IMPLANTABLE CARDIOVERTER DEFIBRILLATORS AND MEMS SENSORS

A LITERATURE SURVEY

COPYRIGHT, FARIBA SIRJANI, 2005

1 ABSTRACT

Sensing in implantable cardioverter defibrillator (ICD) devices is achieved generally with few sensing leads that are lodged in the atria, the ventricles, or both. As such, the electrical signals propagating in the myocardium are sensed in an averaged manner and may include far-field effects. Further, programmed refractory and blanking periods prevent reporting of the signals after a sense, a pace, or a shock event. Last, signals do not reach leads lodged in infarct areas. As a result, statistical algorithms must be applied to the sensed signals in order to determine the existence and range of arrhythmia. Because, the sensing systems of ICD devices cannot yield a high resolution map of electrical activity of the heart, a high density network of MEMS sensors may help create the opportunity for a closer look at the patterns and amplitudes of electrical activity during arrhythmic events. Based on a literature survey, a high density web of magnetometers that may be fabricated using MEMS technology is proposed as a subject for future research as a means of obtaining a high resolution map of electrical activity of the heart.

2

I. INTRODUCTION

Approximately 3 million people worldwide have hearts that beat too fast. Abnormally rapid heartbeats can deteriorate into a life-threatening condition that has been the major cause of sudden cardiac death. Sudden cardiac death kills 450,000 Americans each year, more than lung cancer, breast cancer, AIDS, and stroke combined. [1].

Sudden cardiac death results from an abrupt loss of heart function or cardiac arrest. The victim may or may not have diagnosed heart disease; the time and mode of death are unexpected; and death occurs within minutes after symptoms appear and mere few minutes of delay in treatment significantly impacts survival. [2]. For example, with treatment after 2 minutes, the survival rate is about 30%, after 5 minutes about 20%, and after 12 minutes about 1.4%. [3].

All known heart diseases can lead to cardiac arrest and sudden cardiac death. Most of the cardiac arrests that lead to sudden death occur when the electrical impulses in the diseased heart become irregular. Irregular heart rhythm, called arrhythmia, causes the heart to suddenly stop beating effectively. The electrical impulses in the diseased heart may become rapid, causing ventricular tachycardia (VT), or may become chaotic, causing ventricular fibrillation (VF), or both. Some cardiac arrests are due to extreme slowing of the heart or bradycardia. [2].

VF represents the major arrhythmia in prehospital cardiac arrests afflicting approximately 70% of the patients. In VF, the heart's electrical activity becomes disordered causing the heart's pumping chambers, or ventricles, to contract in a rapid and chaotic way. When the ventricles "flutter" rather than beat, the heart pumps little or no blood resulting in circulatory arrest. [4].

Various events may lead to VF. Myocardial infarction is the underlying precipitant of VF in cardiac arrest in about 50% of victims. VF can also result from worsening of chronic ventricular arrhythmias, from low-voltage electric shock, electrolyte imbalances, freshwater near drowning, profound hypothermia, or certain drugs. [4].

Not only are the causes of VF uncertain, the exact mechanism is unknown. Treatment of VF is called defibrillation and is usually achieved by applying an electric shock to the heart by an external or an implantable defibrillator. Many different explanations have been offered for successful defibrillation. [5]. All the different theories, hypotheses, and postulations regarding the reasons why defibrillation works, originate from the problem that the exact mechanism of fibrillation itself is unknown. One author calls defibrillation an incomplete and imperfect art. [3].

The paper starts with an overview of the anatomy and physiology of the heart and its electrical basis of functioning. Next, the cable theory, that has been used to describe how some cells including some cardiac cells in the body behave like a cable carrying an electrical current, is presented. Then, theories regarding fibrillation mechanisms are

3 reviewed. Basics of pacemakers and ICD devices are presented next, followed by a review of sensing mechanisms and issues. Sensing issue takes the paper into a discussion of leads used for sensing. The current results of research and development regarding the use of MEMS technology in ICD and related devices such as pacemakers are available in the U.S. Patent and Trademark Office databases. This literature review summarizes the results obtained from these databases. Last, one of the methods presented in the reviewed literature is recommended as the method of choice for further research in the area of obtaining a map of electrical activity during fibrillation.

In short, this paper reviews the literature regarding the mechanisms used or proposed for sensing a fibrillation event with an emphasis on MEMS based mechanisms. The paper concludes with a brief commentary regarding the utility of MEMS sensors in obtaining a high resolution map of electrical activity of the heart that may lead to decoding the fibrillation mechanism and future research issues and recommendations for future research in this area.

References 1. http://www.medtronic.com/newsroom/news_20040902a.html. 2. http://www.americanheart.org/presenter.jhtml?identifier=4741. 3. F. M. Charbonnier, “External Defibrillators and Emergency External Pacemakers,” Proceedings of the IEEE, vol. 84, no. 3, p. March 1996. 4. http://www.merck.com/mrkshared/mmanual/section16/chapter206/206a.jsp. 5. R. Plonsey and D.G. Fleming, Bioelectric Phenomena, McGraw-Hill, Inc., New York, 1969.

4 II. ANATOMY, PHYSIOLOGY, AND ELECTROPHYSIOLOGY OF THE HEART

Anatomy describes the form and structure of an organism, or an organ within an organism, and its various parts. Physiology is a study of processes and functions of each part within an organism or within an organ. In this section, the general anatomy of the heart is presented followed by an explanation of the physiological functions of each part. The section ends with a review of the major components of the electrical structure of the heart and their functions.

II.A. Anatomy

The heart weighs between 200 to 425 grams and is slightly larger than the size of a person’s fist. Each day, the average heart beats 100,000 times, pumping about 7,600 liters of blood. A double-layered membrane called the pericardium surrounds the heart like a sac and keeps it in place. The outer layer of the pericardium surrounds the roots of the heart's major blood vessels and is attached by ligaments to the spinal column, diaphragm, and other parts of the body. The inner layer of the pericardium is attached to the heart muscle. A coating of fluid separates the two layers of membrane, letting the heart move as it beats, yet stay attached to the body. The middle muscular layer of the heart is called the myocardium. The inner layer of the heart consists of a membrane that lines the cavities of the heart and forms part of the heart valves and is called the endocardium.

Various parts of the anatomy of the heart, including the blood vessels, the chambers, the valves, and the electrical system are reviewed in this subsection.

II.A.1. Blood Vessels

The main vessels that carry blood to and from the heart are discussed below.

Coronary Arteries: Because the heart is composed primarily of cardiac muscle tissue that continuously contracts and relaxes, it must have a constant supply of oxygen and nutrients. The coronary arteries are the network of blood vessels that carry oxygen- and nutrient-rich blood to the cardiac muscle tissue. The term coronary is applied to these vessels because they surround the heart like a crown. When cholesterol plaque accumulates in a coronary artery to the point of blocking the flow of blood, the cardiac muscle tissue fed by the coronary artery beyond the point of the blockage is deprived of oxygen and nutrients. This area of cardiac muscle tissue ceases to function properly. The condition when a coronary artery becomes blocked causing damage to the cardiac muscle tissue it serves is called a myocardial infarction or heart attack. Because this is a type of failure distinct from VF, coronary arteries are not of interest to this paper that is focused on VF.

5 Superior Vena Cava: The superior vena cava is one of the two main veins bringing de- oxygenated blood from the body to the heart. Veins from the head and upper body feed into the superior vena cava, which empties into the right atrium of the heart.

Inferior Vena Cava: The inferior vena cava is the other of the two main veins bringing de-oxygenated blood from the body to the heart. Veins from the legs and lower torso feed into the inferior vena cava, which empties into the right atrium of the heart.

Aorta: The aorta is the largest single blood vessel in the body, approximately the diameter of one’s thumb. This vessel carries oxygen-rich blood from the left ventricle to the various parts of the body.

Pulmonary Artery: Arteries are vessels carrying blood away from the heart. The pulmonary artery is the vessel transporting de-oxygenated blood from the right ventricle to the lungs.

Pulmonary Vein: Veins as vessels carrying blood to the heart. The pulmonary vein is the vessel transporting oxygen-rich blood from the lungs to the left atrium.

II.A.2. Chambers [12a]

The heart has 4 chambers. The upper chambers are called the left and right atria, and the lower chambers are called the left and right ventricles. The right and left modifiers are from the viewpoint of the person having the heart. Therefore, in the figures depicting the heart, the right atrium is on the left side of the picture and the left atrium on the left side.

Right Atrium: The right atrium receives de-oxygenated blood from the body through the superior vena cava, carrying blood from head and upper body, and inferior vena cava, carrying blood from legs and lower torso. The cardiac muscle tissue of the atrium contracts in a coordinated, wave-like manner and the de-oxygenated blood collected in the right atrium flows into the right ventricle.

Right Ventricle: The right ventricle receives de-oxygenated blood as the right atrium contracts. Once the ventricles are full, they contract. As the right ventricle contracts, blood flows into the pulmonary artery toward the lungs.

Left Atrium: The left atrium receives oxygenated blood from the lungs through the pulmonary vein. As the contraction progresses through the atria, the blood passes into the left ventricle.

Left Ventricle: The left ventricle receives oxygenated blood as the left atrium contracts. Once the ventricle is full, it contracts. As the left ventricle contracts, the blood flows into the aorta and throughout the body. The left ventricle is the largest and strongest chamber in the heart. The left ventricle's chamber walls are only about a half-inch thick, but they have enough force to push blood through the body.

6 An early detour into physiology is required to qualify the role of atrial contraction. Atrial contraction normally makes only a minor contribution to the filling of the two ventricles when the person is at rest. However, the contraction of the atria is a useful safety factor in at least two circumstances. During tachycardia, when the time for passive filling is short, the atrial contraction can provide a much-needed boost. Atrial contraction is also useful in certain pathologic conditions, for example, when a narrowed valve resists the flow of blood from atrium to ventricle. Otherwise, blood from the atria can passively flow into the ventricles.

II.A.3. Valves

Four types of valves regulate blood flow through the heart:

Tricuspid Valve: The tricuspid valve separates the right atrium from the right ventricle. It opens to allow the de-oxygenated blood collected in the right atrium to flow into the right ventricle. It closes as the right ventricle contracts, preventing blood from returning to the right atrium; thereby, forcing it to exit through the pulmonary valve into the pulmonary artery.

Pulmonary Valve: The pulmonary valve separates the right ventricle from the pulmonary artery. As the ventricles contract, it opens to allow the de-oxygenated blood collected in the right ventricle to flow to the lungs. It closes as the ventricles relax, preventing blood from returning to the heart.

Mitral Value: The mitral valve separates the left atrium from the left ventricle. It opens to allow the oxygenated blood collected in the left atrium to flow into the left ventricle. It closes as the left ventricle contracts, preventing blood from returning to the left atrium; thereby, forcing it to exit through the aortic valve into the aorta.

Aortic Valve: The aortic valve separates the left ventricle from the aorta. As the ventricles contract, it opens to allow the oxygenated blood collected in the left ventricle to flow throughout the body. It closes as the ventricles relax, preventing blood from returning to the heart.

II.A.4. Electrical System

The normal heart has three intrinsic pacemaking tissues: The Sinoatrial (SA) node, the Atrioventricular (AV) node, and the Purkinje fibers. Pacemaking is the spontaneous creation of an electrical impulse by a cell. This activity in the heart pacemaker tissues is independent of any signals from the brain. Any cardiac cell with pacemaker activity can initiate the heartbeat. The pacemaker with the highest frequency will be the one to trigger a signal that will propagate throughout the heart. The fastest pacemaker sets the pace of the heart beat and overrides all slower pacemakers. These same specialized cells carry the electric pulses generated to the required areas of the heart. [12a]. The pacemaker cells look like nerves as they are long and branching. However, they are just a different kind of muscle cell. The electric pulses generated by the pacemaker cells are

7 called action potentials. Generation of action potential will be discussed later in this paper.

The three types of pacemaker tissue, the SA node, the AV node, and the Purkinje fibers or the His-Purkinje system are further described below. [7, 12b].

SA node: The SA node is comprised of a bundle of specialized cells in the right atrium. The SA node cells are the heart’s fastest pacemaker and create the electrical pulses that makes the heart beat. The SA node normally produces 60-100 electrical signals per minute creating the heart rate or pulse. It is the smallest electrical region of the heart. SA cells are stable oscillators whose currents are always varying with time. The conduction velocity of the electrical signals in the SA nodes is about 0.05 m/s.

AV node: The AV node is a bundle of specialized cells that bridges between the heart's upper and lower chambers, or between the atria and ventricles. The AV node cells allow electricity to pass through while other cells between the atria and ventricles do not allow this. The AV node is the secondary site of origin of the electrical signal in the heart. Normally, the AV node may be excited by a signal reaching it by way of specialized atrial conduction pathways. Pacemaking rate of AV is slower and about 40 beats per minute. Its pacemaker activity is thus considered secondary. If the SA node fails, the AV node can assume control of the heart. The conduction velocity of electrical signals in the AV node is also 0.05 m/s. The conduction velocity in the atrial pathways between the SA and AV nodes is much faster at 1 m/s.

His-Purkinje system: The His-Purkinje is a special system of electricity conducting cells located in the ventricles. Electricity travels through the His-Purkinje system to make the ventricles contract. The His-Purkinje system includes: the His bundle, the Right bundle, the Left bundle, and the Purkinje fibers where the system ends. The His-Purkinje system originates at the AV node with the bundle of His and splits to form the Left and the Right bundles. The Right bundle conducts the electrical signal to the right ventricle and the Left bundle conducts the signal to the left ventricle. Purkinje fibers have the slowest intrinsic pacemaker rate of about 20 beats per minute or less. They are thus considered tertiary pacemakers that become active only after SA and AV nodes both fail. On the other hand, the bundle of His and the Purkinje fibers are an effective conduction system within the ventricles because they can conduct an electrical signal faster than any other tissue within the heart. Ventricular muscle tissue and the bundle of His both conduct at 1 m/s while the Purkinje system conduct electrical signals at 4 m/s.

Atrial and ventricular muscle cells conduct the electrical signals from cell to cell among the cardiac muscle cells, or myocytes, along special conducting bundles. However, there is no spontaneous pacemaker activity in these cells. [12b].

II.B. Physiology

The functions of the different parts of the anatomy presented above are reviewed in this subsection. Mechanical activity of the heart is presented separate from its electrical

8 activity. Some related concepts and definitions such as the concept of syncytium and the definition of ECG and its components are mentioned at the end.

II.B.1. Mechanical Activity of the Heart

A heartbeat is a two-part pumping action that takes about a second.

Diastole: As blood collects in the upper chambers, namely the right and left atria, a natural electrical signal causes the atria to contract while the ventricles remain relaxed. This contraction pushes blood through the tricuspid and mitral valves into the resting lower chambers, namely the right and left ventricles. This relaxation of the cardiac muscle tissue in the ventricles is the longer of the two parts of the pumping phase and is called the diastole. When the ventricles relax, they make room to accept the blood from the atria.

The decreased pressure due to the relaxation of the ventricles is called diastolic pressure. The pressure in the discharge vessel of the right ventricle, or the pulmonary artery, is about 7 mm Hg while that in the discharge vessel of the left ventricle, or the aorta, is 80 mm Hg. The pulmonary and aortic valves stay closed until sufficient pressure is built up during the next phase.

Systole: The second part of the pumping phase begins when the ventricles are full of blood. Another electrical signal causes the ventricles to contract. This contraction of the cardiac muscle tissue in the ventricles is called systole. When the ventricles contract, they force the blood from their chambers into the arteries leaving the heart. The left ventricle empties into the aorta and the right ventricle into the pulmonary artery. As the tricuspid and mitral valves shut tight to prevent a back flow of blood, the pulmonary and aortic valves are pushed open. While blood is pushed from the right ventricle into the lungs to pick up oxygen, oxygen-rich blood flows from the left ventricle to the heart and other parts of the body.

The increased pressure due to the contraction of the ventricles is called systolic pressure and it is about 125 mm of Hg.

After blood moves into the pulmonary artery and the aorta, the ventricles relax, and the pulmonary and aortic valves close. The lower pressure in the ventricles causes the tricuspid and mitral valves to open, and the cycle begins again. This series of contractions is repeated over and over again, increasing during times of exertion and decreasing during rest.

The heart in reality consists of two pumps, the left heart which is the main pump and the right heart which is the boost pump. [12]. The output of each pump is approximately 5 liters per minute which can increase by several folds during exercise.

II.B.2. Electrical Activity of the Heart

9 As mentioned before, the heart does not depend on a rhythm generator in the brain. [12]. The mechanical activity described above is accompanied by electrical activity in the heart. [6]. Electrical impulses traveling through the middle heart muscle (the myocardium) cause the heart to contract. This electrical activity begins in the SA node and travels through the muscle fibers of the atria and ventricles. An electrical signal passing through a chamber wall causes the chamber to contract. Once the signal has moved out of the wall, the chamber relaxes. [7].

The electrical activity of the heart may be divided into the following steps: First, the SA node, as the natural pacemaker, creates an electrical signal. Second, the electrical signal follows the natural electrical pathways, or conducting bundles, through both atria. The flow of electricity causes the atria to contract pushing the blood into the ventricles. Third, the electrical signal reaches the AV node where the signal pauses to give the ventricles time to fill with blood. Last, the electrical signal spreads through the His- Purkinje system causing the ventricles to contract and push blood out to the lungs and body. The varying conduction velocity of the different cardiac tissues causes the pauses in the conduction at the AV node as well as the almost simultaneous contraction of the ventricles.

In the atria the propagation of the electrical signal causes a contraction from right to left and downward. Because the SA node is located in the right atrium, atrial contraction begins and ends in the right atrium earlier than in the left atrium. The left ventricle, however, contracts before the right. In the ventricles, the middle wall or the septum contracts from left to right. Then the back part of the septum contracts. Then the myocardium contracts from the inside endocardial cells to the outside epicardial cells. The left ventricle contracts from the apex, or the pointed bottom part, toward the base, or the area between the ventricles and the atria. The right ventricle contracts like a bellows, that is used for fanning a fire, while the left ventricle contracts like a hand squeezing a tube of toothpaste. [12].

II.B.3. Heart as a Functional Syncytium

The pacemaker cells, found at the SA node, are self-excitatory. As a consequence, a regular succession of electrical pulses originates from the SA node. These pulses lead to a regular series of heart beats. [9].

Pulses initiated by the pacemaker cells excite neighboring cells and these cells in turn excite their neighbors so that the excitation spreads from cell to cell in the atria. Nonconducting fibrous tissue separates the atria from the ventricles. So the electrical activity in the atria cannot reach the ventricles by going cell to cell. Instead, atria’s excitation reaches the AV node and from the AV node specialized conduction cells carry the impulse into the ventricles. The tissue of the AV node conducts very slowly, hence introducing a latency between the atrial and ventricular excitation and buffering the ventricles to some degree from abnormal excitation of the atria. [9].

10 In the ventricles, the His-Purkinje tissue conducts the electrical excitation to numerous points in the right and left ventricles. From these points, the excitation is conducted from cell to cell within the working ventricular muscle from inside to the outside and from the apex, located at the bottom of the heart, to the base, located at the top of ventricles. The Purkinje tissue located at the endings of the right and left bundles of the His tissue behaves in many ways like nerve axon while Purkinje tissue is in fact a variety of cardiac muscle tissue. [9].

Because the excitation in the myocardium, or the heart muscle, spreads contiguously from cell to cell, the heart behaves functionally as a syncytium. A syncytium is a mass of cell material with numerous nuclei sharing the same protoplasm without membranes in between. A syncytium usually results from fusion of cells. The heart muscle is multicellular with membranes separating the cells. However, the mechanism of cell-to- cell spread of the excitation causes the heart to look functionally like one big cell. [9].

II.B.4. Electrocardiogram (ECG)

An ECG is a recording of the electrical activity in the heart. The hills and valleys on an ECG recording are called waves:

• The P-wave shows the heart's upper chambers (atria) contracting; • The QRS complex shows the heart's lower chambers (ventricles) contracting; and • The T-wave shows the heart's lower chambers (ventricles) relaxing. [7].

II.C. Cellular Basis of Signal Generation and Conduction

Ventricular fibrillation is primarily a problem with the electrical system of the heart. [11]. Problems arise in how the electrical impulse is produced by the myocardial cell and how the impulse is conducted from one cell to the next. [11]. This section discusses generation of an electrical signal by a pacemaker cell and conduction from cell to cell.

II.C.1. Myocardium Cells

The following section explains a general cell and then focuses on the parameters specific to the myocardium cell. [10].

The biological cell is surrounded by a membrane that separates the intracellular fluid, or cytoplasm, from the extracellular fluid. This cell membrane is usually about 100 Å thick and consists of two layers of lipids placed back to back in what is known as a lipid bilayer. Lipid bilayer is the foundation of all biological membranes, and is a prerequisite of cell-based life. [13].

Within a critical range of concentrations, certain kinds of lipids alone in a test tube of water will self-organize to form a "bilayer." The bilayer is composed of two opposing layers of lipid molecules arranged so that their hydrocarbon tails face one another to form an oily core. Thus, the hydrocarbon tails of the lipid molecules, which are hydrophobic,

11 constitute the membrane core. The polar heads of the lipid molecules, which are hydrophilic, face the aqueous solutions on either side of the membrane in the interacellular and extracellular milieus. Because of the oily core, a pure lipid bilayer is permeable to small hydrophobic solutes but has an only very low permeability to inorganic ions and other hydrophilic molecules. Water itself is an exception to this rule, and crosses freely.

Due to its bilayer structure, the membrane can then be thought of as a parallel plate capacitor with a hydrophobic lipid dielectric between its hydrophilic faces. A typical value for the membrane capacitance is 1 µF/cm2. Scattered along the membrane surface are encrusted proteins, some of which may completely traverse the membrane thickness. The proteins completely traversing the membrane sometimes contain small openings that form pores or channels across the membrane. These channels account for membrane’s relatively high conductivity. The capacitance of the membrane is due to the lipid bilayer matrix and its conductance is due to the proteins inserted in the lipid bilayer.

