Clinical Cardiac Pacing, Defibrillation and Resynchronization Therapy 4th Edition: Expert Consult Premium Edition - Enhanced Online Features and Print

Clinical Cardiac Pacing, Defibrillation, and Resynchronization Therapy fourth edition KENNETH A. ELLENBOGEN, MD Kontos Professor of Medicine Chairman, Division of Cardiology Director, Clinical Electrophysiology Laboratory Medical College of Virginia Richmond, Virginia

G. NEAL KAY, MD Professor, Department of Medicine Director, Clinical Electrophysiology Section The University of Alabama at Birmingham Birmingham, Alabama

CHU-PAK LAU, MD Director, Cardiac Health Heart Centre Honorary Clinical Professor Department of Medicine University of Hong Kong Queen Mary Hospital Hong Kong, China

BRUCE L. WILKOFF, MD Director, Cardiac Pacing and Tachyarrhythmia Devices Robert and Suzanne Tomsich Department of Cardiovascular Medicine Institute Medical Information Officer Sydell and Arnold Miller Family Heart and Vascular Institute Professor of Medicine Cleveland Clinic Lerner College of Medicine of Case Western Reserve University Cleveland, Ohio

1600 John F. Kennedy Blvd. Ste 1800 Philadelphia, PA 19103-2899

CLINICAL CARDIAC PACING, DEFIBRILLATION, AND RESYNCHRONIZTION THERAPY Copyright © 2011, 2007, 2000, 1995 by Saunders, an imprint of Elsevier Inc.

978-1-4377-1616-0

All rights reserved. No part of this publication may be reproduced or transmitted in any form or by any means, electronic or mechanical, including photocopying, recording, or any information storage and retrieval system, without permission in writing from the publisher. Details on how to seek permission, further information about the Publisher’s permissions policies and our arrangements with organizations such as the Copyright Clearance Center and the Copyright Licensing Agency, can be found at our website: www.elsevier. com/permissions. This book and the individual contributions contained in it are protected under copyright by the Publisher (other than as may be noted herein).

Notices Knowledge and best practice in this field are constantly changing. As new research and experience broaden our understanding, changes in research methods, professional practices, or medical treatment may become necessary. Practitioners and researchers must always rely on their own experience and knowledge in evaluating and using any information, methods, compounds, or experiments described herein. In using such information or methods they should be mindful of their own safety and the safety of others, including parties for whom they have a professional responsibility. With respect to any drug or pharmaceutical products identified, readers are advised to check the most current information provided (i) on procedures featured or (ii) by the manufacturer of each product to be administered, to verify the recommended dose or formula, the method and duration of administration, and contraindications. It is the responsibility of practitioners, relying on their own experience and knowledge of their patients, to make diagnoses, to determine dosages and the best treatment for each individual patient, and to take all appropriate safety precautions. To the fullest extent of the law, neither the Publisher nor the authors, contributors, or editors, assume any liability for any injury and/or damage to persons or property as a matter of products liability, negligence or otherwise, or from any use or operation of any methods, products, instructions, or ideas contained in the material herein. Library of Congress Cataloging-in-Publication Data Clinical cardiac pacing, defibrillation, and resynchronization therapy / [edited by] Kenneth A. Ellenbogen … [et al.].—4th ed. p. ; cm. Includes bibliographical references and index. ISBN 978-1-4377-1616-0 (hardcover : alk. paper)  1.  Cardiac pacing.  2.  Defibrillators.  I.  Ellenbogen, Kenneth A. [DNLM:  1.  Cardiac Pacing, Artificial.  2.  Cardiac Resynchronization Therapy.  3.  Defibrillators, Implantable.  4.  Pacemaker, Artificial. WG 168] RC684.P3C54 2011 617.4′120645—dc23 2011026279

Executive Publisher: Natasha Andjelkovic Developmental Editor: Janice M. Gaillard Publishing Services Manager: Patricia Tannian Project Manager: Sarah Wunderly Design Direction: Ellen Zanolle

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To my wife and family, Phyllis, Michael, Amy, and Bethany, for their patience, support, and love. To my parents, Roslyn and Leon, who instilled in me a thirst for learning. To my students, teachers, and colleagues, who make each day an absolute delight. KAE

To my teachers, colleagues, and students, who have taught me about cardiac pacing. I am also indebted to the many members of the industry who have dedicated their professional careers to the design and improvement of pacing technology. These individuals have greatly improved the therapy that clinicians can offer to their patients, undoubtedly resulting in an improvement in their lives. Perhaps most important, this book is dedicated to my wife, Linda, for her patience and understanding during its preparation. GNK

To my wife and family, Carven, Yuk-Fai, and Yuk-Ming, for their understanding, support, and love. To my teachers, patients, and colleagues, who are my source of inspiration and encouragement. CPL

To my wife, Ellyn, children Jacob and Margaret, Benjamin and Kara, Ephram and Kay for their godly and inspirational patience and support. To my grandchildren, Isabelle and Tobias, for life, hope, and love. To my parents, Harvey and Glenna, for their unconditional love and insights. To Yeshua, the Messiah, for His salvation, and His sustaining covenant love. And for the inspiration of His words in Proverbs 15:2: “The tongue of the wise makes knowledge acceptable.” May the words of this book prove to be wise and useful to the student of cardiac pacing, defibrillation, and heart failure device therapy. BLW

CONTRIBUTORS Amin Al-Ahmad, MD Assistant Professor Department of Cardiovascular Medicine Stanford University School of Medicine Director, Cardiac Electrophysiology Laboratory Stanford University Medical Center Stanford, California Timing Cycles of Implantable Devices Angelo Auricchio, MD, PhD Professor of Cardiology University of Magdeburg Magdeburg, Germany Director, Clinical Electrophysiology Unit Fondazione Cardiocentro Ticino Lugano, Switzerland Basic Physiology and Hemodynamics of Cardiac Pacing Bryan Baranowski, MD Associate Staff Section of Cardiac Pacing and Electrophysiology Heart and Vascular Institute Cleveland Clinic Cleveland, Ohio Imaging of Implantable Devices Gust Bardy, MD Clinical Professor of Medicine University of Washington Director, Seattle Institute for Cardiac Research Seattle, Washington Subcutaneous Implantable Cardioverter-Defibrillators Peter H. Belott, MD, FACC, FHRS Director of Electrophysiology Sharp Grossmont Hospital La Mesa, California Permanent Pacemaker and Implantable Cardioverter-Defibrillator Implantation Janneke Berecki-Gisolf, MD, PhD Senior Research Fellow Monash University Accident Research Centre Melbourne, Victoria, Australia Pacing in Neurally Mediated Syncope Syndromes Paola Berne, MD Senior Fellow Department of Cardiology Electrophysiology Section Thorax Institute Hospital Clínic de Barcelona Barcelona, Spain ICD Therapy in Channelopathies Josep Brugada, MD, PhD, FESC Professor of Medicine Barcelona University Medical Director, Hospital Clínic de Barcelona Barcelona, Spain ICD Therapy in Channelopathies

