factors are much more complex than can be represented by Ohm’s law alone. Although Ohm’s law (and therefore resistance) is accurate for conduction of current through a metal wire, other factors such as capacitance and inductance are not fixed quantities but change during and after the course of an electrical stimulus. When the pacing pulse begins, several events occur almost immediately. In order for conduction to occur from the electrode to the tissues and blood pool, electron motion in the lead wires and electrodes is converted to ion motion in the interstitial fluid. The electric field between the electrodes results in (1) reversible or irreversible oxidation-reduction processes occurring at the interface between the electrode and the electrolyte and/or (2) the electrode-tissue interface’s acceptance of current without charge crossing the interface. In the latter case, the interface acts as a capacitor. Figure 3-18 shows the result of this sequence of events. The leading-edge voltage/current ratio is referred to as ohmic polarization. However, during the duration that the pacing stimulus is applied, there develops increasing charge on the capacitor resulting in increasing opposition to further current flow from the start to the end of the stimulus. The result is that the voltage is not linearly related to the current. The circuit is said to be reactive. The vector combination of this reactance and the resistive elements (e.g., lead wires) contribute to the impedance in the electrical circuit. Impedance can be defined as the vector sum of all forces opposing the flow of current in an electric circuit. For calculations involving a unidirectional current (DC) (e.g., a unipolar pacemaker stimulus firing into a resistor on a test bench) or an alternating current (AC) circuit that

SECTION 1  Basic Principles

resistance. Impedance has a magnitude and a phase angle, both dependent on the rates of change of the applied voltage. The phase angle represents the difference in timing of sinusoidal current flow peaks compared with sinusoidal voltage peaks when a sinusoidal voltage is applied to a circuit.

Pulse generator pulse

Polarization effect

76

Vp

Leading edge

Trailing edge

Vo

Time (pulse width) Figure 3-18  The effect of polarization (upper panel) has been separated from the capacitively coupled constant-voltage pacing pulse (lower panel) to help clarify the electrical manifestation of electrode polarization. At the leading edge of the pulse, polarization is essentially zero. With time (pulse duration), the voltage resulting from polarization (sometimes called polarization overvoltage, Vp) increases. When the pacing pulse is shut off, polarization overvoltage decays exponentially as a result of diffusion. (From Stokes K, Bornzin G: The electrodebiointerface [stimulation]. In Barold SS, editor: Modern cardiac pacing, Mt. Kisco, NY, 1985, Futura, pp. 33-78.)

contain reactive elements (e.g., electrodes in contact with electrolytes), the concept and term impedance (Z) must be used in place of resistance. Impedance is a vector sum of resistance (R) and reactance (X). A voltage or current pulse of any shape can be broken down mathematically into combinations of sinusoidal components. In a simple series circuit, where R is the resistance and X is the sum of the capacitive and inductive reactances, the magnitude of impedance Z is defined as follows:

CAPACITANCE The difference between the conduction of electric currents by electrolytic fluids and the flow of electrons over metal wires is crucial to pacing. Because a negatively charged pacing electrode in contact with the endocardium is surrounded by blood and interstitial fluid, positively charged ions move toward that electrode during a pacing stimulus. This is time-dependent and results in the phenomenon of polarization, which develops rapidly as the stimulus is applied and dissipates slowly after the end of the stimulus. The opposite effect occurs on the positively charged anode. This effect of ions moving to oppose the flow of electric current has the effect of a capacitor in the circuit. A capacitor is an object that stores energy in an electric field by holding positive charges apart from closely approximated negative charges. A capacitor requires a material or space between the layers of negative and positive charges that is normally nonconducting (the dielectric). A cell membrane, although leaky, acts as a capacitor by separating the negatively charged inside of the cell from the more positively charged outside. Cell membranes have very high capacitance per unit area of cell membrane. The interface between a pacing electrode and the charged electrolytes that surround the electrode at its surface in the myocardial tissue acts, in part, as a capacitor. The terms Helmholtz capacitor and Helmholtz capacitance are used in this chapter for capacitor-like effects that occur at pacemaker and defibrillator electrode-electrolyte interfaces. Capacitance (C) specifies, for a given voltage applied across a capacitor, how much electrical charge (Q) can be stored by the capacitor. If V represents a steady voltage applied across the capacitor, then Q = CV. (If E is used instead of V as the symbol for electric potential, the relationship may be expressed as Q = CE.) The unit for capacitance is the farad. One farad is the capacitance of a capacitor that, on being charged to 1 volt, will have stored 1 coulomb of charge (the amount of charge delivered by 1 ampere flowing for 1 second). Coulombs delivered can be expressed as follows: t

Q t = ∫ i t dt 0

Z = R2 + X2 Because there are multiple tissues, blood pools, interstitial and intracellular spaces between the start of the circuit (the battery anode) and the end of the circuit (the battery cathode), each with a different mix of resistance, capacitance, and inductance, the total impedance (ZT) is the sum of these impedance (ZT = Z1 + Z2 + Z3 …). CAPACITANCE AND REACTANCE: THE COMPONENTS OF REACTANCE Pure inductors and pure capacitors have no energy losses. They store or release energy in or from an electric field (capacitor) or a magnetic field (inductor). They also change the time relationships between varying voltage and current. The effects can be quantified in terms of reactance. Capacitive and inductive reactances oppose each other. If a sine-wave voltage is applied to a pure capacitance, the current peaks occur 90 degrees earlier than the voltage peaks. If a sine-wave voltage is applied to a pure inductance, the current peaks occur 90 degrees later than the voltage peaks. The total reactance in a simple series circuit is the scalar sum of the inductive and capacitive elements, each of which varies with the frequency content of the applied signal. Impedance also varies with signal frequency; it is the vector sum of reactance and

where Qt is the total charge delivered between time 0 and time t, and it is the instantaneous current at each time segment between time 0 t and time t. The integral ∫ i t dt is the net area under the instantaneous 0

current-versus-time plot. INDUCTANCE

When an alternating current flows through a wire, a magnetic field surrounding the wire is induced. An inductor is an object that stores or releases energy in or from a changing magnetic field. The voltage difference across an inductor is proportional to the rate of change of current flowing through the inductor. Energy is stored during the formation of the magnetic field and is released when the magnetic field decreases or disappears. Inductance is the term that specifies the relationship between the voltage across an inductor and the rate of change of current traversing the inductor. The magnitude of the inductance can be represented by the symbol L. If vt represents the instantaneous voltage across the inductor, and iL represents the instantaneous current flowing through the inductor, the relationship is given by the following equation: vt = L

t di L or, i L = ∫ v t dt 0 dt



CHAPTER 3  Stimulation and Excitation of Cardiac Tissues

Note that the voltage across the inductor is directly proportional to the rate of change of current flowing through the inductor. Cell membrane currents have some of the current- and voltage-versus-time characteristics of an inductance in parallel with a capacitance.2 These inductancelike effects are related to the timing and magnitude of potassium ions moving into and out of the cell.81 For reactive elements connected in series, net reactance is the scalar sum of inductive reactance (positive in the mathematical complex plane) and capacitive reactance (negative in complex plane). Pure reactance values depend on the rates of change of current and voltage, whereas pure resistance values do not. The component of reactance that is most relevant to both pacing electrodes and cell membranes is capacitance, with inductance being much less important. For example, the cardiac action potential spreading throughout the heart generates a changing magnetic field that transiently stores a very small amount of energy. However, the changing magnetic field generated by spread of the action potential is so small that it is not clinically significant except in the research setting.82 For circuits with resistance R, capacitance C, and inductance L in series, where qt represents the charge accumulated across the capacitance at any time t, and where the current through these combined elements at time t is it, the voltage vt at time t is described by the following equation: t

q di ∫ i t dt + L di vt = it R + t + L or v t = i t R + 0 C dt C dt t

(remembering that qt = ∫ i t dt, and that the voltage across a capacitor 0 at time t is q t ). C These equations indicate that, for an instantaneous current it, the instantaneous voltage across the series circuit is the sum of the effects at that instant in time of the resistance, capacitance, and inductance of the circuit. Note especially that the instantaneous effects are highly t

related to the net amount of charge, (qt = ∫ i t dt) that has accumulated 0

in the capacitor from time 0 to the instantaneous time t. The equations show that the capacitance effect on the voltage decreases as the capacitance increases. Thus the polarization voltage that produces afterdepolarizations in the electrogram immediately following a pacing stimulus decreases as the electrode capacitance increases. This is especially relevant for automatic capture algorithms which rely on the accurate detection of an evoked response to assess whether capture has occurred. THE HELMHOLTZ CAPACITANCE About 1879, Helmholtz suggested that a layer of ions (inner Helmholtz layer) is attracted to the surface of a charged electrode and that this layer is bounded on its nonelectrode side by a layer of oppositely charged ions (outer Helmholtz layer) in the solution.83 This interface acts as a capacitor. These (and more complicated charge redistributions) occur because of attraction and repulsion interactions between an electrode held at a given electric potential (e.g., 2 V negative during 0.5-msec pacemaker pulse) and the ions and charged molecules in the interstitial fluid. The charge placed on the electrode, by electrostatic attraction, forces the accumulation of a polarized water layer and a second layer of hydrated, oppositely charged ions adjacent to the electrode surface. The oppositely charged ions come from the electrolyte. The interface at the electrode becomes an electrical double layer. This is in effect a charged capacitor. However, the actual interface is not a simple double layer. The Helmholtz model does not take into account other factors, such as absorption on the surface, thermal buffeting, and interaction between solvent dipole moments and the electrode. Models more complicated than the Helmholtz include the Gouy-Chapman and Gouy-Chapman-Stern.84 These models have dissimilar shapes for the plot of electric potential variation with distance from the electrode. The layers have relatively

77

Cc T–

Rw –

Rc Rt

IPG Ca i0 T+

Rw+ Ra

Figure 3-19  Simplified Equivalent Circuit for Cardiac Pacemaker. T− and T+ are, respectively, the negative (cathode) and positive (anode) terminals of the implanted pulse generator (IPG; in a unipolar device, T+ is inside the can). The direction (in the conventional sense, positive to negative) of current flow is shown with i0. RW− is the resistance of the conductor wire leading to the distal tip of the lead (cathode), whereas RW+ is that of the wire leading to the bipolar lead’s proximal ring electrode or the unipolar pulse generator’s can. Rt is the resistance of the tissue between the bipolar lead’s two electrodes, or the lead tip to the pulse generator can in a unipolar system. Rc is the ohmic polarization of the cathode, whereas Ra is that of the bipolar anode or unipolar generator can. Cc is the capacitance of the cathode, and Ca is that of the anode (can). The equivalent circuit is the same for unipolar and bipolar systems, but the values for some of the components differ.

high dielectric constants and form an interface that behaves electrically like a capacitor (Cc and Ca in Fig. 3-19). In a semiconductor or in localized regions of an electrolyte, an excess of positive or of negative charge may be present. If the excess is of positive charge carriers, the positive carriers are the majority carriers and the negative charge carriers the minority carriers. An excess of negative carriers makes these the majority and the positive carriers the minority. The second layer of the double layer is formed of hydrated ions and more water (Fig. 3-20). In physiologic electrolytes, the ions include Na+ and Cl− in major concentrations (majority carriers). Other ions present in lower concentrations (minority carriers) include hydronium (H3O+), hydroxyl (OH−), and phosphate (HPO42−). The ions attracted to or repelled from the electrode during the electrical stimulation pulse make up a separation of charge in the tissue electrolyte. When the pacemaker pulse is applied as a negative voltage to the electrode, electrons accumulate in the electrode. Reversible reactions may form metal-oxide complexes on the surface of the electrode. A primary water layer forms on the electrode. Positive ions surrounded by water molecules—a water shell—make a secondary water layer. This accumulation of positive ions in the electrolyte near the electrode unbalances local electrolyte charge neutrality. Secondary ion rearrangements occur in great complexity, with several names for the various processes. When the pacemaker pulse stops, the ions that have accumulated, being no longer attracted to or repelled from the electrode, gradually rearrange themselves back toward their original, electrically neutral position. Ion rearrangement is not as fast as the transmission of an electric potential in a wire. The result is that after a pacemaker or defibrillator pulse stops, the Helmholtz capacitor acts as if it were a temporary battery of declining voltage discharging itself into the tissue (Fig. 3-21). The decaying voltage gradient in the tissue persists long enough to be detected by a pacemaker or defibrillator and may be great enough to interfere with sensing in autocapture devices in some situations. Other factors being equal, the voltage decay duration is greater

78

SECTION 1  Basic Principles

FARADIC AND NONFARADIC CURRENTS + H3O + Na – Cl + Na

– OH

+ Na

+ Na

A

B

C

Figure 3-20  Hypothesized Structure of the Helmholtz Double Layer. At the left is an uncharged cathodic electrode interface. When the electrode is charged, a layer of surface hydration develops (region A). Region B contains a second, more loosely held hydration layer with hydrated ions. Based on electrostatic considerations, layer B has a high dielectric constant, approaching that of pure water. It has a thickness of less than 10 Å. Region C represents bulk solution. (From Stokes K, Bornzin G: The electrode-biointerface [stimulation]. In Barold SS, editor: Modern cardiac pacing, Mt. Kisco, NY, 1985, Futura, p 63.)

Tek

T Trig’d

M Pos: 1.980ms

2

1

At least two types of processes involving ion flow occur at the inter­ face between a pacemaker or defibrillator electrode and tissue. First, oxidation-reduction reactions, which can be reversible or irreversible, involve electron movement across the interface and constitute faradic current. The second process, nonfaradic current, occurs without transfer of electrons across the interface. It consists of electron flow in or out of the electrode itself and a flow of ions in various layers or “clouds” toward or away from the interface. This nonfaradic process is similar to charging or discharging an electrical capacitor, but at the surface of an electrode in electrolyte there is directed electron drift in or out of the electrode and directed ion drift within the electrolyte. The ion flow and the electron flow each constitute an electric current; yet no charge crosses the interface. Away from the electrode in the body of the electrolyte, the electric potential gradient causes ions to move away from or toward the electrode region, depending on their charge. Ion mobility characteristics, concentration gradients, and temperature gradients also affect ion movement. In the heart the process is more complicated than in an electrolyte solution alone, because of the anisotropic properties of the extracellular and intracellular domains. FARADIC CURRENT Faradic current is so named for Michael Faraday (1791-1867) who quantitatively described the effects of electric current through an electrolyte in 1834, stating that “the chemical decomposing action of a current is constant for a constant quantity of electricity, notwithstanding the greatest variations in its sources, in its intensity, in the size of the electrodes used, in the nature of the conductors (or nonconductors) through which it is passed, or in other circumstances.”85 If one thinks of an electrode-electrolyte interface as resembling a variable resistor and variable capacitor in parallel, faradic current flow would be flow—electron transfer—through the resistor. Whether electron transfer across the interface occurs depends on the properties of the electrode and the electrolyte and on the applied electric pulse characteristics. The current crossing the interface when one is charging a battery is faradic current produced by oxidation-reduction reactions. CAPACITIVE CURRENT FLOW

Ch1 500mV

Ch2 500mVBw 1 msec

Ch1

Figure 3-21  Slow Decay of Pacing Catheter Lead Voltage Between Biphasic Stimulus Phases. The plot shows current (top trace) and induced voltage (lower trace) from a Bloom stimulator delivered between the pacing tip of an ICD lead and a “hot can” electrode in saline. At the end of each phase, the voltage measured at the pulse generator terminals initially goes rapidly toward zero but then, partway toward zero, it begins to change only slowly. This is the effect of discharge of the Helmholtz capacitor into the surrounding electrolyte. This voltage response between phases is a function of the rates of ion movement in the electrolyte.

and the polarization voltage is less when the electrode surface area is larger. Another approach to minimize polarization is to abruptly reverse the polarity of a pacing stimulus during the pulse (biphasic pulses). When biphasic pulses are used, minimal postpulse polarization persists, provided the time between phases is in the microsecond range. If the time between phases is increased into the millisecond range, the duration and amplitude of persisting polarization increases.

This mechanism is not a transfer of electrons in one direction or the other across the electrode-electrolyte interface. Capacitance current flow is the accumulation of charge on the electrode at the interface balanced by a corresponding accumulation of charge of net opposite sign in the electrolyte adjacent to the interface. Charge flows in the lead, and in the electrolyte, but not across the interface. The flow of charge into the capacitor is measurable current in a pacemaker lead, even though in a perfect capacitor, no charge crosses the interface. The resulting arrangement of ion groupings at and near the electrode-electrolyte interface can be very complicated.84 The accumulation of charge separated by charge sign at the interface defines the Helmholtz double-layer capacitance.84 Energy is stored as an electrostatic field that exists between the positive and negative charge layers. When a pacemaker pulse is applied to an electrode in the heart, the charge movement into or out of the Helmholtz capacitance—both electrons in the electrode and ions in the electrolyte—is an electric current, even though no charged particle necessarily crosses the interface. At a pacemaker electrode, the capacitance is not constant. It is dependent on current density, electrolyte composition, and the area and other surface characteristics (e.g., fractal surface) of the electrode material.86 Because of the minute distances separating positive and negative charge layers at an electrode in electrolyte, the capacitance is high, ranging from about 0.2 to 40 mF/mm2 of geometric area of the



CHAPTER 3  Stimulation and Excitation of Cardiac Tissues

electrode.87 The fractal-surface electrode has a much higher capacitance than a smooth-surface electrode. dV states that the current i entering or leaving The equation i t = C dt a capacitor at time t varies directly as the magnitude of the capacitance (C, measured in farads) multiplied by the rate of change of the voltage difference across the capacitor. Because a constant-voltage pulse is nominally a rectangular wave, its rate of change is small except when the pulse is being turned on or off. Therefore current flow into the capacitor (i.e., into the pacemaker electrode-electrolyte interface, Helmholtz capacitance) occurs very rapidly during the leading edge of a nominal constant-voltage pacemaker pulse. Then as charge accumulates at the electrode-electrolyte interface, the accumulating negative charge collection at the electrode surface slows the rate of further accumulation. It does so because the accumulating negative charge opposes the inflow of additional negative charge. Finally, unless the pacemaker output voltage were to be raised during the pulse, no further accumulation can occur. The voltage across the capacitor becomes equal and opposite to the constant-voltage pulse that had been driving current flow into the lead. Note that the Helmholtz capacitance is influenced by the current density at the electrode, the types and numbers of ions in the electrolyte, the material and surface of the electrode, the temperature, and other factors. THE RELATIONSHIP BETWEEN ELECTRODE-TISSUE INTERFACE CAPACITANCE AND POLARIZATION 1 t i t dt, describing the voltage that C ∫0 occurs across a capacitor when current flows into it, the developed voltage varies inversely as the capacitance (C, measured in farads). If both sides of this equation are multiplied by C, the equation becomes Note that in the expression v t =

t

Cv t = ∫ i t dt . The amount of stored charge (Qt) at any time t is 0

t

∫ i dt. 0

t

Therefore Cvt = Qt, and vt = Qt/C. That is, for a given amount of charge put into the defibrillator or pacemaker electrode, the voltage across the interface will be decreased if the surface area of the electrode (and therefore the capacitance) is increased. This voltage is the polarization voltage. During its decay time after the stimulation pulse has ended, the polarization voltage can interfere with autosensing of capture. For a given stimulus voltage, the greater the polarization voltage at the electrode-tissue interface, the lesser the voltage gradient elsewhere in the tissue between the cathode and the anode. A fractal-surface electrode, with its high capacitance,87 has, for the same charge flow in and out, a lesser potential difference across the electrode-electrolyte interface than a smooth-surface electrode. Therefore the fractal surface is preferred. WARBURG IMPEDANCE

In 1899, Warburg described an effect of ion motion caused by a voltage difference between two electrodes in an electrolyte as an “impedance.” The Warburg impedance is an effect of ion diffusion activity under the influence of the potential gradient at the interface between electrode and electrolyte (Fig. 3-22). It is said that “a Warburg impedance can be difficult to recognize because it is nearly always associated with a charge transfer resistance and a double layer capacitance” (B. Rodgers, personal communication). The Helmholtz capacitance concept is based on the separation of charges in the electrode from charges in the electrolyte at angstromdimension levels. Warburg developed his impedance concept to model the effects of diffusion. When a sine-wave voltage is applied across an electrolyte, this impedance manifests itself as a 45-degree phase shift between the voltage and the current if the current density is infinitely low. The Warburg impedance magnitude becomes small compared with other impedances at the electrode-electrolyte interface as the electrolyte concentration is made greater and as the stimulation frequency increases.84

Input current from pulse generator cathode Lead resistance

79

Helmholtz doublelayer capacitance

Electrode interface with electrolyte in adjacent myocardium

Output current to pulse generator anode Myocardial bulk impedance

Faradic charge Warburg transfer resistance diffusion impedance Figure 3-22  Equivalent circuit representation of an electrodeelectrolyte interface, including the Helmholtz capacitance and the Warburg impedance. (Modified from Rodman JE: Solution, surface and solid state assembly of porphyrins [PhD thesis]. Cambridge, UK, 2000, University of Cambridge.)