II.C.1.a. Myocardium Cell Membrane

A myocardium cell membrane, like other cells, is composed of a bimolecular leaflet of phospholipids molecules with protein molecules floating in the lipid bilayer. The nonpolar hydrophobic ends of the phospholipid molecules point toward the middle of the bilayer and the polar hydrophilic ends point toward the outside of the membrane bordering the water phases. This orientation is thermodynamically favorable. The lipid bilayer membrane is about 50-70 Å thick or approximately twice the length of the each phospholipid molecule that is 30-40 Å long. Various protein molecules are inserted in the lipid bilayer matrix of the membrane, some protruding through the entire membrane thickness and others inserted into one of two inner or outer layers of the bilayer membrane. The membrane has fluidity such that the proteins float in the lipid bilayer matrix and can move in the plane of the membrane.

The lipid bilayer forming the cell membrane behaves as a parallel plate capacitor with an approximate capacitance value of 0.7 µF/cm2, corresponding to a dielectric constant of 5, and a dielectric thickness of 60 Å which is also the thickness of the membrane. In general, the dielectric constant of oils is about 3 to 5 while polar materials have larger dielectric constants such as 81 for water.

The specific resistance of the cell membrane is about 1000 Ω-cm2. A bilayer of purely lipid molecules has a resistance many orders of magnitude higher. The relatively high conductivity of the membrane is attributed to the proteins that span across the thickness of the cell membrane.

The low dielectric constant of the membrane allows it to withstand an enormous potential difference, something of the order of 133 KV/cm, across its thin 60 Å-thick span. The cell membrane has fixed negative charges as its outer and inner surfaces due to the phospholipids and the embedded or absorbed protein molecules. Most proteins at a PH

12 near 7 possess a net negative charge. These surface charges impact the electric field near the membrane.

II.C.2. Ion Distributions and Their Maintenance

The compositions of the intracellular and extracellular milieus are very different. For the cell to function, it is necessary to transport substances from one milieu to the other across the cell membrane. Molecules and ions are transported across a cell membrane by diffusion or in bulk or a combination of the two mechanisms.

Molecules such as oxygen, carbon dioxide, alcohol and fatty acids that are uncharged and have a high lipidic solubility diffuse through the lipid matrix under a concentration gradient according to Fick’s laws. Other molecules such as sugars and amino acids exhibit a much higher rate of transport across the membrane than that expected from their membrane solubility and ordinary diffusion. The facilitated diffusion carrying these molecules is likely a result of “simple pores” that are protein-lined channels within the membranes or “simple carriers” that are molecules that bind to the diffusing molecule and facilitate its diffusion across the membrane.

Ions, on the other hand, are driven not only by their ionic concentration but also under the influence of any electric field present. A famous equation for ionic diffusion is the Nernst-Plank equation. Nernst-Plank takes into account both the impact of the electric field as well as the ionic concentration. The charged ions such as Na+, K+, or Cl- are assumed to diffuse through ionic channels rather than through the lipid matrix. It is assumed that the ion species do not interfere with the passage of one another. This assumption is called the “principle of independence” and holds true in dilute solutions. Another assumption regarding ionic movement across membranes is the “principle of electroneutrality” within the cell. As a consequence of electroneutrality, in steady state there can be no continuous transport of net charge from extracellular to intracellular space across a membrane unless a sink and an accompanying source of electric charge are also present in the system. Otherwise a charge buildup would occur acting in opposition to the transport. So, the net electric current across a membrane is zero even though individual ion fluxes occur all the time.

Electroneutrality applies to the space within a cell but not across a membrane. A 100 Å thick membrane with 100 mV across it results in an electric field of 100,000 V/cm which is quite a strong field.

Another form of transport is active transport in which the substance is transported against its concentration gradient. Such transport requires an infusion of energy which in the case of biologic cell is metabolic energy derived from nutrients.

II.C.2.a. Resting Potential

The transmembrane potential in resting atrial and ventricular myocardial cells is about - 80 mV. The resting potential or maximum diastolic potential in Purkinje fibers is

13 somewhat greater about -90 mV and in the nodal cells it is somewhat lower about -60 mV.

The ionic composition of the extracellular fluid bathing the heart cells is similar to that of the blood plasma. It is high in Na+ and Cl- but low in K+. The intracellular fluid, on the other hand, has a low concentration of Na+ and Cl- but a high concentration of K+. The free Ca2+ outside concentration is about 2 mM/kg (mili-Mole per kilogram) while the concentration of free ions inside the cell is about 10-7 mM/kg. The total intracellular concentration of Ca2+ is about 2 mM/kg but most of it is bound while ineracelluar K+ is mostly free.

Thus, under normal conditions the myocardial cells maintain an internal ion concentration that is markedly different from that in the medium bathing the cells. These ion concentration differences underlie the resting potential and excitability.

II.C.1.b. Action Potential [14]

Upon sufficiently strong electrical stimulation, the cell generates a characteristic electrical response known as an “action potential.” [4]. For scientific experimentation, the form of the stimulus is generally a rectangular current pulse, injected into the cell via an intracellular microelectrode. The injected current pulse constitutes a “depolarizing” stimulus. It is depolarizing because it reduces the negative potential across the membrane and renders it less polarized. While the resting potential across the membrane is about - 100 mV with the inside of the cell more negative than the outside, the injected current will cause the inside to be more positive for a short while before going back to its original polarized state.

Following the stimulus-induced depolarization, the membrane conductance to sodium ions increases. The concentration of sodium ions being greater in extracellular space, this increase in conductance leads to an influx of sodium ions that will in turn cause a further depolarization of the cell. Further depolarization increases the sodium conductance further, resulting in even more depolarization. A regenerative process is triggered, leading to the rising phase or upstroke of the action potential. The increase in sodium conductance is quickly followed by an increase of the membrane conductance to potassium ions. As the concentration of the potassium ions is greater inside the cell, the increased conductance causes an outward flux of potassium ions and a repolarization rendering the inside of the cell negative again. This constitutes the falling phase of the action potential. In some cells the membrane potential oscillates around this value before stabilizing resulting in hyperpolarizing or depolarizing afterpotentials. [4].

The action potential is an all or nothing event. Above a certain threshold or liminal intensity, an action potential is triggered. The amplitude of the action potential is fixed independent of the magnitude of the stimulus pulse. A delay or latency between the application of the stimulus and the start of the action potential may occur depending on the amplitude of the stimulus. The delay diminishes with an increase in the stimulus amplitude. [4].

14

Action potential is the cell’s signaling mechanism. Cells communicate when a propagating action potential makes contact with other cells

II.C.1.c. SA Node Action Potentials [14]

Cells within the SA node are the primary pacemaker site within the heart. These cells have no true resting potential. Instead, they generate regular, spontaneous action potentials. The depolarizing current in SA cells is carried primarily by relatively slow, inward Calcium ion currents. Most other cells, that are stimulated by action potentials, are depolarized by fast Sodium ion currents.

SA node action potentials are divided into three phases. The depolarization phase is primarily due to increased Calcium ion conductance. The rate of depolarization is much slower than found in other cardiac cells, such as Purkinje cells, because depolarization in SA cells depends on the slow Calcium ion current. The repolarization phase occurs as Potassium ion conductance increases and Calcium ion conductance decreases. The spontaneous depolarization phase is due to a fall in Potassium ion conductance and a small increase in Calcium ion conductance. A slow inward Sodium ion current also contributes to spontaneous depolarization. This current is thought to be responsible for what is termed the pacemaker or "funny" current. Once spontaneous depolarization reaches threshold, at about -40 mV, a new action potential is triggered.

During depolarization, the membrane potential moves toward the equilibrium potential for Calcium ions, which is about +134 mV. During repolarization, relative Calcium ion conductance decreases and relative Potassium ion conductance increases. This brings the membrane potential closer toward the equilibrium potential for Potassium ions. Therefore, the action potential in SA nodal cells is primarily dependent upon changes in Calcium and Potassium ion conductances.

Although pacemaker activity is spontaneously generated by SA nodal cells, the rate of this activity can be modified significantly by external factors such as drugs and diseases.

II.C.1.d. Non-Pacemaker Action Potentials [14]

Unlike pacemaker cells, non-pacemaker cells have a true resting membrane potential that remains near the equilibrium potential for Potassium ion. When these cells, for example ventricular myocytes (muscle cells) and Purkinje cells, are rapidly depolarized to a threshold voltage of about -70 mV, there is a rapid depolarization that is caused by a transient increase in fast Sodium ion channel conductance. At the same time, Potassium ion conductance falls. These two conductance changes move the membrane potential away from Potassium equilibrium potential and closer toward Sodium potential. An initial repolarization is caused by the opening of a special type of Potassium ion channel. However, because of the large increase in slow inward Calcium ion conductance, the repolarization is delayed and there is a plateau phase in the action potential. This inward calcium movement is through long-lasting Calcium channels that open up when the

15 membrane potential depolarizes to about -40 mV. Repolarization occurs when Potassium ion conductance increases and Calcium ion conductance decreases. The action potential in non-pacemaker cells is, therefore, primarily determined by changes in fast Sodium ion, slow Calcium ion, and Potassium ion conductances.

References 1. http://www.americanheart.org/presenter.jhtml?identifier=4741 2. http://www.merck.com/mrkshared/mmanual/section16/chapter206/206a.jsp 3. W. A. Tacker, editors, Defibrillation of the Heart, Mosby, New York, 1994. 4. Ramesh M. Gulrajani, Bioelectricity and Biomagnetism, John Wiley and Sons, Inc., New York, 1998, pp. 378-79. 5. F. M. Charbonnier, “External Defibrillators and Emergency External Pacemakers,” Proceedings of the IEEE, vol. 84, no. 3, p. March 1996. 6. R. Plonsey and D.G. Fleming, Bioelectric Phenomena, McGraw-Hill, Inc., New York, 1969. 7. http://www.guidant.com/condition/heart/heart_electrical.shtml 8. http://www.math.utah.edu/~veronese/research.html 9. R. Plonsey and R.C. Barr, Bioelectricity: A Quantitative Approach, Plenum Press, New York, 1988. 10. N. Sperelakis, Physiology and Pathophysiology of the Heart, Kluwer Academic Publishers, Boston, 1995, chapters 3-4. 11. S.B. Dunbar, K. Ellenbogen, A.E. Epstein, ed., Sudden Cardiac Death, Futura Publishing Company, Inc., New York, 1997, chapter 5: “The Nature of Activation During Ventricular Fibrillation.” 12. W.F. Boron, E. L. Boulpaep, Medical Physiology: A Cellular and Molecular Approach, Saunders, 2003, chapter 21. 12a. Boron, chapter 20. 12b. Boron, chapter 17. 11. http://www.fact-index.com/l/li/lipid_bilayer.html 12. Richard E. Klabunde, “Cardiovascular Physiology Concepts,” http://www.cvphysiology.com/Arrhythmias/A004.htm,1999-2004.

16 III. ACTION POTENTIAL

An electric stimulus of appropriate strength applied either internally, through a microelectrode, or externally, gives rise to an action potential. The action potential consists of an extremely rapid depolarization and a very slow repolarization returning to the resting membrane potential.

The different stages and some other miscellaneous aspects of action potentials are discussed in this section.

III.A. Effective, Absolute, and Relative Refractory Periods [1]

The duration of the action potential is said to create a refractory period, during which the cell is unexcitable no matter how strong the stimulus. The refractory period, however, is not absolutely unexcitable during its entire length. During the action potential upstroke and the early part of its downstroke, it is impossible to generate a second action potential no matter how strong the stimulus. This interval is called the “absolute refractory period.” This period is also called the “effective refractory period” and applies to the phases of depolarization, initial repolarization, plateau, and part of repolarization, when the cell is refractory to the initiation of new action potentials. During the effective or absolute refractory period, stimulation of the cell does not produce new action potentials. This characteristic allows the action potential to propagate down the cell fiber without interference from a second stimulus. The effective refractory period acts as a protective mechanism in the heart by preventing multiple action potentials before the muscle cell has had the opportunity to recover from the first. The non-excitable period keeps the contractions of the ventricular muscle cells synchronous and strong. The length of the refractory period limits the frequency of action potentials, and the resulting contractions, that can be generated by the heart. The length of the effective refractory period can be altered by antiarrhythmic drugs that block Potassium channels. Increasing the effective refractory period can prevent repeated excitation of the heart that can lead to tachycardia.

Immediately following the absolute refractory period, there exists a “relative refractory period” during which a second stimulus of a stronger magnitude can trigger a second action potential overlapping the one already in progress. That is not the worst news: there may also exist a period of superexcitability during the refractory period when even a stimulus of below threshold magnitude may trigger a new pulse. Depolarizing afterpotentials, that appear at the tail end of the refractory period, may cause this superexcitablity. [2].

III.B. Changing Thresholds: Accommodation

The threshold value of stimulus required to trigger an action potential in the cell is not precise. The membrane can accommodate the potential so that the threshold value can be changing with time rather than being a constant value. Threshold may be raised or lowered based on the history of the stimuli applied to the membrane that have

17 depolarized or hyperpolarized the cell. Change in the threshold level with time is called “accommodation.” [3].

Rise in the threshold, corresponding to a decrease in excitability, may result from a previous depolarization by a slow rising ramp or a subthreshold pulse of current.

If instead of a rectangular current pulse, the membrane is depolarized with a slow ramp of injected current, the threshold value of the current required to trigger the action potential would be higher than that necessary for a pulse. As a result of the gradual increase in the stimulating current, the membrane becomes less excitable. [2].

This decrease in excitability may even occur if the membrane is partially depolarized with a subthreshold pulse prior to the application of a second above-threshold pulse. Subsequent to the subthreshold pulse, a larger than usual pulse is required to trigger an action potential in the cell. [2].

Both ramp and subthreshold depolarization varieties of accommodation can be likened to vaccination. The vulnerability of the cell to the stimulus is reduced by applying a weaker version of the stimulus to the cell ahead of time.

Another species of accommodation is lowering of the threshold potential based on the history of the cell. A counterintuitive phenomenon that leads to a lowering of the threshold level is hyperpolarization. If the membrane is kept hyperpolarized, the threshold potential is expected to increase to compensate for the lower than usual intracellular potential during hyperpolarization. Yet, sustained depolarization can actually cause a decrease in the threshold potential. [2].

The decrease in the threshold resulting from hyperpolarization is often so great that an action potential may be triggered by simply removing the hyperpolarizing current. In other words, bringing the cell from its hyperpolarized state back to its normal resting potential is sufficient to trigger an action potential. This phenomenon is also known as “anode break excitation” for historical reasons. [2].

III.C. Strength-Duration Curves of Action Potential

Upon sufficiently strong electrical stimulation, the cell generates a characteristic electrical response known as the action potential. The threshold intensity of the stimulus required to just trigger an action potential diminishes as the duration of the current pulse stimulus increases. A short-lived stimulus needs to be larger in magnitude in order to produce the same effect as a long-lived stimulus of smaller magnitude. [3]. Similarly, earthquake tremors of lower Richter that last longer can cause a lot more damage than those of a larger Richter of short duration.

The relationship between stimulus strength (in amperes or volts) and duration (in seconds) of the stimulus required to reach the threshold value is called a strength-duration curve. The strength-duration curve is exponential. The curve is characterized by a

18 “rheobase” current (or voltage) value and a “chronaxie” time value. The rheobase is the amplitude of the stimulus below which the membrane cannot be excited, no matter how long the duration. In other words, the rheobase is the value of the stimulus as time goes to infinity. The chronaxie is the duration of time required for a current twice the rheobase value to generate an action potential. The chronaxie is significant as a nominal time period required to reach the threshold voltage and the chronaxies of different membranes or the same membrane under different conditions can be usefully compared. [2,3].

References 1. Richard E. Klabunde, “Cardiovascular Physiology Concepts,” http://www.cvphysiology.com/Arrhythmias/A004.htm,1999-2004. 2. Ramesh M. Gulrajani, Bioelectricity and Biomagnetism, John Wiley and Sons, Inc., New York, 1998, pp. 378-79. 3. R. Plonsey and D.G. Fleming, Bioelectric Phenomena, McGraw-Hill, Inc., New York, 1969.

19 IV. CABLE THEORY

Propagation of the action potential in the excitable tissue of the heart is explained by the “Cable Theory.” This theory was first used to analyze the transmission of signals in underwater cable and was later applied to transmission of signals in nerve cells, muscle cells, or any long cylindrical conducting tissue.

In 1855, Lord Kelvin presented to the Royal Society a theoretical analysis of attenuation of signals in the transatlantic cable that was then being planned. An undersea cable is similar to a nerve or muscle fiber. It has a conducting core covered with an insulating sheath, and is surrounded by sea water. As the insulation is not perfect, there will be a finite leakage resistance through the insulation. [1].

The cable theory was applied to nerve fibers in the 1946 by Hodgkin and Rushton. It was applied to Purkinje fibers in 1962 by Noble. [2].

This section reviews the Cable Theory. The section begins by explaining the characteristics of a cardiac fiber modeled as a cable and continues with the derivation of the cable equation. Then, the section presents the cable theory when applied to conditions where the cell is excited below the threshold required for an action potential to travel from cell to cell and the above threshold conditions where uniform propagation occurs. Next, the section presents the solution of the cable equation for the case of above threshold uniform propagation of action potential.

IV.A. Cardiac Fibers as Cables

Cardiac fiber is treated as a biologic cable which in the case of Purkinje fibers measures about 70 µm and in the ventricular muscle about 16 µm in diameter. The inside and the outside of these cells consist of well conducting solutions. These solutions are separated by the almost insulating lipid bilayer membrane. [2].

It was formerly discussed that the lipid bilayer structure of the cell membrane gives rise to a parallel plate capacitor whose capacitance is proportional to the area of the membrane. The approximate capacitance created by the lipid bilayer of the cell is 1 µF/cm2 and closer to 0.7 µF/cm2 for cardiac cells. The ionic channels in the membrane create a specific resistance or resistivity of the order of 1000 Ω- cm2.

Measurements of passive membrane properties have shown a resistance about 10 to 20 KΩ/cm2 and a capacitance of about 1 to 1.5 µF/cm2 across a membrane for cardiac Purkinje and ventricular muscle fibers. The internal resistance along the cell fiber, Ri, is about 100 to 200Ω for Purkinje and about 400Ω for ventricular muscle fibers. This internal resistance, Ri, is about twice as big as the external resistance, Ro. The internal longitudinal resitance, Ri, determines the propagation velocity and can be regarded as a low resistance path. Ri is not altered by frequency of pacing but it increases by cooling, lack of oxygen, alcohol, and other conditions affecting the cell. [2].

20 While the cable analogy is quite helpful in understanding the cardiac impulse propagation, there are major differences between a cell and a real copper cable. Compared to an insulated copper cable, the biologic cable is an inferior conductor in which the signal attenuates rapidly and dissipates within 2 mm. Moreover, the membrane is leaky compared to a cable insulator. [2].

IV.B. Derivation of the Cable Equation

The external longitudinal resistance of a fiber is Ro and the longitudinal resistance of the fiber interior is Ri. Both Ri and Ro are expressed in Ω/cm or resistance per unit length. This unit results from the assumption of no radial flow where both intracellular and extracellular current flow is in the longitudinal direction only. The intracellular and extracellular potentials, Vi and Vo, depend only on the longitudinal coordinate along the fiber axis. The transmembrane potential, Vm, is then equal to Vi-Vo. The intracellular and extracellular voltage drops along a small length, ∆z, of the fiber are calculated according to Ohm’s law:

∆Vi = - Ri * ∆z * Ii ∆Vo = - Ro * ∆z * Io

Where, Ii and Io are the intra- and extracellular longitudinal currents. In the limit as ∆z approaches zero, the above equations reduce to partial differential equations:

∂Vi/∂z = -Ri*Ii ∂Vo/∂z = -Ro*Io

Subtracting the two results in:

∂Vm/∂z = ∂Vi/∂z - ∂Vo/∂z = - [Ri*Ii - Ro*Io]

The membrane current, Im, is related to Ii and Io and also relates the inside and outside currents. Any change, ∆Ii, in Ii along a length ∆z of the fiber is the algebraic sum of the membrane current, Im, leaking out of the intracellular space and any applied, Ia, stimulus current:

∆Ii = -Im *∆z + Ia * ∆z

Assuming no extracellular stimulus current, any change in the extracellular current, ∆Io, is equal to the membrane current leaking out:

∆Io = Im *∆z

In the limit as ∆z approaches zero the above two equations reduce to:

∂Ii/∂z = -Im + Ia ∂Io/∂z = Im

21

Observe that if the total current is set as: I = Ii + Io, then ∂I/∂z=Ia indicating that I remains constant, in fact zero, unless an applied stimulus Ia exists.

Combining the current and the voltage equations produces:

∂Vm/∂z = - [Ri*Ii - Ro*Io] = -Ri*Ii + Ro*Io = -Ri*(I-Io) + Ro*Io = -Ri*Ii +Ri*Io+Ro*Io = -Ri*Ii + (Ri+Ro)*Io

∂Vi/∂z = Ri/(Ri+Ro) * ∂Vm/∂z – (Ri*Ro)/(Ri+Ro) * I ∂Vo/∂z= Ro/(Ri+Ro) * ∂Vm/∂z – (Ri*Ro)/(Ri+Ro) * I

IV.C. Deviations from Resting Potentials

The deviations of intracellular, extracellular, and transmembrane potentials from their resting values, vi, vo, and vm, can be obtained by integrating the above differential equations with respect to z from z to ∞. These deviations, when larger than a threshold value, are action potentials.

It is assumed that at ∞, the cable is always at rest so that vi(∞,t) = vo(∞,t) = vm(∞,t) = 0. It is also assumed that any stimulus occurs prior to t=0 or occurs outside the area of interest so that I(z)=0 as well. Then, a perturbation is set in motion before time zero and the changes or deviations in intra- and extracellular potential characterizing this perturbation are given by: vi =Vi-Vr, vo=Vo-Vr, vm=Vm-Vr vi(z,t) = Ri/(Ri+Ro) * vm(z,t) vo(z,t) = Ro/(Ri+Ro) * vm(z,t)

So, the end results are simple voltage divider equations that divide the transmembrane potential between internal and external areas of the cell directly proportional to the resistances in the inside and outside of the cell. Measuring the amplitudes of the internal and external potentials, therefore, can provide a ratio between internal and external resistances.

Further assuming that no stimulus is injected at z=0, which is equivalent to setting Ia=0, gives a relationship between the transmembrane current per unit length and transmembrane potential. This is the fundamental equation resulting from the assumptions of one-dimensional cable theory:

Im = 1/(Ri+Ro) * ∂2Vm/∂z2

This equation is valid for both Vm and vm.