Mina K. Chung, MD Associate Professor of Medicine Cleveland Clinic Lerner College of Medicine of Case Western Reserve University Cleveland, Ohio Imaging of Implantable Devices Joshua M. Cooper, MD Assistant Professor of Medicine University of Pennsylvania Attending Cardiac Electrophysiologist Hospital of the University of Pennsylvania Philadelphia, Pennsylvania Engineering and Construction of Pacemaker and ICD Leads Ann M. Crespi, PhD Senior Principal Scientist Department of Battery Research Medtronic, Inc. Minneapolis, Minnesota Power Systems for Implantable Pacemakers, Cardioverters, and Defibrillators George H. Crossley III, MD Clinical Professor of Medicine University of Tennessee College of Medicine Chief, Cardiac Services Baptist Hospital Nashville, Tennessee Pacemaker, Defibrillator, and Lead Codes, and Headers J. Kevin Donahue, MD Associate Professor Department of Medicine, Biomedical Engineering, Physiology & Biophysics Case Western Reserve University Clinical Cardiac Electrophysiologist The MetroHealth System Cleveland, Ohio The Biologic Pacemaker Derek J. Dosdall, PhD Assistant Professor Department of Internal Medicine Division of Cardiology University of Utah School of Medicine Adjunct Assistant Professor Department of Bioengineering University of Utah Adjunct Assistant Professor Department of Animal, Dairy, and Veterinary Sciences Utah State University Salt Lake City, Utah Cardiac Electrical Stimulation Kenneth A. Ellenbogen, MD Kontos Professor of Medicine Chairman, Division of Cardiology Director, Clinical Electrophysiology Laboratory Medical College of Virginia Richmond, Virginia Pacing for Atrioventricular Conduction System Disease; Interventional Techniques for Device Implantation

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Contributors

Andrew E. Epstein, MD, FAHA, FACC, FHRS Professor of Medicine University of Pennsylvania Chief, Cardiology Section Philadelphia VA Medical Center Philadelphia, Pennsylvania Troubleshooting of Implantable Cardioverter-Defibrillators Laurence M. Epstein, MD Associate Professor of Medicine Harvard Medical School Chief, Arrhythmia Service Director, Electrophysiology and Pacing Laboratory Brigham and Women’s Hospital Boston, Massachusetts Engineering and Construction of Pacemaker and ICD Leads Derek V. Exner, MD, MPH, FRCPC, FACC, FHRS Professor University of Calgary Canada Research Chair, Cardiovascular Clinical Trials Medical Director, Arrhythmia Program Libin Cardiovascular Institute of Alberta Calgary, Alberta, Canada Clinical Trials of Defibrillator Therapy Jeffrey M. Gillberg, MSEE Research Director Bakken Fellow Medtronic, Inc. Minneapolis, Minnesota Sensing and Detection Anne M. Gillis, MD, FRCPC Professor of Medicine University of Calgary Calgary, Alberta, Canada Pacing for Sinus Node Disease Andrew A. Grace, MBChB, PhD Research Group Head University of Cambridge Consultant Cardiologist Papworth Hospital Cambridge, United Kingdom Subcutaneous Implantable Cardioverter-Defibrillators Henry Halperin, MD, MA Carver Professor of Medicine Johns Hopkins School of Medicine Baltimore, Maryland Electromagnetic Interference and CIEDs Haris M. Haqqani, MBBS(Hons), PhD Senior Lecturer School of Medicine The University of Queensland Consultant Electrophysiologist The Prince Charles Hospital Brisbane, Queensland, Australia Engineering and Construction of Pacemaker and ICD Leads David Hayes, MD Medical Director, Affiliated Practice Network Mayo Clinic Mayo College of Medicine Rochester, Minnesota Ethical Issues

Margaret Hood, MBChB Cardiologist Auckland City Hospital Auckland, New Zealand Subcutaneous Implantable Cardioverter-Defibrillators Henry H. Hsia, MD Associate Professor of Medicine Stanford University School of Medicine Associate Director, Electrophysiology Laboratory Stanford University Medical Center Stanford, California Timing Cycles of Implantable Devices Raymond E. Ideker, MD, PhD Jeanne V. Marks Professor of Medicine Professor Departments of Biomedical Engineering and Physiology University of Alabama at Birmingham Birmingham, Alabama Principles of Defibrillation: From Cellular Physiology to Fields and Waveforms Carsten W. Israel, MD Associate Professor of Medicine and Cardiology Department of Cardiology Division of Electrophysiology J. W. Goethe University Frankfurt, Germany Chief of Cardiology Evangelical Hospital Bielefeld Bielefeld, Germany Clinical Trials of Atrial and Ventricular Pacing Modes Bharat K. Kantharia, MD, FRCP, FAHA, FACC, FESC, FHRS Professor of Medicine Director, Cardiac Electrophysiology Training Program The University of Texas-Health Science Center at Houston Director, Cardiac Electrophysiology Laboratories Heart and Vascular Institute–Memorial Hermann Hospital Houston, Texas Approach to Pulse Generator Changes Karoly Kaszala, MD, PhD Assistant Professor of Cardiology Virginia Commonwealth University School of Medicine Director, Cardiac Electrophysiology Hunter Holmes McGuire VAMC Richmond, Virginia Electromagnetic Interference and CIEDs G. Neal Kay, MD Professor of Medicine Director, Cardiac Electrophysiology Section Division of Cardiovascular Disease University of Alabama at Birmingham Birmingham, Alabama Cardiac Electrical Stimulation Paul Khairy, MD, PhD, FRCPC Associate Professor Department of Medicine University of Montreal Canada Research Chair, Electrophysiology and Adult Congenital Heart Disease Director, Montreal Heart Institute Adult Congenital Center Montreal, Quebec, Canada Sensing and Detection