Ovadia and Zavitz88 made an impedance spectroscopy study of the interface between a platinum electrode and a metabolically active, perfused living heart. Three impedance spectral components were found: the Warburg impedance, a thin-film impedance, and a single high-angle constant-phase impedance. The impedance spectrum was not single-valued and was not stable in time. Figure 3-22 shows one equivalent circuit representation of an electrode-electrolyte interface, including the Helmholtz capacitance and Warburg impedance. The Helmholtz capacitance-Warburg impedance effects can become especially important in explaining some of the threshold measurements and seeming contradictions found during biventricular pacing procedures,89,90 as discussed later.

ELECTRODE DESIGN: BALANCING HIGH CURRENT DENSITY AND POLARIZATION The total impedance in the circuit of a pacing stimulus includes resistance in the lead conductor, impedance at the electrode-tissue interface, capacitance due to the effects of ions in the electrolytic interface, resistance within the vasculature and inductance. Energy dissipation (as heat) in the lead connecting the pulse generator to the electrode is due to Ohmic resistance and is proportional to the square of the current. To minimize charge drain from the battery, low conductor resistance is desirable. At the same time, having a relatively high impedance at the electrode-tissue interface and a low amount of polarization are also desirable. So, how to increase impedance at the electrode-tissue interface? At the pacing threshold, a certain minimum current density must be achieved. If the interface area is made smaller, the current can be reduced equivalently while maintaining the same current density. However, small electrodes are associated with greater polarization (Fig. 3-23). Polarization is inversely and exponentially proportional to the size of the electrode and the temperature and conductivity of the tissue.91,92 For practical purposes, in vivo temperature and conductivity are essentially constant. One remembers from the equation, vt =

1 t i t dt C ∫0

that the voltage vt across the Helmholtz capacitance at the electrode is proportional to the accumulated charge and inversely proportional to the capacitance. If the electrode contact area with the tissue is made smaller to increase the current density and nothing else is changed, the capacitance will be decreased. For a given applied charge, that action will increase the Helmholtz capacitance voltage,

80

SECTION 1  Basic Principles

Ohmic pacing impedance (Ω × 102)

100 (N = 48 models, 323 canines)

Stimulus

d

50

RMP + + + + + + + + – – – – – – – – – – – – – – – – – – – + + + + + + + + + + +

A

10 5

+ + +

+ + – + – – + – – + – –– –– – + – + + +

O ∆v TP RMP

1 .1

.5 1.0 5.0 10 Geometric surface area (mm2)

50 100

Figure 3-23  Leading-Edge Ohmic Polarization or Electrode Resistance of Chronic Canine Ventricular Leads (12 Weeks After Implantation) Measured With a Medtronic Model 5311 Pacing Systems Analyzer as a Function of the Unipolar Cathode’s Geometric Surface Area. Each data point was obtained from a population of animals testing one lead design. The study included 323 animals and 48 lead models.

because of the inverse relationship between the voltage and the capacitance. For a given electrode material (as a first approximation), capacitance (C) is a function of the interface surface area (Ai): C=

Stimulus

B

d

Figure 3-24  A, The charge distribution along a cylindrical excitable cell with a cathodal stimulating electrode delivering a rectangular current pulse of intensity I and duration d. The mechanism of stimulation in which the current pulse reduces the resting membrane potential (RMP) to the threshold potential (TP) to produce a propagated action potential is shown in B. (From Geddes LA: Accuracy limitations of chronaxie values. IEEE Trans Biomed Eng 51(1):176-181, 2004.)

100%

eA i d

where e is the dielectric constant of the Helmholtz double layer and d is its thickness. Because e and d are essentially constants in vivo, the capacitance of the electrode varies as a function of the interface surface area. The solution to the competing problem of creating a smaller electrode with low polarization is to increase the electrode surface area by fractal or sintered coatings. CHEMICAL REACTIONS AND CORROSION AT THE ELECTRODE-TISSUE INTERFACE In order to avoid irreversible tissue injury at the electrode-tissue interface, the electrode should not be susceptible to corrosion and deposition of toxic products in the adjacent tissue. Electrodes that undergo only minimal irreversible reactions are highly desirable for pacemakers and defibrillators. When an external voltage is applied, electrodes that do not produce irreversible chemical reactions are so-called ideal electrodes. Although metal oxide films form on electrodes, occurrence of other oxidation-reduction reactions and corresponding electron transfer through the interface do not occur or are minimal in ideal electrodes. In this sense, no real electrode can be completely “ideal” in all circumstances. If the magnitude and duration of polarization secondary to occurrence of the electric pulse is brief, electrochemical reactions occurring at the electrode may reverse. However, if the charge redistribution time is long, the effects of charge redistribution on the electrode and on the tissue may become irreversible.

MECHANISM OF ELECTRICAL STIMULATION OF THE MYOCARDIUM Figures 3-24 and 3-25 illustrate that, for a rectangular, constantcurrent cathodal pulse of increasing intensity, the transmembrane

Propagated action potential

TP Cathode response RMP (0%)

0.5

1.0

1.5

msec Anode response

Figure 3-25  Local Potential Changes Under the Cathode and Anode With Increasing Stimulus Intensity. Note that under the cathode, when the stimulus intensity reduced the membrane potential from the resting membrane potential (RMP) to the threshold potential (TP), excitation occurred. Excitation did not occur under  the anode with increasing stimulus intensity. (From Geddes LA:  Accuracy limitations of chronaxie values. IEEE Trans Biomed Eng 51(1): 176-181, 2004.)

81

CHAPTER 3  Stimulation and Excitation of Cardiac Tissues

potential of a myocyte is depolarized before returning to the resting potential with a membrane time constant. As the cathodal pulse rises, the cell membrane resists a change in potential. However, if the stimulus current is increased, the cell membrane cannot overcome the applied potential and the membrane potential will eventually reach threshold with initiation of a self-propagating action potential. With anodal pulses the membrane potential is decreased and the cell is hyperpolarized followed by return of the membrane potential to the resting state. The typical electrical stimulus used for cardiac pacing is a direct current pulse with a constant leading edge voltage and a fixed pulse duration. Although the voltage at the electrode-tissue interface could be either negatively charged (a cathodal stimulus) or positively charged (an anodal stimulus), virtually all permanent pacemakers and ICDs deliver a cathodal pulse from the distal electrode. However, it should be recognized that anodal stimulation often occurs, either from the proximal electrode of bipolar lead or from another electrode of a multipolar lead that is used for cardiac resynchronization therapy (CRT). In order for the stimulus to generate a self-propagating wavefront of cardiac depolarization in the chamber to which the stimulus is applied, the combination of the amplitude and pulse duration must be above a threshold value (see strength-duration relationship section below). The stimulus must also be applied to the myocardium at a time when it is not refractory (see strength-interval relationship section below). The self-propagating wavefront of depolarization is carried by actions potentials produced by the opening and closing of ion channels and intercellular GJs triggered by changes in the transmembrane voltage of the myocytes. Once initiated, the depolarizing wavefront will continue to propagate through cardiac tissues that are excitable. In reality, an applied electrical stimulus above threshold value produces an electric field that initiates cardiac excitation as a result of passive effects on the transmembrane potential (the difference in voltage between the inside and outside of the cell). Because of the much lower impedance within the extracellular space and the high impedance offered by the sarcolemmal membrane, relatively small amounts of current actually flow through the cell membrane. In the heart, two domains of charge flow exist93: extracellular and intracellular. The two domain pathways are different in their anisotropic characteristics, and they are interwoven. Current passes from one domain to the other through cell membranes. Bidomain theory and studies are based on the concept that at every point in the heart, there are two electric potential vectors, one intracellular and one extracellular. Mathematically, each domain is continuous and occupies the entire domain. Membrane current leaves one domain and enters the other at the same coordinates. The analytical and experimental work from these concepts has provided insight about depolarization wavefront passage over the myocardium, current flow directions in relation to wavefront direction, and virtual electrode formation.94-96 Rather, when a single myocyte is exposed to an induced electric field, the extracellular voltage (within the extracellular space) shows significant polarization such that one end of the cell is hyperpolarized while the opposite end is depolarized97 (Fig. 3-26). And because very little current flows into the cell, the intracellular space shows minimal difference in voltage from one end of the cell to the other. The transmembrane voltage (Vm), the difference between potential inside (Vi) of the cell and outside (Ve) of the cell is given by the equation: (Vm = Vi − Ve). Thus Vm varies from hyperpolarized on one side of the cell (adjacent to the positively charged anode with ↑Vm) to depolarized on the other side (adjacent to the negatively charged cathode with ↓Vm). As sodium channels (I Na + ) and calcium channels (ICa2+ ) in the membrane are opened by depolarization of the transmembrane voltage, an action potential may be induced at the depolarized side of the cell. The induced action potential may then spread to adjacent cells by passing ions through GJs that connect the cells and depolarizing their membrane, thereby opening Na+ and Ca2+ channels and initiating a self-propagating wave of depolarization.

Vi Vm+

Vm–

A

+

–

Ve Vm

Voltage



B

Vi Ve

Figure 3-26  Response of a Cardiomyocyte to Electric Field. Relatively little current passes through the relatively high impedance cell membrane. A, Cardiomyocyte exposed to extracellular electric field. The intracellular voltage (Vi) is relatively low impedance and thus does not vary from one side of the cell to the other while the extracellular voltage (Ve) drops from one end of the cell to the other. B, Transmembrane potential (Vm) is defined as Vm = Vi − Ve, so the left side of the cell is hyperpolarized by the external field while the right side of  the cell is depolarized. If the extracellular field gradient is large enough, the right-sided portion of the cell may reach threshold and initiate an action potential that may activate the entire cell and spread to adjacent cells through gap junctions. (From Dosdall DJ, Fast VG, Ideker RE: Mechanisms of defibrillation. Annu Rev Biomed Eng 12:233-258, 2010.)

ELECTRIC POTENTIAL GRADIENTS AND CURRENTS FOR STIMULATION AND DEFIBRILLATION A rectangular pacing stimulus at the electrode-tissue interface may be either negatively charged (cathodal stimulus) or positively charged (anodal stimulus). Cardiac excitation may occur immediately adjacent to the electrode or at a distance by a virtual electrode effect. When the electric field strength exceeds approximately 1 V/cm during diastole, myocardial stimulation may result.98 The electric field of this intensity must be applied to only several dozen myocytes, typically with a current of less than 0.5 mA. An increase in the electric field gradient strength to approximately 6 V/cm may result in ventricular fibrillation if it is applied locally during the vulnerable period of the ventricle which is approximated by the peak of the T-wave.98 A field strength of this intensity usually requires a shock of >20 mA of current. A field strength of the same 6 V/cm applied to nearly the entire ventricular mass will also result in ventricular defibrillation though a shock of approximately 10 A is required to generate a minimum field gradient of this strength.98 Thus the current required for defibrillation is approximately 1000 times that required for stimulation. MAKE AND BREAK ACTIVATION The mechanism by which an electrical stimulus initiates a selfpropagating wave of depolarization (“captures the heart”) has been extensively studied.96,99-106 Rectangular cathodal and anodal stimuli are characterized by an abrupt change in electric potential both at the onset of the pulse and at its termination (Fig. 3-27). Thus for a cathodal pulse, the electric potential applied to the heart suddenly changes from 0 to negative at the start of the pulse, with an abrupt change from negative back to 0 at the end of the pulse. For anodal stimuli, the polarity is opposite so that the electrode potential goes from 0 to positive at the start of the pulse and from positive back to 0 at the end of the pulse. It has been demonstrated that these transient changes in potential either at the start or end of the pulse may stimulate a selfpropagating wave of depolarization (Figs. 3-27 and 3-28).101,107-109 The terminology for capture at the beginning of the pulse is either cathodal or anodal make stimulation, depending on the polarity of the pulse.101 The term cathodal or anodal break stimulation is used when capture is induced by the end of the pulse (see Fig. 3-27). The proof that both make and break stimulation could result in capture was demonstrated by Dekker in 1970 by delivering pulses that either began or ended in

82

SECTION 1  Basic Principles

Cathodal make 0 – mV –100 Anodal break

0 + mV –100

A

Figure 3-27  Action Potentials Are Initiated by Cathodal and Anodal Excitation. Excitation make and break phenomena occur under and near electrodes. Cathodal make occurs when the excitation current directly excites the tissue under the electrode. Anodal break occurs under the excitation electrode on termination of the stimulating pulse. (From Antoni H: Electrical properties of the heart. In Reilly JP, editor: Applied bioelectricity: from electrical stimulation to electropathology, New York, 1998, Springer, p 160.)

the refractory period of the canine ventricle.107 Thus break stimulation was proven by observing myocardial capture despite delivering long duration stimuli with the onset of the pulse within the refractory period of the previous beat.107 In this case, break stimulation occurred at the end of the pulse after the end of the ventricular refractory period. Make stimulation was proven by capture that was induced by a stimulus that was delivered in late diastole with the end of the pulse within the refractory period of the captured beat. If a rectangular pulse caused a QRS response to occur before the break, break stimulation could not have been the cause of capture and make stimulation was proven. It has been recognized that ventricular capture can be induced by all four modes of stimulation, including cathodal make, cathodal break, anodal make and anodal break.99,101,108,109 Each of these modes of stimulation is associated with a somewhat different strength-interval curve. The stimulation threshold at end diastole is lowest for cathodal make excitation, followed by anodal make, cathodal break, and anodal break stimulation (mean threshold values in the canine ventricle with platinum epicardial electrodes of 0.4 mA, 1.3 mA, 2.2 mA, and 3.0 mA, respectively.107 It should be recognized that there are exceptions to this rule. The duration of the ventricular refractory period is also somewhat dependent on the mode of stimulation with anodal break stimulation having the shortest effective refractory period (ERP) (mean 130 msec) and anodal make excitation having the longest ERP (mean 175 msec) in the canine ventricle with very long pulse duration. The ERP measured with cathodal make and break stimuli were of intermediate duration (both with mean of 155 msec).107 The shape of the strength-interval relationship is also determined by the mode of stimulation. Both cathodal break and anodal break stimuli are associated with a dip in the strength-interval curve as stimuli are delivered progressively earlier in diastole before abruptly increasing as the relative refractory period is encroached upon. This dip corresponds to a period of hyperexcitability and is more prominent and steep with anodal break stimulation. The diastolic dip in the strength-interval curve is also more pronounced at longer pulse durations. Thus during the strength-interval curve there may be a shift from make to break stimulation as the coupling interval is progressively decreased and the period of hyperexcitability is encountered.

B 100 msec Figure 3-28  This Figure Demonstrates the Response to Multiple Rectangular Anodal Pulses Delivered During EndDiastole. In the upper half of both figures the myocardial response is shown as obtained from bipolar needle electrodes at a distance of approximately 10 mm from the center of the stimulating electrode. The records of the myocardial responses are placed in the same vertical order as the current pulses shown below them. Current strength was fixed at 2 mA, slightly above the threshold for the longest pulses (1.5 mA). Pulse duration was varied in two different ways. A, Instant of current break was fixed in the cardiac cycle and the current make was moved progressively later in the cardiac cycle. The myocardial response is seen to shift with the current make. B, Instant of current make was fixed in the cardiac cycle and the current break was moved progressively later. The myocardial response arrives with a fixed delay after current make, irrespective of the timing of current break both before and after the response. Such response behavior was interpreted as anodal make stimulation. (From Dekker E: Direct current make and break thresholds for pacemaker electrodes on the canine ventricle.  Circ Res 27(5):811-823, 1970.)

The mechanism of make and break stimulation has been elucidated by optical mapping of transmembrane voltage and action potentials using the voltage-sensitive fluorescent dye, di-4-ANEPPS.101,102,109 These studies have modeled the cardiac syncytium as having two domains or conducting spaces: an intracellular domain that is superimposed on an extracellular domain.109 These intracellular and extracellular spaces are separated by a cell membrane with higher impedance.



CHAPTER 3  Stimulation and Excitation of Cardiac Tissues

Conductivity within both the intracellular and extracellular spaces is highly influenced by the direction of myocardial fiber orientation. From the experimental studies of Clerc,110 the intracellular conductivity in the longitudinal axis of myocardial fibers was found to be 8.9 greater than along the transverse axis. Within the extracellular space, the electrical conductivity shows less anisotropy with a conductivity ratio of 2.5 (longitudinal/transverse). This difference in anisotropy between the extracellular and intracellular domains is responsible for the pattern of cathodal make, cathodal break, anodal make, and anodal break stimulation.101 If the conductivity ratios within the extracellular and intracellular spaces were identical, a cathodal make stimulus would result in a region where the extracellular potential was negative. As the transmembrane potential is the difference between potential outside and inside the myocyte, the region of myocardium under the cathode would be depolarized as both the inside and outside of the cell would be negatively charged (Fig. 3-29A). This would result in an elliptical wavefront that would propagate away from the center of the depolarized region of myocardium. In reality, and as predicted by modeling using the bidomain difference in anisotropic ratios for the extracellular and intracellular spaces, cathodal make stimulation produces a very different pattern of depolarization and hyperpolarization in the myocardium near a stimulating cathode. The bidomain model predicts that, for cathodal make stimulation, the depolarized area under the cathode will assume a “dog bone” shape that is oriented transverse to the direction of myocardial fibers (Fig. 3-29B). This depolarized area is surrounded in the longitudinal axis by two regions of hyperpolarized myocardium. The depolarized region is a virtual cathode whereas the two hyperpolarized regions form virtual anodes. The spread of activation away from the virtual cathode starts slowly as it moves transverse to the direction of fiber orientation and must spread around the virtual anodes before it can propagate in the longitudinal direction (see Fig. 3-29B). In contrast, anodal make stimulation produces the opposite pattern (see Figs. 3-29C and 3-28). The anodal pulse induces a hyperpolarized region under the electrode that extends with a dog bone shape oriented transverse to the direction of fiber orientation (see Fig. 3-29C). In the longitudinal direction, two depolarized virtual cathodes are produced, both of which result in rapid longitudinal spread of activation in the myocardium. Figure 3-30 shows the spread of activation from either cathodal (panel A) or anodal (panel B) make stimulation. Within 2 msec of the cathodal stimulus activation spreads in the transverse direction from the virtual cathode of depolarized tissue (Fig. 3-30A). There is a region of delayed conduction in the longitudinal axis as the wavefront encounters the hyperpolarized virtual anode. By 8 msec, the spread of activation proceeds in both the transverse axis (with slow conduction velocity) and in the longitudinal axis (with rapid conduction velocity). For an anodal make stimulus, the central, dog bone virtual anode of depolarized tissue is surrounded in the longitudinal axis by two smaller virtual cathodes (panel B). Within 2 msec, there is rapid longitudinal and slower transverse conduction away from the site of stimulation. For cathodal break stimulation, the same pattern of a central dog bone-shaped region of depolarization is produced under the electrode that is oriented transverse to the axis of fiber orientation which is surrounded by two smaller virtual anodes in the longitudinal axis (see Figs. 3-30 and 3-31). After an initial delay, break stimulation occurs by activation that spreads away from the virtual anodes in the longitudinal direction. Anodal break stimulation (see Fig. 3-29E) results from spread of activation away from the central virtual anode starting in the transverse axis with slow conduction velocity before it encounters rapidly conducting, nonrefractory longitudinal fibers. However, the mechanism of anodal break stimulation may differ for stronger and weaker anodal stimulus intensity (Figs. 3-32 and 3-33).109 For a strong anodal stimulus, anodal break stimulation is predicted by bidomain modeling to occur starting from the virtual cathodes that surround the dog bone-shaped virtual anode directly under the electrode. This mechanism of anodal break stimulation can be explained by passive effects on anisotropic myocardium. However, when the stimulus

83

Depolarized

Resting

2 mm

Hyperpolarized

A

Fiber direction

B

Cathodal

Make

C

Anodal

Make

D

Cathodal

Break

E

Anodal

Break

Figure 3-29  A schematic representation of the theoretical predictions of action potential propagation for point stimulation of cardiac tissue based on Sepulveda et al96 and Roth.100 A, Cathodal make stimulation in an equal-anisotropy model. An elliptical region of tissue would be directly depolarized (orange) by a strong point stimulus (black dot) and would act as a virtual cathode. An elliptical action potential wave front (yellow lines separated by 2 msec) would propagate away from the edge of the virtual cathode. B, Cathodal make stimulation in a model with differing anisotropic conductivities for the intracellular and extracellular spaces. The virtual cathode is yellow to orange, and the virtual anode is blue. The resulting propagating wave front would initially have the transverse dogbone shape, but because of the greater longitudinal conduction velocity, the wave front would become elliptical by the time it was 5 mm away from the stimulus electrode. C, The same model as in B, but for anodal make stimulation. A pair of action potential wave fronts propagating outward from the virtual cathodes (orange) merge and form an elliptical wave front within 1 cm of the stimulus electrode. D, The same model for cathodal break stimulation. Early activation occurs from the virtual anodes (blue) along the fiber direction. E, Anodal break stimulation, in which initial activation progresses transverse to the fibers from the dog bone-shaped virtual anode (blue). (From Wikswo JP, Lin S-F, Abbas RA: Virtual electrodes in cardiac tissue: a common mechanism for anodal and cathodal stimulation. Biophys J 69(6):21952210, 1995.)

current is decreased, anodal break stimulation occurs from the central dog bone-shaped, virtual anode with a significant delay from the end of the pulse (see Fig. 3-31). This mechanism of anode-break stimulation is likely produced by the active recruitment of hyperpolarizationactivated, If channels in the membrane, which open in response to hyperpolarization, thereby resulting in depolarization of the cell membrane. If the membrane depolarization reaches threshold, Na+ and Ca2+ channels are opened and a new action potential is initiated. Computer simulations for all four types of stimulation (cathode make, anode make, cathode break and anode make) are shown in Figure 3-34.