22 IV.C.1. Passive Membrane: Subthreshold Conduction

In a passive fiber, the voltage transients are kept below threshold and no action potential results. Subthreshold conduction is called “electrotonic” conduction. In medical terminology, electrotonic condition refers to the modified condition of a nerve, when a constant current of electricity passes through any part of it. It is under the electrotonic conduction that the membrane can be represented as a constant resistance Rm, in ohms times unit length of fiber, in parallel with a constant capacitance Cm, in farads per unit length.

Measurements of signal propagation and resistivity have shown that the higher the Rm and the lower the Ri, corresponding to a well-insulated, high-conductivity conductor, the further the signal spreads. [2].

In this subthreshold region, the membrane is linear as opposed to suprathreshold region where the nonlinear membrane conductances come into play. The equivalent circuit for a length ∆z of the membrane becomes the parallel combination of a resistance Rm/∆z and a capacitance Cm*∆z. Further, a voltage, Vr, corresponding to the resting membrane potential is placed in series with Rm. [3]. The cable equation can be written as:

Im *∆z = Cm *∆z * ∂Vm/∂t + (Vm-Vr)/(Rm/ ∆z) or

Im = Cm * ∂Vm/∂t + (Vm-Vr)/Rm

A time constant and a space constant are defined for the membrane as follows:

τm = Rm*Cm ______λ = √ Rm/(Ri+Ro)

The membrane time constant is the time constant corresponding to the capacitance in parallel with the resistance. The membrane space constant has the dimensions of length and determines the tendency for voltage perturbations to spread along the fiber. A large λ indicates greater spread along the cable.

These constants are related as follows:

λ2 * ∂2vm/∂z2 = τm * ∂2vm/∂t2 +vm

The assumption that all currents, except for Im, be longitudinal limits application of the above equation to the case of a small diameter fiber surrounded by an extracellular space of restricted dimensions. Cable equations are not applicable when the extracellular space is large enough so that radial currents may be present. However, if the extracellular

23 space is so large that its resistance is negligibly small, then Ro and Vo are set to zero and cable equations may be used.

A general solution to the above partial differential equation is obtained through Laplace transform. Specific solutions for a current-impulse and a current-step excitation are discussed. The solutions apply, again, when the magnitude of the excitation is below threshold level of generating an action potential.

For a current-impulse input, the transmembrane potential variation, vm, takes the form of a Gaussian distribution whose variance increases in time. In other words, the sharp Gaussian curve flattens as time progresses. At the origin where the pulse is applied, vm is a maximum and decays exponentially with time. At points away from the point where the impulse is applied, the voltage transient increases at first to a maximum and then diminishes to zero.

For a step-current excitation, the steady-state voltage along the cable diminishes exponentially with a space constant λ. In other words, vm=constant*exp(-z/λ) where the constant=(Ri*I0*λ)/2 and I0 is the excitation current step applied at t=0 and z=0. The space constant, λ, is proportional to the square root of the radius of the cable. So, the thicker the cable, the larger the space constant, and the faster the decay of the excitation down the cable. In terms of variation of the vm with time, the transient response of the cable at the origin rises as an error function such that vm rises to 84% of its final value in one time constant. Error function resembles an exponential rise only it rises faster than an exponential curve. At distances further from the origin of the application of the excitation, the rise of the transmembrane voltage is slower and smaller.

Response to a subthreshold step-current excitation indicates that if the applied current to a cable, or cell, is increased to a higher level and kept at that level, the transmembrane voltage increases but drops back to zero with distance from the origin of excitation. Thus, even a sustained increase in excitation current, cannot travel down the fiber indefinitely if the magnitude of excitation is below the required threshold.

IV.C.2. Action Potential: Excitation above Threshold

For action potential propagation, the membrane is excited above the threshold value. Under these circumstances the cable equation is modified to include an iion which is the ionic current per unit length applied to the membrane. Then:

Im = Cm * ∂Vm/∂t + iion (Vm, t)

The equation for the deviation voltage vm yields:

2 2 Im = 1/(Ri+Ro) * ∂ vm/∂z = Cm * ∂vm/∂t + iion (vm, t)

If the cable is circular with a radius of a and an intracellular conductivity of σi, and if the extracellular resistance Ro= 0, then:

24

(a* σi) /2 * ∂2vm/∂z2 = Cm * ∂vm/∂t + Iion(vm, t)

This equation is a partial differential equation for the membrane potential vm(z,t). Here, Iion is the ionic current per unit area of the membrane:

Iion = GK *(vm+Vr-EK) + GNa *(vm+Vr-ENa) + GL *(vm+Vr-EL)

The subscripts correspond to potassium, sodium, and leakage respectively. G values are the conductances and E values are the electrical energy input by each ionic pump.

If all propagation is eliminated or ∂Ii/∂z = 0, then Ia=Im and the above Im equation written for Ia, becomes:

Ia = Im = Cm * dVm/dt + GK *(Vm-EK) + GNa *(Vm-ENa) + GL *(Vm-EL)

This is a nonlinear but ordinary partial differential equation that may be solved for Vm. In this equation, Ia is the applied current pulse per unit area necessary to trigger the action potential. If the amplitude of the applied current pulse is sufficiently large, a propagated or membrane action potential results.

IV.D. Uniform Propagation

In order to find the solution for a propagating action potential, the partial differential equation in vm, repeated below, must be solved in conjunction with equations that express the variability of conductance values, G, with vm.

2 2 (a* σi) /2 * ∂ vm/∂z = Cm * ∂vm/∂t + GK * (vm+Vr-EK) + GNa * (vm+Vr-ENa) + GL * (vm+Vr-EL) (vm Eqn.)

The equations expressing conductance values in terms of vm are derived from the Hudgkin-Huxley model for cell membrane.

IV.D.1. Hudgkin-Huxley Membrane

Hudgkin and Huxley modeled the cell membrane as an equivalent circuit with a variable membrane conductance per unit area, G, in series with a resting potential Vr, where the series combination of the conductance and the resting potential is in parallel with a capacitance per unit area, Cm. The potential across this entire circuit is the transmembrane potential, Vm, defined as intracellular potential minus the extracellular potential.

While the membrane capacitance, Cm, is assumed to be constant, the membrane conductance, G, is a function of the transmembrane potential,Vm, and varies by several orders of magnitude with changes in Vm. As a result, the biological membrane is

25 profoundly nonlinear. Current leaving the cell is defined as positive and the potential Vm is measured from the inside to the outside of the cell.

The series combination of the conductance, G, and the resting potential, Vr, may be divided into a parallel combination of 3 branches. Each of the 3 branches consists of a series combination of a conductance and a resting potential. The 3 branches correspond to sodium, potassium, and leakage currents and are shown by GK, GNa, and GL, in series with EK, ENa, and EL respectively.

The sodium and potassium conductances, GNa and GK, are voltage controlled and vary with Vm. The leakage controlled conductance, GL, is on the other hand a constant ohmic conductance that is independent of Vm and independent of time.

A cell at rest starts in a polarized state. As an action potential, above the required threshold level, excites the cell, the cell is depolarized. Depolarization causes an activation or increase in both sodium and potassium conductivities, GNa and GK. The increase in conductance is almost exponential. Sodium conductance increases faster and stops increasing when the potassium conductance is still increasing. The conductances go through the phases of delay, rise, and fall as the action potential passes through. During fall or deactivation, the conductances return to their resting values.

Hudgkin and Huxley derived a series of first-order ordinary differential equations describing the rise and fall of GNa and GK with variations in Vm. The fundamental assumptions underlying these equations are existence of “distinct and independent channels” for sodium and potassium.

IV.D.2. Hudgkin-Huxley Assumptions

Hudgkin and Huxley did not solve the (vm Eqn.) directly. Rather, they made some simplifying assumptions that agreed with their experimental observations. They assumed that the action potential propagated with constant velocity ν down the cell fiber and its waveform did not degenerate or decrement. The two assumptions of non-degenerating waveform propagating at a constant velocity are together called the assumption of “uniform propagation.” With uniform propagation the (vm Eqn.) is simplified into:

∂2vm/∂z2 = (1/v2)* ∂2vm/∂t2 (wave Eqn.)

This is the well-known wave equation. If at time t=0, the spatial variation of the action potential along the cell is given by ϋm(z), then under the “uniform propagation” assumptions, for any value of z and t, we get: vm(z,t) = ϋm(z-ν*t)

Differentiating vm with respect to z, ∂vm/∂z and ∂2vm/∂z2, gives first and second derivatives of the entity ϋm with respect to its argument z-ν*t. Differentiating vm with

26 respect to t, ∂vm/∂t and ∂2vm/∂t2, gives first and second derivatives of the entity ϋm with respect to its argument z-ν*t, that are multiplied by the velocity –ν once and twice:

∂vm/∂z = ϋm’ and ∂2vm/∂z2= ϋm”

∂vm/∂t = –ν * ϋm’ and ∂2vm/∂t2= ν2 * ϋm”

Inserting these values into the (vm Eqn.) yields an ordinary differential equation in t:

2 2 2 (a* σi) / (2 * ν ) * d vm/dt = Cm * dvm/dt + GK * (vm+Vr-EK) + GNa* (vm+Vr-ENa) + GL * (vm+Vr-EL)

Hudgkin and Huxley solved the above ordinary differential equation by guessing at initial values of velocity ν and calculating the resulting potential vm(t). Assuming that the membrane characteristics, Cm, GK, GNa, GL, remain constant, the value of vm obtained as a solution of the (vm Eqn.) remains constant for all values of propagation velocity, ν, as long as (a* σi) / (2 * ν2) is a constant:

(a* σi) / (2 * ν2) = 1/ K

Thus, under uniform propagation, a new propagation velocity, ν, obtained by a changed in cell conductivity or a change in cell diameter, always satisfies the relationship: ______ν = √ K*(a* σi)/2

As mentioned above, membrane characteristics must remain constant and the change in velocity cannot be obtained by changes in the membrane characteristics. The propagation velocity, ν, is proportional to the square root of the cell radius, a. This same relationship applied to subthreshold propagation as well. An equivalent form of this relationship may be expressed as:

ν2 (Ri + Ro) = constant

IV.D.3. Qualitative Explanation of Propagation

Propagation of action potential is explained by cell membrane currents. Propagation is initiated by local currents that initiate from the site of excitation of action potential. When a cell is at rest potential, the sodium membrane current is generally outward pumping the sodium ions out against their chemical concentration and against the electrostatic field across the membrane. When the action potential reaches a point along the cell cable, the membrane currents start to change to an inward sodium current as a result of partial depolarization and opening of the sodium channels. A portion of this inward current of sodium ions loops backward and outward, joining the potassium ions that are exiting the membrane under concentration gradient effects. The outward potassium current, together with the joining sodium ions, help re-polarize the upstream

27 portion of the cable as the action potential moves downstream. The portion of the inward sodium current that has not gone backward and outward, flows downstream helping to depolarize the cell as the action potential moves forward. This cycle repeats.

IV.D.4. Exact Solution without the Hudgkin-Huxley Assumptions

The exact solution of the suprathreshold cable equation, the (vm Eqn.), was obtained later by a finite difference solution of the partial differential equation. Comparing the exact solution with the solution obtained using Hudgkin-Huxley assumptions, showed that the assumption of uniform propagation is valid for distances of greater than 1 cm from the site of the stimulation. It further showed that for z > 1 cm and t > 1 ms, when a constant propagation velocity has been obtained, the action potential waveforms stay constant and nondecremental.

References 1. http://www.genesis-sim.org/GENESIS/cnsweb/Cabletut.html. 2. Borys Surawicz, Electrophysiologic Basis of ECG and Cardiac Arrhythmias, Willams and Wilkins, Baltimore, 1995. 3. http://hebb.mit.edu/courses/8.515/lecture1/tsld020.htm.

28 V. MECHANISMS OF FIBRILLATION

This section begins with a finer view of the fibrous structure of heart muscle cells. Next, the section reviews the conduction velocities of the action potential. Once the foregoing groundwork is set, the section ends with a review of factors affecting reentry as the main mechanism suspected of inducing fibrillation.

V.A. Microscopic Structure of the Relevant Heart Muscles and Cells [1a]

Cardiac muscle is composed of fiber bundles. The middle part of the cardiac muscle, the myocardium, is a functional syncytium. The basic cell in the myocardium, the myocyte, is 50 to 100 µm in length with a diameter of 5 to 20 µm. The cells form end-to-end connections by means of intercalated discs. These discs are a continuum of sarcolemma, which is the plasma membrane of a muscle fiber, and consists of two apposing cell membranes and interposed intracellular space.

The plasma membrane of the discs consists of the normal lipid bilayer that in the case of the discs is covered by a thin external coating containing fixed negative charges with an affinity for various ions including calcium. The membranes of the two neighboring intercalated discs are separated by an interstitial gap of 200 to 300 Å. At certain nexus points or gap junctions, the two adjoining membranes approach to become separated by only 20 to 30 Å. At the nexus points or gap junctions, the low electrical impedance facilitates rapid spread of electrical activity from cell to cell.

There are channels in the gap junctions formed by the members of a family of proteins called connexins. These proteins show unique conductances and voltage sensitivities. Purkinje and ventricular muscle fibers express distinct tissue-specific patterns of connexins. The difference in patterns of the connexins in each fiber contributes to the differences in the conduction properties of the fiber.

The plasma membranes of the muscles fibers, the sarcolemma, fold to form a system of transverse tubules that enables the extracellular fluid to penetrate deep within the cytoplasm. These transverse tubules are called the T system. The T tubules do not actually open into the cytoplasm, but they bring action potential and ions closer to the sites of electromechanical coupling. The T tubules do not exist in the fibers of the specialized conducting system and in most atrial cells.

Another system in the plasma membrane of the muscle fibers is called the SR system for sarcoplasmic reticulum. Reticulum is a term applied to any fine network. SR is a fine network of tubules surrounding the myofibrils but it is not in communication with the extracellular space. The SR contains discrete reservoirs, called junctional SR, close to the membrane of the muscle fiber in the region of T tubules. These reservoirs are believed to be the site of storage and release of calcium-initiating contractions.

A simplified sequence for contraction of heart muscle fibers follows. Changes in membrane potential during activation spread rapidly from cell to cell in the preferential

29 direction across the gap junctions and also reach deep into the cell through the T system. The junctional SR reservoirs release calcium that causes the contraction of the fibers. SR subsequently, takes back the calcium leading to relaxation of the muscle.

The continuity of the atrial and ventricular muscles is broken up by atrio-ventricular fibrous tissue that consists of long narrow grooves disconnecting the atria from the ventricles. As a result, the atria and the ventricles remain connected only through the bundle of His which penetrates the tunnels of the fibrous tissue. The atrial and ventricular fibers attach to the fibrous skeleton of the heart only at certain points. In short, the heart muscle fibers are arranged and connected such that mere contiguity of tissue does not imply structural continuity; and, cardiac muscle between the atria and the ventricles does not constitute a conducting tract.

V.B. Conduction Velocity of Action Potential

At a threshold potential, the displacement of the membrane potential increases out of proportion to the applied current and generates an action potential in an all or none fashion. Repolarization consists of an initial phase that is fast in Purkinje and ventricular fibers followed by a slow plateau and ends with a fast terminal repolarization. Even the fast part of the repolarization, however, is still a 100 times slower than the process of depolarization. [1b].

Action potential in various cardiac fibers takes different shapes based on the mechanism of depolarization and the presence or absence of spontaneous diastolic depolarization. Atrial and ventricular myocardial fibers have stable resting membrane potential while the specialized conducting fibers consisting of the SA node, specialized atrial fibers, AV node, and Purkinje fibers do not have a stable resting membrane potential. This means that after reaching the maximum diastolic potential, the cells without a stable resting membrane potential, start to depolarize again. This automatic depolarization may bring the membrane back to the threshold potential. [1b]. Slow diastolic depolarization, which brings the membrane potential to the threshold potential for rapid depolarization, is caused by the action of one or several cardiac membrane currents and the mechanisms differ among different types of automatic fibers. [1d].

It was mentioned that once initiated, depolarization spreads at an all-or-none fashion at a velocity closely paralleling the maximum velocity of depolarization (Vmax). In fact, in a cable-like structure, Vmax often parallels the square of the conduction velocity. Thus determination of Vmax has been useful in explaining conduction changes. [1c].

Absent changes in threshold potential, the value of Vmax depends on the membrane potential at the onset of depolarization. The membrane potential at the onset of depolarization is called the take-off potential. Vmax increases with more negative take- off potential with the maximum Vmax occurring at -85 mV take-off potential. With take- off potentials more negative than -85 mV, Vmax shows no further increase. At take-off potentials less negative than -85 mV, however, Vmax falls. At a take-off potential of -55 mV, there is no take-off; action potential cannot be elicited at a potential this high. [1c].

30

During the relative refractory period, the action potentials have reduced Vmax and the resulting impulses propagate at decreased conduction velocity. [1c].

In addition to depending on the take-off potential, Vmax also depends on duration of preceding depolarization. When diastolic depolarization is of short duration, i.e., when it rises steeply during stimulation, Vmax may remain normal. However, as the duration of depolarization is increased beyond 20 ms, Vmax keeps decreasing. This is called time- dependent inactivation. The inactivation also increases with an increasing level of sub- threshold depolarization. Inactivation is complete after 250 ms at the subthreshold level of depolarization. This interval represents steady-state inactivation time. For diastolic intervals that are much longer than 200 ms, slow conduction dependent on diastolic depolarization is manifest. [1c].

The resting membrane potential for cardiac Purkinje, atrial, or ventricular muscle fibers is about -80 to -90 mV. The depolarization in the cardiac fibers shoots up to about 30 mV resulting in a total upshoot of about 120 MV. The excitation wave, resulting from depolarization, propagates along both the interior and the exterior of the polarized membrane. [1b]. The velocity of the excitation wave depends on the velocity of depolarization; the upstroke velocity of the cardiac action potential is the most important determinant of propagation velocity. For example, the duration of upstroke in the ventricular muscle fiber is about 1 ms and in the Purkinje fiber 0.3 ms. This translates into an upstroke velocity of about 100 V/s for the ventricular muscle and 300 V/s for the Purkinje fiber. Upstroke is often shown as a straight line. In reality, upstroke is in the shape of a sigmoid with the maximum velocity, Vmax, occurring at the middle portion. Vmax is about 200 to 300 V/s for the ventricular and about 600 to 1000 V/s for the Purkinje fiber. So, the upstroke velocities are average velocities. The corresponding conduction velocities are about 0.5 to 1.0 m/s in the ventricular muscle fiber and about 2 to 3 m/s in the Purkinje fiber. [1b].

Conduction velocity also decreases when the membrane is hyperpolarized or when its potential becomes more negative than -85 mV. So, changes in conduction velocity parallel changes in Vmax during depolarization at rest, during repolarization, and during diastolic depolarization. Conduction velocity can also change by altering the relationship between resting potential and threshold potential. [1c].

V.C. Reentry Mechanisms

Reentry is a common mechanism responsible for out of place beats and rhythms and is believed to play a dominant role in the precipitation and maintenance of life-threatening tachyarrhythmias. Reentry can occur in the ventricles, the atria, the AV node, and along the accessory atrioventricular pathways. This part focuses on ventricular tachycardia. [1e].

Circus movement is defined as a contraction or excitation wave traveling continuously in circular fashion around a ring of muscle or through the wall of the heart. Circus

31 movement forms the basis of reentrant currents in humans. The following are important in the maintenance of the circus movement: unidirectional block in the path of the signal and slow conduction around the block causing re-excitation of the previously blocked site. Physical interruption of the circuit terminates the circuit movement and prevents its reinitiaiton. [1e].

Connections of Purkinje fibers to ventricular muscle are thought to be weak links in the path of propagation of an impulse. At these restricted junctions, for example, electrotonic conduction may be of importance. Also, octanol has been shown to slow down conduction across the Purkinje-ventricular muscle junction by 80% whereas conduction in the ventricular layer is slowed by only 10%. [1c].

A common example of reentry occurs at the Purkinje-ventricular junction where a loop of Purkinje fiber is attached to ventricular muscle. It is hypothesized that a blockage in one of the two paths of the Purkinje loop, slows down travel of the signal in that strand. In the meanwhile, the signal that has traveled down the other strand gets a chance to travel through the muscle fiber and come back up the blocked strand. The blocked strand is blocked in one way and not the other and the reentrant signal travels quickly up the strand interfering with the signal that was traveling down through the unidirectional blockage. In two adjacent muscle fibers, one of the fibers may have a unidirectional block. The signal traveling down the adjacent fiber can make a U-turn in the blocked fiber and interfere with the original signal traveling in the direction of the unidirectional block.

The foregoing are unproven hypotheses and models trying to explain reentry. [1e]. The impetus for this paper is that reentry, or any other mechanisms that may underlie fibrillation, are not known to a desirable degree of certainty. Obtaining detailed maps of fibrillation activity is hoped to provide better insight into reentry mechanisms or alternative mechanisms giving rise to fibrillation.

V.C.1. Factors Present for Reentry [1e]

Several factors must be present for reentry. The slowly conducting reentrant pathway must have excitable entry and exit points and must be isolated laterally from the rest of the heart. The isolation protects the pathway from invasion. The isolation may be anatomical or functional. Functional isolation means that the boundaries are refractory during reentrant propagation. An anatomical pathway is a pathway incorporating well- defined anatomic structures, such as an accessory pathway or a His bundle, and anatomical barriers are like an AV valve ring or a scar.

V.C.2. Wavelength and Length of Reentry Pathway [1e]

The wavelength of the electrical impulse is defined as the distance traveled by the depolarization wave during the time the tissue restores excitability sufficiently to propagate another impulse. Wavelength equals conduction velocity (m/s) multiplied by the duration of the functional refractory period (s). For example, in a Purkinje fiber, wave travels at 3 m/s with a refractory period of 0.4 s resulting in a wavelength of 1.2 m.

32 This wavelength is too large to be able to reenter the fiber. However, in a depressed depolarized myocardium cell, the conduction velocity can drop to 0.05 m/s; also, the refractory period of a premature beat is 0.2 s; resulting in a wavelength of 1 cm. This short wavelength can reenter heart muscle fibers which together are longer than 1 cm.

In an anatomically defined circuit, a wavelength shorter than the circuit implies the presence of an excitable gap. So prolonging the wavelength by increasing conduction velocity or lengthening the refractory period will suppress smaller circuits. Smaller circuits create shorts in the main circuit and short circuits are expected to facilitate fragmentation and fibrillation. When wavelengths are short, a small region of block is sufficient to cause arrhythmias.