Daniel B. Kramer Teaching Fellow in Medicine Harvard Medical School Beth Israel Deaconess Medical Center Boston, Massachusetts Guidelines for Managing Pacemaker and Implantable Defibrillator Advisories Steven P. Kutalek, MD Associate Professor of Medicine Director, Cardiac Electrophysiology Associate Chief, Division of Cardiology Drexel University College of Medicine Philadelphia, Pennsylvania Approach to Pulse Generator Changes Rachel Lampert, MD Associate Professor Yale University School of Medicine Attending Physician Yale New Haven Hospital New Haven, Connecticut Ethical Issues Chu-Pak Lau, MD Director, Cardiac Health Heart Centre Honorary Clinical Professor Department of Medicine University of Hong Kong Queen Mary Hospital Hong Kong, China Implantable Sensors for Rate Adaptation and Hemodynamic Monitoring; Leadless Pacing Concepts Kathy L. Lee, MBBS, FRCP, FACC Honorary Clinical Assistant Professor Medical School University of Hong Kong Hong Kong, China Leadless Pacing Concepts Charles J. Love, MD Professor of Medicine Director, Cardiac Rhythm Device Services Division of Cardiovascular Medicine The Ohio State University Medical Center Columbus, Ohio Pacemaker Troubleshooting and Follow-up William H. Maisel, MD, MPH Associate Professor of Medicine Harvard Medical School Director, Pacemaker and ICD Service Beth Israel Deaconess Medical Center Boston, Massachusetts Guidelines for Managing Pacemaker and Implantable Defibrillator Advisories Saman Nazarian, MD Assistant Professor of Medicine Johns Hopkins University School of Medicine Director, Ventricular Arrhythmia Ablation Service Johns Hopkins Hospital Baltimore, Maryland Electromagnetic Interference and CIEDs

Contributors

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Mark J. Niebauer, MD, PhD Assistant Professor Cleveland Clinic Lerner College of Medicine of Case Western Reserve University Staff Physician Cleveland Clinic Cleveland, Ohio Defibrillation Testing, Implant Testing, And Relation to Empiric ICD Programming Marco V. Perez, MD Clinical Instructor Stanford University School of Medicine Stanford University Medical Center Stanford, California Timing Cycles of Implantable Devices Robert Andrew Pickett, MD Physician St. Thomas Research Institute St. Thomas Heart at Baptist Hospital Nashville, Tennessee Pacemaker, Defibrillator, and Lead Codes, and Headers Stephen M. Pogwizd, MD Featheringill Endowed Professor in Cardiac Arrhythmia Research Professor of Medicine, Physiology & Biophysics, and Biomedical Engineering Director, Center for Cardiovascular Biology Associate Director, Cardiac Rhythm Management Laboratory University of Alabama at Birmingham Birmingham, Alabama Principles of Defibrillation: From Cellular Physiology to Fields and Waveforms Frits W. Prinzen, PhD Professor of Physiology Maastricht University Maastricht, the Netherlands Basic Physiology and Hemodynamics of Cardiac Pacing François Regoli, MD, PhD Attending Physician, Cardiologist Fondazione Cardiocentro Ticino Lugano, Switzerland Basic Physiology and Hemodynamics of Cardiac Pacing Dwight W. Reynolds, MD Professor of Medicine Chief, Cardiovascular Section University of Oklahoma Health Sciences Center Oklahoma City, Oklahoma Permanent Pacemaker and Implantable Cardioverter-Defibrillator Implantation Michael P. Riley, MD, PhD Assistant Professor of Medicine Hospital of the University of Pennsylvania Philadelphia, Pennsylvania Troubleshooting of Implantable Cardioverter-Defibrillators Anthony Rorvick, BS Medtronic, Inc. Minneapolis, Minnesota Power Systems for Implantable Pacemakers, Cardioverters, and Defibrillators

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Contributors

Elizabeth Vickers Saarel, MD Associate Professor University of Utah Director, Electrophysiology Primary Children’s Medical Center Salt Lake City, Utah Imaging of Implantable Devices

Warren M. Smith, MBChB Honorary Clinical Associate Professor in Medicine University of Auckland Cardiologist Auckland City Hospital Auckland, New Zealand Subcutaneous Implantable Cardioverter-Defibrillators

Leslie A. Saxon, MD Clinical Scholar and Chief Division of Cardiovascular Medicine Keck School of Medicine University of Southern California Los Angeles, California Clinical Trials of Cardiac Resynchronization Therapy: Pacemakers and Defibrillators

Bruce S. Stambler, MD Professor of Medicine Case Western Reserve University Cleveland, Ohio Pacing for Atrioventricular Conduction System Disease

Craig L. Schmidt, PhD Senior Director, Energy Systems Research Medtronic Energy and Component Center Medtronic, Inc. Minneapolis, Minnesota Power Systems for Implantable Pacemakers, Cardioverters, and Defibrillators Gerald A. Serwer, MD Professor of Pediatrics University of Michigan Attending Pediatric Cardiologist University of Michigan Congenital Heart Center Ann Arbor, Michigan Pediatric Pacing and Defibrillator Use Robert S. Sheldon, BSc, MD, PhD Professor of Cardiac Sciences University of Calgary Senior Vice President of Research Alberta Health Services Calgary, Alberta, Canada Pacing in Neurally Mediated Syncope Syndromes Richard B. Shepard Emeritus Professor, Cardiovascular Surgery University of Alabama at Birmingham Birmingham, Alabama Cardiac Electrical Stimulation Ira Shetty, MD Clinical Assistant Professor Loyola University Medical Center Maywood, Illinois Pediatric Pacing and Defibrillator Use Chung-Wah Siu, MD Clinical Assistant Professor Department of Medicine University of Hong Kong Hong Kong, China Implantable Sensors for Rate Adaptation and Hemodynamic Monitoring Paul M. Skarstad, PhD Senior Director of Research (Retired) Medtronic Energy and Component Center Medtronic, Inc. Minneapolis, Minnesota Power Systems for Implantable Pacemakers, Cardioverters, and Defibrillators