84

SECTION 1  Basic Principles

Depolarized <–2 –1

Fiber direction

0

∆F/F (%)

1 1 mm

>2 Hyperpolarized

A

B

C

Figure 3-30  False-Color Images of the Transmembrane Potential Associated With Injection of Current Into Refractory Cardiac Tissue. A, The image for a −10 mA, 2-msec cathodal S2 stimulus applied at a point electrode. Note the dog bone-shaped virtual cathode (orange) and the pair of adjacent virtual anodes (blue). The fiber orientation is from lower right to upper left. The color bar shows the fractional change in fluorescence. B, The complementary image for a + 10 mA, 2-msec anodal S2 stimulus at the same location on the heart. Note that the dog bone is now the virtual anode (blue), whereas the adjacent areas are virtual cathodes (orange). C, Direction of fiber orientation and color scale ranging from depolarized (yellow) to hyperpolarized (blue). (From Wikswo JP, Lin S-F, Abbas RA: Virtual electrodes in cardiac tissue: a common mechanism for anodal and cathodal stimulation. Biophys J 69(6):2195-2210, 1995.)

0

2

4

6

8

0

2

4

6

8

0

3

6

9

12

0

3

6

9

12

Cathode make –10 mA

A

Anode make +10 mA

B

Cathode break –2 mA

C

Anode break +3 mA

D 1 mm Figure 3-31  False-Color Images of the Transmembrane Potential Associated With Injection of Current Into Fully Repolarized, Excitable Cardiac Tissue. The number inside each frame is the time in milliseconds. Upper two rows: make stimulation. A, Cathodal make stimulation with a 1-msec, −10-mA SI stimulus current; B, 1-msec, +10-mA SI anodal make stimulation of the same heart. Lower two rows: break stimulation. The direction of the epicardial fibers from lower right to upper left is evidenced by the shape of the virtual electrode pattern and the elongation of the propagating wave front in the last two images of each sequence. C, 180-msec, −2-mA cathodal break S2 stimulation of another heart; D, 150-msec, +3-mA anodal break S2 stimulation of a third heart. For each row, the leftmost images are at the end of the S2 stimulus (0 msec), and the other images were at 2-msec (A and B) or 3-msec (C or D) intervals thereafter. The direction of the epicardial fibers from lower right to upper left is evidenced by the shape of the virtual electrode pattern and the elongation of the propagating wave front in the last two images of each sequence. The color scale is the same as in the previous figure. (From Wikswo JP, Lin S-F, Abbas RA: Virtual electrodes in cardiac tissue: a common mechanism for anodal and cathodal stimulation. Biophys J 69(6):2195-2210, 1995.)



CHAPTER 3  Stimulation and Excitation of Cardiac Tissues

1

1

2

3

40

0

0

–40

–40

–80

–80

–120

–120

10

40

0

0

–40

–40

–80

–80

–120

–120

8

10

15

40

18

40

2.5

5

40

4

6

19

0

–40

–40

–80

–80

–120

–120 2.5

20

21

40

0

40

0

0

1.25

2.5

5.0

40

0

12

1.25

85

–40

–40

–80

–80

–120

Figure 3-32  Transmembrane potential was calculated using the bidomain model with the modified action potential model for a 0.3-mA anodal stimulus pulse of 10-msec duration. The time (in msec) since the beginning of the stimulus pulse is marked in the top left-hand corner. The color bar on the right side shows the color scale used in the panels (in mV). Cardiac fibers are aligned along the x-axis in the figure. The peak hyperpolarization induced underneath the stimulating electrode is −1170 mV, and the peak depolarization is 53 mV. (From Ranjan R, Tomaselli GF, Marbán E: A novel mechanism of  anode-break stimulation predicted by bidomain modeling. Circ Res 84(2):153-156, 1999.)

0

2.5

5.0

–120

Figure 3-33  Transmembrane potential was calculated using the bidomain model with the modified action potential model for a 0.15-mA anodal stimulus pulse of 10-msec duration. The time (in msec) since the beginning of the stimulus pulse is marked in the top left-hand corner. The color bar on the right side shows the color scale used in the panels (in mV). Cardiac fibers are aligned along the x-axis in the figure. The peak hyperpolarization is −404 mV, and the peak depolarization is 47 mV. (From Ranjan R, Tomaselli GF, Marbán E: A novel mechanism of anode-break stimulation predicted by bidomain modeling. Circ Res 84(2):153-156, 1999.)

86

SECTION 1  Basic Principles

0

3

6

9

12

15

CM

A

15 mV

AM

B

–85 mV

CB

–185 mV

C

AB

D Figure 3-34  Computer Simulation of Four Types of Excitation With Stimulating Electrode (Black Rectangle). Cathode make (CM) (A), anode make (AM) (B), cathode break (CB) (C), and anode break (AB) (D) excitation are shown at various times; columns are time (t) in milliseconds. In A and B, t = 0 when the stimulation pulse turns on, and in C and D, t = 0 when the stimulation pulse is turned off. (From Wikswo JP, Roth BJ: Virtual electrode theory and pacing. In Efimov IR, Kroll MW, Tchou PJ, editors: Cardiac bioelectric therapy: mechanisms and practical implications. New York, 2009, Springer Science, pp 283-330.)

STRENGTH-DURATION RELATIONSHIP Early experiments by Hoorweg in 1892111 and Weiss in 1901112 suggested that there was a linear relationship between the intensity of a stimulus (measured in volts) and the duration of the stimulus that would excite (capture) nerves and muscles.113,114 Using a voltage source, a galvanometer, and a low-leakage capacitors to conduct quantitative stimulation studies, Hoorweg111 found that the voltage at which the capacitor must be charged to cause depolarization of nerves and muscles is inversely related to the capacitance of the capacitor: Vc = aR +

b C

where Vc is the threshold voltage to which a capacitor of capacitance C must be charged to produce stimulation, R is the resistance of the circuit through which the capacitor is discharged, and a and b are coefficients that varied with the experimental tissue. Hoorweg determined that there was only one specific capacitance value for which the threshold charge was a minimum. Weiss determined that for relatively long duration stimulation pulses, there seemed to be a lower limit of stimulus intensity that would result in stimulation.112,114 In fact, as demonstrated by Lapicque in 1909,115 the relationship between stimulus intensity and pulse duration is an exponential curve (Fig. 3-35A). The typical model of a cell membrane is that of a resistor and capacitor in parallel.116 When the current flowing from a pacing stimulus (I) encounters the myocyte membrane, it is divided by the resistance (R) and the capacitance (C) of the membrane. The total current (IT) equals the current flowing through the resistor (IR) plus the current flowing through the capacitor (IC), I T = I R + IC

As current that flows in the circuit is determined by the voltage of the pulse (V), the resistance of the membrane (R), and the capacitance (C) of the membrane, a parameter that increases exponentially over the duration of the pulse (t), the total current is given by the equation: I=

V dV +C R dt

By integrating these parameters, the following equation can be determined: −t

V = IR(1 − e τ ), where τ is the time constant of the membrane. The interaction of the stimulus intensity (measured in mA) and the duration of the pulse (measured in msec) was described by Lapicque,115 who demonstrated that the current required to stimulate myocardium increased exponentially as the pulse duration was decreased. At the other end of the curve, as the pulse duration is increased there exists a minimum current that will result in stimulation of the tissue.117 This exponential relationship is known as the strength-duration curve. The mathematical description of this relationship is given by the equation: I(t) =

Irheobase −t

1− e τ

where I is the stimulus current at pulse duration (t), Irheobase is the threshold current at an infinitely long pulse duration (known as rheobase), and τ is the membrane time constant. From the above equation, when the threshold current (It) is equal to twice the rheobase current



CHAPTER 3  Stimulation and Excitation of Cardiac Tissues

I–

rge

a Ch

Current Charge

d I = b(l+c/d)

turns upward, as seen at stimulus duration 2 msec. The effect of polarization on reducing current flow in constant-voltage pacing does not occur with constant-current pacing because the pulse generator develops whatever voltage is necessary to force the set level of current into the circuit (Fig. 3-37). The effect of polarization on reducing current flow in constant-voltage pacing does not occur with constantcurrent pacing, because with the latter the pulse generator develops whatever voltage is necessary to force the set level of current into the circuit. The slope of the strength-duration curves for passively fixed atraumatic leads does not change with time after implantation (Fig. 3-38). The acute, peak, and chronic mean thresholds of five ventricular and three atrial, passively fixed, atraumatic leads in 77 canines had a (log10) slope of about 0.60 ± 0.07 V/msec.119 The correlation coefficient for these relationships was typically more than 0.99 at pulse durations of less than 0.5 msec for polished electrodes and at durations of less than 1.0 msec for low-polarizing designs. Traumatic electrodes, such as those with active fixation, typically had a lower slope immediately after implantation, which shifted to about the same value of 0.60 V/msec within days as the acute trauma to the electrode-tissue interface healed. Because strength-duration curves are typically plotted on linear rather than logarithmic coordinates, changes in voltage and current threshold after implantation are usually seen as a shift upward and to the right. In most cases, because it is related to the slope of the curve (the time constant of the tissue), chronaxie does not change significantly with time.

)

(Q

Q = b(d+c)

2b b

0

Current (I)

Rheobase

c = Chronaxie

A

Current (I), Charge (Q), Energy(U)

Duration (d) I, (xb) Q, (xbc) U, (xb2rc) 100 50 I = b(I+ c ) d 20 Q = bd(I+ c ) 10 d 5 U = b2rd(I+ c )2 d 2 Charge (Q) 1 0.5 0.001 0.01

B

I– d

STIMULUS CHARGE-DURATION RELATIONSHIP

Energy (U) Current (I) 0.1

87

1.0

10

100

It should always be remembered that implantable pacemakers and ICDs are powered by a battery with a fixed quantity of available charge. Therefore minimizing the amount of charge used for stimulation of the heart is an important determinant of device longevity. For a pacing stimulus, the charge (Q) that is delivered is the product of current (I), measured in charges per second, and pulse duration (d), measured in msec, such that:

Duration (d/c)

Figure 3-35  A, Lapicque hyperbolic strength-duration (d) curve for current (I) and the Weiss linear strength-duration relationship for charge (Q) = Id. B, Universal strength-duration curves, plotted logarithmically, for current energy and charge, with the duration axis divided by c, the chronaxie. (From Geddes LA: Accuracy limitations of chronaxie values. IEEE Trans Biomed Eng 51(1):176-181, 2004.)

−t

(2 × Irheobase), the term e τ = 0.5 . Also from these equations, it follows that by defining these two values of rheobase (the threshold current at infinitely long pulse duration) and chronaxie (the pulse duration at twice rheobase current, or 0.693 τ), the entire strength-duration curve can be determined. The strength-duration relationship predicted by these equations is graphically presented in Figure 3-36). The exponential strength-duration relationship is such that increasing the pulse duration to 10 times chronaxie yields a stimulus intensity that is only 10% higher than rheobase. Coates and Thwaites estimated acute strength-duration curves for 101 passive fixation atrial leads and 224 passive fixation ventricular leads at the time of permanent pacemaker implantation in 229 consecutive patients.118 The mean value for chronaxie was 0.24 ± 0.07 msec in the atrium and 0.25 ± 0.07 msec in the ventricle. The observed rheobase voltage was 0.51 ± 0.2 V in the atrium and 0.35 ± 0.13 V in the ventricle. Although the voltage and current strength-duration relationships are usually plotted on linear coordinates, the relationship is really logarithmic119 (see Figs. 3-35 and 3-36). When plotted in this manner, the strength-duration curve for constant-current stimuli is now seen to be a straight line up to rheobase, whereas the constant-voltage line

Q = Id =

Irheobase −d

d

1− e τ

The graphical plot of charge versus pulse duration shows that the charge with each stimulus decreases rapidly as the pulse duration is shortened before approaching a minimum asymptote (Qmin) (see Fig. 3-36).116 This minimum charge is related to the membrane time constant, and therefore chronaxie, by the equation: Q min = τIrheobase Thus in terms of charge, and therefore battery drain, a pulse duration of less than τ does not save significant charge drained from the battery. Reducing the pulse duration to only 1/10th of chronaxie produces a charge that is only 10% above the minimum. The charge equation can also be rearranged to determine the energy at each point on the strength-duration curve. As energy (E) is equal to the product of voltage (V), current (I) and pulse duration (d): E = VId =

I2 d R

The minimum energy on the strength-duration curve is predicted to be 1.25 τ. This is also the point where the strength-duration curve for capture and the charge versus pulse duration curve intersect (see Fig. 3-35B). As a practical point, when programming the pulse duration of a pulse generator, chronaxie is an excellent choice to minimize energy, limit charge drained from the battery, and provide a stimulus

88

SECTION 1  Basic Principles

1000.0

I–

500.0

d

I=

b

; Q = Id

I – e–d/τ

U = I2 Rd

200.0 100.0

I, Q U

50.0

20.0 U

10.0

=

En

er

gy

Udmin= I 25τ

5.0

2.0 1.0 0.5 0.001 0.002

τ b = Qmin

I = Current

Q = Charge

0.005 0.01 0.02

0.05

0.1

0.2

0.5

1.0

2.0

5.0

b = Rheobase

50.0 100.0 200.0 500.0

10.0 20.0

Normalized duration d/τ Figure 3-36  Theoretical Strength-Duration Curve for Current (I), Energy (U), and Charge (Q). The current asymptote is b; the charge asymptote is τb (where τ is the membrane time constant) and the duration for minimum energy is τb. (From Pearce JA, Bourland JD, Neilsen W, et al: Myocardial stimulation with ultrashort duration current pulses. Pacing Clin Electrophysiol 5(1):52-58, 1982.)

Constant voltage

B

Constant current

Figure 3-37  The voltage and current waveforms obtained with a constant-voltage stimulator (small resistance) (A). Note the peak on the current waveform. In B are shown the voltage and current waveforms when using a constant-current stimulator (large resistance) delivering a rectangular current waveform. (From Geddes LA: Accuracy limitations of chronaxie values. IEEE Trans Biomed Eng 51(1):176-181, 2004.)

amplitude within the limits that can be provided by the pulse generator (see Fig. 3-35). There are several clinically relevant aspects to the strength-duration relationship: first, whereas true rheobase occurs at a pulse duration of approximately 10 msec, the practical value of increasing the pulse duration beyond approximately 1.5 msec is minimal because this is on the flat portion of the curve and there is very little decrease in the threshold amplitude beyond this point; second, when the pulse duration is shortened below approximately 0.20 msec, the stimulus amplitude encounters the steeply rising portion of the exponential curve such that the stimulus amplitude required to capture may begin to exceed the maximum amplitude that the pulse generator can deliver. Because of these facts, the capture threshold should always be noted as the combination of amplitude and pulse duration that were used to determine

Threshold (mA)

Voltage

Current

10 8 6 4

10 8 6 4

Threshold (volts)

A

2 1 .8 .6 .4

x x

2 1 .8 .6 .4

x x

.1

.3

.5 .8 1.0 2.0

Chronic

x x

x x

Peak (3 weeks)

x x

x

x

Implant

.2

.2

x x

x

.1

.3

.5 .8 1.0 2.0

Pulse width (msec) Figure 3-38  Logarithmic plots of canine ventricular constant-current (left) and constant-voltage (right) strength-duration curves for a passively fixed, atraumatic, unipolar lead with a polished platinum, 8-mm2 ring tip at various times after implementation (three different animals). (From Stokes K, Bornzin G: The electrode-biointerface [stimulation].  In Barold SS, editor: Modern cardiac pacing, Mt. Kisco, NY, 1985, Futura, p 41.)

myocardial capture, (e.g., 0.5 V at 0.3 msec, or 0.4 V at 0.5 msec, etc.). From a clinical standpoint, determining the voltage threshold at a pulse duration of 0.3 to 0.5 msec and programming the amplitude twice the threshold value is an effective method for managing stimulus amplitude while providing an excellent safety margin. It should also be recognized that the value of chronaxie, and therefore the strength-duration curve, depends of the homogeneity of the tissue and the proximity of



CHAPTER 3  Stimulation and Excitation of Cardiac Tissues

CONSTANT-CURRENT VERSUS CONSTANT-VOLTAGE PULSES There are two basic types of output circuits that have been used for artificial electrical stimulation: constant-voltage and constant-current. For a constant-voltage stimulator with large output capacitance, the leading edge voltage will be constant regardless of the impedance of the load connected to it. In other words, a constant-voltage stimulator has low output impedance. In contrast, a constant-current stimulator has high output impedance and generates the same current regardless of the load (see Fig. 3-37). The constant-voltage pulse generator produces a current waveform that is different in shape from the voltage waveform (Figs. 3-37 and 3-39). The current waveform depends on the impedance of the electrode and tissues which contain nonlinear resistive and capacitive elements. A constant-voltage generator produces a spike in current at the start of the pulse and a reversal of current at the end of the pulse. These distortions in the current waveform are due to the capacitive elements at the electrode-electrolyte interface. In contrast, the current waveform of a constant-current pulse is not distorted (see Figs 3-37 and 3-38). However, the voltage waveform of a constantcurrent pulse demonstrates an exponential rise in voltage at the onset of the pulse and an exponential decay at the end of the pulse due to the capacitive nature of the tissue-electrolyte interface. Impedance is not constant during an electrical stimulus as ions of opposite polarity increase during the pulse (polarization) (Fig. 3-40). Modern electrodes are designed to minimize polarization by the use of such materials as iridium oxide, platinized platinum, or activated carbon surfaces. The greater the polarization of the electrode during the pulse, the higher value of chronaxie that will be measured. In contrast to the discussion above for external stimulators with large output capacitors, almost all permanent pacemakers and ICDs use a capacitively coupled, constant-voltage waveform and much smaller output capacitance. Because the output capacitance is much smaller than for external stimulators, the charge on the capacitor declines during the pulse as current is delivered into the tissue. Thus over the course of the pulse there will be a decrease in the voltage such that the leading edge voltage (Vle) will be greater than the trailing edge voltage (Vte). This is associated with a fall in current from the leading edge to the trailing edge. And although the leading edge voltage is constant, the droop in voltage and current from leading edge to trailing edge is inversely related to the impedance of the circuit (see Fig. 3-39). The difference between leading edge and trailing edge voltage is directly related to the amount of charge transferred out of the capacitor: Vle − Vte = Q* C pg where Vle is the leading-edge voltage, Vte is the trailing edge voltage, Q* is the amount of charge removed from the output capacitor (measured in coulombs), and Cpg is the capacitance of the output capacitor (measured in farads). Thus the term “constant-voltage” pulse is not technically accurate. Because less charge is removed from the output capacitor, the voltage droop is less steep for a lead with high impedance at the electrode-tissue interface than it is for a lower impedance lead. In addition, the higher capacitance of the output capacitor, the lower the droop in voltage over the pulse.