V.C.3. Unidirectional Block [1e]

Unidirectional block can be continuous or transient. Blocks may result from depression of conduction or prolonged refractoriness. Among the two, depression of conduction caused by cellular uncoupling and abnormal anisotropy may be the main cause of blocks. A computer model simulating reentry has shown that reentry can be initiated in a homogeneous one-dimensional sheet of coupled elements by changing the sequence of activation and the amount of coupling among individual elements. This reentry happens in absence of any obstacles or any dispersion of the effective refractory periods. So mere cellular uncoupling is enough to initiate reentry.

Nonuniform anisotropy can also create reentrant ventricular tachycardias. In one experiment, activation proceeded along the lines of functional conduction block. Here, the block was created by nonuniform anisotropy. Side-by-side uncoupling of the muscle bundles oriented parallel to the long axis was due to slow activation in direction transverse to the myocardial fibers. In this experiment, abnormalities of the action potential or a prolonged refractory period were not thought to be behind reentry.

In another experiment, reentry was elicited within a small 1 to 2 mm2 area as a result of discontinuous anisotropic propagation alone or in combination with nonuniform repolarization.

In a computer model of a ring-shaped, one dimensional cardiac fiber, cellular uncoupling played an important role in the genesis and maintenance of unidirectional block and reentry along with nonhomogenous refractoriness, excitability, fiber cross-sectional area, and gap junction resistance.

V.C.4. Excitable Gap and Duration of Revolution

The excitable gap is the time interval between the end of the refractory period which coincides with the recovery of excitability and the arrival of next impulse. The excitable gap depends on the length of the pathway, conduction velocity, and duration of refractory period.

33 The duration of revolution depends on the length of the pathway, conduction velocity, and refractoriness. The transit time can be lengthened by uniform reduction of conduction velocity or by delayed transmission in one or more regions of the circuit. If the excitable gap is present, lengthening of refractoriness will reduce the excitable gap but will not prolong the duration of revolution unless the fully excitable gap is eliminated.

V.C.5. Documenting Reentry

In small tissue strands, such as segments of Purkinje fibers, reentry can be documented with a limited number of strategically placed recording electrodes. In the entire heart, reentrant activity appears as continuous electrical activity spanning the duration of the cardiac cycle. Such activity can be recorded with a composite electrode overlying a portion of the epicardium.

References 1. Borys Surawicz, Electrophysiologic Basis of ECG and Cardiac Arrhythmias, Willams and Wilkins, Baltimore, 1995. 1a. Surawicz, Chapter 2. 1b. Surawicz, Chapter 3. 1c. Surawicz, Chapter 5. 1d. Surawicz, Chapter 6. 1e. Surawicz, Chapter 10.

34 VI. PACEMAKER AND ICD BASICS

Permanent pacemakers are devices that provide electrical stimuli to cause cardiac contraction during periods when intrinsic cardiac electrical activity is inappropriately slow or absent. They function by sensing intrinsic cardiac electric potentials. If these potentials are too infrequent or absent, electric impulses are transmitted to the heart, thereby stimulating myocardial contraction.

An ICD is a specialized device designed to directly treat a cardiac tachydysrhythmia. If a patient has a ventricular ICD and the device senses a ventricular rate that exceeds the programmed cut-off rate of the ICD, the device performs cardioversion/defibrillation. Alternatively, the device, if so programmed, may attempt to pace rapidly for a number of pulses to attempt pace-termination of the ventricular tachycardia.

Newer devices are a combination of ICD and pacemaker in one unit. These combination ICD/pacemakers are implanted in patients who require both devices.

This section reviews the basic components of pacemaker and ICD devices and the considerations underlying their design and operation. It reviews the required terminology and presents the complications faced by use of each device. This section, in short, provides a basis for understanding pacing and defibrillation.

VI.A. Pacemakers [1, 2]

Pacing systems consist of a pulse generator and pacing leads. With permanent systems, endocardial leads are inserted into the venous system, usually via a vein, and advanced to the right or left ventricle and/or atrium. Permanent leads are unipolar or bipolar. To complete the electrical circuit in a unipolar system, the metallic pulse generator serves as part of the circuit.

The pulse generator is placed subcutaneously or submuscularly and is connected to the leads. Pulse generators contain, among other things, a battery, an output circuit, a sensing circuit, and a timing circuit. The battery most commonly used in permanent pacers is a lithium-iodide type and has a life span of 5-8 years.

Pacemaker energy output is dependent upon the signal amplitude and pulse width. Signal amplitude is measured in electrical units of volts or milliamperes. Pulse width is a measure of output duration and is measured in milliseconds. For proper permanent pacer operation, signal amplitude and width are set high enough to reliably achieve capture of the myocardium, yet low enough to prolong battery life.

Pulse generators can be set to a fixed-rate (asynchronous) or demand (synchronous) mode. In the fixed-rate mode, an impulse is produced at a set rate and has no relationship to the patient's intrinsic cardiac activity. This mode carries a small but inherent danger of producing lethal dysrhythmias should the impulse coincide with the vulnerable period of

35 the T wave. In the demand mode, the sensing circuit searches for an intrinsic depolarization potential. If this is absent, a pacing response is generated.

VI.A.1. Pacing Codes

The North American Society of Pacing and Electrophysiology (NASPE) and the British Pacing and Electrophysiology Group (BPEG) have developed a code to describe various pacing modes. This is illustrated in the following table

Pacemaker code used to describe various pacing modes Chamber Chamber Response to Programmability Paced Sensed Sensed Beat Rate Modulated 1st Letter 2nd Letter 3rd Letter

V V T P

A A I M

D D D C

O O O R

O

V = Ventricle A = Atrium D = Dual (Both Chambers) O = None T = Triggered I = Inhibited D = Double (Atrial Triggered and Ventricular Inhibited) P = Simple Programmability M = Multiprogrammable C = Communicating (Telemetry)

A pacemaker in VVI mode denotes that it paces and senses the ventricle and is inhibited by a sensed ventricular event. Alternatively, AAT mode represents pacing and sensing in the atrium, and each sensed event triggers the generator to fire within the P wave.

The DDD mode denotes that both chambers are capable of being sensed and paced. This requires two functioning leads, one in the atrium and the other in the ventricle. It should be noted that a two-wired system need not necessarily be in DDD mode, since the atrial or ventricular leads can be programmed off. Additionally, single tripolar lead systems are available that can sense atrial impulses and either sense or pace the ventricle. Thus, this system provides for atrial tracking without the capability for atrial pacing and is suitable for use in patients with certain kinds of heart malfunction.

36

Pacemaker programming can be performed noninvasively by an electrophysiologist or cardiologist. Because of the myriad of pacemaker types, patients should carry a card with them providing information about their particular model. Most pacemaker generators, however, have an x-ray code that can be seen on a standard chest x-ray. The markings, along with the shape of the generator, may assist with deciphering the manufacturer of the generator and pacemaker battery. This may be helpful in the event a patient neither recalls the company nor has the permanent pacemaker card.

VI.A.2. Magnet Inhibition

Placing a magnet over a permanent pacemaker causes sensing to be inhibited by closing an internal reed switch. This temporarily "reprograms" the pacer into the asynchronous mode, where pacing is initiated at a set rate. Each pacemaker type has a unique asynchronous rate for beginning-of-life (BOL), elective replacement indicator (ERI), and end-of-life (EOL). Therefore, application of a magnet can determine if the pacer’s battery needs to be replaced.

VI.A.3. Pacemaker Complications

Major pacemaker complications include failure to output, failure to capture, and failure to sense correctly.

Failure to output occurs when no pacing spike is present despite an indication to pace. This may be due to battery failure, lead fracture, a break in lead insulation, oversensing (inhibiting pacer output), poor lead connection at the takeoff from the pacer, and "cross- talk" (when atrial output is sensed by a ventricular lead in a dual-chamber pacer).

Pseudomalfunction is a type of output failure characterized by a phenomenon termed hysteresis. This occurs when a pacer is set to sense below the lower pacing rate limit. For example, if the lowest pacing rate programmed is 60 beats per minute (bpm), a pacer set to sense down to an intrinsic rate of 50 bpm, a hysteresis rate, begins to pace at 60 bpm when the patient’s intrinsic rate falls below 50 bpm. It continues to pace at the lower rate limit of the pacemaker, in this example 60 bpm, until it again senses intrinsic activity. The intrinsic activity sensed by the pacemaker may be at above 50 bpm but below 60. This sensed event inhibits pacing, even though the intrinsic activity sensed is below the necessary 60 bpm. The pacemaker permits the intrinsic rate to go down to 50 bpm before pacing at 60 bpm.

Failure to capture occurs when a pacing spike is not followed by either an atrial or a ventricular complex. This may be due to lead fracture, lead dislodgement, a break in lead insulation, an elevated pacing threshold, myocardial infarction at the lead tip, certain drugs, metabolic abnormalities, cardiac perforation, poor lead connection at the takeoff from the generator, and improper amplitude or pulse width settings.

37 Oversensing occurs when a pacer incorrectly senses electrical activity and is inhibited from correctly pacing. This may be due to muscular activity, particularly oversensing of the diaphragm or muscles, electromagnetic interference, or lead insulation breakage.

Undersensing occurs when a pacer incorrectly misses intrinsic depolarization and paces despite intrinsic activity. This may be due to poor lead positioning, lead dislodgment, magnet application, low battery states, or myocardial infarction.

A final category of pacer failures is termed operative failures. This includes malfunction due to mechanical factors, such as infection, skin erosion, internal bleeding, lead dislodgment, blood clots in the veins, and a number of other physiological disorders.

VI.B. ICD [1, 3, 4]

This subsection reviews the implantable cardioverter-defibrillator technology. The implantable defibrillator was first tested in the 1970s. The Johns Hopkins University Hospital tested an implantable defibrillator in humans by putting a lead at the apex of the right ventricle and the other in the superior vena cava applying a shock of 5-15 joules. This same hospital performed the first human implantation in 1980. The early system used epicardial patches. The early electrode surfaces were large and sandwiched the heart. ICD received approval from the FDA in 1985 and was first marketed by Cardiac Pacemakers Inc. (CPI) which was a subsidiary of Eli Lily and Company.

VI.B.1. General Principles Underlying ICDs [3]

ICDs must distinguish between a VF and a VT and apply the corresponding shock. Also key in successful defibrillation are creation of sufficient voltage gradient and depolarization of sufficient amount of tissue. The shock should be able to depolarize all tissue and uniformly prolong all refractoriness.

To understand defibrillation must first understand fibrillation. In order to have sustained fibrillation, a critical mass of the heart must be involved. Fibrillation is initiated and sustained by continuous reentry of its wavelets within a critical mass of the heart. As a result, small hearts are difficult to fibrillate while large hearts are difficult to defibrillate. To be successful, a defibrillation shock must reduce the excitable volume of the heart to below this critical mass. The defibrillation shock depolarizes a critical volume or mass of the heart. When a sufficient mass is depolarized, there is not enough mass left for the fibrillatory wavefront to propagate.

ICD is meant to provide a safe and effective therapy for VT and VF. Thus, even early in its development, the concept of measuring a drop in the right ventricular pressure as a means of detecting VF was abandoned in favor of a more reliable electrical signal monitoring. An epicardial sensing system replaced the early intravascular system.

In the early ICDs the functions were nonprogrammable but reliable. Sensing of VT was rate-based and VF was sensed using a probability density function. The rate cut-off was

38 fixed. All therapies were high-energy shocks at 30 or 35 joules. Second and third generations of ICD added pacing and tier-therapy features.

VI.B.2 Optimal Defibrillation Methods [3]

Early experiments indicated a simple dose-response relationship for successful defibrillation. They also indicated that the energy required for defibrillation correlated with the size of the heart. At the same time, a shock of excessive strength would fail to defibrillate and a shock of certain strength that would be successful in one occasion would fail in another.

Passing an electrical current of sufficient intensity for a sufficient period of time can end ventricular fibrillation. For a defibrillation shock to be effective, it must create a sufficient voltage gradient and current density throughout the applied area of the heart. Achieving this result requires overcoming the limitations that are created by the variation of current distribution in the heart and other parts of the body. Factors affecting the current distribution are: distance between the electrode source of energy and the tissue, tissue resistivity which is both tissue specific and anisotropic, and the electrical state of the tissue because for generation of action potential, partially depolarized tissue requires a larger potential gradient than a fully repolarized tissue.

It is desirable to prolong the refractoriness period. A prolonged refractory period promotes uniformity of tissue recovery and prevents propagation of any remaining activation front and reentry of post-shock responses. With a defibrillation shock of inappropriate intensity, post-shock responses may send the heart back into fibrillation, while, with a sufficiently prolonged refractory period, re-fibrillation would not occur.

The difference in potential gradient between the highest and the lowest sites can vary about 20 to 1 in the case of epicardial or intracardiac defibrillation. So, the ICD must create a voltage gradient of 80 to 120 V/cm in the nearest tissue to get a gradient of 4 to 6 V/cm in distant tissues. At the same time, a gradient exceeding 200 V/cm can create cellular damage.

VI.B.2.a. Upper Limit of Vulnerability

As mentioned above, a weak shock resulting in an unsuccessful defibrillation creates a new activation front. This front when encountering a partially refractory tissue causes re- entry and initiates re-fibrillation. This re-fibrillation is similar to inducing fibrillation during the vulnerable period when the cells are neither in full polarization nor in the absolute refractory period.

A shock of sufficient strength, on the other hand, will not induce re-fibrillation. The shock strength above which fibrillation is not induced is called the “upper limit of vulnerability” or ULV. No matter when during the depolarization-polarization cycle the shock is applied, it will not cause fibrillation when its intensity is above the ULV. Another definition for ULV is the highest energy that can induce fibrillation when

39 applied during the vulnerable period of repolarization. Slight changes in the metabolic environment or in the automaticity can change the value of the ULV and turn an otherwise successful shock into an unsuccessful shock. Therefore, the defibrillation threshold is of a probabilistic nature and requires the use of a probability density function (PDF).

VI.B.2.b. Waveform

The defibrillation waveform of an ICD is also critical in the delivery of an effective shock. For implantable devices, the best waveform is a low-tilt truncated pulse. This waveform is clinically advantageous because it avoids the delivery of unnecessarily high initial current which may cause tissue damage and also limits the delivery of low current at the terminal portion. Such low currents are not only ineffective in terminating fibrillation but also may re-initiate fibrillation.

The waveform is further biphasic. A single truncated waveform is called monophasic and was the form used in early ICDs. In a biphasic waveform, the first and second parts of the waveform are delivered with opposite polarities. With a biphasic wave, the energy required for defibrillation is significantly reduced, approximately halved, because biphasic waves can extend refractoriness with a lower intensity. The biphasic wave can stimulate the potassium-induced depolarized cells in addition to prolonging refractoriness in partially repolarized tissue.

To obtain the optimal waveform, the initial polarity, total and individual pulse amplitude and duration, and degree of tilt are among the variables that have been studied. Some of these factors are programmable in the current ICDs. For example, many ICD units use a 65% tilt meaning that the wave form is reversed in polarity after its voltage drops to 65% of the highest voltage value. Recent ICD units use a 50% drop. The smaller drop results in shorter phase duration as well.

VI.B.3 Defibrillation Electrodes and Leads [3]

Early ICD defibrillation and lead designs used a single endovascular system. The lead was placed in the right ventricle and performed both sensing and shocking. The early leads required a high current, voltage, and energy. Subsequent models used one or more epicardial patches. Epicardial patches assured optimal energy delivery without tissue damage. Thereafter, the electrode and lead system evolved into multiple endocardial electrodes and patches and eventually into a single lead with separate high-integrity sensing and defibrillation.

At the beginning a combination of several electrodes had to be used for delivering the defibrillation pulse. For example, Medtronic units used a combination of a right ventricular apical coil, a superior vena caval coil, a coronary sinus coil, and a subcutaneous patch. Guidant systems used a combination of a right ventricular coil, a superior vena caval coil, and a subcutaneous patch or array. With the advent of biphasic

40 waveform, defibrillation energy was significantly lowered eliminating the need for multiple electrode combinations.

Integration of sensing and defibrillating function in the same electrodes was also found to have many disadvantages. A strong defibrillation shock would create sensing and pacing problems if applied by the same electrode. Also, the large electrode surface needed for defibrillation was not ideal for sensing and pacing purposes.

For an epicardial system, the defibrillation electrodes were comprised of small or large epicardial patches while the sensing electrodes were a pair of epicardial screw-in leads or a bipolar endocardial lead. In later models, the sensing and defibrillation electrodes of the endocardial lead would run through the same system but would be positioned at different locations. In some models the sense and pace electrodes are combined into one and defibrillation electrodes are separate.

VI.B.3.a. Material

Electrode materials must be highly conductive, biostable and biocompatible with a low corrosive tendency. Therefore, ICD lead material is similar to that used in the pacing industry with similar advantages and disadvantages. The electrode material is either titanium, in the case of patches, or platinum/iridium alloy in most endocardial leads. The defibrillation electrode patches are insulated with silicone. For the endocardial lead, both silicone and polyurethane are used.

Factors that impact ICD leads are similar to those impacting pacemaker leads. Advances in both areas are closely related.

VI.B.4 Batteries and Capacitors [3]

One clear difference between pacemaker and ICD technology is in energy storage and delivery. An ICD must deliver up to 30 joules of energy in a matter of milliseconds while a pacemaker typically delivers 1 to 10 microjoules.

ICD batteries are lithium-based because this kind of battery has advantages in terms of stability, high energy density, and fast discharge kinetics. These batteries must have low impedance which is accomplished by using lithium-vanadium or lithium-silver-vanadium material in contrast to the high-impedance lithium-iodine batteries used for pacemakers. Lowering internal impedance is also accomplished by increasing the anodal and cathodal surface areas. As a result, the ICD battery is usually larger than its pacemaker counterpart. Further, in order to deliver a high voltage over a short period of time, two batteries may be used in series.

The beginning of life voltage of an ICD may be 3.2 V or 6.4 V depending on whether one or two batteries are used. To deliver up to 750 volts with a 3.2 V battery, ICD uses capacitors. The energy stored in a capacitor is dependent upon its material and volume. Typically, two or three capacitors are connected in series. The addition of these

41 capacitors that are usually large and their insulating material is one reason for the large size of the ICD unit. The large battery surface and multiple capacitors comprise the bulk of the pulse generator volume of the ICD.

With time, the battery voltage drops and the time to fully charge the capacitors increases. These two parameters are used to warn that the ICD is near its end of life.

VI.B.5 Cardioversion and Pacing [3]

Cardioversion is the therapy for terminating VT. VT can be terminated using a much smaller energy than that required to end VF. ICD units allow very low energy levels for treating VT. The lowest effective energy is known as the Cardioversion Energy Requirment (CER) is determined by testing. With each test shock, both the efficacy of cardioversoin and its potential accelerating effect should be analyzed.

The key purpose for applying a cardioversion shock instead of defibrillation is the lower pain inflicted by the lower shock. However, only very low energy shocks that are less than one joule in energy are perceived significantly less painful by the patient. So, with a shock of above 5 or 10 joules the patient gains only minimally in terms of comfort. Further, determination of the shock value is imprecise. So, low-energy shock for VT therapy may be backed up with a high-energy shock.

Pacing in ICD is used as a therapy for both bradycardia and tachycardia. The concept of using pacing for treating tachycardia is not new. The objective of anti-tachycardia pacing is to deliver pacing stimulus such that the pacing stimulus captures the excitable gap and renders it refractory to the spontaneous depolarization caused by VT. This is achieved by delivering a train of pacing stimuli.

VI.B.6. ICD Complications [1]

Major ICD complications include operative failures, sensing and/or pacing failures, inappropriate cardioversion, ineffective cardioversion/defibrillation, and device deactivation. Operative failures are identical to those found in regular pacemakers.

Sensing problems similar to those seen with pacers may occur with ICDs. An example of appropriate failure to treat is when a device has a cut-off rate of 180 bpm. If ventricular tachycardia occurs at 160 bpm, the device, appropriately, fails to cardiovert the patient since the rate of the dysrhythmia is below the programmed rate cut-off.

Inappropriate cardioversion is the most frequent complication associated with ICDs. This is an issue especially in patients with a tendency for atrial fibrillation or in patients that have received multiple shocks in rapid succession without prior symptoms. For example, if, in the aforementioned example, the patient develops atrial fibrillation with a ventricular response of greater than or equal to 180 bpm, the device delivers therapy which would be superfluous because the problem is in the atrium not in the ventricles. Newer devices have certain enhancements that allow discrimination between such

42 rhythms. Causes, other than a supraventricular dysrhythmia, include T-wave oversensing, lead fracture, lead insulation breakage, electrocautery (application of a needle heated by an electric current to destroy tissue for example to remove warts), MRI, and electromagnetic interference.

Failure to deliver cardioversion is caused by failure to sense, lead fracture, electromagnetic interference, and inadvertent ICD deactivation.

Ineffective cardioversion may be due to inadequate energy output, rise in the defibrillation threshold (which may be due to antiarrhythmic medications), myocardial infarction at the lead site, lead fracture, insulation breakage, and dislodgment of the leads or of the myocardial cardioversion patches. The dislodgment of the myocardial cardioversion patches still are seen sometimes in patients who had ICDs implanted during open chest surgery prior to approximately 1993.

Many ICDs deliver a programmed set of therapies per dysrhythmic episode. The number of therapies per episode is manufacturer specific. If a delivered therapy does not terminate an arrhythmia, the device goes to the next programmed therapy. For example, a total of 6 attempts at defibrillation are attempted per episode of ventricular fibrillation. The device attempts defibrillation and then reevaluates the cardiac rhythm. If the arrhythmia persists, it delivers therapy number two and so on, until all 6 attempts have been delivered. Once this occurs, the device does not deliver therapy until a new episode is declared. Of course, an initial therapy for ventricular tachycardia may be pacing rather than cardioversion.

VI.B.7. Examples of Manufacturer’s Advertising Material on ICD [4]

The following material was taken from St. Jude Medical’s web site, and presents an example of the type of features advertised for an ICD and the language used.

Atlas®+ HF ICD Model V-340, 42 J stored/36 J delivered energy, 81 g, 40 cc, 1.4 cm:

The St. Jude Medical (SJM) Atlas+ HF systems are intended to provide ventricular antitachycardia pacing and ventricular defibrillation for automated treatment of life- threatening ventricular arrhythmias. The Systems are also intended to provide a reduction of the symptoms of moderate to severe heart failure (NYHA Functional Class III or IV) in those patients who remain symptomatic despite stable, optimal medical therapy (as defined in the clinical trials section of the manual), and have a left ventricular ejection fraction less than or equal to 35% and a prolonged QRS duration. Contraindications for use of the pulse generator system include ventricular tachyarrhythmias resulting from transient or correctable factors such as drug toxicity, electrolyte imbalance, or acute myocardial infarction.