Marc Strik, MD PhD Student Department of Physiology Maastricht University Maastricht, the Netherlands Basic Physiology and Hemodynamics of Cardiac Pacing Michael O. Sweeney, MD Associate Professor of Medicine Harvard Medical School Cardiac Pacing and Defibrillation Cardiac Arrhythmia Service Brigham and Women’s Hospital Boston, Massachusetts Troubleshooting of Biventricular Devices Charles D. Swerdlow, MD, FACC, FAHA, FHRS Clinical Professor of Medicine Department of Cardiology Cedars Sinai Heart Institute David Geffen School of Medicine University of California Los Angeles Los Angeles, California Sensing and Detection Sandeep Talwar, MD, PhD, FRCP Senior Electrophysiology Fellow Division of Cardiology University of Utah Salt Lake City, Utah Clinical Trials of Cardiac Resynchronization Therapy: Pacemakers and Defibrillators Patrick J. Tchou, MD Associate Section Head, Cardiac Electrophysiology and Pacing Department of Cardiovascular Medicine Heart and Vascular Institute Cleveland Clinic Cleveland, Ohio Defibrillation Testing, Implant Testing, And Relation to Empiric ICD Programming Hung-Fat Tse, MD William MW Mong Professorship in Cardiology Academic Chief, Cardiology Division Department of Medicine Queen Mary Hospital The University of Hong Kong Hong Kong, China Implantable Sensors for Rate Adaptation and Hemodynamic Monitoring



Mintu P. Turakhia, MD, MAS Instructor of Medicine Stanford University School of Medicine Director of Cardiac Electrophysiology Veterans Affairs Palo Alto Health Care System Stanford, California Timing Cycles of Implantable Devices Darrel F. Untereker, PhD Vice President of Corporate Research and Technology Strategic and Scientific Operations Medtronic, Inc. Minneapolis, Minnesota Power Systems for Implantable Pacemakers, Cardioverters, and Defibrillators Niraj Varma, MA, DM, FRCP Consultant Cardiac Electrophysiologist Heart and Vascular Institute Cleveland Clinic Cleveland, Ohio Follow-up Monitoring of Cardiac Implantable Electronic Devices Gregory P. Walcott, MD Associate Professor of Medicine University of Alabama at Birmingham Birmingham, Alabama Principles of Defibrillation: From Cellular Physiology to Fields and Waveforms Paul J. Wang, MD Professor of Medicine Stanford University School of Medicine Director, Cardiac Arrhythmia Service and Cardiac Electrophysiology Stanford University Medical Center Stanford, California Timing Cycles of Implantable Devices

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Oussama Wazni, MD Assistant Professor of Medicine Cleveland Clinic Lerner College of Medicine of Case Western Reserve University Director, Research Development Director, Out Patient Department, Cardiac Electrophysiology Cleveland Clinic Cleveland, Ohio Prevention and Management of Procedural Complications; Techniques and Devices for Lead Extraction Bruce L. Wilkoff, MD Professor of Medicine Cleveland Clinic Lerner College of Medicine of Case Western Reserve University Director, Cardiac Pacing and Tachyarrhythmia Devices Medical Information Officer Heart and Vascular Institute Cleveland, Ohio Prevention and Management of Procedural Complications; Techniques and Devices for Lead Extraction Seth J. Worley, MD Section Chief for Research Lancaster General Health President and Medical Director Lancaster Heart and Stroke Foundation Lancaster, Pennsylvania Left Ventricular Lead Implantation; Interventional Techniques for Device Implantation Paul C. Zei, MD, PhD Clinical Associate Professor of Medicine Stanford University School of Medicine Director, Cardiovascular Medicine Outpatient Clinics Chief, Clinic Advisory Council Stanford University Medical Center Stanford, California Timing Cycles of Implantable Devices

PREFACE It all started with cardiac pacemakers in 1958, but now we have car-

diovascular implantable electronic devices (CIED). Investigations of the electrical stimulation of the heart began long ago, but the truly impressive technological developments of this technology over the past more than 60 years has been the result of intense collaborative efforts of physicians, engineers, allied professionals, and heart rhythm device manufacturers. When we began to collect the information to mentor heart rhythm device professionals for the initial 1995 edition of Clinical Cardiac Pacing we did not have a chapter on implantable defibrillators. We added implantable defibrillators in the second edition (2000), and in the third edition (2007) we added resynchronization therapy. Now we have a fourth edition. Although there is not a new form of therapy, there are still important changes in the technologies and our insights into how they are applied. There has been a huge increment in the evidence base of how and in whom these devices should be used, and a great maturation of the techniques of implantation and extraction of devices, and management of patients with specific conditions. Many new technologies are discussed, including leadless pacing, and the subcutaneous ICD. Our philosophy in putting together the fourth edition remains the same as that of our first three editions. We have planned this book to emphasize the science of cardiac pacing, implantable defibrillation, and cardiac resynchronization therapy, and to underline the importance of the fact that it is an interdisciplinary field. Physicians are part of a large web of health professionals who need increasing amounts of information about implantable devices. We have a web site with this edition that includes figures and movies not included in the paper version of our text, as well as much additional material. All of the figures from the text are included and available for download from this web site. We have sought to meet the needs of many with this textbook. Clinicians, scientists, nurses, technicians, and engineers will find the information in these pages practical, authoritative, and helpful in better understanding this therapy. We are excited about the opportunity to present this material in a comprehensive scientific manner, and in doing so we stand on the shoulders of giants in the field of heart rhythm devices. Sy Furman wrote the initial edition’s foreword, and for that affirmation we are extremely thankful. We are more in debt to his legacy of innovation, problem solving, always asking more questions, and, finally, his commitment to educating any and all who