STRENGTH-INTERVAL RELATIONSHIP The voltage and current thresholds vary as a function of the coupling interval between the stimulus and a train of preceding beats.107,123-125 Figure 3-41 shows a typical ventricular constant-current strength interval curve. At relatively long coupling intervals (>270 msec), the

Delivered voltage

Resultant voltage

5V

Resultant current 20mA

5V

Delivered current 20mA Constant current

Constant voltage

Figure 3-39  Voltage and current waveforms from a capacitively coupled, constant-voltage implantable pulse generator, the Medtronic Model 7000A (left), and the constant-current mode output of an external research stimulator, Medtronic Model 1356 (right). A unipolar lead with a polished platinum, 8-mm2 ring tip is used with a 900-mm2 titanium anode in 0.18% NaCl solution. The voltage across the lead is measured against a 100-mm2 Ag/AgCl electrode to eliminate the anode’s polarization from the waveform. (From Stokes K, Bornzin G: The electrode-biointerface [stimulation]. In Barold SS, editor: Modern cardiac pacing, Mt. Kisco, NY, 1985, Futura, pp 33-78.)

40 36 32 28 Volts and milliamperes

the electrode to excitable tissue.120 The chronaxie value is different for muscle and nerves as well as for complex tissues such as atrial and ventricular myocardium, the AV node, and Purkinje fibers.121,122

89

24 Leading edge

20 16 12 8

Trailing edge Leading edge

Voltage

4

Current

Trailing edge

0 100

1000

10,000

Resistance (ohms) Figure 3-40  Effect of Pacing Impedance on Capacitively Coupled, Constant-Voltage Leading-Edge and Trailing-Edge Waveform Amplitudes. The leading-edge voltage remains constant as a function of pacing impedance at values of 200 Ω or greater (Medtronic Model 5950 pulse generator). The trailing edge of the voltage waveform changes slightly with impedance up to about  1000 Ω. Current falls significantly with increasing impedance. Any constant-voltage source may no longer be constant at very low pacing impedances, because the battery becomes unable to supply enough current to maintain a steady voltage. These very low impedance values are not likely to be encountered clinically in a properly functioning pacing system. (From Barold SS, Winner JA: Techniques and significance of threshold measurements for cardiac pacing. Chest 70(6): 760, 1976.)

90

SECTION 1  Basic Principles

7

9 8

6

7

5

5

Distal unipolar cathodal

4

Threshold (mA)

Current (mA)

6

Bipolar

3

Proximal unipolar anodal

2

4

3

Anode break

2

Cathode break

1 0 240

280

320

360

400

Delay (msec)

intensity of the stimulus required to capture the myocardium is relatively constant, approaching an asymptotic minimum. As the coupling interval is decreased, the intensity of the stimulus must be increased in order for capture to occur. This region where an increased intensity of the stimulus is required to capture the myocardium is the relative refractory period of the myocardium and approximates the terminal portion of the repolarization phase (phase 3) of the ventricular action potential. As the coupling interval is decreased further, a coupling interval is reached where a stimulus of any intensity will no longer capture the myocardium (the absolute refractory period). The shape of the strength-interval curve is different for cathodal and anodal stimuli (Figs. 3-42, 3-43, 3-44). Thus the late diastolic threshold is typically lower for cathodal than for anodal stimuli.107 However, there is a characteristic dip in the strength-interval curve during the relative refractory period for anodal stimuli such that at short coupling intervals the anodal threshold may be less than that of a cathodal stimulus.126 This dip in the anodal strength interval curve represents a brief period of hyperexcitability and has been associated with the induction of ventricular fibrillation when stimuli are delivered just after the peak of the T-wave.123,127,128 The strength-interval curves are different for make and break stimulation such that both cathodal and anodal make stimuli have a lower late diastolic threshold than cathodal and anodal break stimuli.107 However, the dip in the strength interval curve can be particular marked for anodal break stimuli.107 The coupling interval where the relative refractory period is encountered is directly related to the underlying basic drive cycle length. Thus at short pacing cycle lengths there is a shortening of the relative and absolute refractory periods in both the atrium and ventricle. For clinical purposes, the term effective refractory period refers to the longest coupling interval that will not capture the myocardium using a stimulus of 2 msec pulse duration and an amplitude that is twice late diastolic current or voltage threshold. WEDENSKY EFFECT In 1886, Wedensky demonstrated in a neuromuscular preparation that a strong stimulus produced a prolonged period of enhanced excitability

1

Anode make Cathode make 150

200

250

300

Coupling interval Figure 3-42  Strength-interval curves plotted as current (mA) as a function of the coupling interval after the preceding, normally conducted QRS. Curves are shown for anode break, cathode break, anode make, and cathode make stimuli. (Modified from Dekker E: Direct current make and break thresholds for pacemaker electrodes on the canine ventricle. Circ Res 27(5):811-823, 1970.)

Pulse duration (msec) 10 5 1.0 0.5 7 6 Threshold (mA)

Figure 3-41  Strength-Interval Curves. Unipolar distal cathodal, unipolar proximal anodal, and bipolar strength-interval curves during an acute study in a patient with a temporary bipolar lead (equal-sized cathode and anode). The bipolar and unipolar anodal refractory periods are equal and are shorter than the unipolar cathodal refractory period. (From Mehra R, Furman S: Comparison of cathodal, anodal, and bipolar strength-interval curves with temporary and permanent pacing electrodes. Br Heart J 41(4):468, 1979.)

5 4 3

Anode break

2 1

Anode make

100

150

200

250

Coupling interval msec Figure 3-43  Strength interval curves plotted for short rectangular anodal pulses of 0.5, 1.0, 5, and 10 msec pulse duration. Notice the marked dip in the anode make strength interval curve at coupling  intervals less than 170 msec, indicating a period of hyperexcitability.  The dotted line indicates, for each particular coupling interval, the lowest thresholds of either the anodal make or anodal break stimuli. (Modified from Dekker E: Direct current make and break thresholds  for pacemaker electrodes on the canine ventricle. Circ Res 27(5): 811-823, 1970.)



91

CHAPTER 3  Stimulation and Excitation of Cardiac Tissues

Pulse duration (msec) 50 20 4 1

1.0 0.9 0.8 Rheobase (V)

7

Threshold (mA)

6 5 4

0.7 0.6 0.5

Increment

0.4

Decrement

0.3 0.2 0.1

3

0.0 2250 2000 1750 1500 1250 1000

2

750

500

250

Cycle length (msec)

1

150

200

250

Coupling interval msec Figure 3-44  Strength-Interval Curves for Unipolar Cathodal Make Stimuli of Different Pulse Durations. Note the dip in the curve at short coupling intervals prior to an abrupt rise in threshold during the refractory period. (Modified from Dekker E: Direct current make and break thresholds for pacemaker electrodes on the canine ventricle. Circ Res 27(5):811-823, 1970.)

Figure 3-45  Pacing thresholds determined by gradually incrementing and decrementing the pulse amplitude until gain and loss of capture, respectively, are demonstrated in a patient with complete atrioventricular block. The pacing threshold was determined at cycle lengths of 2000, 1500, 1000, 750, and 500 msec and with a constant pulse duration of 2.0 msec. To prevent variation in cycle length during incrementing and decrementing pulse amplitudes, a backup pulse was delivered at 25 msec. Note that the threshold values determined in this manner, with increments and decrements of the stimulus amplitude, are similar. Therefore the Wedensky effect is marginal when the pacing cycle length is maintained at a constant value.

EFFECT OF PACING RATE ON STIMULATION THRESHOLD 129

of the nerve. Ventricular pacing stimuli that are subthreshold during late diastole may encounter a period of enhanced (supernormal) excitability when delivered near the end of the T-wave in patients with complete AV block.130 Although this is a demonstration of supernormal excitability and not a true Wedensky effect, Castellanos et al130 convincingly showed that a strong stimulus (15 times threshold) which captured the ventricle would sometimes allow subsequent stimuli that had previously been of subthreshold intensity to capture for several beats thereafter. Thus the effect described by Wedensky has been demonstrated in the human ventricle using strong stimuli and is not explained by supernormal excitability because it persists long past the end of the T-wave. In contrast, Engel et al were unable to demonstrate the Wedensky effect in patients with bradycardia.131 During clinical measurements of stimulation threshold, the pacing threshold that is measured when the stimulus amplitude is gradually decremented until loss of capture occurs may be lower than when it is measured by gradually increasing the amplitude from a subthreshold value until capture is achieved (Fig. 3-45). This hysteresis phenomenon has also been termed the Wedensky effect, though it is not what Wedensky described. In addition, the clinical observation in patients with bradycardias has been that very small decrements in stimulus intensity produce very little or no capture hysteresis. Langberg et al132 found no difference between the threshold to maintain capture as the stimulus amplitude is decremented and that measured when increasing the stimulus amplitude from a subthreshold value in patients at pacing cycle lengths greater than 400 msec. They concluded that the apparent Wedensky effect is most likely explained by asynchronous pacing into the relative refractory period.132 Swerdlow et al studied 60 Hz alternating current (AC) pulses applied through an electrophysiologic catheter that had been positioned at the right ventricular apex in patients undergoing ICD implantation.133 Using this AC current, intermittent right ventricular capture with bipolar AC current occurred at a mean of 55 ± 16 µA, whereas continuous capture required 68 ± 18 µA.133 It is likely that the intermittent capture in this model is the result of stimulation during the supernormal period of excitability.

The estimated rheobase voltage in the atrium increases at very rapid atrial rates (>225 bpm). Thus a train of antitachycardia pacing stimuli will often require a higher stimulus intensity in order to consistently capture when delivered at high rates.134 This increase in stimulation threshold at high pacing rates is likely to occur because the amount of action potential shortening at higher rates is not sufficient to prevent stimuli from encountering the relative refractory period. Similarly, Hook et al reported a significant increase in the ventricular pacing threshold (at least 1 V) as compared with the late diastolic threshold in 10 of 16 patients at a cycle length of 400 msec; when the cycle length was decreased to 300 msec, an increase in threshold was observed in 15 of 16 patients.135 The magnitude of the increase in voltage threshold was equal to 1 V at a cycle length of 400 msec but increased by as much as 4 to 9 V at a cycle length of 300 msec in 29% of patients. Because of the increase in threshold at very high pacing rates, the intensity of antitachycardia stimuli is usually programmed at a much higher voltage than is necessary for antibradycardia pacing. Furthermore, because antiarrhythmic drugs having use dependence are commonly prescribed for patients implanted with an ICD, programming the stimulus amplitude to the maximum is especially prudent for anti­ tachycardia pacing in these individuals. MONOPHASIC, BIPHASIC, AND TRIPHASIC WAVEFORMS Biphasic defibrillation shocks require a lower voltage, energy, and potential gradient across the myocardium for successful termination of ventricular fibrillation than do monophasic shocks.136 The effects of monophasic and biphasic waveforms on stimulation threshold in frog and rabbit myocardium have been studied by Knisley et al.137 At very long pulse durations (>20 msec) there is no difference in rheobase voltage between monophasic and biphasic waveforms. However, at shorter pulse durations (<5 msec), chronaxie is significantly lower for monophasic than for biphasic waveforms (Fig. 3-46). In addition, the energy requirements for excitation are significantly greater for biphasic waveforms. Despite their higher stimulation thresholds, biphasic waveforms have been extensively modeled for atrial pacing with floating

92

SECTION 1  Basic Principles

1.4

TABLE Pathologic Features of Chronically Implanted 3-2  Electrodes

*

Threshold (V/cm)

1.2

M B

2.52.5

1.0

2.5 *

0.8

55

5

0.6

10

20

0.4

1010

Acute Reaction Capillary dilation Increased endothelial permeability Interstitial edema PMN infiltrate Myocyte necrosis

Chronic Reaction Macrophage and giant cell infiltration Histiocyte infiltration Collagen deposition Myofibrillatory disarray Fatty infiltration Calcification

PMN, Polymorphonuclear leukocyte.

0.2 2

1

0.0 0

5

10

15

20

Threshold (V)

Duration (msec) Figure 3-46  Monophasic (M) and biphasic (B) strength-duration curves obtained in strips of frog ventricular myocardium superfused with a solution containing 3 mmol/L of potassium. At pulse duration of 20 msec, there is no significant difference in the rheobase threshold. Note that the M waveform produces a lower threshold than the B waveform at pulse durations of 5 msec or less. Therefore although the M and B waveforms have a similar rheobase, chronaxie is less with M than with B stimuli. (From Knisley SB, Smith WM, Ideker RE: Effect of intrastimulus polarity reversal on electric field stimulation thresholds  in frog and rabbit myocardium. J Cardiovasc Electrophysiol 3(3): 239, 1992.)

x x 3

x

1

x

x

xx

x

4

2 x 3 x 4 5 6

xx

0 0 138

intra-atrial electrodes. These studies suggest that, as compared with a bipolar monophasic waveform, an overlapping biphasic waveform applied to closely spaced floating atrial electrodes produces similar maximum current density. However, the region of the highest current density is confined to a smaller region. This suggests that an overlapping biphasic configuration may be a strategy to reduce the risk of bystander stimulation of the right phrenic nerve which has been a problem for single lead DDD pacing.139 Triphasic waveforms have been studied as a method to reduce after-potentials on the electrode and to improve detection of the evoked response.140 Using this “chargebalancing” approach with a positively charged first phase, a negatively charged second phase, and a third positively charged phase significantly improved detection of the evoked response and discrimination from after-potentials.140 The Electrode-Tissue Interface and Foreign Body Response It is well known that stimulation thresholds change as a function of time after lead implantation related to foreign body reaction in the myocardium (Table 3-2), typically rising to a peak value after several weeks (Fig. 3-47).116,141,142 With older polished electrodes, some patients had thresholds that evolved over longer periods, up to 6 months.143 Luceri et al144 observed that, after the acute rise in stimulation threshold, 43% of 120 patients had stable chronic thresholds for up to 8 years; 17% had chronic thresholds that decreased at a rate of 5% per year, and 19% had thresholds that rose at a rate of 14% per year; 20% had thresholds that varied widely around a stable mean. Also, in children temporary increases in pacing threshold during minor childhood illnesses have been documented.145 Modern pacing electrodes have a porous or microporous surface structure. Most also elute a glucocorticosteroid. The technological progression from polished to porous to microporous to steroid-eluting electrodes has significantly reduced the evolution of stimulation threshold as a function of time after implantation.146 In fact, as can be determined from the follow-up data available, the thresholds of porous and microporous steroid-eluting leads do not change significantly with increasing time after implantation.147-149 The reasons for threshold

2

4

8

12

Implantation time (wk) Figure 3-47  Canine ventricular voltage thresholds at 0.5 msec for 8-mm2 unipolar, transvenous, tined leads as a function of time after implantation. 1, Polished platinum ring-tip (manufacturer A); 2, polished platinum ring-tip (manufacturer B); 3, porous-surface platinum hemisphere (manufacturer C); 4, porous-surface titanium hemisphere (manufacturer B); 5, platinized Target Tip (manufacturer B); 6, steroid-eluting titanium porous-surface electrode (manufacturer B).

changes as a function of time and electrode design are to be found in the foreign body response to the electrodes. The inflammatory response begins with dilation of blood vessels and results in increased endothelial permeability. This increased perfusion and permeability of the blood vessel walls allows plasma to leak into the surrounding tissue, producing edema. Because plasma and serum are more conductive than myocardium, pacing impedance decreases almost immediately after lead implantation. After 2 to 3 days, a mixed cellular inflammation begins to appear in the tissues surrounding the electrode, and the stimulation threshold begins to rise. Tissue inflammation reaches its peak in about 1 week, with clearly evident interstitial edema and cellular necrosis. Pacing impedance typically reaches its minimum value at the inflammatory peak (about 1 week after implantation), after which it increases as the edema resolves. The stimulation threshold may or may not reach its peak at the same time as the nadir of pacing impedance. After the early cellular reaction, the inflammatory response is characterized by a gradual accumulation of macrophages that reach the electrode surface, adhere to the electrode, and become activated. Macrophages attempt to dispose of dead cells or foreign material by the combined processes of endocytosis and exocytosis. This results in the release of enzymes and oxidants into the tissue at the electrode-tissue interface. Macrophages differentiate into giant cells that surround the electrode surface. The process of exocytosis is replaced in 3 to 4 weeks by the development of a collagenous capsule that surrounds the electrode. If the electrode material is biocompatible, there is a stabilization of the stimulation threshold in the 4-6-week period.



CHAPTER 3  Stimulation and Excitation of Cardiac Tissues

A

B

C

D

93

Figure 3-48  Photomicrographs of Chronic Electrode-Tissue Interfaces. A, Capsule surrounding a polished electrode (upper left) has a relatively thick layer of activated phagocytic cells, including foreign body giant cells on the surface and macrophages further out. Collagenous material encapsulates these cells, with fibrous stringers extending outward into the myocardium. B, Myofibrillar disarray is seen between the collagenous capsule and normal myocardium. The capsule surrounding a porous electrode (upper right) appears to be thinner. However, the active cellular component of the capsule and some of the collagen have been removed with the electrode to facilitate trimming of the tissues. C, Microporous (Target Tip) electrode-tissue interface is shown at lower left. Note the thin, fibrous capsule over the external surfaces and at the ridges.  D, Some macrophages are seen deep in the Target Tip electrode’s grooves at higher magnification.

Chronic atraumatic polished electrodes typically have a layer of foreign body giant cells on their surface (Fig. 3-48). These are covered by a layer (or layers) of macrophages that can be surprisingly thick. This cellular component of the capsule is not an acute, transient phenomenon. It has been observed on electrodes up to 13 years after implantation in Stokes’ canine studies.149,150 These cells are covered by a layer of collagen that is oriented along the surface of the electrode. The myocytes adjacent to the collagen layers are disarrayed and are interspersed with collagen fibers that are radially oriented with respect to the surface of the electrode. There may also be circular “holes” in this layer of disoriented myocytes, which appear to be infiltrated with fatty material. Groups of macrophages with fused membranes (foreign body giant cells) enveloping bundles of myocardial fibers adjacent to the electrode were observed 1 to 4 weeks after implantation. The presence of this fatty infiltration, its location, and its severity depend on the stability, myocardial location, and shape of the electrode relative to the vector of myocardial contraction. Outside the disarrayed myofiber zone, one finds normally oriented myocardium. Therefore it is clear that the concept of a simple fibrous capsule (Irnich’s “d”)151,152 separating the electrode surface from viable myocardium does not completely describe this complex biologic response and its effect on pacing. The histologic findings of pacemaker and ICD leads in 70 patients coming to autopsy a mean of 4 years after implant have been studied by Dvorak et al.153 Atrial electrodes were most frequently anchored in the endocardium or at the endocardium/myocardium border whereas ventricular pacemaker (PM) or ICD electrodes were most frequently in the myocardium. Ventricular ICD electrodes were more often situated in the myocardium than pacemaker electrodes. A chronic inflammatory cell infiltration was usually present in the sheath around the electrodes, including leukocytes, erythrocytes, histiocytes, and macrophages (Fig. 3-49). Foreign body giant cells, particles of foreign material, and sporadic calcifications could sometimes be found. The sheath

was thicker around ventricular ICD electrodes (1.08 ± 0.60, 0.302.50 mm) than around ventricular PM electrodes (0.76 ± 0.43, 0.102.30 mm, P < 0.04). Interestingly, the thickness of the sheath around steroid-eluting electrodes was no different than around nonsteroideluting electrodes (0.72 ± 0.28 vs. 0.79 ± 0.51). Effect of Corticosteroids on the Evolution of Pacing Thresholds Glucocorticosteroids have important antiinflammatory effects by stabilizing the membranes of phagocytes through the interaction with surface receptors, and the release of lysosomal contents.154 Dexamethasone has no significant direct electrical effects on myocyte membranes.155,156 Thus the effect of glucocorticosteroids on stimulation threshold is likely the direct effect of its inhibition of the release of inflammatory mediators from the cellular infiltrate surrounding pacing electrodes. Oral or parenteral corticosteroids lower pacing threshold within hours of administration though the effect is reversible. The development of steroid-eluting electrodes or collars around the electrode has been a very significant advance in cardiac pacing that has markedly reduced the incidence of exit block as a clinically significant problem.150 Steroid-eluting electrodes prevent the acute inflammatory response and minimize the rise of threshold from implant to the chronic phase. However, it should be recognized that electrodes that produce myocardial trauma may still produce a significant rise in pacing threshold. The clinical experience with the Telectronics 330-801 active fixation atrial lead demonstrated that a chronic rise in stimulation threshold was frequently observed,157 probably as a result of a large, heavy distal electrode and a stiff, J-shaped retention wire that produced significant trauma to the atrial endocardium. Steroid-eluting electrodes have a long and impressive track record in cardiac pacing for markedly limiting the rise in threshold over time.119,149,150,157-164 In early studies of both atrial and ventricular