With 36 joules delivered energy, the Atlas+ HF ICD is designed to provide the best in patient safety, programming flexibility, comfort, and quality of life. The Atlas+ HF ICD

43 is the world’s most powerful ICD offering cardiac resynchronization therapy (CRT), and it packs its power into a compact, 81 gram and 40 cc physiologic-shaped can. In addition to its high energy, the device offers high performance, advanced SVT discriminators, storage of up to 30 minutes of fully annotated electrograms, programmable waveforms for effective DFT management and more. Other features of this ICD include: programmable sense refractory periods to help ensure proper sensing of wider QRS complexes, independently programmable RV and LV amplitude and pulse width, negative AV/PV hysteresis with search, independent RV and LV threshold testing during follow up, impressive options for biventricular pacing therapy—right ventricular sensing along with right ventricular pacing and simultaneous pacing, clear, informative stored EGMs offer quick and easy interpretation of device therapy episodes for diagnostic and programming purposes, potential for reduced implant time with DC Fibber™ VF Induction, comprehensive tachycardia and bradycardia diagnostic data, and proven dual- chamber bradycardia therapy.

Epic™ HF ICD At 36 cc and 73 grams, the Epic HF model V-338 ICD offers superior patient care and simplified follow-up. The Epic HF ICD is the world's smallest and lightest 30 joule (delivered energy) biventricular ICD. Small size doesn't mean sacrifice, though. The Epic HF ICD features advanced resynchronization parameters along with 34 joules stored/30 joules delivered energy and a 7.5 second charge time. This is the world’s smallest 30 joules delivered CRT-D (Cardiac Resynchronization Therapy Defibrillator) with fully independent RV and LV outputs, helps to ensure effective resynchronization therapy with negative AV/PV hysteresis and unmatched flexibility in DFT management.

The St. Jude Medical (SJM) Epic HF systems are intended to provide ventricular antitachycardia pacing and ventricular defibrillation for automated treatment of life- threatening ventricular arrhythmias. The Systems are also intended to provide a reduction of the symptoms of moderate to severe heart failure (NYHA Functional Class III or IV) in those patients who remain symptomatic despite stable, optimal medical therapy (as defined in the clinical trials section of the manual), and have a left ventricular ejection fraction less than or equal to 35% and a prolonged QRS duration. The physician should be familiar with all components of the system and the material in the User's Manual before beginning the procedure. Do not implant the pulse generator if the acute defibrillation lead impedance is less than 20 ohms or the impedance of chronic leads is less than 15 ohms. Damage to the device may result if high voltage therapy is delivered into an impedance less than 15 ohms. The pulse generator provides dual-chamber bradycardia pacing with ventricular resynchronization therapy. If another pacemaker is used, it should have a bipolar pacing reset mode and be programmed for bipolar pacing to minimize the possibility of the output pulses being detected by the device. Ensure that a separate standby external defibrillator is immediately available during the implant procedure or induction testing. The patient should avoid strong magnetic fields since they are potentially capable of inhibiting tachyarrhythmia therapies. Avoid MRI devices because of the magnitude of the magnetic fields and the strength of the radiofrequency (RF) fields they produce. Avoid lithotripsy unless the therapy site is not near the pulse generator and leads as it may damage the pulse generator. Use devices emitting ionizing

44 radiation with caution as they can damage CMOS circuitry in the pulse generator. Avoid diathermy, even if the device is programmed off, as it may damage tissue around the implanted electrodes or may permanently damage the pulse generator. Changes in the patient's disease and/or medication may necessitate reevaluation of the patient's clinical arrhythmias and require reprogramming of the device. Patients may experience psychological effects such as imagined pulsing, dependency, fear of inappropriate shocks, and fear that pacing capability may be lost.

Potential adverse events associated with the Epic HF CRT-Ds, include, but are not limited to the following: acceleration of arrhythmias (caused by device), air embolism, allergic reaction, bleeding, cardiac tamponade, chronic nerve damage, death, erosion, exacerbation of heart failure, excessive fibrotic tissue growth, extracardiac stimulation (phrenic nerve, diaphragm, chest wall), extrusion, fluid accumulation, formation of hematomas or cysts, inappropriate shocks, infection, keloid formation, lead abrasion and discontinuity, lead migration/ dislodgement, myocardial damage, pneumothorax, shunting current or insulating myocardium during defibrillation with internal or external paddles, potential mortality due to inability to defibrillate or pace, thromboemboli, venous occlusion, venous or cardiac perforation. Patients susceptible to frequent shocks despite antiarrhythmic medical management may develop psychological intolerance to an ICD system that may include the following: dependency, depression, fear of premature battery depletion, fear of shocking while conscious, fear that shocking capability may be lost and imagined shocking (phantom shock).

References 1. Barry M Weinberger, Lawrence C Brilliant, “Pacemaker and Automatic Internal Cardiac Defibrillator,” http://www.emedicine.com/emerg/topic805.htm, last updated, July 25, 2001, 2. S. Serge Barold, Ronald X. Stroobandt, Alfons F. Sinnaeve, Cardiac Pacemakers Step by Step: An Illustrated Guide, Blackwell Publishing, 2004. 3. Liem, L. Bing, Implantable Cardioverter-Defibrillator, a Practical Manual, Kluwer Academic Publishers, Boston, 2001. 4. www.sjm.com.

45 VII. SENSING

This section reviews the basics of sensing in both pacemakers and ICD devices. The section explains asynchronous and synchronous pacing, the three-letter codes used for pacemakers indicating the chambers sensed and paced, oversensing and undersensing consequences, filtering, similarities and differences in sensing in pacemakers as opposed to sensing in ICD devices, programming the refractory and blanking periods, distinguishing VT from VF, the difference between dual chamber and single chamber devices and problems such as crosstalk in dual chamber devices.

VII.A. Sensing in Pacemakers [1]

When the heart is incapable of generating the sinus rhythm by itself and misses beats, namely in the case of bradycardia, the pacemaker generates stimuli which cause the heart to beat and create the QRS complex. In asynchronous or automatic pacing, when the pacemaker automatically creates stimuli without sensing the heartbeat, then there is a danger of causing ventricular fibrillation. Because, in asynchronous pacing, even if the heart generates a beat of its own, notwithstanding the heart beat the pacemaker also creates a stimulus. If this pacemaker stimulus falls in the non-refractory vulnerable period, for example the end part of T wave, the untimely stimulus is capable of throwing the heart into fibrillation. So, asynchronous or automatic pacing runs the risk of causing fibrillation. While this risk is quite small in usual clinical situations, it demonstrates the need for sensing.

The voltage can be sensed in the ventricles or the atria and the sensing of a voltage may either inhibit or trigger a stimulus. Three-letter codes for single chamber pacemakers such as VVT, AAT, VVI, AAI, VAT, VAI, AVT, or AVI indicate the chambers paced and sensed (ventricle or atrium) and the mode of response (triggered or inhibited).

VVI or demand ventricular pacing responds to the danger of fibrillation created by an untimely stimulus. With this mode, a competitive rhythm is not possible. Moreover, the battery power is saved by not wasting its pulses. With this type of pacemaker, sensing a QRS complex generated by the heart causes the pacemaker to reset and delays the next pacemaker-generated pulse. Sensing a QRS complex gives rise to an “escape interval” which is the interval between the sensed complex and the next pacemaker stimulus while the interval between two pacemaker stimuli, uninterrupted by a reset, is an “automatic interval.” The escape interval is equal to the automatic interval. However, the escape interval measured by a surface ECG is slightly longer than an automatic interval because intracardiac sensing occurs later than the beginning of the surface QRS complex.

What the pacemaker senses determines the demand pacing. The pacemaker detects what goes on inside the heart itself by measuring the voltage between the two electrodes used for pacing. The voltage associated with intracardiac depolarization is larger than the voltage observed at the surface of the body by the ECG. The intracardiac voltage is called an EGM for electrogram as opposed to ECG which stands for electrocardiogram. The intracardiac voltage, EGM, is behind in time with respect to ECG. This is because

46 the ECG can sense the voltage as soon as it starts in the sinoatrial node (SA) but the EGM has to wait for this signal to reach the sensing electrodes that are located, for example, at the apex of the ventricle. So the sensing electrodes sense the intracardiac potential or EGM in order to determine whether pacing is required or not. The sensors in the ventricle sense the R wave or the refractory period.

The intracardiac electrogram, EGM, senses the difference between the two signals traveling through the heart tissue at the pacing electrodes. In a unipolar system, where only one electrode is used for sensing, the voltage difference is measured between the electrode and an indifferent plate at the pacemaker. The signal arriving at the plate is so small that the voltage sensed is the difference between the spike sensed at the electrode and a flat line. In the bipolar system, the two electrodes are both within the heart and both see the same spike but with a phase difference. A bipolar system usually consists of a ring electrode and a tip electrode where the tip electrode extends further than the ring. The spike of voltage generated by the SA node and traveling through the heart arrives at the ring electrode first and at the tip electrode later generating identical spikes of voltage in both but at a delay. The difference between the two voltages observed creates the bipolar EGM. It is the timing or phase difference that generates the bipolar EGM. The problem with bipolar sensing is that if a signal is generated from a location within the heart that travels perpendicular to the line connecting the two electrodes, this signal arrives at both electrodes simultaneously and there will be no difference in the phase of the two spikes. As a result, the difference between the two will be zero and the activity will go unsensed. If such signals are considered likely, unipolar sensing would be superior.

Oversensing occurs when the threshold for sensing an action potential is set low and the sensors may mistake a T wave for a QRS complex. Undersensing occurs when the threshold is set too high and some QRS complexes may go undetected. A sensitivity of 4 mV means that the pacemaker can only sense a signal equal or greater than 4 mV and it will not detect a 3 mV signal. Application of a magnet on a pacemaker closes a reed switch within the pacemaker that turns the pacemaker into an asynchronous pacemaker that generates pulse irrespective of the sensed signals. Usually, application of a magnet on an ICD deactivates the tachycardia therapy but not its pacing and sensing function for bradycardia therapy.

Sensing and stimulation are two different functions. Sometimes, like in ICD devices, sensing and stimulation are carried out by different electrodes. The heart tissue where the sensing or stimulation occurs may also differ. So, a site within the heart tissue that yields excellent capture does not necessarily provide a good signal for sensing.

VII.A.1. Filtering

A pacemaker unit includes a battery, a timer, an output circuit, and a sensing circuit. The sensing circuit includes an amplifier, a filter that is usually a bandpass filter, and a level detector. In a VVI pacemaker, rate, pulse width, voltage amplitude, sensitivity of the sensing circuit, the refractory period, the effect of hysteresis, and the mode of pacing may

47 be programmable. The programming and reading of the data may be achieved by telemetry.

The sensor electrode senses the slope of the fast rising part of the biphasic endocardial electrogram. The fast rising part occurs as the myocardium underlying the sensor electrode depolarizes. The sensor, thus, measures the slew rate which is the rate of change in signal amplitude per unit of time. A large intracardiac signal (EGM) almost always has a sharp rising slew rate that is easily sensed. In small signals, the slew rate is small and the slope is slow-rising. The sensor may not be able to sense an EGM with a slow-rising slope.

The sensed endocaridal signal is then filtered through a bandpass filter that eliminates both high and low frequencies. Frequency is the same as slew rate. High frequency signals likely originate from an external source such as myopotentials from skeletal muscles or electromagnetic interference. These signals are filtered out so that they can no longer affect the timer. Low frequency signals correspond to T waves, that result from cardiac repolarization, and are also filtered out. Signals with the mid-band frequency, corresponding to the myocardial slew rate, pass through the filter without attenuation.

In a graph of amplitude versus frequency (slew rate), each type of potential existing or occurring in the heart, that can generate a signal, occupies a region of its own that may overlap regions belonging to other types of heart potentials. For example, premature ventricular complexes (PVC) occupy a region with low to medium slew rate and low to high amplitude, myopotentials occupy a region corresponding to medium to high slew rates but very low amplitudes, P and R waves have the same range of frequency but R waves can have much larger amplitudes, and T waves occupy the corner corresponding to low frequency and low amplitude.

Because the various heart signals overlap in their slew rates, discrimination by slew rate alone is not possible. Detecting the amplitude, or level, as well can help further discriminate the signals from one another. The filter must allow in the R waves and the PVC signals but block out the T waves and the myopotentials. Adjusting the band-pass frequency and amplitude to prevent oversensing, sensing T waves and myopotentials, and undersensing, filtering out R waves and PVCs, is difficult.

In order to avoid sensing of the generated stimulus, the created QRS complex and the T wave, afterpotentials, or the combination of the T wave and the afterpotential, the sensor function enters a refractory period after sensing or after pacing. The refractory period blocks the signals received during the period that the heart is generating its own QRS complex or the paced QRS complex.

VII.B. Sensing in ICD [1]

The sensing systems of ICDs and pacemakers have similarities and differences. Both sensing mechanisms have to distinguish true cardiac depolarization from most

48 electromagnetic noise and repolarization signal. Thus, ICD sensing algorithms implement blanking and refractory periods as well. To avoid double sensing, which may cause false diagnosis of a tachyarrhythmia, a non-programmable blanking period is usually set.

ICD sensors have to be sensitive to ventricular fibrillation signals before and after a shock given by the ICD. Filters used by both pacemaker and ICD sensing mechanisms generally reject signals below 10 Hz and above 60 Hz. The frequencies allowed by the filter are subsequently rectified and amplified.

The sensing mechanism of ICD must be able to detect the small and variable depolarization amplitudes associated with VF. The depolarization amplitude during fibrillation may be significantly smaller than the baseline signal. The ICD sensing system has to adjust to ensure detection of small signals. The system has to be able to reject now larger repolarization signals while sensing small ventricular depolarization activity.

Sensing of small and fluctuating VF signals is achieved by a self-adjusting mechanism. One way to self-adjust, used in Medtronic devices, is by adjusting the sensitivity threshold. After a sense or pace ventricular signal, the sensitivity is lowered (the threshold is raised) to avoid sensing a T wave. The threshold gradually comes back down to the programmed value which may be between 0.03 mV to 0.06 mV. Another way of detecting lower amplitude signals of VF while filtering out the higher amplitude signals of T waves, used in CPI and Ventritex devices, is by controlling the gain of the amplifier. The auto gain control, AGC, devices adjust the sensitivity to the size of the R wave. The amplification of incoming signals is changed to assure the sensing of small ventricular depolarizations. When the R wave is adequately large, there is little likelihood of T-wave or P-wave oversensing. However, VF can have R waves as small as 1 mV or less and setting the sensitivity high enough, namely the threshold low enough, to sense such R waves can result in oversensing.

Undersensing or underdetection of fibrillatory signals is not a problem because it is typically corrected automatically by the ICD self-adjust mechanism. If the intracardiac signal changes significantly, some small amplitude R waves may go undetected but when consecutive R waves remain small, the auto-adjust sensitivity system or the automatic gain control system will have sufficient time to adjust. The adverse effect of undersensing that occurs before the system self corrects is that the detection of arrhythmia and the delivery of shock is delayed. To remedy this delay, most ICD models implement an algorithm that takes into account such drop-outs.

VII.B.1. ICD Sensing Systems

The EGM signal generated by a sensor depends on the type of electrodes used and electrode pairs used for sensing vary from manufacturer to manufacturer. In some models (CPI) sensing is performed based on the voltage between the tip electrode and the distal defibrillation coil; these systems are called “integrated bipolar” systems. The

49 integrated systems have larger electrode surfaces and can sense far-field signals. At the same time, integrated bipolar systems are more susceptible to noise. The EGM signals generated by these coil electrograms resemble the surface ECG because they show both the P and R waves and distinguish normal QRS complexes from wide QRS complex.

In other models (Medtronic) sensing is done through the tip and a distal ring electrode; these systems are called “true bipolar” systems. These systems sense only local electrical activity and generate a narrow EGM.

VII.B.1.a. Refractory and Blanking Periods

The refractory period, in pacing, is an interval or timing cycle following a sensed or paced event during which the sense amplifier will not respond to incoming signals. Dual- chamber pacemakers have separate refractory periods for each chamber (atrial and ventricular). In most modern pacemakers, the refractory periods are programmable values.

Blanking period, also known as the ventricular blanking period, is an interval initiated by the delivery of an output pulse during which the sense amplifier of the pulse generator is temporarily disabled. Blanking period is a fraction of the refractory period and occupies the first part of a refractory period. In dual-chamber pulse generators, the blanking period is intended to prevent the inappropriate detection of signals from the opposite chamber (crosstalk). Blanking periods are not available in all pulse generators. The blanking period, stated in milliseconds, may be programmable but is usually preset.

All ICD models identify a tachyarrhythmia based on the cardiac cycle length. The cardiac cycle length between sensed consecutive depolarizations are measured and checked against the programmed length. An ICD like a pacemaker has several refractory periods in its sensing. During these refractory periods the cardiac events are ignored. These refractory periods affect the measured cycle length.

The refractory periods of ICD are similar to those used in the pacemaker. Oversensing is more serious in ICD. Ventricular oversensing causes false detection of a ventricular tachyarrhythmia and in delay of ventricular pacing. To avoid double counting of a ventricular complex that arrives with conduction delay, the ICD has a programmable blanking period of 100-120 miliseconds after a ventricular sensed event. A refractory period is also programmed after a paced ventricular event to avoid T-wave oversensing that would also cause false tachyarrhythmia detection.

Of the two problems of R-wave double counting and T-wave oversensing, the later can be more problematic. To overcome T-wave oversensing, the refractory period may have to extend to a value that would significantly limit the programming of VT detection rate. For example if 400 ms is blanked to eliminate T-wave oversensing, then a VT with a rate of [60,000 ms/min]/ [400 ms] = 150 per min or 150 bpm(beats per minute) would be at the edge of detection and VT events with lower rates would go undetected. The ICD

50 must also incorporate blanking periods after a shock High intensity shocks adversely influence the sensing system.

In dual-chamber ICD devices, appropriate blanking and refractory periods are even more important. For example, atrial oversensing can cause inappropriate inhibition and inappropriate assumption of atrial tachyarrhythmia that will in turn cause inhibition of ventricular therapy. Tracking is a dual-chamber pacing function in which atrial activity is sensed and results in a paced ventricular response after a predefined delay (the AV interval). Inappropriate tracking caused by atrial oversensing can cause rapid ventricular pacing that can in turn induce ventricular tachycardia.

Blanking periods are usually nonprogrammable and refractory periods are usually programmable. Refractory periods must be chosen short enough to allow wide range of tracking but not so short as to allow rapid ventricular pacing in a patient that was susceptible to VT.

VII.B.2. Detection Enhancement in ICD

As stated above, in all ICD models the initial detection criterion is based on the R-R rate or interval. Additional detection enhancement criteria may be programmed. All models count the number of beats per given time period. However, there is a duration criterion in addition to this rate criterion. The fast rate must be sustained for certain duration. The duration criterion can be programmed for as short as 1 second which is preferred for the VF zone when prompt treatment is usually desired. The objective is to allow sufficient time for those arrhythmias that disappear by themselves yet not the kind of delay that delivers the therapy too late. The other objective is to require only a percentage of R-R cycles or heart rate as an indicator of arrhythmia in order to accommodate undersensing.

VT must be distinguished from other types of arrhythmia in order not to mistake a sinus tachycardia or an atrial fibrillation for a VT and apply a shock to the ventricles where the problem is in fact in the SA node or the atria.

In single chamber devices, onset/acceleration, stability, and EGM width or morphology criteria are used for distinguishing a VT from other arrhythmias that do not require a shock.

Dual chamber devices are better tuned in discriminating tachyarrhythmia and are able to screen out supraventricular tachyarrhythimas (SVT) without compromising VT or VF detection. Supraventricular tachyarrhythmias are defined as those that are situated or occurring above the ventricles, especially in an atrium or in the atrioventricular node. The incorporation of the electrical information from the atrial sensors in dual chamber devices causes them to be more accurate than single chamber devices.

VII.B.2.a. Distinguishing VT in Single Chamber Devices

51 In the absence of dual-chamber sensing, VT is discriminated from SVT by using the characteristics of the R-R interval and the morphology of the EGM. Reentry VT has an abrupt and sudden change in its R-R interval right from the onset and thereon it continues to have a stable R-R interval. Whereas, sinus tachycardia does not cause a sudden change and atrial fibrillation does not have stable R-R intervals.

Therefore, the onset criterion is used to distinguish VT from sinus tachycardia and is useful for patients with slow VT or those whose exercise sinus rate may exceed the programmed VT zone. The stability criterion is used to distinguish VT from atrial fibrillation and is useful for patients who have rapid ventricular rates during atrial fibrillation.

Abrupt onset/acceleration of reentry VT and its regular R-R interval do not distinguish a reentry VT from a number of reentry SVT events, such as reentry atrial tachycardia or atrioventricular node (AV) reentrant tachycardia or AV reciprocating tachycardia, all of which are likely to have a sudden onset and regular R-R intervals. Yet, the most common situations that must be distinguished from reentrant tachycardia are sinus tachycardia and atrial fibrillation and using the onset and R-R interval criteria can offer a significant reduction in inappropriate shocks.

When onset or stability criteria are programmed into an ICD, there is always a risk that therapy is withheld for too long when it was actually necessary because some VT will always fail to satisfy the onset and stability criteria. Further, VT and AT can occur simultaneously. If, for example, VT develops after the onset of sinus tachycardia it will have a gradual onset and may be missed. Or, a VT that occurs during an AF with rapid conduction shows an irregular R-R interval. Therefore, a safety mechanism is programmed to prevent withholding therapy for VT.

As an additional safeguard, some devices like the CPI device program a Sustained Rate Duration requirement that must be met in addition to the requirements of onset and stability in order for the shock to be delivered. This feature can be programmed between 10 seconds to one hour.

Electrogram features such as the width, used by Medtronic, or morphology, used by St. Jude/Ventritex, help to determine whether a tachyarrhythmia is wider or dissimilar to a baseline in order to declare it VT. The morphology factors that are compared include the number of peaks, sequence of peaks, peak polarity, amplitude, and width.