demonstrate interest. He will always be greatly missed and remind us of our debt to all our colleagues, physicians, engineers, nurses, and technicians. We also owe a great debt of gratitude to Serge Barold and particularly to his singularly important text, Modern Cardiac Pacing, published in 1985 by the Futura Publishing Company. Each of us learned more about cardiac pacing from this textbook than from any mentor. It is hard to give credit to some and not others, but clearly there are so many others and also the giants that came before them. It is our hope that this edition will continue to mentor the next generation of heart rhythm professionals. We gratefully acknowledge the invaluable assistance and encouragement of Natasha Andjelkovic and Janice Gaillard of the Health Sciences Division of Elsevier for all their help in keeping the fourth edition on track. We owe a great debt of gratitude to our colleagues and fellows from the Medical College of Virginia and the McGuire Veterans Affairs Medical Center, the University of Alabama, The University of Hong Kong and Queen Mary Hospital, and the Cleveland Clinic for their patience and support in shouldering the extra workload that allowed us to finish our chapters and editing on time. Most important, we cannot thank enough our many contributors and their colleagues, who labored extensively, often taking time from family and other projects, to finish their chapters. This large group of individuals deserves all the credit and thanks for making the fourth edition possible. Our wonderful secretaries, Vera Wilkerson (Virginia Commonwealth University/Medical College of Virginia), Monica Crosby (Cleveland Clinic), Jenny To (The University of Hong Kong), and Dorothy Welch (University of Alabama) were invaluable for their contributions to help complete this project. This textbook is designed to be a functional tool and reference, helping clinicians, scientists, and engineers make the decisions that improve patients’ lives every day. It is our desire that this book serve as a valuable resource to all of these people for many years to come. Kenneth A. Ellenbogen, MD G. Neal Kay, MD Chu-Pak Lau, MD Bruce L. Wilkoff, MD

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VIDEO CONTENTS Introduction Bruce L. Wilkoff, MD 9 Basic Physiology and Hemodynamics of Cardiac Pacing 1. 3D electroanatomical mapping in a patient with left bundle branch block QRS morphology 2. Animated 3D reconstruction of the left ventricular wall with myocardial strain 21 Permanent Pacemaker and Implantable CardioverterDefibrillator Implantation 3. Catheter Delivery of a Fixed Extended Screw in Lead to Coronary Sinus Os 4. Catheter Delivery of a Fixed Extended Screw in Lead 5. Removal of a Fixed Extended Screw in Lead in the Atrium 22 Left Ventricular Lead Implantation 6. Intracardiac View of Coronary Sinus Os and Thebesian Valve 7. Intracardiac View of the Thebesian Valve and Its Fenestrations 8. Intracardiac View of the Right Atrium from the IVC Including the Coronary Sinus Os and Thebesian Valve 9. Rotating Intracardiac View of the Right Atrium from the IVC Including the Coronary Sinus Os and Thebesian Valve 10. Intracardiac View of a White and Floppy Thebesian Valve 11. Intracardiac View of a Tattered Thebesian Valve 12. Intracardiac View of an Occlusive Balloon Inflation in the Coronary Sinus Os 13. Intracardiac View of Insertion and Occlusion of the Coronary Sinus with a Venography Catheter 14. Intracardiac View of a Double Coronary Sinus OS and Thebesian Valve 15. Intracardiac View of Coronary Sinus Os and Thebesian Valve with Only Free Access from Inferior Direction 16. Intracardiac View of Coronary Sinus Os and Tricuspid and Thebesian Valves with Attempted EP Catheter Access from Above 17. Intracardiac View of Coronary Sinus Os and Tricuspid and Thebesian Valves with Successful EP Catheter Access from Below 18. Intracardiac View of Coronary Sinus OS and Thebesian Valve with Attempted EP Catheter Access from Above 19. Intracardiac View of a Complicated Thebesian and Tricuspid Valve 20. Intracardiac View of Coronary Sinus Os with Early Branching 21. Intracardiac View of Coronary Sinus Os with Early Branching

22. Intracardiac View of Catheters in the Coronary Sinus and Through the Tricuspid Valve 23. Fluoroscopic Venogram of the Main Coronary Sinus and Mid-CS Valve 24. Fluoroscopic Advancement of a Decapolar EP Catheter into the Coronary Sinus 25. 3D image of the Cardiac Venous Anatomy 26. Occlusive Venography of Coronary Sinus LAO 30 27. Occlusive Venography of Coronary Sinus Anteroposterior Projection 28. Occlusive Venography of Coronary Sinus RAO 30 Projection 29. Sheath and Wire Subselected into Anterolateral Branch 30. Sheath and Wire Subselected into Posterior Vein Branch 31. Subvalvular Pouch Mimicking Coronary Sinus Os 32. True Coronary Sinus Venogram 33. Final Lead Position in Posterior-Lateral Branch 24 Approach to Pulse Generator Change 34. Abdominal ICD Pulse Generator with Subcutaneous Patch Electrode 35. Abdominal ICD Pulse Generator with Previously Cut and Capped Additional Electrode 36. Patient from Previous Case After Extraction and Reimplantation 37. Alternative Approach to an Occluded Subclavian Vein 38. Total Occlusion of the Left Subclavian Vein with Reconstitution via the Right Side 25 Prevention and Management of Procedural Complications 39. Transesophageal Echo of Large Right Atrial Vegetation on an ICD Lead 40. Transesophageal Echo of Large Right Atrial Vegetation on a Pacemaker Lead 41. Subclavian Vein Venography and Lead Fractures 26 Techniques and Devices for Lead Extraction 42. Needles-Eye-Snare Extraction of Fractured Ventriculo-Jugular Shunt 43. Fluoroscopy of Embolized Telectronics Accufix Wire 44. Laser Lead Extraction of Single Coil ICD Lead Special Presentation 45. ICD Implantation: Implantation and Management of Complications (PowerPoint presentation)

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Cardiac Electrical Stimulation G. NEAL KAY  |  DEREK J. DOSDALL  |  RICHARD B. SHEPARD

Pathologic cardiac conditions such as asystole, bradycardia, conduc-

tion block, and dyssynchronous contraction may lead to significant morbidity and mortality. One of the most influential and effective medical device therapies developed in the last century is the cardiac pacemaker and implantable cardioverter-defibrillator (ICD). The lives of millions of patients have been improved and extended through the use of pacing techniques that regulate heart rate and improve cardiac output. Direct electrical stimulation of excitable cardiac tissue causes a change in the transmembrane potential of resting cardiac cells. If the transmembrane potential is raised to a certain level, an action potential is initiated in the cells and spreads to adjacent working myocardial cells. The action potential induces mechanical contraction of the cardiac cells that spreads throughout the heart as a self-propagating wavefront. This chapter reviews the fundamental concepts of artificial electrical cardiac stimulation, including the cellular aspects of myocardial stimulation, the influence of external current on cardiac tissue, waveform and electrode considerations, clinical applications and considerations, and ongoing research regarding cardiac stimulation.