94

SECTION 1  Basic Principles

M M E

EL EL

M

E

F

A

C

B

M

F

D

Figure 3-49  A, Fibrous sheath (F) around a pacemaker ventricular electrode (EL) anchored in the myocardium (M) composed of dense to hyaline connective tissue. Two-headed arrow indicates the maximum width of the fibrous sheath. Haematoxylin–eosin stain, magnification ×40. B, Fibrous sheath (F) around an atrial pacemaker electrode (EL) in the endocardium (E). C, Thrombus from an electrode surface with foreign body giant cells (arrows) surrounding an atrial electrode 8 days after implantation. Fibrin–Weigert stain, magnification ×400. D, Necrotic muscle fibers (arrows) and mixed inflammatory infiltrate near a ventricular pacemaker electrode 4 days after implantation. Haematoxylin–eosin stain. Magnification ×400. (From Dvorak P, Novak M, Kamaryt P, et al: Histological findings around electrodes in pacemaker and implantable cardioverter-defibrillator patients: comparison of steroid-eluting and non-steroid-eluting electrodes. Europace 14(1):117-123, 2012.)

steroid-eluting leads, the acute rise in threshold that is characteristic of nonsteroid leads was markedly reduced.149,157,159,165-169 Mond et al demonstrated that the threshold plotted versus time was flat for 98% of steroid-eluting leads.157 A randomized trial compared Biotronik Siello T steroid-eluting ventricular leads (n = 42) with similar nonsteroid leads (n = 46) and Siello JT steroid-eluting atrial leads (n = 24) with similar nonsteroid leads (n = 27).163 Patients with nonsteroid atrial leads had significantly higher threshold compared with steroid leads at 1-week and at 1-, 3-, 6-month follow-up with a peak at 1 month (1 month 1.4 ± 0.6 vs. 0.7 ± 0.3 V at 0.4 msec, P < 0.001; 6 month 0.3 ± 0.5 vs. 0.2 ± 0.3 V at 0.4 msec, P = 0.002). Patients with nonsteroid ventricular leads had a significantly higher threshold compared with steroid leads at 1-, 3-, 6-month follow-up (6 month 1.0 ± 0.3 vs. 0.6 ± 0.2 V at 0.4 msec, P < 0.001). In a multicenter, controlled trial that examined the performance of standard steroid-eluting pacing leads in the atrium (Medtronic model 5524) and ventricle (Medtronic model 5024) versus high impedance steroid-eluting pacing leads in the atrium (Medtronic model 5534) and ventricle (Medtronic model 5034) among 609 patients, there were no significant differences in pacing thresholds for the control and high impedance during a follow-up of12 months.169 The mean impedance of the high impedance pacing leads in the atrium and ventricle at 12 months was 992 ± 175 and 1,080 ± 220 Ω, compared with 522 ± 69 and 600 ± 89 Ω for the standard pacing leads in the atrium and ventricle (P ≤ 0.001 for the high impedance leads compared with standard leads in each chamber). High impedance leads were associated with lower lead current drain

than standard pacing leads in the atrium and ventricle for up to 1 year. Bipolar, steroid-eluting epicardial pacing electrodes have demonstrated excellent performance in pediatric patients, a group prone to leadrelated complications.164,170-173 Among 82 consecutive pediatric and adult patients with epicardial, bipolar, steroid-eluting leads, the longterm threshold was a mean of 0.75 V at 0.4 msec in the atrium and 1.0 V on both the left and right ventricles.164 As a result of these dramatic effects on stimulation threshold, steroid elution has become the standard of care in cardiac pacing.150

PHYSIOLOGIC AND PHARMACOLOGIC EFFECTS ON STIMULATION THRESHOLD Hypoxemia and acidosis produce an increase in stimulation threshold such that pacing stimuli may not capture. This is seen clinically in patients who have respiratory arrest, where pacing stimuli may not capture the myocardium until adequate ventilation and pH balance are restored (Table 3-3). Ischemia produces variable effects on stimulation threshold, depending on the location of the pacing electrode relative to the ischemic myocardium. In the presence of acute myocardial ischemia, the resting membrane potential decreases (cells become partially depolarized), the action potential upstroke velocity decreases, and the action potential duration dramatically shortens. In the presence of metabolic blockade with 2,4-dinitrophenol, there is an upward shift in the strength-duration curve.174 Therefore if the stimulating



CHAPTER 3  Stimulation and Excitation of Cardiac Tissues

ventricle), activation of the sympathetic nervous system may actually reduce the pacing threshold. Hyperkalemia increases the stimulation threshold when the serum potassium concentration exceeds 7 mEq/L.175-177 The reduced excitability during hyperkalemia can be reversed by the intravenous administration of calcium.175 Hypokalemia may decrease the stimulation threshold and restore capture to a subthreshold stimulus.178 The infusion of 3% NaCl solution to achieve a mean serum Na+ concentration of 158 mEq/L produced no changes in atrial or ventricular threshold while current increased 21% to 32%.177 This resulted in a 40% to 55% increase in energy due to a 20% decrease in impedance. Hyperglycemia that is severe (>600 mg/dL) may raise the pacing threshold by as much as 60%.179 Therefore patients with renal failure or diabetic ketoacidosis may experience loss of capture due to the combined effects of hyperglycemia, hyperkalemia, and acidosis. Hypothyroi­dism has also been demonstrated to raise the stimulation threshold, an effect that is reversible with thyroid hormone administration.180,181 Glucocorticosteroids may have a dramatic effect to lower stimulation threshold (Fig. 3-50), either during the acute phase after lead implantation or chronically.74,182 In contrast, the mineralocorticoid aldosterone increases threshold.183 Catecholamines, whether endogenous or synthetic, may significantly lower the stimulation threshold.183-185 Preston et al183 reported that oral ephedrine produces a 10% to 20% reduction in stimulation threshold. Intravenous or even oral isoproterenol can be used to increase excitability in the acute management of exit block or to reverse the effects of antiarrhythmic drugs.183,186 Propranolol has been demonstrated to increase the stimulation threshold when administered to a physically active patient but not when the patient is in the resting state.183 Several antiarrhythmic drugs have been shown to increase the pacing threshold, including ibutilide,187 procainamide,188 quinidine,189 propafenone,190,191 flecainide,192 encainide,193 moricizine,194 amiodarone,195 and β-blockers.196 In a canine right atrial preparation, amiodarone increased the diastolic threshold from 0.17 ± 0.02 mA to 0.31 ± 0.03 mA.197 The effect of amiodarone on the ventricular stimulation threshold has been reported to be less significant than in the

electrode is located in an ischemic region, the stimulation threshold would be expected to increase. With further ischemia and infarction, the myocardial threshold may rise dramatically. This may be seen clinically in patients who develop an acute inferior myocardial infarction (MI) with right ventricular (RV) infarction; in such patients, a previously implanted pacemaker may suddenly lose capture. However, if the stimulating electrode is located in a nonischemic region (e.g., right

TABLE Clinical Factors Affecting Stimulation 3-3  Threshold Factor Electrode radius Output capacitor Hypoxemia Acidosis Glucocorticosteroids Mineralocorticoids Type 1A drugs Type 1B Type 1C Sotalol and dofetilide Amiodarone Ibutilide Type 1C Type 1B Sotalol and dofetilide Eating Sleeping Exercise Catecholamines Hyperkalemia Hypokalemia Hyperglycemia

95

Effect on Threshold Larger radius +, Smaller radius − Smaller capacitance +, Larger capacitance − ++ or +++ ++ or +++ − + ↑ +/− ++ Slight ↓ or +/− + + + +/− − or +/− + or +/− + or +/− − − ++ − + or ++

% 30 25 d-aldosterone, 500 µg

Percent change in threshold

20 15

I.V. injection

10 5 0 –1

0

1

2

3

4

5

6

7

8

9

10

11

–5

12

13

14 15 Hours

–10 –15

Methylprednisolone, 40 mg

–20 –25 Figure 3-50  Effect of Methylprednisolone 40 mg Orally on Ventricular Stimulation Threshold. Note that the threshold decreases by 25% within 2 hours. The administration of aldosterone 500 µg increases the stimulation threshold with a peak effect at approximately 5.5 hours. (From Preston TA, Fletcher RD, Lucchesi BR, et al: Changes in myocardial threshold: physiologic and pharmacologic factors in patients with implanted pacemakers. Am Heart J 74:235-242, 1967.)

96

SECTION 1  Basic Principles

atrium.195,198 The effect of type 1C drugs on pacing threshold can be clinically significant even with a medium dose of these drugs. In a study of patients receiving oral propafenone, a relatively moderate propafenone dosage of 450 mg a day increased the pulse duration threshold (measured at 2.5 V) from 0.14 ± 0.10 to 0.21 ± 0.16 msec (+55%) in the atrium and from 0.10 ± 0.08 to 0.15 ± 0.09 msec (+63%) in the ventricle.191 Among patients who received propafenone 900 mg a day, the mean atrial and ventricular threshold increased from 0.12 ± 0.10 to 0.17 ± 0.14 msec with the lower dose and to 0.27 ± 0.22 msec (+125%). The rise in threshold with type 1C drugs may be at least partially reversed by the intravenous administration of sodium bicarbonate,199,200 insulin,201 and by 3% NaCl. The ventricular pacing threshold has not been found to increase and may slightly decrease with IKr blocking drugs such as sotalol202 or dofetilide.203 Whether stimulation thresholds demonstrate significant variation throughout the day has been debated. An increase in threshold during such activities as eating or sleeping has been reported.183 Grendahl and Schaanning reported minimal variation in threshold during the day, after meals, or during sleep and physical activity.204 In addition, others have found that diurnal variation in threshold is minimal.205 Vagal stimulation significantly shortens the effective refractory period in both the right atrium and left atrium206 but atropine has been reported to leave the stimulation threshold unchanged.183 The clinical experience with automatic threshold detection and amplitude adjustment algorithms has shown minimal variation in threshold, suggesting that diurnal effects are small.207

DESIGN FACTORS THAT AFFECT THE PERFORMANCE OF PACING ELECTRODES The performance of the electrodes used for pacing is an important determinant of stimulation threshold, impedance, sensing, and the long-term evolution of these parameters. The size, geometric shape, surface structure, materials, and elution of pharmacologic agents each have a major impact on lead performance over time. GEOMETRIC SURFACE AREA Stimulation threshold is inversely related to the size of the stimulating electrode.151,208,209 The stimulation threshold of a spherical electrode decreases as it radius and geometric surface area decreases. The leading edge ohmic resistance at the electrode tissue interface is inversely related to the surface area of the electrode. The electric field strength (E) necessary for stimulation is a nonlinear function of the electrode radius (r) and of the potential (V) applied. The tissue in contact with the electrode undergoes a foreign body inflammatory response that results in the development of a fibrotic capsule around the electrode.210,211 This fibrotic capsule is conductive but is not excitable.212 Thus for a chronically implanted lead, an electrode of radius (r) is usually surrounded by a fibrotic capsule of thickness (d) separating it from the excitable myocardium. This forms a “virtual electrode” of radius equal to r + d.212,213 With a voltage that is constant, the field strength of an electrical stimulus in the excitable myocardium varies as an inverse function of the inexcitable capsule thickness. Thus the intensity of the field in the excitable myocardium declines as the thickness of the fibrotic capsule increases. Irnich152 modeled this theory as follows: E=

V r 2 r r+d

Irnich based this equation on measurements taken at a fixed pulse duration of 1 msec. Although this equation is missing the crucial parameter of the pulse duration for the strength-duration relationship, it does highlight the effect of the fibrotic capsule on the field strength seen by the underlying myocardium. If two electrodes having different radii (0.5 and 5 mm) were to develop a fibrotic capsule 1 mm thick

(d), the electrode with the smaller actual radius will be associated with a greater percentage increase ( r + d ) in virtual electrode size (300% r versus 20%). This equation helps to explain the early clinical observation that smaller electrodes were associated with a lower acute threshold than larger electrodes but tended to develop a greater rise in threshold over time.92,214 For polished platinum, spherical electrodes, when the thickness (d) exceeds the electrode radius (r), the threshold increases. The combination of steroid elution and a small geometric surface area has been effective in reducing chronic stimulation threshold while markedly increasing the impedance at the electrode-tissue interface.169,215 As compared with steroid-eluting electrodes with larger surface area (5.8 mm2), the Medtronic Capsure Z leads with surface area of 1.2 mm2 result in similar stimulation thresholds but significantly reduced current drain as a result of higher electrode-tissue impedance.169 The mean current was significantly lower with the smaller electrode than that of the other types of leads (0.42 µA for CapSure Z ventricular lead vs. 0.85 µA for CapSure SP, 1.42 µA for CapSure, and 1.54 µA for Target Tip).215 Scherer et al216 compared impedance and current drain with the Medtronic CapSure Z (1.2 mm2) and the CapSure SP leads (5.8 mm2). The CapSure Z leads provided a significant reduction in current drain, resulting in a reduction of mean energy consumption at the 12-month follow-up from 10.4 ± 5.0 µJ in the CapSure SP group to 6.6 ± 1.4 µJ in the CapSure Z group (median from 9.9 µJ to 6.9 µJ respectively). This provided an estimated increase in mean longevity of more than 1 year from 81.1 ± 23.5 months in the CapSure SP group to 94.5 ± 13.4 months in the CapSure Z group (median: 76.5 months to 95.0 months, respectively). ELECTRODE SURFACE STRUCTURE Another design consideration is the relationship between electrode size and polarization. As the surface area of an electrode increases there is an increase in Helmholtz capacitance. Thus for a stimulus of a given charge (Q), the polarization (V) is related to the capacitance (C) of the electrode-electrolyte interface by the equation: V=Q C Increasing the Helmholtz capacitance of the electrode decreases the polarization. This is particular important for sensing of the evoked response and increases the reliability of autocapture algorithms. Thus an ideal electrode would have a small geometric size in order to increase current density, increase resistance at the electrode-tissue interface, and yet, have higher Helmholtz capacitance to minimize polarization. Modern electrodes are designed not to be spherical but to have porous or fractal coatings that greatly increase the surface area without increasing the overall radius of the electrode (Fig. 3-51).147 These highly engineered surface coatings produce areas of very high current density and have changed the relationship between electrode radius and threshold. The complex surface structure with microporous or fractal coatings greatly reduces polarization while maintaining an overall small geometric radius. Electrodes with a porous surface structure that allows tissue ingrowth have been shown in animal studies to have a thinner fibrous capsule than electrodes with a solid or polished surface.217-221 Although the value of a microporous electrode surface structure as a fixation mechanism is doubtful, the long-term performance of these leads for both stimulation and sensing has been excellent.221-223 Furthermore, modern electrodes (or collars around the electrode) often elute a corticosteroid which markedly reduces the inflammatory foreign body response and the thickness of the fibrotic capsule. These design features allow the benefits of a smaller electrode (increased electrode-tissue interface resistance) to be achieved without a marked rise in chronic threshold. Although this discussion has emphasized design features that reduce stimulation threshold at the cathode in contact with the myocardium, there are some electrodes, the anodes, which are designed to reduce the probability of unwanted stimulation. Whether the reference



CHAPTER 3  Stimulation and Excitation of Cardiac Tissues

A

97

B

C Figure 3-51  Electron Micrographs of Microporous Electrode Surfaces. A, Activated carbon surface of Siemens Model 412S/60 electrode at about 8000× magnification. B, Medtronic Model 4011 platinized surface at 6900× magnification. C, Polished platinum surface of Medtronic Model 6971 at 7000× magnification. The polished platinum surface has little microstructure. Its actual microscopic surface area is similar to its apparent or geometric surface area (8 mm2). The platinized platinum surface (B) is composed of particles so small that they absorb visible light, and the surface appears black. The true surface area of the interface is many orders of magnitude greater than the geometric surface area. (From Seeger W: A scanning electron microscopic study on explanted electrode tips. In Aubert AE, Ector H, editors: Pacemaker leads, Amsterdam, 1985, Elsevier Science, pp. 417-432.)

electrode (the anode) is the pulse generator casing (in the case of unipolar stimulation) or located more proximally on the lead (in the case of bipolar stimulation), the anode is made significantly larger than the cathode in order to reduce the anodal current density. In the case of unipolar stimulation, the very large surface area of the pulse generator casing minimizes the chances of stimulating muscle and nerve near the pocket. In the case of bipolar pacing, the larger surface area of the proximal ring electrode reduces, but may not eliminate, the possibility of anodal stimulation of the myocardium. ELECTRODE SIZE AND SENSING PERFORMANCE The sensing performance of an electrode is crucial for the function of pacemakers and ICDs. As discussed in detail in Chapter 4, a selfpropagating wavefront of depolarization in the myocardium generates a small amplitude electrical field that is sensed as the potential difference between an electrode in contact with the myocardium and a reference electrode located more proximally on the lead or the pulse generator casing itself. The amplitude of the electrogram that is sensed by the pulse generator is a function of the ratio of the source impedance (at the electrode-myocardial interface) to the input impedance of the sensing amplifier. In order to maintain as large a sensed electrogram as possible, the input impedance of the sensing amplifier is very high. Because small electrodes may be associated with low Helmholtz capacitance, electrodes are designed to have a very large surface area of using low polarization microporous or fractal coatings.

AUTOMATIC CAPTURE DETECTION ALGORITHMS Ventricular Capture Detection Algorithms The concept of an automatic capture algorithm that detects the ventricular evoked response following a pacing stimulus was proposed by Preston and Bowers in 1974.224 This approach utilizes a sensing window after the pacing stimulus to determine capture by analyzing whether the stimulus is followed by an evoked response (ER) (Fig. 3-52). The earliest iterations of automatic capture verification and adjustment of the pacing stimulus amplitude were hampered by the challenge of discriminating the ER from after-potentials. Thus if an after-potential was misinterpreted as an ER, the stimulus amplitude could be lowered below the actual pacing threshold with loss of capture.225 Similarly, fusion between the pacing stimulus and a conducted ventricular beat could also create difficulties accurately detecting whether capture had occurred.226,227 Indeed, if a premature ventricular depolarization occurred in the ventricular blanking period, a backup stimulus with increased amplitude could result in a closely coupled ventricular couplet that induced ventricular fibrillation (Fig. 3-53).228 Several automatic ventricular threshold algorithms have been introduced (Table 3-4). The first highly successful automatic ventricular capture algorithm (St. Jude Medical Autocapture) measured the depolarization in the electrogram within the first 100 msec after a 4.5 V and 0.5 msec unipolar test stimulus (Figs. 3-54 and 3-55). A second test stimulus of 4.5 V and 0.5 msec was delivered 110 msec after the first stimulus that was predicted to be within the refractory period of the ventricle. Any depolarization on the electrogram within 100 msec after

98

SECTION 1  Basic Principles

Capture

Non-capture Polarization artifact Evoked response

Figure 3-52  Concept of the Evoked Response (ER). Capture of the myocardium results in a combined evoked response that is a negatively charged deflection in the electrogram superimposed on polarization artifact. Noncapture results in polarization artifact without the ER. (Courtesy Biotronik, Inc., Berlin, Germany.)

N

VPD AP VP1 VP2 VF1

V

V

V

V

V

V

Figure 3-53  Ventricular Fibrillation Induced by a Backup Pacing Stimulus With the Autocapture Algorithm. A ventricular premature depolarization (VPD) occurs simultaneously with an atrial pacing stimulus such that it falls within the postatrial ventricular blanking period. A ventricular pacing stimulus falls within the physiologic refractory period of the ventricle following the VPD. As no evoked response (ER) was detected, a backup pulse of 4.5 V and 0.49 msec was then delivered at the peak of the T-wave resulting in polymorphous ventricular tachycardia (VT) that degenerated to ventricular fibrillation (VF). (From Sweeney MO, Porkolab FL: Ventricular fibrillation induced by double premature ventricular pacing stimuli in a dual-chamber pacemaker with AutoCapture. Heart Rhythm 6(3):429-432, 2009.)