VII.B.2.b. Distinguishing VT in Dual Chamber Devices

Current ICD devices are capable of delivering a full spectrum of therapy for ventricular tachyarrhythmia as well as a physiologic brady therapy with dual chamber sensing and pacing. These units are superior not only in delivering the required therapy but also in their sensing capabilities. Thus, these devices are more fine-tuned in screening out SVT without compromising VT or VF detection.

52 Here, SVT can be discriminated from VT using the information from the atrial channel as well. Different models incorporate the dual chamber information differently and many use the Onset and Stability criteria as well. It is also important to prevent a therapy inhibition when the device is facing a true VT and therapy is in fact needed.

In Guidant and CPI models, two parameters Vrate>Arate and AFib Rate Threshold are used to avoid therapy inhibition by Onset and Stability criteria. So, in the case that an irregular tachycardia is detected, therapy is delivered unless the existence of an AFib is confirmed. If a slow-onset tachycardia is in progress, therapy is inhibited only if the V rate is less than the A rate. In Medtronic models, the logic algorithm assesses the pattern and relationship between the atrial and ventricular EGM and distinguishes between VT and various types of SVT such as atrial fibrillation or flutter and sinus tachycardia. In St. Jude/Ventritex models, the logic algorithm first compares the V and A rates and if VPP, stability and n:1 P/R ratio are used for assessment. When RR=PP, RR stability, PR stability, PP stability, and onset criteria are used for differentiation.

VII.B.3. Crosstalk

Crosstalk is a complication of dual-chamber devices. Crosstalk or cross-chamber sensing is defined as the inappropriate detection of an event in one channel by the sensor of the other channel. Crosstalk in ICD can cause a potentially lethal effect. Ventricular oversensing by the atrial channel can cause false sensing of an atrial tachyarrhythmia and may cause erroneous tracking by the ventricular channel. Atrial oversensing by the ventricular channel, a rarer occurrence, can cause false VT detection. Or, atrial flutter can be interpreted as VT by the device

Crosstalk depends on the amplitude of the atrial electrical stimulus and the sensitivity of the ventricular channel. In pacemakers, crosstalk occurs only in dual-chamber devices that are programmed for synchronous AV pacing. The atrial pacing spike is sensed by the ventricular channel as a ventricular event and causes inhibition of the ventricular pacing. This results in loss of left ventricular output, and in the case of a pacer-dependent patient, in syncope or asystole.

Crosstalk depends on the positioning of the leads. The leads should be positioned such that they minimize the possibility of crosstalk or contact with the shocking coil.

VII.B.4. Far Field Sensing [2]

In the study of diffraction and antenna design, the near field is that part of the radiated field that is within a small number of wavelengths of the diffracting edge or antenna. The part of the radiated field that is beyond this is known as the far field.

53

In telecommunication, the term near-field region can mean of the following: (1) the close-in region of an antenna wherein the angular field distribution is dependent upon distance from the antenna; (2) in the study of diffraction and antenna design, the near field is that part of the radiated field that is within a small number of wavelengths of the diffracting edge or antenna; (3) in optical fiber communications, the region close to a source or aperture. The diffraction pattern in this region typically differs significantly from that observed at infinity and varies with distance from the source. In telecommunication, far-field region is the region outside the near-field region where the angular field distribution is essentially independent of distance from the source.

If the source has a maximum overall dimension D that is large compared to the wavelength, the far-field region is commonly taken to exist at distances greater than 2D 2/λ from the source, λ being the wavelength. For a beam focused at infinity, the far- field region is sometimes referred to as the Fraunhofer region.

In ICD and pacemaker sensing, farfiled signals that interfere with the proper sensing of the cardiac electrical event are those that are not seen by the sensor as a dipole event. These farfield signals must be filtered or suppressed.

References 1. Liem, L. Bing, Implantable Cardioverter-Defibrillator, a Practical Manual, Kluwer Academic Publishers, Boston, 2001. 2. http://en.wikipedia.org/wiki/Far_field.

54 VIII. SENSING LEADS

This section presents the history of sensing leads and general considerations of their design and manufacture. The section ends with a review of the sensing leads available as advertised by major pacemaker and ICD manufacturers. Many of the devices use the same leads for sensing, pacing, and at times even delivering a defibrillating shock.

The leads of an ICD or a pacemaker device are insulated wires that carry the heart electrical signal to the pulse generator. They also carry energy from the pulse generator to the heart. One end of each lead connects to the pulse generator. The other end is connected to a heart chamber or is placed under the skin near the heart. Some leads allow a device to sense arrhythmias and deliver electrical energy to correct the heart rhythm. Other leads just allow the device to sense signals or just deliver energy to the heart. An ICD or a pacemaker device may have more than one lead. [1].

The electrical leads of a device, which is outside the chest cavity, must connect to the heart, which is inside. Most leads are threaded through a vein into the chambers of the heart. The leads must be small in diameter, slippery enough to thread through a vein, highly resistant to corrosion, and capable of withstanding flexing for years without breaking. They must sense the very small electrical signals associated with heart’s natural activities while remaining immune to noise introduced by outside sources. Most leads are made of titanium and silicone, and some have an extra-slippery coating of a different type of silicone to make them easy to thread through a catheter. Each has a mechanism at one end for attachment to a heart muscle, by a surgeon who makes the attachment while manipulating the other end. Some leads administer small amounts of a steroid at the attachment point to reduce the inflammation caused by the attachment. [2].

Before embarking on a review of literature regarding leads, this section must introduce the “French unit” which is commonly used for expressing the diameter of leads. French is a scale used for denoting the size of catheters, sounds, and other tubular instruments. One French unit (symbol F) is 0.33 mm in diameter, so that an 18 French or 18F needle has a diameter of 6 mm.

VIII.A. History

Since the start of electrocardiography at the beginning of the 20th century, cardiac pacing has evolved rapidly. At the centennial, it was recognized among the ten best engineering achievements of the last hundred years. Pacemakers were introduced in 1958 and their various aspects of power sources, circuits, leads and electrodes have been continuously improved. [3].

Fundamental characteristics of leads are sensing, waveform morphology, lead geometry, and lead diagnostics.

Pacemakers were initially designed for asynchronous delivery of the stimulation irrespective of heart activity. The electrodes were configured to provide the lowest

55 stimulation threshold. To deliver cardiac stimulation, at least one electrode had to be in contact with the myocardium so endocardial leads were supplied with a tip electrode designed for intimate contact with the cardiac muscle. The position of the second electrode required to complete the pulse delivery circuit was irrelevant and in unipolar systems, the second electrode was provided by the large surface area of the pulse generator itself. In bipolar systems, the second electrode was comprised of a ring that was located 1 to 2 cm away from the distal pacing tip. When sensing was added to pacemakers, this exact electrode design was used for the sensing circuit. This sensing circuit was useful for ventricular sensing but inadequate for atrial sensing. [4].

Atrial sensing is very difficult for two main reasons: positioning and fixation of leads in the atrium is harder and the atrium has intrinsic electrical properties. The earliest attempts for atrial sensing involved the use of epicardial leads that were abandoned because they needed a thoracotomy and because they caused signal attenuation with passage of time. Later, atrial electrodes were used between the atrial wall and esophagus. These electrodes had the same problem of declining P-wave amplitudes. Eventually, there was a complete switch to transvenous systems. The leads were placed within the coronary sinus in an attempt to stabilize and fix the lead. The atrial sensing was poor in this area as well. There was no clear morphologic distinction between P-waves and QRS complexes. Moreover, the ventricular complexes in areas of the coronary sinus were larger than the signals from the atrial myocardium. [4].

Atrial sensing, thus, has two interrelated problems: intrinsically low signal amplitude and difficulties with signal discrimination. Atrial myocardium, which has a significantly smaller muscle mass than the ventricle, generates an electrical signal of lower amplitude (2.5 mV versus 8 to 10 mV). Further, this P signal can not be discriminated from the QRS complex observed from the atrium. [4].

Later observations indicated that: contact between the atrial electrode and the endocardium was not required for P-waves to be seen; signals that were derived by sensors perpendicular to the plane of tissue depolarization were twice the amplitude of those parallel to the wave front; and small surface area electrodes were superior to those with larger surface areas. [4].

The endocardial leads started out as stainless steel and were changed to platinum and platinum alloys in the 1970s. In 1980s carbon materials were used to minimize the stimulation threshold. The insulation material started out as Teflon or polyethylene and was replaced by silicon rubber, which is biostable, and later by polyurethane that requires a thinner layer. [4].

The early atrial leads, like the Smyth J lead in 1969, had a simply exposed conductor coil as its electrode. This lead was implanted in the right atrial appendage and had a 20% rate of incidence of dislodgment, loss of sensing, or loss of capture. In 1976, a J lead with 9- 10 mm long silicone rubber thines was used by Smyth that had a 7% rate of problems requiring reoperation. This lead had an effective 16-French circumference at the tines with a 10-French lead body. The tines pushed against one side of the atrial appendage to

56 force the canted electrode, forming an angle with the body of the lead, into contact with the opposite appendage wall. Later, the tines were trimmed to form the shape of a Christmas tree such that the tines were shorter at the tip, near the electrode, and longer further down the lead. Later, they questioned the value of the extra tines and removed all but the most distal 3 electrodes which were also shortened. At the same time, short-tined ventricular leads had become available also. By 1979, the leads were insulate by slippery polyurethane material, had 3 short tines close to the distal tip, a 6-French lead body, and an insertion body of 9-French. Thereafter, the improvements focused on the electrode tips which are now microporous and steroid eluting. [5].

Active atrial fixation leads began to appear in the early 1970s to solve the dislodgment problems associated with transvenous ventricular leads. The electrodes of these leads were made of metallic barbs or were anchored with metallic or nylon wires. One kind of these early active fixation leads used in the atrium had 10-mm long metal wires. The lead was introduced through a catheter and when the lead was in place, it was pushed out allowing the wires to spring apart. Other types had nylon barbs that were retracted by pulling on a filament in the lumen of the stylet and were extended by pushing on the stylet. These designs lost favor because of implant difficulties, design issues, and safety concerns. Endocardial exposed fixed screw-in leads were more successful. These leads use a helical corkscrew for both fixation and as the active electrode. Once the lead is in place, the helix is extended by rotating the terminal pin. Complications associated with insertion of this type of lead ranged from 2% to 5% which was considered a significant improvement. These kinds of leads can be fixed almost anywhere in the atrium and are used for multisite pacing. [5].

VIII.B. Leads Today

The leads used today are transvenous, directed through a vein from the implant area to the inside of the heart, entering the chambers and attaching to the endocardium. Epicardial or epicardial/myocardial leads are usually used only for pediatric purposes where no suitable transvenous path is available. [6].

Pacemaker leads usually have one of two mechanisms for holding their electrodes in place against a heart muscle. One type of lead has a screw-in tip that simply screws a short distance into the muscle. The other type of lead has small, flexible, rubber tines that press outward against the sides of muscle crevices in the heart’s interior. Either type can have a steroid-eluting tip that administers a small, controlled dose of a steroid to minimize inflammation and scarring of the heart at the attachment point. The steroid dose, typically 1 mg, is administered over a period of several weeks. The steroid can help extend the pacemaker battery life, because less scar tissue means lower resistance and thus less current required for the shock that stimulates a heartbeat. [6].

Medtronic, Inc., one of the main manufacturers of implantable medical devices such as pacemakers and defibrillators, makes the Sprint Fidelis™ family of defibrillation leads. According to the Medtronic web site, at 6.6 French in size, Sprint Fidelis leads are the world’s smallest right ventricular defibrillation leads, allowing for compatibility with 7 French introducers. An introducer is the instrument by which a physician inserts a lead

57 into the chamber of the heart where therapy is typically delivered. The small size of the Sprint Fidelis helps improve passage into a patient’s venous system for an easier implant, and minimizes venous obstruction. ICD leads tend to be heavy and more difficult to implant than pacemaker leads, but the small size of Sprint Fidelis makes it more similar to a pacemaker lead. The Sprint Fidelis family comes in four models and will complement the Medtronic Sprint Quattro™ lead family, which are the industry’s most implanted defibrillation leads. Sprint Fidelis leads will be available in both quadripolar and tripolar configurations, with active and passive fixation options. [7].

Medtronic, Inc. also marketed the first bipolar, steroid-eluting epicardial lead on the market, the Medtronic CapSureTM Epi Model 4968. Epicardial leads are generally used for infants and small children. This bipolar lead provides the benefit of steroid for reduced pacing thresholds. This lead was approved in 1999 for use in children in the USA. This Medtronic lead is designed to be attached to the exterior heart wall in a gentler way than earlier leads, which are attached to the heart with a probe. The Medtronic Model 4968 lead can be fixed in place with sutures and incorporates a steroid- eluting electrode that minimizes inflammation at its contact point. [7].

Saint Jude Medical is another manufacturer of heart valves, pacemakers and ICD devices. QuickSite® LV Lead and Aescula™ LV Left-Heart Lead are manufactured by Saint Jude. The QuickSite model 1056K LV lead is a left ventricular lead, including such features as S-shaped fixation and over-the-wire or stylet-driven compatibility. The steroid-eluting lead is implantable with a 7 Fr introducer. The Aescula LV Left-Heart Lead is a stylet-driven, silicone insulated lead with a titanium nitride (TiN) coated platinum-iridium electrode. The lead’s preshaped S curve distal end tip is designed to allow optimal steerability through the left-ventricular vasculature and is also a means for passive fixation. When a stylet is inserted, the lead straightens to enhance maneuverability. When the stylet is withdrawn slightly, the resultant curve is designed to help the lead pass through acute bends in the vasculature. Complete withdrawal of the stylet allows the preshaped S curve to press against the walls of the vein and provide a means of fixation. The Aescula LV lead also uses a step-ladder tip, featuring ridges designed to prevent microdislodgment of the lead, while at the same time optimizing capture and sensing thresholds. [8].

The Aescula LV lead is an unipolar ventricular lead 75 cm long and 4.7 French in diameter. It is designed for use in combination with a compatible pulse generator to provide permanent pacing and sensing of the left ventricle via the great cardiac vein. The use of unipolar left ventricular pacing is contraindicated in patients who have an implanted ICD or are currently being considered for implantation of an ICD. [8].

Guidant makes the pacing leads FLEXTEND which are steroid-eluting extendable/retractable helix pace/sense leads. FINELINE II Sterox The FINELINE® II Sterox family of steroid-eluting pacing leads is designed for permanent implantation, for either atrial or ventricular applications. The ENDOTAK RELIANCE® G/SG leads are made for ICD devices and are steroid-eluting, endocardial cardioversion/ defibrillation, and pace/sense leads that are available in

58 extendable/retractable and tined models. The electrode coils are covered with GORE™ expanded polytetrafluoroethylene (ePTFE) to prevent tissue ingrowth around and between the shocking coils. The ENDOTAK RELIANCE® and ENDOTAK RELIANCE S leads are steroid-eluting, endocardial cardioversion/defibrillation, and pace/sense leads available in extendable/retractable and tined models. [9].

The EASYTRAK® coronary venous leads are cardiac resynchronization therapy leads intended for use with Guidant's heart failure devices to provide cardiac resynchronization therapy. [9].

References 1. www.guidant.com. 2. http://www.web-ee.com/primers/files/pacemaker.pdf. 3. Andre E. Aubert and Hugo Ector, editors, Progress in Biomedical Engineering, 2: Pacemaker Leads, Proceedings of the International Symposium on Pacemaker Leads, Leuven Belgium, Elsevier, September 5-7, 1984. 4. B. N. Goldreyer, M. G. Wyman, D. S. Cannom, P. A. Kruth, Endocardial Pacing Leads With an Emphasis on Atrial Sensing, Elsevier, 1984. 5. Casten W. Israel, S. Serge Barold, editors, Advances in the Treatment of Atrial Tachyarrhythmias: Pacing, Cardioversion, and Defibrillation, chapter 14: Rick McVenes, S. Serge Barold, Kenneth B. Stokes, Evolving Lead Technology for the Treatment of Atrial Tachyarrhythmias, Futura Publishing, Armonk, NY, 2002. 6. http://www.web-ee.com/primers/files/pacemaker.pdf. 7. http://www.medtronic.com/newsroom/news_20040902a.html. 8. www.sjm.com. 9. www.guidant.com.

59 IX. PATENTS ON MEMS, ICD, AND SENSING

This section reviews the patents available on the United States Patent and Trademark Office (PTO) that may be related to sensing aspects of ICD or pacemaker devices and are implemented in MEMS. The search terms used were “MEMS and Heart.” The term “Heart” was anticipated to cause the search to identify both ICD and pacemaker devices as well as any of their components including the sensing leads.

In addition to the search terms, the pertinent PTO classifications can help identify the patents in an area of interest. The section begins with an explanation of the system of classification used by the PTO and ends with a summary of the patents or patent applications found on the PTO web site that disclose some aspect of sensing using MEMS devices.

IX.A. PTO Classification

PTO has a classification system dividing the entire set of U.S. patents into searchable collections based on the technology claimed in the patents. The primary groupings of patents in this classification system are called classes. Utility classes are based on (1) the technology associated with a particular industry or (2) the subject matter having similar function, use, or structure. Classes are subdivided into relatively small, ordered collections of patents called subclasses. A subclass is the smallest searchable collection of documents in the PTO database. The ordered listing of subclasses that make up a class is called a class schedule. [1].

In most modern class schedules, subclasses are generally arranged in order of diminishing complexity, with the most complex inventions positioned at the top of the class schedule. Combined machines or processes, those which perform diverse operations, will be found higher in the class schedule than single operation machines or processes, which in turn are located higher than the individual parts of the machines. Minor details or accessories are normally found near the bottom of the class schedule, and “miscellaneous” subject matter is placed at the very bottom of the schedule. However, “special” subclasses for inventions having a common unique feature may be positioned above subclasses for more complex inventions. [1].

Certain classes have a potential of being related to inventions regarding sensing of ventricular fibrillation. Inventions regarding treatment and care of the body are under classes 128 and 600-607: 128- Surgery, 600- Surgery, 601- Surgery: Kinesitherapy, 602- Surgery: Splint, Brace, or Bandage, 604- Surgery, 606- Surgery, and 607- Surgery: Light, Thermal, and Electrical Application.

Most of the inventions of interest to us are classified under 607 and then 600. Pacemakers and defibrillators all fall under the electrical application class 607. They may also fall under 73, which is the generic class for devices for making a measurement or a test, that takes all such subject matter not provided for in other classes or they my fall under 324 for measuring, testing (or sensing) of electric properties. Class 424 pertains to

60 drugs and chemicals and may contain coatings of medical devices. Classes in the 900’s pertain to some other aspects of the medical field such as substance abuse and cancer detection. Classes regarding electricity and electronics are in the 300’s and would not be the primary class when dealing with medical devices. [2].

Within these main classes, many different subclasses may pertain to implantable devices sensing and correcting cardiac problems.

IX.B. PTO Searches

Patent applications are published on the PTO site after 18 months from the filing date unless the applicant requests nonpublication and pays the corresponding fee. The published patent applications, when available, are a better representative of the current state of the technology in a given field than issued patents.

Published patents are older and their corresponding applications are usually filed at least 3 years prior to the issue and publication of the patent. On the other hand, almost all issued patents are published on the PTO website irrespective of the wishes of the patentee. Therefore, they represent a more complete list.

Several searches were conducted to find the relevant published patents and patent applications available on the PTO site. The abstracts of the applications were searched for the search terms. Search terms and the results follow.

IX.C. MEMS and Heart

A search was conducted for a combination of the two search words “MEMS” and “Heart” in both the applications database and the published patents database. In order to identify applications or the patents that have a MEMS focus, the term MEMS was searched in the abstract while the term Heart was searched anywhere in the application. The searches were conducted in September of 2004.

IX.C.1. Published Applications

Thirteen applications were listed that contain the terms MEMS in their abstract and the term heart somewhere in the application. These applications were classified under 600, 604, 607, 324, and 73:

1. Self-powered implantable element 2. Implantable microscale pressure sensor system for pressure monitoring and management 3. Systems and methods for the transport of fluids through a biological barrier and production techniques for such systems 4. Remote controlled transdermal medication delivery device 5. Ultra-miniature optical pressure sensing system 6. Flexible MEMS actuated controlled expansion stent

61 7. Pacing channel isolation in multi-site cardiac pacing systems 8. Devices including protein matrix materials and methods of making and using thereof 9. Wireless MEMS capacitive sensor for physiologic parameter measurement 10. Mems switching circuit and method for an implantable medical device 11. Implantable medical device incorporating integrated circuit notch filters 12. Apparatus and method for manufacturing an intracutaneous microneedle array 13. Apparatus and methods using mechanical resonators to enhance sensitivity in lorentz force magnetometers

All the applications that were produced by the search were reviewed and are briefly summarized below. The subject matter of some of the applications was irrelevant as they mentioned the term heart incidentally. Others were more relevant. Few, numbers 5, 9, 11, and 13 above, were completely on point. The irrelevant applications are briefly discussed before being set aside. The most relevant applications, that are revealing regarding the use of MEMS devices for sensing heart related parameters, are discussed in more detail below.

1. In the “Self-powered implantable element” a MEMS accelerometer is used next to a cardiac tissue and generates energy based on the natural bobbing, jiggling, rocking, or twisting motion of the beating heart. The MEMS accelerometer then converts a portion of this motion into electrical energy, which is stored in a capacitor and used to power another element.

2. The “Implantable microscale pressure sensor system for pressure monitoring and management” describes a MEMS pressure sensor based upon detection of an induced inductance. This pressure sensor is implanted in the eye and will provide wireless measurements of fluid pressures in the eyeball.

3. The “Systems and methods for the transport of fluids through a biological barrier and production techniques for such systems” describes a group of microneedles that can be combined with a MEMS pumping system.

4. The “Remote controlled transdermal medication delivery device” is about micro tubes and MEMS valves that control medication delivery.

5. The “Ultra-miniature optical pressure sensing system” describes a MEMS fiber optic pressure sensor embedded in an angioplasty guidewire. This application claims to improve on a pressure sensor introduced in the US in 1999 by RADI of Sweden called PressureWire.TM. This application has the number 20030159518 and was supported by NIH funding.

6. The “Flexible MEMS actuated controlled expansion stent” uses MEMS motors to gradually expand stents that are used inside veins. The MEMS motor described includes moving parts with a pinion gear mounted on a shaft.