Concepts Related to Electrical Stimulation of the Heart STATIC ELECTRIC CHARGE AND ELECTRIC FIELDS Physical objects may acquire an electric charge when they have a net excess or deficit of electrons relative to the number of protons. When the number of electrons exceeds the number of protons, the object is said to have a negative charge, whereas a net deficit of electrons results in the object acquiring a positive charge. An electrically charged object is surrounded by an electric field such that charged objects act at a distance on other objects having similar or opposite charge. The strength of the electric field is related to the magnitude of its charge. A gravitational field also acts at a distance, but an electric field has an important difference, polarity. Thus, electric fields have directionality with the convention that field lines are drawn away from positively charged and toward negatively charged objects. The electric fields surrounding charged objects interact with each other such that the presence of two electrically charged objects results in a force that either attracts the two objects (if they have opposite charges) or repels the two objects (if they have the same qualitative charge). This attractive or repulsive force acts at a distance, not requiring that the charged objects be in contact, and the attractive or repulsive force (F) is given by Coulomb’s law:

F=k

Q1 Q2 r2

where Q1 and Q2 are the magnitude of the charges (measured in coulombs), r is the distance separating the charged objects, and k is a 1 constant equal to with ε being the permittivity of free space. 4πε According to Coulomb’s law, the force attracting or repelling charged objects increases with the magnitude of charge and decreases with the square of the distance that separates them. Potential Difference Voltage is the electrical potential energy possessed by a charge by virtue of its position in space. A charged object in space has potential energy

because of the forces that other charges exert on that object. When two charged objects repel one another, energy is required to move either object closer to the other, thereby increasing the potential energy. However, if the charged object is attracted to another charged object, potential energy will be converted into kinetic energy when either charge moves closer to the other. Thus, the force generated by the interaction of electrical charge is electromotive, tending to move charged objects. Voltage is the potential energy per unit charge for a charge in an electric force field. Thus, in order to move a charge from point A to point B, the work required (measured in joules) is calculated by multiplying the charge, Q, by the voltage difference between points A and B, VAB. Because voltage is potential energy per unit of charge, it is measured in terms of joules/coulomb, as follows:

1 volt = 1 joule/coulomb

When one refers to “voltage,” the reference is to a difference in potential between two points in space. Electric Current An electric current, I, is present when there is a movement of electric charge. Although electric charge may move through several mechanisms, for clinical purposes, electric charge is usually carried by the flow of electrons through a wire (e.g., pacemaker lead) or by the movement of ions in blood, interstitial fluid, across cell membranes, or within the cytoplasm of cells. By historical convention, current is considered to flow in the direction that positive charges would move. In reality, however, an electric current in a wire is carried by the movement of electrons that are negatively charged. For clarity in this chapter, current is stated in terms of electron or ion motion. Because electric current is the movement of charge, it is measured in terms of coulombs per second (It = dQ/dt), with 1 ampere of current equal to the movement of 1 coulomb/sec. The electromotive force for cardiac pacemakers or ICDs is determined by the chemistry of a battery. For lithium-iodine batteries at beginning of life, the chemical reaction generates approximately 2.8 volts of electromotive force. The total amount of charge that is available in the battery is measured in terms of the amount of current that can be provided for a given unit of time. For pacemaker or ICD batteries the amount of charge that can be stored in the battery is measured in terms of ampere-hours, with 1 A-hr equal to 3600 coulombs of charge. ELECTRIC CIRCUIT An electric circuit is an electric charge–conducting pathway that ends at its beginning. For electric circuits involved in myocardial stimulation by pacemakers or ICDs, the voltage difference in the circuit is provided by a battery. The difference in voltage generated by the battery results in a flow of charge (current) from the pulse generator through the conductors in the leads, electrodes, extracellular electrolytes, cell membranes with highly regulated transmission of charged ions in both directions through the membrane, and intracellular ions and charged molecules. As current flows through the complete electric circuit, the net voltage change must be zero (Kirchoff ’s voltage law). Thus, the electromotive force (voltage) generated by the battery (an increase in potential energy) must be completely dissipated (a decrease in potential energy) as current flows through all the elements of the circuit to end at the battery.

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SECTION 1  Basic Principles of Device Therapy

Electric circuits in the clinical practice of pacing have multiple elements, including the pulse generator battery, the lead conductor(s), the electrodes, the myocardium, and blood within the great veins and cardiac chambers. All these elements introduce opposition to the flow of current. Series Circuit In a series circuit, or circuit module, the elements are connected so that current must flow sequentially through each element in the circuit. Thus, the current flowing through all elements in series is the same, with the voltage difference decreasing sequentially as it passes through each element in the circuit. Parallel Circuit In a parallel circuit, two or more elements are joined at each end to a common conductor (node in the circuit). Therefore, the potential difference is the same before each element and after each element, and current will flow from one of the common conductors to the other through any or all of the elements. The degree of current flow in each element is inversely related to the factors that oppose the flow of electric charge in that element. At any node (junction) in an electric circuit, the sum of currents flowing into that node is equal to the sum of currents flowing out of that node (Kirchoff ’s current law). Most biologic circuits are made of various combinations of series and parallel modules or subcircuits. For example, because of electrochemical effects, an electrode placed in the heart may act as a capacitor in parallel with a resistor, both in series with the lead joining the pulse generator to the electrode. ELECTRODE POLARITY All defibrillator and pacemaker electric circuits have both a positively charged electrode (the anode) and a negatively charged electrode (the cathode). The negatively charged cathode is typically the tip electrode on a pacing lead. Electrons from the pulse generator flow through the cathode-tissue interface and return to the anode, which may be located on a pacing lead or the pulse generator casing. The terminology used for electrode polarity may be confusing as applied to lead electrodes and the electrodes of a battery. The electrode in a battery at which oxidation occurs (e.g., oxidation of lithium to yield Li+ plus an electron, e−) is the battery anode. The battery anode, by continuing oxidation, furnishes electrons to the circuit external to it. Therefore, in contradistinction to the terminology used for pacing leads, the terminal of the battery where electrons are provided to the circuit is the battery anode. From the battery anode, electrons flow through the circuitry and eventually enter the pacemaker lead electrode that is in contact with the myocardium. This electrode, receiving electrons from the pulse generator and furnishing electrons to the tissue, is the lead cathode. The return electrode located in the heart or on the pulse generator casing is the lead anode. It collects electrons from the tissue and returns them through the pulse generator circuitry to the positive electrode of the battery, the battery cathode, where reduction occurs (e.g., I2 + 2e− yields 2I−). The consistency in the terminology is that, when oxidation occurs, it occurs at an anode, and in the circuitry, an anode connects to a cathode that subsequently connects to another anode, and so on. OPPOSITION TO FLOW OF ELECTRIC CURRENT In cardiac pacing and defibrillation, several elements in the electric circuit oppose the flow of electric current. As current flows through some of these elements, such as the conductor wires in the leads and pulse generator, the opposition to current flow results in energy being lost as heat. These elements are known as ohmic, and the opposition to current flow is called resistance (R). The instantaneous voltage developed across a perfect resistor is linearly proportional to the instantaneous current flow through the resistor. If a steady voltage across the