TABLE Automatic Ventricular Threshold Features 3-4  Ventricular automated threshold Beat-to-beat capture determination Capture determination method Threshold search interval Automatic stimulus amplitude

St. Jude Medical Yes Yes PDI 8 hrs or LOC 0.25-0.30 V > Thr

Medtronic Yes No ER 15 min-7 days 2× Thr

Boston Scientific Yes Yes ER 21 hrs or LOC 0.5 V > Thr

Biotronik Yes Yes ER 10 min-24 hrs 0.3-1.2 V > Thr

Sorin Yes No ER 6 hrs 2× Thr

ER, Evoked response; LOC, loss of capture; PDI, paced depolarization interval; Thr, threshold voltage.

the second stimulus was presumed to be the result of polarization on the electrode and not a true evoked response.229 AutoCapture was not recommended if the ER amplitude was small (<2.5 mV), the polarization signal amplitude was large (>4 mV), the ER sensitivity/polarization ratio was <1.7, or the ER amplitude/ER sensitivity ratio was <1.8.230 If the ER was of sufficient amplitude and the polarization voltage did not exceed these limits, then the AutoCapture algorithm could be programmed to detect capture on a beat-to-beat basis and the amplitude was programmed at 0.3 V greater than the last automatically measured threshold. The AutoCapture system utilizes two sense amplifiers, including a standard amplifier that senses R-waves and a second, evoked response amplifier. After a pacing stimulus, the ER amplifier is blanked for the next 14 msec. Following this blanking period, the ER amplifier sensing

window is open from 14 to 62.5 msec. If no evoked response is detected within this ER sensing window, a backup stimulus with amplitude of 4.5 V and 0.49 msec is delivered 62.5 msec following the first pacing stimulus. A threshold search is initiated whenever the ER sense amplifier fails to detect an ER for two consecutive beats, every 8 hours in the absence of loss of capture, or manually as commanded by the programmer. The Microny and Regency pacemakers reduced the stimulus amplitude by 0.3 V until two consecutive beats where no ER is detected (each followed by a backup stimulus). The output is then increased by 0.3 V until capture is confirmed for two consecutive stimuli, which is defined as the stimulation threshold. For the dual chamber pacemakers including Affinity and Identity, the AV interval is shortened during capture testing to 25 msec for the sensed AV delay and 50 msec for a paced AV delay in order to reduce the chance that fusion with an



CHAPTER 3  Stimulation and Excitation of Cardiac Tissues

intrinsically conducted ventricular beat would be misinterpreted as capture. The pacing output is reduced stepwise by 0.25 V until loss of capture (again with a backup stimulus of 4.5 V at 0.49 msec). After two consecutive subthreshold stimuli, the stimulus amplitude is increased by 0.125 V until two consecutive stimuli with the same amplitude are followed by an ER (defined as the stimulation threshold). The stimulus amplitude is maintained at a value 0.3 V higher than the measured threshold for single chamber pacemakers and 0.25 V higher than threshold for dual chamber models (with beat-to-beat confirmation of capture and high voltage backup stimuli). The SJM Zephyr and later models calculate a paced depolarization integral to differentiate the ER plus polarization from polarization alone P/R detector Absolute refractory period ER detector

Open

Sense amplifier

Open 15 ms

47.5 ms

Figure 3-54  St. Jude Medical Microny AutoCapture Evoked Response Detection Algorithm. This device used separate sensing and evoked response (ER) detection amplifiers. Following the pacing stimulus, the ER detector was blanked for 15 msec. An ER detection window was open from 15 to 62.5 msec. The sense amplifier was blanked for the first 15 msec but then alert. (From Lau C, Cameron DA, Nishimura SC, et al: A cardiac evoked response algorithm providing threshold tracking: a North American multicenter study. Pacing Clin Electrophysiol 23(6):953-959, 2000.)

Rate timer starts

P/R detector

99

(Fig. 3-56). The ratio of the ER sensitivity to the maximum measured polarization must be >1.333 in order for the algorithm to operate successfully. Lau and colleagues studied the effect of five different electrode designs on detection of the ER and polarization with permanent bipolar pacing leads at the time of implantation.230 The tip electrode area ranged from 1.2 mm2 to 8.3 mm2 and the materials included titanium nitride (Ti-N), platinized platinum alloy, noncoated platinum, and iridium oxide (Irox). There was a moderate amount of variability in the sensed ER for any of the five electrode materials (ranging from 7.8 to 18.7 mV) that was not statistically significantly different. However, Ti-N electrodes (both active and passive fixation) had significantly lower mean polarization voltage than platinum active fixation helix leads (1.33 vs. 8.17 mV). There was no statistically significant difference in the polarization with the other electrode materials. As a result of these differences in polarization, AutoCapture was recommended for 91.7% of Ti-N, 60% of the 1.2 mm2 platinized platinum, 38% of Irox, and 0% of platinum helix leads. There was no correlation between the sensed R-wave, stimulation threshold, ER, and polarization voltage. The modern leads of most manufacturers are compatible with acceptable AutoCapture function.231,232 Among 207 leads from three different lead manufacturers, AutoCapture was recommended for 91% of leads at implant, 96% at 1 month, and 95% at 3-months follow-up.232 Despite these encouraging results, some leads may not be compatible with AutoCapture because of significant polarization. Based on the evoked response/polarization signal measurements, Kacet and colleagues231 reported that AutoCapture was recommended for use in 11% of patients with a Biotronik SX60BP lead, 67% with a Biotronik Y60BP lead, 63% with an Ela BT46 lead, 80% with a SJM 1388T lead, and 88% with a SJM 1450T lead. Although these leads are no longer manufactured, there are many patients with older leads who will undergo pulse generator replacement. For them, automatic capture verification may not be feasible. After a 20-msec blanking period following the ventricular pacing stimulus, the Biotronik Evia pacemaker attempts to differentiate the ventricular ER from polarization artifact during a 60-msec window. Analysis of the electrogram is performed in two separate phases. In both phases, the AV delay is shortened to 50 msec after a paced atrial event and to 15 msec after a sensed atrial event to ensure ventricular pacing. First, five ventricular pacing pulses are delivered at the programmable maximum voltage setting (2.4, 3.0, 3.6, 4.2, 4.8 V). If

Rate timer restarts

Absolute ref period

Absolute refractory period

Open

ER detector 15 ms

47.5 ms

4.5 volt 0.49 ms

Figure 3-55  St. Jude Medical Microny AutoCapture Evoked Response Detection Algorithm Backup Pace. Following a subthreshold test pacing stimulus, the evoked response (ER) detection window failed to detect an ER during the open window extending from 15 to 62.5 msec. A backup stimulus was delivered at 4.5 V at 0.49 msec that was followed by capture. (From Lau C, Cameron DA, Nishimura SC,  et al: A cardiac evoked response algorithm providing threshold tracking: a North American multicenter study. Pacing Clin Electrophysiol 23(6): 953-959, 2000.)

100

SECTION 1  Basic Principles

Closed

Open

11 ms

74 ms

ER detection

PDI-area of ER in V. EGM channel Blanking Evoked response detection window

A Capture confirmed

Loss of capture Open

ER sensitivity = 500 PDI = 750

B

PDI ≥ ER sensitivity value

ER sensitivity = 500 PDI = 150

PDI < ER sensitivity value

Figure 3-56  The St. Jude Medical Paced Depolarization Integral. A, The pacing stimulus is followed by a blanking period of 11 msec. The evoked response (ER) detection window then extends for the next 74 msec. The area of the negative deflection during the ER detection window is integrated to determine the paced depolarization integral (PDI). B, The PDI following a stimulus that is suprathreshold measures 750 msec. The ER sensitivity is 500 msec, and the PDI is 150% of the ER sensitivity, thereby recognized as capture. Noncapture is recognized by a PDI of only 150 which is less than the ER sensitivity of 500 msec, thereby recognized as noncapture. (Redrawn with permission of St. Jude Medical, St. Paul, MN.)

noncapture is detected at the maximum voltage setting, the second phase of signal analysis is aborted and the test is classified as unsuccessful. In the next phase, five “double” pacing pulses are delivered, one pacing pulse followed by another pacing pulse delivered 100 msec later during the absolute refractory period. These pulses are used to verify that the polarization artifact is small enough to distinguish capture from noncapture. The pacing amplitude decrements with every paced beat by 0.6 V, until the first noncaptured beat. The algorithm then decrements by smaller increments of 0.1 V until the first failed capture, at which point it determines that the capture threshold is the preceding value. The pacing amplitude is then automatically adjusted to a programmable safety margin above the measured threshold (0.5 V through 1.2 V in steps of 0.1 V). The long-term performance of the St. Jude AutoCapture algorithm has been reported by several authors.229,233-242 In a large, North American multicenter trial,229 398 patients were studied with the Auto­ Capture algorithm over a mean follow-up of 1 year. The automatic threshold feature was operative in 94.5% of patients at predischarge and 91.5% at 12 months follow-up with effective backup pacing in all patients and no adverse events. Backup pulse was delivered in 0.4% of beats and the stimulation threshold increased from implant to 12 months (0.54 ± 0.43 V at 0.31 msec vs. 0.89 ± 0.34 V at 0.31 msec pulse duration). The ER measured 8.92 mV at implant and increased slightly to 9.60 mV at 12 months whereas the polarization voltage remained stable (1.12 mV at implant and 1.18 mV at 12 months). Backup pulses were delivered due to fusion with native conduction in 0.3% of beats.

Backup pulses delivered after undersensing of effective capture was rare (0.001% of beats). Biffi and colleagues241 studied the effectiveness of automatic threshold algorithms from three pacemaker manufacturers in 321 patients with a mean age of 73 ± 12 years. This study group all had modern, steroid-eluting leads with either active or passive fixation mechanisms. Among 321 patients, there were 41 patients (12.8%) with a chronic threshold of greater than 1.5 V at 0.4 msec. The threshold was between 1.6 and 2.5 V in 9%, between 2.6 and 3.5 V in 2.2%, and greater than 3.5 V in 1.5% of these patients. Remarkably, there was no loss of capture in this population as all high threshold leads were successfully managed by automatic capture algorithms. Thus automatic capture algorithms provide a margin of safety even for patients with steroid-eluting leads. The AutoCapture algorithm has also been shown to be effective in children and in patients with epicardial pacing leads.173,234,237,242 The AutoCapture algorithm could be activated in 100% of 16 newborns with complete AV block and epicardial pacing leads at device implantation.242 However, during a mean follow-up of over 5 years, Auto­ Capture was discontinued in 47% of patients, primarily because of a decrease in the ER amplitude. The mean ER amplitude decreased from 9.3 mV at implant to 4.6 mV at 12 months follow-up. Among 56 patients with bipolar, steroid-eluting epicardial pacing leads, AutoCapture detected a mean daily fluctuation in stimulation threshold of 0.58 V for RV leads and 0.63 V for LV leads.173 The range in daily threshold fluctuation was between 0.16 and 2.92 V. In addition, the greatest variation in threshold was correlated with the highest baseline



CHAPTER 3  Stimulation and Excitation of Cardiac Tissues

threshold and lower pacing impedance.173 Thus automatic capture detection with beat-to-beat backup pulses may be especially important for children with epicardial leads as the stimulation threshold may demonstrate considerable daily variability. The Boston Scientific RV Automatic Capture (RVAC) algorithm is designed to detect ventricular capture on a beat-to-beat basis with the RV amplitude automatically adjusted to a value of 0.5 V higher than threshold at 0.4 msec pulse duration. These devices reduce polarization and increase the chances of accurately detecting the ER by reducing the coupling capacitance of the output capacitors to 2 µF.243 A stimulus that fails to capture is followed by a backup stimulus within 70 msec. If there is no loss of capture, an automatic threshold determination is made every 21 hours. These pulse generators utilize a separate unipolar (RV tip to can) sensing amplifier that is dedicated for sensing of the ER. Filtering of the ER signal is 3 to 65 Hz, less than for the normal sensing amplifier (25 to 85 Hz). The ER signal can be measured from 1 to 32 mV, though the ER must be >2 mV for RVAC to be programmed “on.” The ER is assessed for noise or fusion and if there is no noise or fusion, the peak positive amplitude is measured in a time window extending from 10 to 70 msec after the pacing stimulus. If the peak positive amplitude exceeds the capture detection threshold (CDT), capture is verified. As the ER is sensitive to respiratory motion, the change in ER due to respiration is calibrated at the initiation of RVAC. The CDT is averaged over 12 beats on a beat-to-beat basis and is set to 55% of the difference between the average ER amplitude (after adjustment for respiratory variance) and the polarization artifact. As fusion between an intrinsic R-wave and a paced beat can interfere with accurate ER detection, the RVAC algorithm actively attempts to differentiate this fusion from capture (Fig. 3-57). If a negative ER amplitude peak occurs within a noise window extending from 27 msec to 48 msec after a pacing stimulus, noise or fusion is suspected and the AV delay is extended (for dual chamber devices) or the interventricular (V-V) interval is lengthened (for single chamber devices) by 64 ms for the next cycle to search for intrinsic R waves. Thus the AV delay can be extended up to 64 ms longer than programmed maximum. If R-waves are sensed during the extended AV delay, fusion is confirmed. The AV delay (or V-V interval) then remains at the increased interval until a ventricular pacing stimulus occurs. If there is no sensed R-wave during the extended AV delay (or V-V interval), a ventricular pacing stimulus is delivered and the ER is evaluated for capture. This

101

morphology-based fusion detection cycle will occur four times before either testing fails or the beat-to-beat RVAC algorithm is suspended. The utility of the fusion management algorithm to discriminate fusion from capture was evaluated in 45 patients with the RVAC algorithm.227 Among over 4.4 million beats analyzed by Holter monitoring, there were no fusion beats classified as loss of capture and there were no unnecessary backup pacing stimuli delivered. Thus the positive predictive value of the fusion algorithm was 100%. The utility of the RVAC algorithm to minimize the need for reprogramming of the stimulation amplitude or pulse duration was assessed among 960 patients with the Boston Scientific Insignia pacemaker.244 This study was remarkable in that the ventricular leads were from nine different manufacturers, including 27% that had been chronically implanted. Over the first 12 months of follow-up, 95.9% of patients did not have reprogramming of their stimulation parameters.244,245 The need for reprogramming because of an inadequately functioning RVAC algorithm was only 0.4%. Medtronic pacemakers offer Ventricular Capture Management (VCM), an automatic capture detection and threshold algorithm that also detects the ER after a pacing stimulus. VCM searches for the capture threshold using a series of support cycles and test paces. Each series has three sets of support cycles, with each set followed by a test pace and an automatic backup stimulus. The support cycles are use the programmed amplitude and pulse duration which may or may not include ventricular paced events. A test pace follows each set of support cycles and is delivered at a test amplitude or pulse duration. A backup pace (at the programmed amplitude and pulse duration of 1.0 msec) is automatically delivered 110 msec after each test pace regardless of capture or loss-of-capture. One to three of the test paces in a series are used to determine if a particular stimulus amplitude or pulse duration is above or below the capture threshold. When the first of the three test paces indicates capture or the last two test paces indicate capture following loss-of-capture (LOC) on the first stimulus, the series is determined to be above the threshold. When two of the three test paces indicate loss-of-capture, the amplitude and pulse duration are determined to be subthreshold. When modifying first the amplitude and then pulse duration, the pacemaker is looking for two points that lie on the strength-duration curve. These points define the boundary between settings that capture the myocardium and those that do not. The test amplitude is based on the result from the previous pacing Back-up pace

Vpace 10 ms

Vpace

Noise window Capture

Vpace

70 ms

Minimum value

Noise window

Noise window

Non-capture

Fusion

Figure 3-57  The Boston Scientific Fusion Management Algorithm. This is offered in both single and dual chamber modes during testing and during beat-to-beat right ventricular autocapture mode. A noise window occurs after each ventricular pacing stimulus (VP) between 27 and 48 msec. If the negative evoked response (ER) amplitude peak occurs within the window, noise or fusion is suspected and the atrioventricular (AV) delay is extended (dual chamber) or the V-V interval is lengthened (single chamber) by 64 msec on the next cycle to search for intrinsic R waves. This extension can result in an AVD that is 64 msec longer than programmable maximum. If R-waves are sensed at the extended AV delay, fusion is confirmed and the AV delay (or V-V interval) remains at the increased interval until a VP occurs, similar to AV hysteresis. If nothing is sensed at the extended AV delay (or V-V interval), a VP is issued, and the ER is evaluated for capture for this cycle. When capture is being evaluated, early detection applies if the brady mode is nontracking and sensing is Fixed. This morphology-based noise/fusion detection will occur four times before either testing fails or RVAC’s Beat to Beat will go to Suspension. (Courtesy Boston Scientific, Marlborough, MA.)

102

SECTION 1  Basic Principles

Amplitude (V)

threshold search (or 0.75 V if no previous search data have been collected). During VCM amplitude testing, the pulse duration is 0.4 msec. A series of support cycles and test paces determines whether the amplitude is above or below threshold. If the test amplitude is above threshold, the amplitude is reduced by one setting and the test series is repeated until the amplitude results in loss of capture. The amplitude is then increased and retested until it is above threshold for three consecutive test series (defined as the amplitude threshold). Following determination of the amplitude threshold at 0.4 msec, the pulse duration threshold is determined starting at the pulse duration from the previous search or 0.21 msec if there has been no previous search. The stimulus amplitude is set at twice the measured amplitude threshold. The pulse duration is then decremented until loss of capture and then incremented until suprathreshold. The pulse duration determined to be above threshold is designated the pulse width threshold. When operating in a tracking mode, the pacemaker shortens the AV interval for each support cycle and test pace based on calculations using the shortest AV interval measured during a stable rhythm check. The test pace is always followed by a backup stimulus after 110 msec. When operating in a nontracking mode, the lower rate is not changed on support cycles while the lower rate on test paces is set to the shortest

V-V interval observed on any support cycle or seen during a stable rhythm check plus 15 bpm or minus 150 msec, whichever is faster. When a pacing threshold search cannot be completed, another search is automatically initiated within 30 minutes. The VCM algorithm first uses bipolar sensing (if programmed bipolar) but can automatically switch to the unipolar configuration if the bipolar ER is inadequate. Ventricular Capture Management can be programmed to provide automatic adaptation of ventricular amplitude and pulse duration based on pacing threshold search results (Figs. 3-58 and 3-59). Following each threshold search, a target output is determined by applying a programmable safety margin (amplitude margin parameter). The target is always rounded up to the next programmable setting. Adaptation can take place only within an output range that is defined by a programmable lower limit (Minimum Adapted Amplitude parameter) and an upper threshold limit of 5.0 V and 1.0 msec. The minimum pulse width for ventricular capture management is 0.4 msec. Capture Management will not program ventricular outputs above 5.0 V or 1.0 msec. The clinical performance of the VCM algorithm has been evaluated by several authors.171,246 Chen et al246 measured the ventricular pacing threshold in 155 patients with the Medtronic EnPulse pacemaker and found the mean threshold was 0.57 ± 0.21 V at 0.4 msec when

6

Target (when 3X margin applied)

5

Output adapted to 5.0 V, 1.0 ms limit

4 Operating amplitude and pulse width.

3

Amplitude threshold

2

Note: The bottom line indicates the programmed minimum adapted amplitude

1

.2

.4 .6 .8 Pulse width (ms)

1.0

1.2

Output range

Figure 3-58  Medtronic Ventricular Capture Management Operating Range. The bottom line indicates the programmed minimum adapted amplitude and the upper line represents the maximum amplitude and pulse duration (5.0 V and 1.0 msec). In this example, an automatically determined threshold was 2.0 V at 0.4 msec. Using a programmable safety margin of 1 V, the operating amplitude was set to 3.0 V at 0.4 msec. (Courtesy Medtronic, Inc., Minneapolis, MN.) 5

A = Amplitude (V)

4 Target (when 2.0X margins applied) 3 Operating amplitude and pulse width. 2

P Output range

1

A

.2

.4

Note: The bottom line indicates the programmed minimum adapted amplitude .6

.8

1.0

P = Pulse width (ms) Figure 3-59  Ventricular Capture Management. The output range of the ventricular amplitude and pulse width are shown relative to the automatically determined strength-duration curve. The pulse width and amplitude are doubled from the determined threshold. (Courtesy Medtronic, Inc., Minneapolis, MN.)