62 7. In “Pacing channel isolation in multi-site cardiac pacing systems” a multi-site or multi- chamber pacing is disclosed that uses miniaturized electrical isolation circuitry to minimize the effects of leakage currents generated during delivery of a pacing pulse from affecting sense amplifiers in the other pacing channels. Isolation of the pace/sense electrodes from leakage currents is achieved by using monolithic isolation circuits. The monolithic isolation circuits may be formed using a MEMS isolation transformer comprising low-loss input and output coils separated by an insulation layer.

8. “Devices including protein matrix materials and methods of making and using thereof” is primarily chemically oriented. It discusses a compound that may be used for coating MEMS devices that are implanted or implantable in the heart.

9. The “Wireless MEMS capacitive sensor for physiologic parameter measurement” is an application filed by Dr. Najafi of Ann Arbor Michigan. This application discloses a wireless MEMS capacitive sensor for implantation into the body of a patient to measure blood pressure, blood flow, intracranial pressure, intraocular pressure, and glucose levels. The application number is 20020151816.

10. The “Mems switching circuit and method for an implantable medical device” uses MEMS switches in place of transistor switches formerly used on implantable devices.

11. In “Implantable medical device incorporating integrated circuit notch filters”, the disclosed device improves the ability to sense low level cardiac signals in the presence of high level electromagnetic interference. This application discusses sense amplifiers for sensing physiologic signals and parameters, RF telemetry capabilities for uplink transmitting patient data and downlink receiving programming, and other electronic circuitry susceptible to electromagnetic interference. It focuses on miniaturized, integrated circuit notch filters for use in such circuitry. This application is by Medtronic and the application number is 20020026224.

12. In “Apparatus and method for manufacturing an intracutaneous microneedle array” a microneedle array is constructed of silicon and silicon dioxide compounds using MEMS technology and standard microfabrication techniques to create hollow cylindrical individual microneedles.

13. In “Apparatus and methods using mechanical resonators to enhance sensitivity in lorentz force magnetometers” a magnetometer is disclosed that can be manufactured in arrays with the resonant frequency for each magnetometer being rapidly, sequentially, and dynamically varied through the use of piezo/MEMS elements. Catheter magnetic fields sensors are used in heart catheterization for cardiac ablation by monitoring the local current distributions on the heart at the positions of the ablating catheter. The application was filed by The Johns Hopkins University and the application number is 20010035750.

IX.C.2. Issued Patents

63 The same search in the database of issued patents yielded the following 11 patents:

1. 6,742,873 Inkjet printhead construction 2. 6,723,086 Remote controlled transdermal medication delivery device 3. 6,711,437 Pacing channel isolation in multi-site cardiac pacing systems 4. 6,668,109 Method and system for ultra-fast switching of optical signals 5. 6,667,725 Radio frequency telemetry system for sensors and actuators 6. 6,639,313 Hermetic seals for large optical packages and the like 7. 6,558,361 Systems and methods for the transport of fluids through a biological barrier and production techniques for such systems 8. 6,539,253 Implantable medical device incorporating integrated circuit notch filters 9. 6,379,324 Intracutaneous microneedle array apparatus 10. 6,312,612 Apparatus and method for manufacturing an intracutaneous microneedle array 11. 6,173,650 MEMS emergetic actuator with integrated safety and arming system for a slapper/EFI detonator

The only relevant patent, aside from the ones related to the applications already discussed, is Patent number 6,667,725, “Radio frequency telemetry system for sensors and actuators” relates to combining Radio Frequency (RF) technology with micro- inductor antennas and signal processing circuits for RF telemetry of real time, measured data, from MEMS sensors, through electromagnetic coupling with a remote powering/receiving device for use in the biomedical area. The microminiaturized inductor/antenna system is used for contact-less powering of an oscillator circuit providing an RF telemetry signal from bio-MEMS systems.

IX.D. MEMS and Heart Related Applications and Patents in Detail

The more relevant applications and patents are discussed in length in this section.

IX.D.1 Sawatari: Ultra-Miniature Optical Pressure Sensing System

This application proposes a device for measuring blood pressure during angiography and angiplasty. It mentions the current technology of using Doppler guidewires for measuring velocity and thus flow rate and asserts that measuring the pressure directly is preferable. The device of the application purports to improve a pressure sensor by RADI of Sweden called PressureWire™. The RADI device has a 360 micron outer diameter and measures pressure directly inside a blood vessel. The application dismisses the RADI device as an electronic sensor with drift problems and with a narrow and high impedance cable that must be adequately shielded to reduce RF interference making it hard to achieve the desired flexibility.

Instead, the application proposes a fiber-optic sensor similar to the one proposed in US Patent 5,987,995. The device of this patent was proposed for continuous monitoring of both the diastolic and systolic blood pressure inside the patient's heart chambers to

64 monitor for possible clogging of the arteries for periods typically as long as 3 days. This patent discloses a fiber optic pressure catheter that includes a light source, an optical fiber receiving light from the light source and a sensor head that optically impacts the optical fiber. The optical fiber is connected to the light source at one end and to the pressure sensor at the other end. The pressure sensor has a membrane that deflects in response to the pressure outside the sensor. Up to this point the device of the issued patent and the application are the same. The issued patent discloses a resilient ribbon connected to the inside of the sensor chamber at one end. The other end of the ribbon moves in front of the optical fiber end that is in the sensor chamber. The application uses a pedestal looking part that protrudes from the membrane inward and can move with the deflections of the membrane. The ribbon and the pedestal both serve the same function: they block the end of the fiber optic fiber in proportion to the deflection of the membrane which is in turn proportional to the pressure outside. Thus, various amounts of light are reflected back into the optical fiber based on the amount of pressure at the membrane. A detection system at the other end determines the pressure at the sensor head from the contrast of spectral fringes of light created by an optical coating on the end of the optical fiber and the light reflecting from the ribbon.

The device disclosed in the patent is small but it is not a MEMS device. It mentions electronic pressure catheters such as “Millar's 9 French” which is 3 mm in diameter. However, the specification fails to mention the dimension of the disclosed device. The application mentions the RADI PressureWire with an outer diameter of 360 microns. It is not clear from the application but seems that the proposed device of the application has parts in the pressure chamber that are 1000 microns (1 mm) wide and 3000 microns (3 mm) long. The outer diameter of the guidewire seems to be 300 microns but that may not include the actual pressure sensor end.

More importantly, this application and the associated patent disclose pressure sensor catheters that cannot be left inside permanently because the wire has to be connected to an outside device.

X.D.2 Najafi: Wireless MEMS Capacitive Sensor for Physiologic Parameter Measurement

This application discusses an inductor-capacitor, or LC, tank resonator where the connection of the capacitor and inductor has a specific resonant frequency, 1/√LC, which can be detected from the impedance of the circuit. If the value of the capacitance varies with pressure, while the inductance stays constant, then the new value of the capacitance may be calculated from the new value of the resonance frequency and the sensed pressure may be calculated from the new capacitance value if the relationship between the pressure and the capacitance is known.

The impedance of the LC tank resonator is measured remotely from the impedance of a readout coil that is magnetically coupled to the internal coil. The readout coil may also be used to excite the resonator of the implanted device and sense the reflected impedance. This mechanism helps the implanted device conserve energy. The application claims that

65 prior art has been limited to implant-readout separation of 1-2 cm rendering deep implants impractical. Heart ventricle pressure monitoring would require a separation distance of 5-10 cm.

This application discloses a two-part sensor where the implantable part is a monolithic MEMS capacitive sensor integrated on the same chip with an inductor. The power source is located on the external read-out part. The implantable part is discussed in detail and is the heart of the alleged invention. The read-out is supposed to be conventional.

The authors suggest implanting the device near a heart valve to monitor the pressure on either side of the valve or inside a vascular stent or even outside the wall of a blood vessel. The device is also suggested for monitoring pressure in the left ventricle, left atrium, and the pulmonary artery.

This application refers to a number of patents and publications that mention methods of monitoring pressure or flow in the heart. Some of these references or references mentioned in turn within one of the references are elaborated below.

6,328,699- “Permanently implantable system and method for detecting, diagnosing and treating congestive heart failure.” This patent discloses a pressure transducer permanently implantable within the left atrium of the patient's heart for sensing fluid pressures within the patient's left atrium.

5,368,040- “Apparatus and method for determining a plurality of hemodynamic variables from a single chronically implanted absolute pressure sensor.” This patent discusses a device for monitoring and measuring the hemodynamic status of a patient's pulmonary pressure and right atrial pressure. The device uses an implanted absolute pressure sensor located in the right ventricle, coupled to an implantable monitoring device, which records pressure values in response to a combination of sensed electrical depolarizations of the atrium and ventricle. This is a class 600 Medtronic patent that only cites to a number of other US patents as references.

6,053,873- “Pressure-sensing stent.” This patent discloses an implantable stent measuring a parameter relating to a rate of blood flow through the stent. A transmitter transmits signals responsive to the measured parameter to a receiver outside the body. This is a class 600 patent.

6,201,980- “Implantable medical sensor system.” This patent describes a chemical sensor system made of an expandable polymer incorporated into an electronic circuit that changes its frequency when the polymer changes dimension. This invention was proposed for monitoring of blood glucose levels in diabetic patients. It is a class 600 invention assigned to UC Berkeley.

The references cited by the Najafi patent include Harpster et al., Proc. 14.sup.th IEEE Int'l. Conf. Microelectromech. Sys. (2001), pp. 553-557). The Harpster article presents a

66 single-chip integrated humidity sensor capable of wireless operation through inductive coupling with a remote transmitter.

X.D.3 Thompson: Implantable Medical Device Incorporating Integrated Circuit Notch Filters

An electronic filter eliminates unwanted frequencies from an electronic signal. A low- pass filter passes low frequencies. A high-pass filter passes high frequencies. A band- pass filter passes a limited range of frequencies. A band-stop filter passes all frequencies except a limited range. Band-stop and band-pass filters can be constructed by combining low-pass and high-pass filters. A notch filter is a type of band-stop filter that acts on a particularly narrow range of frequencies. A notch filter is typically used when the high frequency and the low frequency are less than 1 to 2 decades apart (that is, the high frequency is less than 10 to 20 times the low frequency). [4].

The sensing of these physiologic parameters involves the detection of minute electrical signals in an inherently electrically noisy environment that may include both ambient electromagnetic interference (EMI) in the patient's environment as well as electrical signals either generated in the body or in the sensor due to patient motion or respiration or the like. Sources of EMI include metal detectors such as those used in airports, welders, radio transmitters (broadcast and two-way), cellular phones, microwave ovens, electronic article surveillance systems, and the like. To avoid sensing EMI, it has been the practice to low pass filter the sense amplifier input and to adjust the sensitivity of the sense amplifier to a level that renders it insensitive to low level EMI but capable of sensing the peak voltages of the signal of interest. Descriptions of various sensing windows and noise rejection techniques for implantable and external pacemakers that have been employed or proposed are set forth in U.S. Pat. Nos. 4,357,943, 4,390,020, 4,401,119, 4,436,093, 4,596,292, and 5,188,117.

The expansion of wireless communication modes occupying more and more frequency bands will continue to cause new EMI caused oversensing, undersensing, and inhibition problems to develop. Potential EMI sources transmit in the 820-896 MHz, 1.93-1.99 GHz, and 2.402-2.48 GHz bands. Therefore, there is a need to improve the ability to sense low level cardiac signals or other physiologic signals in the presence of high level EMI.

Both non-physiologic and physiologic data can be transmitted by uplink RF telemetry from the implantable device to the external programmer or through the patient's body to another device. The RF telemetry transmission system that evolved into current common usage relies upon magnetic field coupling through the patient's skin of an antenna with a closely spaced programmer antenna. Low amplitude magnetic fields are generated by current oscillating in an LC circuit of an RF telemetry antenna of the device or programmer in a transmitting mode. The currents induced in the closely spaced RF telemetry antenna of the programmer or device are detected and decoded in a receiving mode. In the MEDTRONIC.RTM. product line, the RF carrier frequency is set at 175 kHz.

67

There are a number of limitations in the current MEDTRONIC.RTM. telemetry system employing the 175 kHz carrier frequency. Off-chip components are required to interface with integrated components at the board level, which constitutes an important bottleneck to miniaturization and the performance of heterodyning transceivers. Recent attempts to achieve single chip transceivers have used direct conversion architectures, rather than heterodyning and have suffered in overall performance. The continued growth of micromachining technologies, which yield high-Q on-chip vibrating mechanical resonators now make miniaturized, single-chip heterodyning transceivers possible. MEMS resonators yield ultra high Qs of over 80,000 under vacuum and center frequency temperature coefficients less than 10 ppm/°C and serve well as a substitute for crystal filters in a variety of high-Q oscillator and filtering applications. MEMS resonators are capable of frequency operation to GHz levels and filtering operation up to the 6th order.

This application discloses a notch filter that is developed using MEMS technology and is therefore small and implantable.

X.D.4. Murphy: Apparatus and Methods Using Mechanical Resonators to Enhance Sensitivity in Lorentz Force Magnetometers

Lorentz Force magnetometers are magnetic field sensors that use mechanical resonators to enhance sensitivity. Toward the goals of smaller size, lower power consumption, and lower cost for similar performance, recent developments have included the use of piezoresistive cantilevers and µ-magnetometers. These devices require extensive and intricate processing and their sensitivities, defined as the minimum detectable field change, are generally in the range of 1 mT to 1 µT. Other devices such as the pendulum clock or quartz crystal resonator controlled watches use mechanical resonators to enhance detection. In these devices, the accuracy is directly linked to the quality factor or "Q" of the resonator.

This application is based on another type of mechanical resonator magnetometer that works on excitation of a resonant bar configured in a xylophone geometry (U.S. Pat. No. 5,959,452). Only, the application uses strings instead of a bar. The xylophone magnetometer using a bar measures the vector component of the magnetic field which lies in the plane of the xylophone and is perpendicular to its major axis. But, it cannot measure the magnetic field in any other direction. The response of the xylophone magnetometer is linear in low frequency magnetic fields over 7 decades of range and has a noise floor below 1 nanoT.

The magnetometer proposed in this application can measure multiple vector magnetic fields and its resonant frequency can be varied. It can also be manufactured particularly in arrays for use in catheters. This mechanical resonator magnetometer is made of an electrically conducting string or a light-conducting fiber coated with an electrically conducting material that is pulled in tension between two anchors. When a current flows in the string and the magnetometer is placed in a magnetic field, the resulting Lorentz Force will cause the string to deflect along an axis perpendicular to the plane formed by

68 the current vector and the magnetic field vector. The Lorentz Force can be detected and is an indicator of the magnetic field strength. Tension of the string, and thus its resonance frequency, can be varied using piezo or MEMS elements. The deflection may be detected by conducting light through the string, provided it is made of a light conducting fiber, and detecting the deflection of the light.

The application proposes the magnetometer to be incorporated into catheters for cardiac catheterization. The device can be used as a single sensor, as a two magnetic field axis device, or as a catheter-based linear multisensor array. In cardiac ablation, aberrant current pathways on the heart are removed by electrical destruction of selected regions of the heart tissue. The precise location of the region to be destroyed must be determined and then the degree of success of the destruction must be assessed. Current practice successively induces heart fibrillation followed by ablation. The sensor of this application monitors the local current distributions on the heart at the positions of the ablating catheter and locates the aberrant current paths to guide the ablation. Presumably, though not stated expressly in the application, this sensor saves the step of inducing fibrillation.

The application suggests approaching the heart with the catheter axis perpendicular to the heart surface. Then, the magnetic field directions produced by the magnetometer current and by the heart currents both lie in the plane of the heart. The direction and magnitude of local surface currents on the heart, including re-entrant loops and other anomalous current paths which produce arrhythmias, can be detected.

The application claims that if a linear array consisting of several magnetometers are used along a single axis inside a single catheter, second and higher order magnetic field gradients can be determined. Arrays of magnetometers lying on the chest can be used to monitor vector currents on the heart. Biological currents in the brain produce very small magnetic fields that are significantly smaller than in the case of the heart. Position and depth of these currents can be measured too. For the brain, a flexible net of sensors can be made that like a "helmet" of magnetometers can conform closely to the skull. A two- dimensional array placed on the surface of the chest or elsewhere on the body could be used to analyze the heart and other organs as well; a multilayer array or net would permit vertical gradient detection.

This application, while mentioning MEMS technology as a means of fabrication of the magnetometer, provides no measure of the size of the device.

X.D.5. Simons: Radio Frequency Telemetry System for Sensors and Actuators

This document is an issued patent with patent number 6,667,725 that discloses a microminiaturized inductor/antenna system for contactless powering of an oscillator circuit providing an RF telemetry signal from bio-MEMS systems, sensors, or actuators. A miniaturized circuit inductor coil is printed on a dielectric substrate. The inductor coil behaves both as an inductor, which acts to charge a capacitive device as well as an antenna for transmitting a RF signal indicative of the level of charge of the capacitive

69 device. The circuit operates in two modes. In the first mode, the inductance coil forms a series resonant circuit with the capacitance of a capacitive MEMS device such as a pressure-sensing diaphragm of a MEMS pressure sensor device. In the second mode, the capacitive device produces an oscillating electrical current flow through a planar printed inductor coil. The inductor coil is equivalent to a helical antenna and hence loses power through RF radiation from the inductor. A remote RF receiving device may be used to receive the RF radiation, from the inductor coil, as a RF telemetry signal. The functional operation begins when an electromagnetic coupling energizes the circuit with a remote- transmitting device followed by oscillation of the circuit. Thus there is no direct or hard connection to the circuit by any power source.

Fabrication of the test chips comprised coating a high resistivity silicon wafer with a thin insulating layer of SOG (likely meant to refer to silicon on germanium) to isolate the printed circuit from substrate losses. Typically the thickness of the insulating SOG layer was about 1 to 2 microns. Following application of the SOG layer, the wafer was patterned using photo resist and the inductor coils were fabricated using standard "lift-off techniques. Inductor thickness was in the range of 1.5 to 2.25 microns to minimize resistive losses in the circuit.

MEMS pressure sensors typically measure as little as 0.350 mm in width making them small enough for use in many in-vivo medical applications. For example, with one implanted MEMS pressure sensor it is possible to measure the internal pressure of body organs or wounds. With two MEMS pressure sensors it is possible to measure the pressure drop across an obstruction in an artery or newly implanted heart valve. With three MEMS sensors it is possible to characterize the flow across a long section of arteries, along the esophagus or through the small intestines.

The point different about the device of this patent is that it represents a "contactless" MEMS pressure measuring system, requiring no directly connected power source such as a battery etc. The circuit is energized by a remotely generated magnetic field from an electromagnetic source acting through the inductor thereby charging the capacitive sensor to an electrical energy state commensurate with the real time pressure being measured by sensor.

IX.E. Earlier Referenced Patents (Not MEMS)

Some of the earlier technologies found in the patents and publications that are referenced, pursuant to the Information Disclosure Statement (IDS) requirements, are reviewed in this section. The reference patents are related to the patents reviewed above and represent the state of the art that the new application tries to improve upon. Most of the patents reviewed here are drawn to various types of pressure sensors. Blood flow and pressure in the vessels and chambers of the heart may be used as an indicator of whether the electrical activity is sufficiently healthy to pump the blood at the requisite pressure. Blood pressure and pressure in a heart chamber are, however, secondary and averaged indicators and do not provide the resolution required for understanding ventricular fibrillation.

70

3,568,661 by Franklin is a 1971 patent disclosing a device that measures blood stream velocity using frequency modulated ultrasound signals. The device works around a surgically exposed blood vessel. Two piezoelectric crystal mounted diagonally on the blood vessel receive the ultrasound signal. One crystal receives the signal downstream and the other the signal that has traveled upstream in the blood vessel. The difference in the frequency of the two signals detected is an indicator of the blood flow velocity between the two.

This device was not implemented in MEMS and was disclosed as applied to an exposed blood vessel. However, with a reduction in size, the same principle may be applied to an implantable device.

3,757,770 by Brayshaw is a 1973 patent disclosing an implantable pressure sensor that radios its readings by an RF transmitter. The pressure sensor is implanted inside the skull and measures the cerebrospinal fluid pressure. The pressure sensor in this device is a bipolar junction transistor whose collector-to-emitter resistance is sensitive to pressure.

4,127,110 by Bullara is a 1978 patent disclosing an implantable pressure sensor for measuring pressure in a brain ventricle. The pressure sensor is a pressure transducer formed of an LC circuit. The change in pressure induces a change in capacitance. The transducer receives its operating energy from an external oscillator source.

4,227,407 by Drost is a 1980 patent disclosing a flow meter for measuring flow volume using interferometery. The device uses transducers for producing and receiving waves that are modified by the flow through a conduit such as a blood vessel. This device does not place its probes directly on the walls of the blood vessel and is thus nonconstrictive.

4,519,401 by Ko is a 1985 patent disclosing a pressure sensor for use inside the brain. The pressure transducer is piezoresistive.

4,541,431 by Ibrahim is a 1985 patent disclosing a telemetry system combined with a magnetic field sensor for use in implantable medical devices such as pacemakers. The sensor is meant to be used as a replacement for a conventional reed switch that is activated in response to a magnetic field.

4,593,703 by Cosman is a 1986 patent disclosing an implantable differential pressure sensor that reports the difference of two bodily pressures on its opposite sides via telemetry. Telemetry is based on the familiar mechanism of a change in the resonant frequency of an LC circuit in response to a change in pressure. This particular device was proposed for measuring intracranial pressure. The normal pressure inside the skull is about 15 mm/Hg or 150 mm - 200 mm of water.

4,600,855 by Strachan is a 1986 patent disclosing a piezoelectric pressure sensor that works within a conduit and can report the change in pressure telemetrically. Fluid passes

71 through a ring containing piezoelectric components and causes a change in the resonance frequency of these components.

4,781,715 by Wurzel is a 1988 patent disclosing a pressure sensor used in a cardiac prosthesis. This patent is not drawn to an implantable pressure sensor although the prosthesis as a whole is implantable.

4,846,191 by Brockway is a 1989 patent disclosing a pressure sensing catheter within a blood vessel. This catheter is of course not implantable. It works using low viscosity fluids and gels transferring the pressure from inside the body to a solid-state transducer.

4,988,752 by Cohen is a 1990 patent disclosing an implantable pacemaker device. The pacemaker includes a pressure sensor that may be implanted in various parts of the heart and based on the deviations of the pressure from a required pressure, triggers the pacemaker into creating a pulse. For example, a pressure sensor in the right atrium creates a signal corresponding to the right atrial pressure which is in turn compared with a fixed signal corresponding to a command right atrial pressure. The difference between the two creates an error signal that is used to control the impulse frequency of the pacemaker. Instead of the right atrial pressure, the right ventricular systolic pressure may be used.