resistor is represented by V, the current by I, and an unchanging resistance by R, the relationship is expressed as follows in Ohm’s law:

V = IR

In reality, when all the elements of a cardiac pacing circuit are considered, these factors are much more complex than can be represented by Ohm’s law alone. These other factors include capacitance and inductance. For calculations involving a pulsed (e.g., pacemaker stimulus) or an alternating current (AC) circuit that contain reactive elements (e.g., electrodes in contact with electrolytes), the term impedance (Z) must be used in place of resistance. Impedance is a vector sum of resistance (R) and reactance (X). The following sections discuss the major components of reactance. CAPACITANCE As mentioned earlier, current can be carried in different ways. Electrolyte is used as a generic term for the electrically conductive extracellular and intracellular fluids near pacemaker or defibrillator electrodes and elsewhere in the heart and blood vessels. The electrolyte conducts ions but not electrons. The difference between the conduction of electric currents by electrolytic fluids and the flow of electrons over metal wires is crucial to pacing. Because a negatively charged pacing electrode in contact with the endocardium is surrounded by blood and interstitial fluid, positively charged ions move toward that electrode during a pacing stimulus. This results in the phenomenon of polarization, which develops rapidly and dissipates slowly after the end of the stimulus. The opposite effect occurs on the positively charged anode. This effect of ions moving to oppose the flow of electric current has the effect of a capacitor in the circuit. A capacitor is an object that stores energy in an electric field by holding positive charges apart from closely approximated negative charges. A capacitor requires a material or space between the layers of negative and positive charges that is normally nonconducting (the dielectric). A cell membrane, although leaky, acts as a capacitor by separating the negatively charged inside of the cell from the more positively charged outside. Cell membranes have very high capacitance per unit area of cell membrane. The interface between a pacing electrode and the charged electrolytes that surround the electrode at its surface in the myocardial tissue acts, in part, as a capacitor. The terms Helmholtz capacitor and Helmholtz capacitance are used in this chapter for capacitor-like effects that occur at pacemaker and defibrillator electrode-electrolyte interfaces. Capacitance (C) is the term that specifies, for a given voltage applied across a capacitor, how much electrical charge (Q) can be stored by the capacitor. If V represents a steady voltage applied across the capacitor, then Q = CV. (If E is used instead of V as the symbol for electric potential, the relationship may be expressed as Q = CE.) The unit for capacitance is the farad. One farad is the capacitance of a capacitor that, on being charged to 1 volt, will have stored 1 coulomb of charge. Again, a coulomb is the amount of charge delivered by 1 ampere flowing for 1 second. Coulombs delivered can be expressed as follows:

t

Q t = ∫ i t dt 0

in which Qt is the total charge delivered between time 0 and time t, and it is the instantaneous current at each time segment between time t 0 and time t. The integral ∫ i t dt is the net area under the instantaneous 0 current-versus-time plot. INDUCTANCE When an alternating current flows through a wire, a magnetic field surrounding the wire is induced. An inductor is an object that stores or releases energy in or from a changing magnetic field. The voltage difference across an inductor is proportional to the rate of change of current flowing through the inductor. Energy is stored during the formation of the magnetic field and is released when the magnetic field decreases or disappears. Inductance is the term that specifies the



1  Cardiac Electrical Stimulation

relationship between the voltage across an inductor and the rate of change of current traversing the inductor. The magnitude of the inductance can be represented by the symbol L. If Vt represents the instantaneous voltage across the inductor, and iL represents the instantaneous current flowing through the inductor, the relationship is given by the following equation: di L 1 t or i L = ∫ v t dt dt L 0 Note that the voltage across the inductor is directly proportional to the rate of change of current flowing through the inductor. Cell membrane currents have some of the current- and voltage-versus-time characteristics of an inductance in parallel with a capacitance.1 These inductancelike effects are related to the timing and magnitude of potassium ions moving into and out of the cell.2

vt = L

REACTANCE (CAPACITANCE AND INDUCTANCE) For reactive elements connected in series, net reactance is the scalar sum of inductive reactance (positive in the mathematical complex plane) and capacitive reactance (negative in complex plane). Pure reactance values depend on the rates of change of current and voltage, whereas pure resistance values do not. Phenomena of this type are caused by differences in the timing of the peaks (phase angles) of the voltages or currents in the various reactive components. These reactive effects can be important in biventricular pacing threshold measurements (see later discussion). The component of reactance that is most relevant to both pacing electrodes and cell membranes is capacitance, with inductance being much less important. For example, the cardiac action potential spreading throughout the heart generates a changing magnetic field that transiently stores a very small amount of energy. However, the changing magnetic field generated by spread of the action potential is so small that it is not clinically significant except in the research setting.3 For circuits with resistance R, capacitance C, and inductance L in series, where qt represents the charge accumulated across the capacitance at any time t, and where the current through these combined elements at time t is it, the voltage Vt at time t is described by the following equation: t idt qt di di ∫ 0 v t = i tR + + L or v t = i t R + +L C dt C dt t (remembering that q t = ∫ i t dt and that the voltage across a capacitor 0 qt at time t is ). C These equations indicate that, for an instantaneous current it, the instantaneous voltage across the series circuit is the sum of the effects at that instant in time of the resistance, capacitance, and inductance of the circuit. Note especially that the instantaneous effects are highly t related to the net amount of charge, q t = ∫ i t dt , that has accumulated 0 in the capacitor from time 0 to the instantaneous time t. The equations show that the capacitance effect on the voltage decreases as the capacitance increases. This has clinical relevance in that, for example, the polarization voltage that interferes with autosensing pulse generators decreases as the electrode capacitance increases.