CHAPTER 3  Stimulation and Excitation of Cardiac Tissues

ACP

80 60 40 20 Log rank: Chi-square 28.78, p <0.0001 0 0

A Freedom from replacement (%)

FOP

50 75 Months

100

80 AVB patients

40 20 Log rank: Chi-square 19.33, p <0.0001 0 50 75 Months

FOP ACP

80 60

SSS patients

40 20 Log rank: Chi-square 6.93, p = 0.0085

0 0

ACP

25

100

25

B FOP

60

Atrial automatic detection of capture and adjustment of stimulus output is now standard for many manufacturers of dual chamber pacemakers. The low amplitude of the ER in the atrium and the presence of far-field R-waves in the atrial electrogram make discrimination of the ER from polarization challenging.251,252 The St. Jude Medical dual chamber pacemakers offer atrial automatic capture detection (ACap Confirm) which detects the atrial ER from the unipolar tip electrode during either unipolar or bipolar pacing. The ACap Confirm algorithm tests for atrial capture every 8 to 23 hours and adjusts the atrial stimulus amplitude to a programmed safety margin (at least 1.7 : 1). The algorithm makes use of the fact that the ER is typically negative in polarity whereas the polarization afterpotentials are positive in polarity. It uses template matching of the electrogram to differentiate atrial capture with an ER from noncapture with polarization only by calculating a Kendall tau rank correlation coefficient that ranges from +1 for a perfect match to −1 for a perfect nonmatch (Fig. 3-61). The algorithm functions well with low polarization electrodes but is not well suited to leads with high polarization properties.252 The threshold search is performed during a set-up procedure with the programmer

125

100

0

C

25

Atrial Automatic Capture Detection Algorithms

Freedom from replacement (%)

100

“low-amplitude standard” programming (2.5 V at 0.4 msec). Compared with “standard” programming, AutoCapture provided a 55% to 60% increase in device longevity. In contrast, AutoCapture provided only 5% to 6% additional device longevity when compared with lowamplitude standard programming. Schuchert and Meinertz250 found that for patients with steroid-eluting ventricular leads and a pulse duration threshold <0.15 msec at 2.5 V, programming to an output of 2.5 V and 0.4 msec significantly prolonged pulse generator longevity compared with the nominal setting of 3.5 V and 0.4 msec (112 months vs. 98 months). Thus although automatic capture algorithms provide only small increases in longevity as compared with low-output settings, their true value is to provide the simultaneous benefit of patient safety.

100

Freedom from replacement (%)

Freedom from replacement (%)

measured manually and 0.54 ± 0.21 V at 0.4 msec when measured by the VCM algorithm. The VCM algorithm was accurate in 93.3% of patients. The use of VCM in pediatric patients has demonstrated that it is highly successful with endocardial leads (99%) but much less so with epicardial leads (31%).171 Because the AV delay is shortened during VCM testing, the actual ventricular threshold during the longer, programmed AV delay may be higher than measured by VCM.247 Thus if the ventricular threshold is to be measured automatically by this algorithm, the VCM should be programmed “Adaptive” to avoid underestimation of the RV pacing threshold during normal function with the programmed AV delay. The value of automatic capture algorithms to increase pacemaker longevity has been studied by several authors.241,246,248,249 Biffi et al241 compared dual chamber pacemaker longevity with standard programming with twice-yearly clinic follow-up (239 patients) to devices with automatic threshold algorithms programmed “on” (61 patients). For standard follow-up patients, the stimulus amplitude was maintained at 3.5 V at 0.4 msec if the measured threshold was <2.5 V and was programmed to 5.0 V if the threshold was between 2.5 and 3.5 V. For the automatic capture algorithms, the stimulus amplitude was maintained 0.25 V greater than the measured threshold at 0.4 msec. Over a mean follow-up interval of 99 months, pulse generator replacement was required for 65% of standard programming patients at a mean of 82 months of service life compared with only 2.9% with automatic capture at a mean of 105 months service life (P < 0.001) (Fig. 3-60). Although the choice of a standard stimulus amplitude of 3.5 V at 0.4 msec may be higher than necessary for most patients with modern, steroid-eluting leads, the value of automatic stimulus management algorithms is nonetheless highlighted by the excellent longevity observed when these algorithms were programmed. A more realistic comparison of AutoCapture versus standard programming has been reported by Boriani et al249 who compared DDDR pacemaker longevity with AutoCapture, “standard” programming (3.5 V at 0.4 msec), and

103

125

100

125

100

FOP ACP

80 60

NMS patients

40 20 Log rank: Chi-square 3.76, p = 0.052 0 0

D

50 75 Months

25

50 75 Months

100

125

Figure 3-60  Freedom from pacemaker replacement in the whole population (A), in sick sinus syndrome (SSS) patients (B), in atrioventricular block (AVB) patients (C), and in neurally mediated syncope (NMS) patients (D). ACP, Autocapture pacing, FOP, fixed-output pacing. (From Biffi M, Bertini M, Saporito D, et al: Actual pacemaker longevity: the benefit of stimulation by automatic capture verification. Pacing Clin Electrophysiol 33(7):873-881, 2010.)

104

SECTION 1  Basic Principles

at a default rate of 90 bpm (or 40 msec less than the mean atrial cycle length but at a rate less than 120 bpm) with four paced test beats at the same amplitude. If an ER is present for at least three beats, capture is confirmed. The amplitude of the atrial stimulus is decremented in steps of 0.25 V until loss of capture with a 5-V backup pulse delivered 40 msec after the test stimulus. The amplitude is then incremented in steps of 0.125 V until capture of two consecutive stimuli is confirmed. Once ACap Confirm is programmed on, the atrial threshold is automatically determined by the pacemaker at a programmable interval. The atrial threshold search is performed over 16 beats if there is no intervening sensed P-wave. In the presence of sensed P-waves, the threshold search is performed after calculating the mean atrial rate interval over 10 beats. If an atrial cycle length greater than the mean is not detected during the subsequent 10 intervals a threshold search is

2

2

1

1 2 1

1 2

Both slopes are positive.

Slopes do not match.

Kendall tau score +1

Kendall tau score +0

Figure 3-61  The St. Jude Medical CapConfirm Atrial Template Matching Algorithm. On the left, the atrial electrogram during capture is used as a template (upper tracing). During capture of a pacing stimulus in the atrium, the electrogram closely matches the template with a Kendall tau correlation coefficient of +1. On the right, loss of atrial capture is recognized by an electrogram that is markedly different than the template with a Kendall tau score of 0. (Courtesy St. Jude Medical, St. Paul, MN.)

AP

postponed and a new 16-beat monitoring interval is started searching for consistent atrial pacing. Boston Scientific offers the PaceSafe atrial capture algorithm on the Vitalio and Intua pacemakers. These devices determine automatic atrial capture with a bipolar atrial lead with pacing from the tip electrode to the can while the atrial ER is detected from the ring electrode to the can. The coupling capacitance of the output capacitors is reduced to 2 µF in order to reduce polarization artifact and increase ER sensing.253 The use of a bipolar atrial lead with interelectrode spacing ≤1 mm may reduce the chances of accurately detecting the atrial ER. The algorithm attempts to measure the atrial threshold every 21 hours. The right ventricular pacing stimulus is used as backup so that atrial PaceSafe is not available in the AAI or AAIR modes. The response to test stimuli is classified as capture, noncapture, fusion, or noise. Once the atrial threshold has been determined, the atrial output is automatically adjusted to an amplitude twice the highest of the last seven threshold measurements at 0.4 msec. If the atrial threshold cannot be determined after 4 days, the algorithm is suspended and the atrial output is programmed to 5.0 V at 0.4 msec. The test sequence raises the atrial pacing rate by 10 bpm with a maximum of 110 bpm. Rather than attempting to distinguish the ER from polarization in the atrium, the Medtronic Atrial Capture Management (ACM) algorithm observes the timing of sensed P and R waves to evaluate capture (Fig. 3-62). Atrial capture is detected by either reset of the underlying atrial rhythm by premature atrial stimuli or by observing the AV delay. The ACM algorithm first determines whether the patient is pacing or sensing as a low rate is desirable during the pacing threshold search to reduce the risk of competition from forced pacing with a fast intrinsic rhythm. A Pacing Threshold Search (PTS) is performed when a stable atrial rhythm is observed for eight pacing cycles and the sensor rate is less than the ADL rate. ACM is available in the DDD and DDDR modes and does not operate if mode switching is operative. Similar to VCM, the ACM performs a threshold search with a series of support and backup stimuli to attempt to distinguish capture from noncapture. Atrial Chamber Reset runs during stable sinus rhythm. It evaluates capture by observing the response of the intrinsic rhythm to the atrial test pace. If the test pace does not capture, the sinus node is not reset,

Test AP Backup AP

AP Overdrive rate interval

Overdrive rate interval - prematurity

AP-VS VS

VS

Test AP

Backup AP

(Expected VS from Test AP)

VP

(Expected VS from Backup AP)

(70 ms) Capture Loss of capture

Scheduled VP Last support AP-VS interval

85 ms

Figure 3-62  Medtronic Atrial Autocapture using atrioventricular (AV) interval response (see text for description of the algorithm). (Courtesy Medtronic, Inc., Minneapolis, MN.)



CHAPTER 3  Stimulation and Excitation of Cardiac Tissues

and an atrial refractory sensed event (AR) is observed after the test pace. If no AR is observed within the AV interval, ACR concludes that the test pace captured the myocardium. The alternative method to detect atrial capture involves observing the AV interval in response to atrial pacing. Atrial-Ventricular Conduction method runs when stable 1 : 1 AV conduction is observed with atrial pacing. The atrial pacing rate is increased by 15 ppm (but no faster than 101 ppm) and the AV interval is lengthened to try to achieve a stable AP-VS rhythm. AVC evaluates capture by observing the conducted ventricular response to the atrial test pace. Each atrial test pace is followed by a backup pace at programmed amplitude and a 1.0 msec pulse width to maintain rhythm stability during the test. If a conducted VS event is observed at approximately the expected AP-VS interval following the atrial test pace, AVC concludes that the test pace captured the myocardium. Beginning at one setting below the last measured value (or 0.75 V if no previous search has been done), the amplitude is reduced in one setting decrements at 0.4 msec pulse width until loss-of-capture is detected. The pacemaker then increments and retests the amplitude one setting at a time until the setting is above the stimulation threshold for three consecutive test series. The setting at which capture is recovered is determined to be the amplitude threshold. ACM can be programmed to an adaptive setting that provides automatic adaptation of atrial amplitude based on periodic pacing threshold search results. The pacemaker applies a programmable amplitude safety margin to the amplitude threshold value measured at a pulse duration of 0.4 msec to determine the target amplitude. If the operating amplitude is above the target, the pacemaker adapts the amplitude down toward the target in one-step decrements. If the operating amplitude is below the target, the amplitude is immediately adapted to the target. ACM will not provide atrial outputs greater than 5.0 V or 1.0 msec. Atrial capture management has been shown to work well with the automatically determined threshold of 0.68 ± 0.35 V at 0.4 msec compared with 0.69 ± 0.38 V at 0.4 msec measured manually.254,255 Importantly, atrial capture management worked equally well with epicardial and endocardial leads, probably because it does not rely on distinguishing an ER from polarization.255 The Biotronik Evia dual chamber pacemakers offer Atrial Capture Control (ACC) which does not rely on detection of the atrial ER. The threshold search is based on the presence or absence of atrial sensing markers generated by the device. The atrium is stimulated at a rate higher than the intrinsic rate to suppress atrial intrinsic events. As soon as the pacing output is lower than the atrial threshold, sensed atrial events will be detected either due to the emerging intrinsic rhythm or due to retrograde conducted events caused by ventricular paces. The detection of sensed atrial activity is used to discriminate between atrial capture and noncapture. ACC determines the actual atrial rate immediately before it starts the threshold search. Automatic measurements are allowed if the atrial and ventricular rate are below 110 ppm. If these conditions are met, the capture control algorithm switches to the DDI mode with an atrial pacing rate 20% higher than the underlying atrial rate with an AV interval of 50 msec. The atrial stimulus amplitude is stepwise decreased until loss of capture occurs. Loss of capture is determined when two or more intrinsic atrial events are sensed. The pacing amplitude is then adjusted by adding a programmed safety margin to the determined threshold. Left Ventricular Autocapture Algorithms As capture of the left ventricle is essential for providing the hemodynamic benefits of CRT, several manufacturers provide automatic algorithms for adjusting LV stimulus amplitude. Medtronic CRT devices detect LV capture by measuring the LV-RV conduction time after an LV pacing stimulus delivered with a shortened AV delay (Figs. 3-63, 3-64, 3-65). In order to distinguish a sensed R-wave in the RV that results from LV capture from an R-wave that is due to intrinsic AV conduction, the interval from the LV stimulus to a sensed R-wave in the RV is compared with the interval from an atrial pacing stimulus to a sensed R-wave. The device first determines if the atrial and ventricular rhythms are stable and compatible with the LV capture testing

Measure V-V conduction interval

105

AP RV sense

LV pace

AP Check A-V conduction 50 ms BV Figure 3-63  Medtronic Left Ventricle (LV) Capture Management Algorithm. The algorithm first checks the interval from the LV pacing stimulus to the right ventricular (RV) sensed event. This interval is used to determine LV capture. The algorithm then lengthens the atrioventricular (A-V) delay to determine if any RV sensed event follows during the expected interventricular (V-V) interval window. If no sensed event occurs during the measured V-V interval plus 80 msec, a biventricular pacing stimulus is delivered. See text for description. BV, Biventricular. (Courtesy Medtronic, Inc., Minneapolis, MN.)

protocol (R-R variability <200 msec measured over 12 cycles). Next, the device determines the LV-RV conduction time (VV interval) and the RA-RV conduction time (the AV interval). If the V-V and AV intervals are similar, testing cannot proceed. However, if the V-V interval is at least 80 msec shorter than the AV interval, the LV capture management algorithm can use these intervals and testing will proceed. These intervals must be consistent over four consecutive cycles for testing to continue. The LV stimulus amplitude is then decremented using three support paces followed by an LV pacing stimulus only. Once loss of capture is determined, the LV stimulus amplitude is incremented until the LV captures on three consecutive trials. The accuracy of the LV Capture Management algorithm was studied in 307 patients with a Medtronic CRT device.256 The proportion of automatic LV threshold tests within one programming step of a manual threshold test was 99.7%. The algorithm measured the threshold successfully in 96% and 97% of patients after 1 and 3 months, respectively.256 The stability of LV capture thresholds was studied in 282 patients with a Medtronic Concerto CRT-D device.207 The LV threshold demonstrated a maximum increase of ≤1.0 V in 91% of patients between 1 and 6 month clinic visits and in 94% between the 6 and 12 months. However, the increase in LV threshold was significantly related to the baseline threshold amplitude. Of the 170 patients with a 1 month threshold of ≤2.0 V, 94% had increases of ≤1.0 V at their 6 month visit; in contrast, 21% of patients with a baseline threshold of >2.0 V demonstrated an LV threshold increase of >1.0 V.207 The SJM Accel CRT, Anthem CRT, and Unify CRT devices offer automatic LV and RV (LVCap Confirm) and RV capture confirmation. The threshold is measured in both the RV and LV leads every 8 or 24 hours. The LV capture confirmation will occur before the RV capture confirmation whenever the LV is programmed before or simultaneous with the RV stimulus. Backup pulses are delivered from the RV at 5.0 V during LV threshold testing. This algorithm works with bipolar, integrated bipolar, and unipolar lead configurations. Capture of the LV is determined from the paced depolarization integral in the LV tip to can configuration (if programmed in that configuration) or from the LV ring to can configuration when programmed LV ring to can. The LV stimulus amplitude can be programmed so that it is adjusted to be 0.25 to 2.5 V greater than the measured threshold. The Biotronik Evia pacemakers and Ilesto ICDs offer automatic detection of the ER in the LV channel to detect capture. The stimulation threshold is measured after a 20-msec blanking period following the ventricular pacing stimulus,

106

SECTION 1  Basic Principles

LV-RV conduction test

AV conduction test

A P

A P

A P

A P

U U P S

U U P S

U U P S

U U P S

180 ms

180 ms

180 ms

A P

A P

A P

B U

180 ms

B U

B U

B U

260 ms

260 ms

260 ms

A P

260 ms

PAV setting (LV-RV + 80 ms)

Measured LV-RV interval

Figure 3-64  Medtronic LV capture management algorithm demonstrating the left ventricular-right ventricular (LV-RV) conduction test during which the atrioventricular (AV) delay is shortened followed by LV pacing. The interval from the LV pace (Vp) and RV sensed events (VS) is measured to be 180 msec. In order to check for fusion, the AV delay is lengthened to 260 msec (the interventricular [V-V] interval + 80 msec) and no intrinsic ventricular sensing is observed followed by biventricular pacing. PAV, Paced AV interval. (From Crossley GH, Mead H, Kleckner K, et al: Automated left ventricular capture management. Pacing Clin Electrophysiol 30(10):1190-1200, 2007.) 3V 0.5ms Loss

3V 0.5ms Loss

Test

Test

Support

A P

B U

A P

A P

U P

U S

Loss

B U

A P

A P

B U

B U

A P

U P

Loss

Test

Support

A P

U S

3.5V 0.5ms Capture

B U

A P

A P

B U

B U

A P

UU P S

A P

A P

B U

B U

Capt

Figure 3-65  Automatic left ventricular (LV) Capture Management algorithm determines loss of capture in the LV based on a long LV paced to right ventricle (RV) sensed interval. Capture (Capt) is determined by a short LV paced to RV sensed interval during the expected window. (Courtesy Medtronic, Inc., Minneapolis, MN.)

these devices attempt to differentiate the ventricular ER from polarization artifact during a 60-msec window. Analysis of the electrogram is performed in two separate phases. In both phases, the AV delay is shortened to 50 msec after a paced atrial event and to 15 msec after a sensed atrial event to ensure ventricular pacing. First, five ventricular pacing pulses are delivered at the programmable maximum voltage setting (2.4, 3.0, 3.6, 4.2, 4.8 V). If noncapture is detected at the maximum voltage setting, the second phase of signal analysis is aborted and the test is classified as unsuccessful. In the next phase, five “double” pacing pulses are delivered, one pacing pulse followed by another pacing pulse delivered 100 msec later during the absolute refractory period. These pulses are used to verify that the polarization artifact is small enough to distinguish capture from noncapture. The pacing amplitude decrements with every paced beat by 0.6 V, until the first noncaptured beat. The algorithm then decrements by smaller increments of 0.1 V until the first failed capture, at which point it

determines that the capture threshold is the preceding value. The pacing amplitude is then automatically adjusted to a programmable safety margin above the measured threshold (0.5 V through 1.2 V in steps of 0.1 V). The Boston Scientific Autogen CRT-D devices offer PaceSafe Left Ventricular Automatic Threshold (LVAT) which is designed to dynamically adjust the left ventricular pacing output to ensure capture of the LV using a programmable safety margin. LVAT measures pacing thresholds from 0.2 V up to the programmable maximum of 7.5 V. LVAT is available in all bipolar and unipolar LV pacing configurations but is not available in devices with a quadripolar header configuration. The LVAT algorithm then measures the ventricular pacing threshold each day and adjusts the voltage output. During testing, LVAT uses the evoked response signal to confirm that each ventricular pacing output captures the ventricle (Figs. 3-66 and 3-67). If testing is successful, the ventricular amplitude is adjusted by adding the programmable safety



CHAPTER 3  Stimulation and Excitation of Cardiac Tissues

15 Capture Loss of capture

Amplitude (mV)

10 5 0 –5 –10 0

50 Time (ms)

100

150

Figure 3-66  Schematic drawing of the Boston Scientific LV Automatic Threshold algorithm demonstrating an evoked response following capture (solid line) and polarization only caused by lack of capture (dashed line). (From Kalahasty G, Giudici M, Lobban J, et al: Acute clinical evaluation of a left ventricular automatic threshold determination algorithm based on evoked response sensing. Pacing Clin Electrophysiol 35(3):348-356, 2012.)