5,178,153 by Einzig is a 1993 patent disclosing a fiber optics fluid flow meter for measuring blood flow in arteries. Two fiber optic pressure sensors are placed in the artery along the direction of flow. The differential pressure flow meter works based on the Bernoulli principle that the sum of pressure head, velocity head, and gravitational head must stay constant along the path of the flow. Differential pressure is created as a result of natural taper in the artery or as a result of a constriction. Differential pressure yield the flow velocity based on the cross-sectional area of the artery. The cross- sectional area is found from a third pressure measurement or by other methods such as thermal, dye dilution, ultrasonics, x-rays and etc.

References 1. http://www.uspto.gov/web/offices/opc/documents/overview_dec02.pdf. 2. http://www.uspto.gov/web/offices/opc/documents/combined_cabrsm_dec_02.pdf. 3. http://en.wikipedia.org/wiki/Fabry-Perot. 4. http://www.fact-index.com/n/no/notch_filter.html. 5. http://hyperphysics.phy-astr.gsu.edu/hbase/magnetic/magfor.html. 6. http://encyclopedia.thefreedictionary.com/Quality+factor.

72 X. FURTURE RESEARCH ISSUES

This section recommends one of the methods described in the previous section for mapping a VF event and lists the issues requiring further research regarding the chosen method. The issues presented start from the current state of the art and whether the state of the art is sufficient to provide the desired spatial resolution. Other issues requiring further research include factors impacting the spatial resolution of the mapping system, whether it is feasible to improve the resolution, and what improvements to the state of the art are required to achieve the desired resolution.

X.A. Magnetic Mapping as the Recommended Method

After reviewing internal, implantable, and external methods of sensing the electrical activity of the heart, this paper proposes further research regarding the use of an external web of miniature magnetometers as the method of choice for mapping a VF event.

Sensing mechanisms used in ICD devices are content with sensing a fibrillation event on the average and do not strive to develop a map of the electrical activity of the entire heart. While one magnetometer is used for sensing an event on a localized basis, a network of magnetometers may be used to develop a map of the overall activity.

Use of an external device also circumvents the issues of readout through tissue, including antenna size and RF design overcoming attenuation in tissue. Electromagnetic noise from the medium or from the electrical activity of skeletal muscles must be taken into account for an external device which will sense all magnetic fields irrespective of their source.

Filtering of the signal, similar to that done in the sensing circuits of pacemakers and ICD devices can be used to suppress the outlying frequencies corresponding to muscles or ambient magnetic activity. During normal heart activity, only the high frequencies of skeletal muscles and electromagnetic interference must be filtered out. In addition to the mid-band frequencies corresponding to myocardial activity, low frequencies corresponding to T-wave repolarization may provide useful information regarding the heart activity. So, a bandpass filter, similar to that used in an ICD or a pacemaker, can be used to isolate the myocardial waves while a second low pass filter is used to capture the T wave repolarization effect. During a fibrillation event, there is no regular QRS complex and all activity is likely to occur at relatively high frequencies. Alternatively, a series of bandpass filters can be used corresponding to various frequencies to get a break down of the frequencies of fibrillation.

A Lorentz force magnetometer was disclosed in Murphy’s patent application that has since been issued as U.S. patent number 6,812,696 on November 2, 2004. Murphy proposes a Lorentz force magnetometer based on a resonator string. The patent states that this magnetometer may be manufactured using standard MEMS fabrication techniques.

73 Murphy draws from small catheter magnetic field sensors that are used to locate the sites suitable for cardiac ablation and proposes the use of a network or web of such miniaturized magnetometers to map the magnetic fields resulting from the electrical activity of the heart from outside the body. Such a web could also be used around a heart that has been taken out of the body, for example the heart of a test dog, and is being studied for fibrillation mechanisms. When a web of sensors is used all around an in-vitro heart or all around the body, a back-projection algorithm similar to that used in computed tomography (CT) can be used to develop a three-dimensional map of the magnetic field activity and thus the electrical activity of the heart. According to Murphy, layers of these magnetometers placed at different distances from the heart can yield the gradient of the magnetic field generated by the electrical activity of the heart.

X.A.1. String Magnetometers

This subsection briefly presents string magnetometers.

An earlier patent by Givens, patent number 5,959,452, listing Murphy as the second inventor, discloses a Lorentz force magnetometer that can measure magnetic fields as low as 1 nano-Tesla (nT) with a dynamic range of 1 mili-Tesla (mT) to 1 T. The magnetometer includes a resonator. The resonator can be a bar supported by two wires placed at nodal points of the fundamental resonance frequency. These supporting wires also supply an alternating current of resonance frequency to the resonator bar.

If the resonator enters a magnetic field, the combination of the magnetic moment and the alternating current generate a Lorentz force on the bar that keeps changing direction with the changes in the direction of the current. The changing Lorentz force causes the bar to vibrate. The bar is fixed at two ends to the current-carrying wires. Therefore, it may virbrate only in one direction that is perpendicular to the plane formed by the two fixing wires. Then the amplitude of the vibration is proportional to the component of the magnetic field located in the plane of the bar and perpendicular to the bar. The amplitude of the vibration or the deflection of the beam is detected by an optical beam deflection method and related to the magnitude of the magnetic field. The device is linear: an increase in the deflection of the beam is linearly proportional to the increase in the magnetic field. Changing the current, changes the sensitivity of the bar from Teslas to nanoTeslas.

The Murphy magnetometer uses a wire or a string as opposed to the bar of Givens. A string has many more degrees of freedom and can vibrate in any direction in the plane perpendicular to the string. The bar could only vibrate along a line perpendicular to the plane formed by the surface of the bar. Therefore, this magnetometer can measure magnetic filed components in any direction in the plane perpendicular to the string. Varying the length of the string changes the resonance frequency of the string and allows it to detect various magnitudes of magnetic fields.

X.B. Effectiveness of the Magnetic Mapping

74 The effectiveness will depend on the how small of a distance can be resolved by the system or, in other words, how high of a spatial resolution can be achieved. Fibrous heart muscle cells are long and narrow and can be 90 microns long. A spatial resolution of the order of 500 microns or 0.5 millimeter was considered sufficient. At a certain limit, spatial resolution is no longer limited by the size of the magnetometer, rather by the distance of the magnetometer from the current source. Therefore, effectiveness also depends on sensitivity of the magnetometer indicating how low a current can be measured how far from the magnetometer.

While a magnetometer can be used to map a VF event, there are theoretical limits which limit such use and must be explored in future research.

X.C. State of Art in Magnetic Mapping [1]

In addition to the string magnetometers explained above, many types of devices have been used to measure a magnetic field. Other magnetometers include: other types of induction sensors, fluxgates, magnetoresistors, Hall-Effect magnetic sensors, magneto- optical sensors, resonance magnetometers, and Superconducting Quantum Interference Detectors (SQUID). Some of these devices are briefly explained below.

X.C.1. SQUID

According to Murphy, SQUID has been used to measure the magnitude and direction of biological currents in the heart from outside the body using the magnetic fields that heart currents produce. These magnetometers, however, operate at liquid helium temperatures requiring complex facilities and significant power for cooling. Murphy suggests that arrays of the string magnetometers proposed by his invention have the potential of being able to monitor vector currents on the heart at comparable resolutions but at significantly reduced costs.

X.C.2. CARTO Magnetic Mapping Procedure

A 2004 healthcare coverage position document by the insurance company Cigna approves a magnetic mapping method used in cardiac ablation for coverage. The position paper sets forth the cases where Cigna covers a procedure called CARTO for electrophysiological mapping. Cigna covers this procedure when it is used to guide radiofrequency ablation for treatment of supraventricular arrhythmias. Cigna does not cover the use of this mapping system for diagnosis, treatment, or management of ventricular arrhythmias. In supraventricular arrhythmias, the irregular heart rhythm originates in the atrium and in ventricular arrhythmias the irregular rhythm originates in the ventricles. The reason stated for coverage of supraventricular arrhythmias is that a significant number of studies have shown its effectiveness in a large percentage of cases. The reason stated for denying coverage for ventricular arrhythmias is not that the procedure is ineffective for these cases. Rather, Cigna concludes that the number of case studies for ventricular cases is insufficient to make definitive conclusions regarding the use of CARTO in ventricular arrhythmias. A 2003 Blue Cross coverage position paper

75 approves magnetic mapping for all supraventricular and some specific types of ventricular arrhythmias.

Cigna defines CARTO as an electroanatomical cardiac mapping system that provides simultaneous electrophysiological and spatial information and displays a three- dimensional reconstruction of the cardiac chambers. According to Cigna, this information can be utilized to detect the originating site of certain arrhythmias and guide radiofrequency ablation of regions of myocardium that trigger or propagate the abnormal cardiac rhythm. CARTO is short for CARTO EP Navigation System and is manufactured by Biosense Webster, Inc. in Diamond Bar, California. Biosense is a division of Johnson and Johnson. This device is approved by the FDA for catheter-based cardiac mapping. CARTO is comprised of miniature passive magnetic field sensors, an external ultra-low magnetic field emitter or location pad, and a CARTO processing unit. The description of this device and the pictures shown on the Biosense website all indicate a catheter as opposed to a network or a web.

X.C.3. Miniature Magnetometers

Miniature magnetometers are used for mapping magnetic fields found in various applications. The trend has been toward smaller size, lower power consumption, and lower cost for similar performances. Piezoresistive cantilevers and µ-magnetometers were developed towards this end.

X.D. Is the State of Art Sufficient?

To achieve a three-dimensional image of the electrical activity at any point in time, a catheter magnetometer is not sufficient. It can produce a three-dimensional map only over the period of time that takes the catheter to travel the heart. The SQUID magnetometer mentioned in Murphy is obviously expensive to operate as it requires very low temperatures. The µ-magnetometer is not sufficiently small (1 cm x 1 cm x 0.6 cm). [1].

X.D.1. Proposed Improvements in the State of Art

A web of catheter-sized magnetometers was not found in the literature. Such a web may be sufficient for producing a three-dimensional image of the electrical activity of the heart at a cellular level. Future research is required to evaluate the feasibility of developing a web of catheter-sized magnetometers as one of the options studied.

X.E. Achieving a Cellular Resolution of VF

The spatial resolution of the magnetometer will depend on the size of the magnetometers used and their distance from the heart. The smaller the magnetometer, the higher the spatial resolution. The Biosense Webster’s catheters that can contain magnetometers can be as small as 4 French in diameter which translates into 4x0.33 = 1.32 mm (1320 micron). A typical heart cell has about 20 microns as its largest diameter while the

76 fibrous cells of conducting tissue can be as long as 90 microns. A device with a resolution of 500 micron would represent an improvement over the current state of the art.

The size of the magnetometer depends on the length of the string used and is defined by MEMS fabrication technology. Once the spatial resolution is determined by the size of the magnetometer, the sensitivity of the device must be taken into consideration.

Given a certain maximum length that defines the spatial variability of the web of magnetometers, each string length can be modified using piezoelectric elements. Passing a current in such elements cause the string to undergo strain and shorten or lengthen. When the string length is such that the resonance frequency of the string and that of the alternating current passing through the string are the same, and the external magnetic field is not changing with time, the string will vibrate with maximum amplitude at the frequency of the alternating current. If the magnetic field changes with time, which is the case with currents passing through the heart, then the string will vibrate at a frequency equal to the sum of the frequencies of the string due to the current passing through it and the frequency of the magnetic field.

When the string vibrates with its highest amplitude, highest sensitivity is achieved. So, highest sensitivity is achieved when the frequency of the current passing through is equal to the resonance frequency of the string. We can conclude that the sensitivity of the magnetometer is defined by how effectively we can tune the length of the string, thus its resonance frequency, with the current passing through.

Considering the above factors affecting the sensitivity of a string magnetometer, future research is required to evaluate whether a map of VF can be obtained with a resolution at a cellular level.

References 1. L.M. Miller, J.A. Podosek, E. Kruglick, T.W. Kenny, J.A. Kovacich, and W.J. Kaiser, “A µ-Magnetometer Based on Electron Tunneling,” Micro Electro Mechanical Systems, 1996, MEMS ’96 Proceedings, An Investigation of Micro Structures, Sensors, Actuators, Machines and Systems, IEEE, The Ninth, pp. 467-472.

77 XI. REFERENCES

Andre E. Aubert and Hugo Ector, editors, Progress in Biomedical Engineering, 2: Pacemaker Leads, Proceedings of the International Symposium on Pacemaker Leads, Leuven Belgium, Elsevier, September 5-7, 1984.

S. Serge Barold, Ronald X. Stroobandt, Alfons F. Sinnaeve, Cardiac Pacemakers Step by Step: An Illustrated Guide, Blackwell Publishing, 2004.

Liem, L. Bing, Implantable Cardioverter-Defibrillator, a Practical Manual, Kluwer Academic Publishers, Boston, 2001.

W.F. Boron, E. L. Boulpaep, Medical Physiology: A Cellular and Molecular Approach, Saunders, 2003.

F. M. Charbonnier, “External Defibrillators and Emergency External Pacemakers,” Proceedings of the IEEE, vol. 84, no. 3, p. March 1996.

S.B. Dunbar, K. Ellenbogen, A.E. Epstein, ed., Sudden Cardiac Death, Futura Publishing Company, Inc., New York, 1997, chapter 5: “The Nature of Activation During Ventricular Fibrillation.”

B. N. Goldreyer, M. G. Wyman, D. S. Cannom, P. A. Kruth, Endocardial Pacing Leads With an Emphasis on Atrial Sensing, Proceedings of the International Symposium on Pacemaker Leads, Leuven Belgium, Elsevier, September 5-7, 1984.

Ramesh M. Gulrajani, Bioelectricity and Biomagnetism, John Wiley and Sons, Inc., New York, 1998.

Casten W. Israel, S. Serge Barold, editors, Advances in the Treatment of Atrial Tachyarrhythmias: Pacing, Cardioversion, and Defibrillation, chapter 14: Rick McVenes, S. Serge Barold, Kenneth B. Stokes, Evolving Lead Technology for the Treatment of Atrial Tachyarrhythmias, Futura Publishing, Armonk, NY, 2002.

Richard E. Klabunde, “Cardiovascular Physiology Concepts,” http://www.cvphysiology.com/Arrhythmias/A004.htm,1999-2004.

L.M. Miller, J.A. Podosek, E. Kruglick, T.W. Kenny, J.A. Kovacich, and W.J. Kaiser, “A µ-Magnetometer Based on Electron Tunneling,” Micro Electro Mechanical Systems, 1996, MEMS ’96 Proceedings, An Investigation of Micro Structures, Sensors, Actuators, Machines and Systems, IEEE, The Ninth, pp. 467-472.

R. Plonsey and D.G. Fleming, Bioelectric Phenomena, McGraw-Hill, Inc., New York, 1969.

78 R. Plonsey and R.C. Barr, Bioelectricity: A Quantitative Approach, Plenum Press, New York, 1988.

N. Sperelakis, Physiology and Pathophysiology of the Heart, Kluwer Academic Publishers, Boston, 1995.

Borys Surawicz, Electrophysiologic Basis of ECG and Cardiac Arrhythmias, Willams and Wilkins, Baltimore, 1995.

W. A. Tacker, editors, Defibrillation of the Heart, Mosby, New York, 1994.

Barry M Weinberger, Lawrence C Brilliant, “Pacemaker and Automatic Internal Cardiac Defibrillator,” http://www.emedicine.com/emerg/topic805.htm, last updated, July 25, 2001. http://hebb.mit.edu/courses/8.515/lecture1/tsld020.htm http://www.web-ee.com/primers/files/pacemaker.pdf http://www.medtronic.com/newsroom/news_20040902a.html www.sjm.com www.guidant.com http://www.uspto.gov/web/offices/opc/documents/overview_dec02.pdf http://www.uspto.gov/web/offices/opc/documents/combined_cabrsm_dec_02.pdf http://en.wikipedia.org/wiki/Fabry-Perot http://www.fact-index.com/n/no/notch_filter.html http://hyperphysics.phy-astr.gsu.edu/hbase/magnetic/magfor.html http://encyclopedia.thefreedictionary.com/Quality+factor http://www.americanheart.org/presenter.jhtml?identifier=4741 http://www.merck.com/mrkshared/mmanual/section16/chapter206/206a.jsp http://www.guidant.com/condition/heart/heart_electrical.shtml http://www.math.utah.edu/~veronese/research.html http://en.wikipedia.org/wiki/Far_field

79 http://www.web-ee.com/primers/files/pacemaker.pdf http://www.genesis-sim.org/GENESIS/cnsweb/Cabletut.html. http://users.rcn.com/jkimball.ma.ultranet/BiologyPages/E/ExcitableCells.html. http://www.oucom.ohiou.edu/CVPhysiology/A007.htm. http://users.rcn.com/jkimball.ma.ultranet/BiologyPages/C/CellMembranes.html. http://www.cvphysiology.com/Arrhythmias/A004.htm. http://www.cvphysiology.com/Arrhythmias/A006.htm. http://www.cvphysiology.com/Arrhythmias/A010.htm.

80 TABLE OF CONTENTS

ABSTRACT 2

I. INTRODUCTION 3

II. ANATOMY, PHYSIOLOGY, AND ELECTROPHYSIOLOGY OF THE HEART 5

II.A. Anatomy 5

II.A.1. Blood Vessels 5 II.A.2. Chambers 6 II.A.3. Valves 7 II.A.4. Electrical System 7

II.B. Physiology 8

II.B.1. Mechanical Activity of the Heart 9 II.B.2. Electrical Activity of the Heart 9 II.B.3. Heart as a Functional Syncytium 10 II.B.4. Electrocardiogram (ECG) 11

II.C. Cellular Basis of Signal Generation and Conduction 11

II.C.1. Myocardium Cells 11 II.C.1.a. Myocardium Cell Membrane 12 II.C.2. Ion Distributions and Their Maintenance 13 II.C.2.a. Resting Potential 13 II.C.1.b. Action Potential 14 II.C.1.c. SA Node Action Potentials 15 II.C.1.d. Non-Pacemaker Action Potentials 15

III. ACTION POTENTIAL 17

III.A. Effective, Absolute, and Relative Refractory Periods 17

III.B. Changing Thresholds: Accommodation 17

III.C. Strength-Duration Curves of Action Potential 18

IV. CABLE THEORY 20

IV.A. Cardiac Fibers as Cables 20

IV.B. Derivation of the Cable Equation 21

81

IV.C. Deviations from Resting Potentials 22

IV.C.1. Passive Membrane: Subthreshold Conduction 23 IV.C.2. Action Potential: Excitation above Threshold 24

IV.D. Uniform Propagation 25

IV.D.1. Hudgkin-Huxley Membrane 25 IV.D.2. Hudgkin-Huxley Assumptions 26 IV.D.3. Qualitative Explanation of Propagation 27 IV.D.4. Exact Solution without the Hudgkin-Huxley Assumptions 28

V. MECHANISMS OF FIBRILLATION 29

V.A. Microscopic Structure of the Relevant Heart Muscles and Cells 29

V.B. Conduction Velocity of Action Potential 30

V.C. Reentry Mechanisms 31

V.C.1. Factors Present for Reentry 32 V.C.2. Wavelength and Length of Reentry Pathway 32 V.C.3. Unidirectional Block 33 V.C.4. Excitable Gap and Duration of Revolution 33 V.C.5. Documenting Reentry 34

VI. PACEMAKER AND ICD BASICS 35

VI.A. Pacemakers 35

VI.A.1. Pacing Codes 36 VI.A.2. Magnet Inhibition 37 VI.A.3. Pacemaker Complications 38

VI.B. ICD 38

VI.B.1. General Principles Underlying ICDs 38 VI.B.2 Optimal Defibrillation Methods 39 VI.B.2.a. Upper Limit of Vulnerability 39 VI.B.2.b. Waveform 40 VI.B.3 Defibrillation Electrodes and Leads 40 VI.B.3.a. Material 41 VI.B.4 Batteries and Capacitors 41 VI.B.5 Cardioversion and Pacing 42 VI.B.6. ICD Complications 42

82 VI.B.7. Examples of Manufacturer’s Advertising Material on ICD 43

VII. SENSING 46

VII.A. Sensing in Pacemakers 46

VII.A.1. Filtering 47

VII.B. Sensing in ICD 48

VII.B.1. ICD Sensing Systems 49 VII.B.1.a. Refractory and Blanking Periods 50 VII.B.2. Detection Enhancement in ICD 51 VII.B.2.a. Distinguishing VT in Single Chamber Devices 51 VII.B.2.b. Distinguishing VT in Dual Chamber Devices 52 VII.B.3. Crosstalk 53 VII.B.4. Far Field Sensing 53

VIII. SENSING LEADS 55

VIII.A. History 55

VIII.B. Leads Today 57

IX. PATENTS ON MEMS, ICD, AND SENSING 60

IX.A. PTO Classification 60

IX.B. PTO Searches 61

IX.C. MEMS and Heart 61

IX.C.1. Published Applications 61 IX.C.2. Issued Patents 63

IX.D. MEMS and Heart Related Applications and Patents in Detail 64

IX.D.1 Sawatari: Ultra-Miniature Optical Pressure Sensing System 64 X.D.2 Najafi: Wireless MEMS Capacitive Sensor for Physiologic Parameter Measurement 65 X.D.3 Thompson: Implantable Medical Device Incorporating Integrated Circuit Notch Filters 69 X.D.4. Murphy: Apparatus and Methods Using Mechanical Resonators to Enhance Sensitivity in Lorentz Force Magnetometers 70 X.D.5. Simons: Radio Frequency Telemetry System for Sensors and Actuators 73

83 IX.E. Earlier Referenced Patents (Not MEMS) 74

X. FURTURE RESEARCH ISSUES 77

X.A. Magnetic Mapping as the Recommended Method 77

X.A.1. String Magnetometers 78

X.B. Effectiveness of the Magnetic Mapping 78

X.C. State of Art in Magnetic Mapping 79

X.C.1. SQUID 79 X.C.2. CARTO Magnetic Mapping Procedure 80 X.C.3. Miniature Magnetometers 80

X.D. Is the State of Art Sufficient? 80

X.D.1. Proposed Improvements in the State of Art 80

X.E. Achieving a Cellular Resolution of VF 80

XI. REFERENCES 82

84