Amplitude (mV)

Capture Fusion Loss of capture

Time (ms) Figure 3-67  Boston Scientific LV Automatic Capture algorithm demonstrating left ventricular (LV) capture (solid line) from fusion (dotted  line), and loss of capture (dashed line). (From Kalahasty G, Giudici M,  Lobban J, et al: Acute clinical evaluation of a left ventricular automatic threshold determination algorithm based on evoked response sensing. Pacing Clin Electrophysiol 35(3):348-356, 2012.)

margin to the highest measured threshold of the last seven successful ambulatory tests (between 1.0 V and the programmable Maximum Amplitude). Seven tests are used to account for circadian cycle effects on threshold and ensure an adequate safety margin. This also allows for a rapid increase in output caused by a sudden rise in threshold while requiring consistently lower threshold measurements to decrease output. When LVAT is set to Auto or Daily Trend, ambulatory ventricular automatic threshold measurements are conducted every 21 hours. In atrial tracking modes, the automatic threshold measurement adjusts the paced AV Delay to 140 msec and the sensed AV Delay to 110 msec. RV pacing is provided as backup throughout LV testing with an applied

107

LV Offset of −80 msec. The starting ventricular pacing output amplitude is the programmable Maximum Amplitude and the ventricular pacing rate is increased by 10 bpm above the current rate. The pulse generator will decrement the ventricular output every three paces until a threshold is determined. If daily testing is unsuccessful, LVAT will return to the previously determined output and the device will perform up to three reattempts at hourly intervals. In an acute study comparing the LVAT measured threshold compared with a manually determined threshold, the algorithm was able to determine an LV threshold in 234 of 261 tests (89.6%).257 The automatically determined threshold was equal to the manually determined threshold in 233 of 234 cases (99.5%) whereas no algorithm-determined threshold was lower than the manually determined threshold. Importantly, the algorithm functioned well in a variety of LV pacing configurations (Unipolar LVTip→Can, Unipolar LVRing→Can, LVTip→RV, and LVRing→RV).257

BIVENTRICULAR PACING The hemodynamic improvement from a synchronized sequence of LV contraction is well documented. Several pacing configurations have been used to achieve stimulation of the left ventricle, including using a standard dual chamber DDDR pacemaker with a right atrial lead in the atrial port of the header and a left ventricular lead in the ventricular port. The LV lead can be epicardial, endocardial, or in a branch of the coronary venous circulation. This configuration can be especially useful in the presence of left bundle branch block (LBBB) with intact AV conduction as timing of the AV interval can achieve fusion between intrinsic conduction and LV stimulation.258 However, there is no ability to sequentially stimulate both ventricles with this configuration. A second, early approach that was utilized to stimulate both ventricles involved placing leads in the RV and on the LV epicardium or in a branch of the coronary venous circulation which were joined with the use of a Y-adapter and connected to the ventricular port of a DDDR pulse generator. In this configuration, the distal electrode of the RV lead was the anode and the distal electrode of the LV lead functioned as a cathode (or vice versa). The advantage of this approach was that a standard DDDR pulse generator could be used and the impedance of the ventricular circuit was quite high with two small electrodes, thereby limiting current drain from the battery. The disadvantages included the fact that the stimulus amplitude and pulse duration could not be independently programmed and for the RV and LV and there was no option to adjust the V-V timing. In addition, the ventricular electrogram was a composite of the unipolar RV and LV signals so that this configuration could not be used for an ICD because of the very high chances of double counting the ventricular rate. The first devices approved for biventricular pacing utilized a shared cathodal configuration in which the current was delivered in parallel to the RV and LV distal electrodes serving as cathodes with the anode being the RV ring electrode or the pulse generator can. The ventricular electrogram was a composite of the RV and LV electrograms, leading to double counting if the interventricular conduction time was prolonged. Because current was delivered in parallel, the current delivered to each cathode was reduced and was inversely related to the LV and RV tip impedances, respectively. Thus if the impedance at the LV tip electrode was high while the impedance at the RV tip electrode was low, more current would flow through the RV electrode than through the LV electrode. The apparent threshold in the LV was dramatically affected by this split cathodal configuration. For example, Mayhew et al89 found that when the unipolar LV threshold was measured (versus the pulse generator casing), the mean LV threshold was 0.7 ± 0.5 V at 0.5 msec. Using the pulse generator casing as the anode and splitting the current between the LV tip and RV tip electrodes in a unipolar shared-cathodal configuration increased the apparent LV threshold to 1.0 ± 0.8 V. When the RV ring electrode was used as the anode (bipolar shared-cathodal configuration), the apparent LV threshold increased to 1.3 ± 0.9 V. Thus splitting the current between the RV and LV produced nearly a doubling of the apparent

108

SECTION 1  Basic Principles

LV threshold when a bipolar shared-cathodal configuration was programmed. Another observation with the shared cathodal configuration was that the impedance measured with a pacing system analyzer (PSA) did not behave as predicted by the simple application of Ohm’s law. For example, when using the RV ring electrode as the anode, the RV tip to RV ring impedance measured 705 Ωs whereas the LV tip to RV ring impedance measured 874 Ω. Using a shared cathodal configuration with the RV ring as the anode and both the LV and RV tip electrodes as cathodes, Ohm’s law would predict a combined impedance 1 1 1 = + (RT) of 390 Ω ( ). In fact, the total impedance R T R RV R LV measured 516 Ω.89 The higher than predicted impedance is explained by the fact that the size and shape of the combined cathodes was significantly different than either electrode alone. When the RV and LV cathodes were combined, the shared cathode results in a Warburg resistance and capacitance. Combining electrodes of similar size essentially doubles the cathodal surface area. Because the electrode impedance of a hemispherical electrode is about proportional to the square root of the electrode surface area, doubling the size of the shared cathode by combining the RV and LV tip electrodes decreases the 1 Warburg resis tance by a factor of . Combining the electrodes in 2 parallel also increases the voltage droop of the constant-voltage waveform. And because coronary venous leads tend to have higher impedance than RV endocardial leads (based on a lower blood pool surrounding the electrode), this tends to further shunt current away from the LV cathode toward the RV cathode when the LV and RV are stimulated in parallel and results in an apparently higher LV threshold. Present CRT devices use separate RV and LV output and sensing circuits with multiple options for programming the stimulation vector. LV pacing leads may have one, two, or four electrodes with independent programmability of the stimulation configuration. Thus the LV can be stimulated by a single electrode in the coronary venous system with the electrode of opposite polarity being the pulse generator casing (unipolar) or in a bipolar configuration using two electrodes in the coronary veins or on the epicardium. Alternatively, any of up to four electrodes in the coronary venous system can be stimulated as the cathode with either the RV ring electrode or the pulse generator as the anode (Fig. 3-68). The advantages of these multiple options for programmable stimulation configurations include minimizing the LV threshold, reducing the chances of left phrenic nerve stimulation, independent timing of LV and RV stimulation, and the option of using 17 mm

10 mm

20 mm

1 mm Figure 3-68  Close Up View of the St. Jude Medical Quartet Quadripolar Lead. This lead has four electrodes measuring 1 mm in width with a total distance of 47 mm from the tip electrode to the most proximal electrode. (From Sperzel J, Dänschel W, Gutleben KJ, et al: First prospective, multi-centre clinical experience with a novel left ventricular quadripolar lead. Europace 14(3):365-372, 2012.)

automatic capture algorithms in both ventricular chambers. As the impedance of electrodes within the coronary venous system tends to be high, by using a bipolar pacing configuration with both the anode and cathode on the LV lead, the current drain on the battery can be reduced compared with using the RV ring electrode or the pulse generator casing as the alternate electrode. The use of two coronary venous leads to stimulate widely separated regions of the LV has been reported to improve the hemodynamic results as compared with single site LV pacing.259,260 This, so-called “triple-site” pacing configuration uses two standard coronary venous leads with the distal electrodes combined as a common cathode using a Y-adapter and the RV distal electrode as the anode. As expected, the impedance of the shared-cathodal, “triplesite” configuration was lower (median 266 Ω vs. 496 Ω, P < 0.001) and the LV threshold was higher than for conventional single LV site pacing (1.5 V per 0.4 msec vs. 1.0 V per 0.4 msec, P < 0.001).260 The development of a quadripolar coronary venous lead allowing multisite LV pacing has been reported to improve the hemodynamic results of CRT as compared with conventional singe LV site pacing.261,262 The St. Jude Medical Quartet quadripolar coronary venous lead has four electrodes covering a distance of 4.7 cm (see Fig. 3-68). Using this quadripolar lead, any two of the four LV electrodes (1 = distal, 2, 3, or 4 = proximal) can be used as the cathode and anode, respectively, or any one of the LV electrodes can be programmed as the cathode with the RV ring electrode used as the anode (Table 3-5).263 Thus there are 10 possible LV stimulation configurations (Fig. 3-69).264 Among 16 patients implanted with the quadripolar lead (160 possible vectors), 10 vectors did not provide LV capture because of high threshold and six resulted in left phrenic nerve stimulation.263 The best hemodynamic response is difficult to predict for an individual patient, though the distal-proximal configuration was most likely to provide the best LV dp/dtmax.264 The LV capture threshold with the conventional LV pacing vector was 1.2 ± 0.7 V at implant and 1.4 ± 1.0 V at 3-month follow-up (P = 0.32). In patients who had multisite LV pacing, the LV threshold using the distal pole (LV1) and the LV threshold using a more proximal pole (LV2) were 1.3 ± 0.8 V and 1.8 ± 1.0 V at implant and 1.3 ± 0.5 V (P = 0.89) and 2.1 ± 1.4 V (P = 0.44) at 3-month follow-up, respectively.261 In general, the lowest pacing thresholds are obtained with the most distal electrode as the cathode and the threshold progressively rises as the more proximal electrodes are stimulated.265 As expected, the highest impedance was obtained with the distal electrode and lower impedance when using more proximal electrodes (Table 3-5). Thus the LV threshold at 0.5 msec was 1.4 ± 1.3 V with the LV1-LV2 vector, 2.7 ± 2.2 V with the LV3-LV2 vector, and 3.1 ± 2.5 V with the LV3-LV4 vector.265

SUMMARY Myocardial stimulation is the fundamental principle underlying artificial cardiac pacing. Perhaps the most important concept for programming of an implantable pacing system is a thorough understanding of the strength-duration relationship. Pulse generators allow the clinician to program both the pulse amplitude (in volts) and the pulse duration (in milliseconds). The stimulation threshold is a function of both these parameters. The exponential shape of the strength-duration curve must always be considered when programming the output pulse to ensure an adequate margin of safety between the delivered stimulus and the capture threshold. For example, pulse durations of 1 msec or greater are located on the flat portion of the strength-duration curve, whereas pulse durations of less than 0.15 msec are on the steeply rising portion of the curve. The practical importance of understanding the exponential relation between pulse duration and amplitude is illustrated by Figures 3-70 and 3-71. If the clinician determines the threshold to occur at point A (2 V and 0.5 msec) by decrementing the stimulus voltage at a constant pulse duration, programming of the pulse duration to 1 msec (point B) would provide very little margin of safety. Similarly, if the threshold is measured to be at point C (3.5 V and 0.15 msec) by decrementing the pulse duration at a constant voltage, doubling of the stimulation



CHAPTER 3  Stimulation and Excitation of Cardiac Tissues

109

TABLE Capture Threshold and Impedances With the St. Jude Medical Quartet Quadripolar Coronary Venous Lead Meas3-5  ured Acutely and 1 Month After Implantation Prehospital Discharge

1 Month

Configuration Capture Threshold at 0.5 msec (V) Conventional Configurations Distal 1 to mid 2 1.4 ± 1.3 Distal 1 to RV coil* 1.3 ± 1.3 Mid 2 to RV coil* 1.5 ± 1.2 VectSelect Quartet Configurations Distal 1 to proximal 4 1.4 ± 1.4 Mid 2 to proximal 4 2.1 ± 1.8 Mid 3 to mid 2 2.7 ± 2.2 Mid 3 to proximal 4 3.1 ± 2.5 Proximal 4 to mid 2 3.7 ± 2.4 Mid 3 to RV coil* 2.5 ± 2.4 Proximal 4 to RV coil* 3.5 ± 2.7

Impedance (Ω)

Capture Threshold at 0.5 msec (V)

Impedance (Ω)

933 ± 224 583 ± 123 431 ± 149

1.3 ± 1.2 1.0 ± 1.1 1.5 ± 1.1

977 ± 237 576 ± 130 499 ± 119

909 ± 181 749 ± 193 738 ± 216 720 ± 163 761 ± 202 419 ± 99 400 ± 109

1.3 ± 1.3 2.3 ± 1.6 2.7 ± 1.6 3.2 ± 2.2 3.7 ± 2.2 2.4 ± 2.1 3.3 ± 2.5

942 ± 203 871 ± 186 883 ± 192 869 ± 191 888 ± 197 495 ± 120 443 ± 106

From Sperzel J, Dänschel W, Gutleben KJ, et al: First prospective, multi-centre clinical experience with a novel left ventricular quadripolar lead. Europace 14(3):365-372, 2012. *Denotes right ventricle (RV) coil cathode.

Vector

Description

Cathode

Anode

Vector 1

Distal Tip to Mid 2

D1

M2

Vector 2

Distal Tip to Proximal 4

D1

P4

Vector 3

Distal Tip to RV coil

D1

RV Coil

Vector 4

Mid 2 to Proximal 4

M2

P4

Vector 5

Mid 2 to RV coil

M2

RV Coil

Vector 6

Mid 3 to Mid 2

M3

M2

Vector 7

Mid 3 to Proximal 4

M3

P4

Vector 8

Mid 3 to RV Coil

M3

P4

Vector 9

Proximal 4 to Mid 2

P4

M2

Vector 10

Proximal 4 to RV Coil

P4

RV Coil

P4 M3 M2 D1

Figure 3-69  Schematic depiction of the St. Jude Medical Quartet quadripolar lead with the distal electrode (D1), second most distal electrode (M2), third most distal (M3), and proximal electrode (P4). There are 10 possible stimulation vectors that are programmable. (From Asbach S,  Hartmann M, Wengenmayer T, et al: Vector selection of a quadripolar left ventricular pacing lead affects acute hemodynamic response to  cardiac resynchronization therapy: a randomized cross-over trial. PLoS One 8(6):e67235, 2013.)

Threshold (V, µJ, µC)

5

x

Energy

4 x x

3

x x

x

x

x

x

x

Charge

x

2 Potential

Rheobase

1 Chronaxie 0 0

.2

.4

.6

.8

1.0

Pulse width (msec)

1.5

Figure 3-70  The stimulation threshold can be described by two points, Rheobase (the lowest voltage that will capture the myocardium at any pulse width (duration), and Chronaxie (the lowest pulse width (duration) that captures the myocardium at an amplitude twice Rheobase. Note  that the chronaxie duration approximates the lowest stimulus energy on the strength-duration curve. (From Stokes K, Bornzin G: The electrodebiointerface [stimulation]. In Barold SS, editor: Modern cardiac pacing, Mt. Kisco, NY, 1985, Futura, pp. 33-78.)

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10 9

Volts

8 7

D

6

2×V E

5 4

C

3 2

A

2×V 2 × PW

B

1 0 0.0 0.1 0.2 0.3 0.4 0.5 0.6 0.7 0.8 0.9 1.0 1.1 1.2 1.3 1.4 1.5 1.6 1.7 1.8 1.9 2.0

Pulse duration (msec) Figure 3-71  Practical Measurement of Stimulation Threshold. If the pulse duration threshold is measured by decrementing the pulse duration at a constant voltage of 2.5 V, a threshold of 0.3 msec is determined (point A). Doubling the pulse duration to 0.6 msec (point B) offers a minimal safety margin. In contrast, if the voltage threshold is determined at a constant pulse duration of 0.15 msec, (point C), doubling the amplitude to point D offers a minimal margin of safety. A more useful method is to determine the voltage threshold at a pulse duration that is near chronaxie (point A) and then doubling the amplitude (point E).

voltage to 7 V (point D) also would provide a poor safety margin. When one considers the shape of the strength-duration curve, a more appropriate programmed setting would be provided by doubling the threshold voltage at a pulse duration of 0.5 msec (point E, 4 V and 0.5 msec). As a general rule, if the threshold is determined by decrementing the stimulus voltage, an adequate margin of safety can be assumed by doubling the voltage if the pulse duration used was greater than 0.3 msec. The two most important points on the strength-duration curve (rheobase and chronaxie) are easily estimated with modern pulse generators (see Fig. 3-70). Rheobase can be estimated by decrementing the output voltage at a pulse duration of 1.5 to 2 msec. Chronaxie can then be estimated by determining the threshold pulse duration at twice the rheobase voltage. If instead the threshold is determined by decrementing the pulse duration, an adequate safety margin can be assumed by tripling the pulse duration only if the threshold is 0.15 msec or less. If the rheobase and chronaxie are measured, doubling of the threshold voltage at the chronaxie pulse duration provides an excellent method for programming a pacing system. In most circumstances, however, experienced clinicians will measure the pacing threshold at an initial stimulus duration of 0.4 or 0.5 msec, then make decisions based on their knowledge of the patient’s problems and medications. What constitutes an adequate safety factor depends on each patient considered individually. How dependent is the patient on the pacemaker? How medically stable is the patient? Are there likely to be significant changes in serum K+ or Ca2+? What medicines is this patient taking that can influence the pacing threshold? Have important threshold variations been seen, or are they anticipated in this patient after the initial stabilization period? Are there problems with the leads? How old is the pulse generator, and what is the known history of this pulse generator model and leads in other patients? These factors determine clinical judgments about voltage and current safety, frequency of pacing and medical status monitoring, and early or late replacement of the pulse generator and leads. When programming the pulse generator at implantation, the clinician must also consider the acute to chronic evolution of the stimulation threshold. Because an acute rise in threshold typically occurs

during the first several weeks after lead implantation, voltage and pulse duration may need to be programmed to higher values than needed for chronic pacing. The physician should reevaluate the stimulation threshold after the acute rise (and sometimes subsequent fall). For most patients, the pacing system can be programmed to chronic output settings at a follow-up evaluation about 6 weeks after lead implantation. Although these recommendations may not be as applicable to patients receiving a steroid-eluting lead, caution is still warranted. The importance of drug and electrolyte effects on the strengthduration curve should also be appreciated. For patients requiring antiarrhythmic therapy, the stimulation threshold should be measured a number of times after drug initiation to ensure an adequate margin of safety for pacing. Similarly, patients who are more likely to experience alterations in electrolyte concentration (e.g., patients with renal failure, patients taking potassium-wasting diuretics) may need their pacemakers to be programmed with a greater margin of safety. Perhaps most important, the degree to which the heart depends on pacing to sustain life or to prevent severe symptoms must be factored into the choice of a programmed margin of safety. For pacemakerdependent patients, a pacing amplitude at least 2.5 times the chronic capture threshold is usually recommended. In contrast, patients unlikely to experience significant symptoms should failure to capture occur may have their pacemaker programmed to a lower margin of safety, perhaps 1.5 to 2 times threshold. The effect of pacing rate on the stimulation threshold should also be considered for patients who require antitachycardia pacing, with the pacing threshold measured at all rates likely to be used. For many patients, the use of automatic threshold measurement and output adjustment algorithms, especially if capture detection occurs on a beat-to-beat basis, significantly reduces these safety concerns and allows the pulse amplitude to be delivered only slightly above threshold. For leads with a very low chronic stimulation threshold, these algorithms may significantly prolong battery longevity. In the presence of high impedance caused by lead fracture, the current output of a constant-voltage pulse generator decreases and loss of capture may occur. If lead insulation failure occurs, the impedance as seen by the pulse generator will decrease as current is shunted between the anode and cathode without flowing through cardiac tissue. This results in an increase in the current from the pulse generator without a change in the nominal output voltage. This change may not be detected early if threshold is determined only by the voltage required for pacing capture. Measuring the voltage/current ratio allows detection of the nominal impedance and alterations in lead insulation or in wire continuity. Because some wire fractures intermittently make and break contact, a normal impedance measurement does not always ensure that the lead is intact. Pacing impedance is determined by four factors: (1) resistance in the conductor wire pathways, (2) polarization at the electrode-tissue interfaces, (3) resistance (small geometric size for high resistance) at the electrode-tissue interface, and (4) impedance/resistance of the tissues between the electrodes. The first two factors are energy inefficient, decreasing the current available for stimulation, whereas the third factor decreases current drain without decreasing the efficiency of stimulation. An ideal electrode would have, among other attributes, high resistance and high capacitance (low polarization voltage) at the electrode-tissue interface. Pacing with a monophasic stimulus is more energy efficient than pacing with a bipolar stimulus. The pacing threshold is greater at normal stimulus durations for biphasic stimuli than uniphasic stimuli with the same total duration. In contrast, biphasic stimuli are more energy efficient for successful defibrillation. A biphasic stimulus with proper characteristics reduces the postpulse ion rearrangements. Biphasic stimuli also may reverse continuing local and undesirable chemical processes at the electrode. Cardiac resynchronization devices may use a variety of stimulation configurations, each with its own advantages and disadvantages in terms of stimulation threshold, device longevity, and optimization of ventricular contraction sequence.



CHAPTER 3  Stimulation and Excitation of Cardiac Tissues

CONCLUSION An applied electrical stimulus above threshold value produces an electric field that initiates cardiac excitation as a result of passive effects on the transmembrane potential (the difference in voltage between the inside and outside of the cell). A rectangular pacing stimulus at the electrode-tissue interface may be either negatively charged (cathodal stimulus) or positively charged (anodal stimulus). Cardiac excitation

111

may occur immediately adjacent to the electrode or at a distance by a virtual electrode effect. A basic understanding of the relationship between stimulus amplitude and duration as well as the effects of drugs, electrolytes, and physiologic changes are important for the safe and effective programming of cardiac implantable electronic devices. In addition, clinicians should understand the basic function of leads and electrodes in order to manage patients whose lives depend on the proper functioning of these devices.

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CHAPTER 3  Stimulation and Excitation of Cardiac Tissues

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