(

)

Cellular Aspects of Myocardial Stimulation THE PHOSPHOLIPID BIMEMBRANE Living cells maintain or regenerate a difference in electric potential across the cell membrane. Excitable tissues such as myocardium respond to electrical stimulation of one or more cells with a wave of electrical depolarization—a transient reversal of the voltage gradient across the membrane—that can propagate from cell to cell. Excitable cells respond to a relatively small applied change in electric potential

5

Phospholipid

Lipid “tail” Phosphate “head” Ion channel

Figure 1-1  The phosphomembrane bilayer of a cardiomyocyte contains ion channels and structures that maintain distinct ion concentrations inside and outside of the cell. Charge pumps maintain an ion concentration gradient across the membrane while ion channels allow for charge in the form of ions to pass through the membrane. The ion current flow causes changes in transmembrane potential and cardiac activation.

difference across the cell membrane by triggering a series of biochemical and biophysical events (described later). These depolarization events result in myocyte contraction. Depolarization propagates along excitable cell membranes, as in the transmission of an action potential along a squid axon, and from myocyte to myocyte through gap junctions. CELL MEMBRANE CHARACTERISTICS Cell membrane characteristics are major determinants of tissue excitability. The membrane of the cardiac myocyte is composed principally of phospholipids, cholesterol, and proteins.4 The membrane phospholipids have a charged polar headgroup and two long hydrocarbon chains arranged as shown in Figure 1-1. The cell membrane comprises two layers of phospholipids with their hydrophobic aliphatic chains oriented toward the central portion of the bilayer membrane and their polar headgroup regions toward the outside boundaries of the membrane. Because the membrane is composed of two layers of phospholipids, the polar regions of the phospholipid molecules interface with the aqueous environments inside and outside the cell. The lipid-soluble hydrocarbon chains are forced away from the aqueous phase to form a nonpolar interior. DETERMINANTS OF THE RESTING TRANSMEMBRANE POTENTIAL Relatively large gradients of individual ion concentrations exist across the cardiac cell membrane.5 The gradient of sodium ions (Na+) across the membrane is approximately 145 millimoles per liter (mmol/L) outside to 10 mmol/L inside. In contrast, the potassium ion (K+) concentration outside the cell is approximately 4.5 mmol/L, whereas the inside concentration is 140 mmol/L. In the absence of a cell membrane, both Na+ and K+ would rapidly move in a direction determined by the concentration gradient. The diffusion force tending to move K+ out of and Na+ into the cell is proportional to the concentration gradients of those ions. The potential energy attributable to the diffusion force (PEd) tending to move K+ out of the cell is given by the following equation:   [K + ]   PEd = RT  ln  + i   [1-1]   [K ]o   where R is the gas constant, T is the absolute temperature, ln is the natural logarithm operator, [K+]i is the concentration of potassium ion inside the cell, and [K+]o is the concentration of potassium ion outside the cell. If the ratio of [K+]i to [K+]o is large, the potential energy across the membrane is large. For each ion species, the difference in concentration between the inside and the outside of the cell results in that ion’s contribution to the difference in electric potential across the cell membrane. In resting cardiac cells, the intracellular cytoplasm has a measured potential of about −90 millivolts (mV), relative to the extracellular

6

SECTION 1  Basic Principles of Device Therapy

fluid. This electric force tends to move positively charged ions such as K+ and Na+ to the inside of the cell and negatively charged ions such as chloride (Cl−) to the outside of the cell in proportion to the potential gradient. The potential energy attributable to the electric force (PEe) tending to move K+ into the cell is expressed as follows: PE e = zFVm



Na+

K+

A

Protein molecules embedded within the cell membrane have numerous functions, including those of being ion channels and signal transducers. The concept of ion channels was proposed in the 1950s by Hodgkin and Huxley.7 However, it was not until the introduction of the patch clamp technique by Neher and Sakmann in 1976 that the properties of these channels could be directly studied.8,9 There are two basic types of ion channels, distinguished by the factors that control opening and closing of the channel. Ion channels at muscle fiber end plates are chemically gated by specific transmitters. The opening of these channels is triggered by the binding of acetylcholine, and their closing is induced by its unbinding. In neuronal axons, conduction is mediated by faster, voltage-gated channels. These channels respond to differences in electric potential between the inside and outside of the

CACNA1C

INa/Ca

NCX1

0

i

Ion Channels

Ca2+

Probable gene SCN5A

ICa.L

Using known values for extracellular K+, Vm(K+) = −90 mV. When Equation 1-3 is solved using Na+ concentrations, a Vm(Na+) of +50 mV is obtained. Therefore, it is the equilibrium potential for potassium ion (not sodium ion) that is the major factor responsible for the resting transmembrane potential. This suggests that the resting membrane is more permeable to K+ than to Na+. To calculate the transmembrane potential when multiple ionic species exist in different concentrations across the membrane, the Goldman constant field equation (modified by Hodgkin and Katz6) is used: + + + + − − − RT   PK [K ]i + PNa [ Na ]i + PCl [Cl ]o + …  Vm =  ln  + + [1-4]  + + −  F   PK [K ] + PNa [ Na ] + PCl [Cl − ] + … where PK+, PNa+, and PCl− are the cell membrane permeabilities for the respective ions. At physiologic concentrations, this equation yields a transmembrane potential of −90 mV (the equilibrium potential for K+). Equation 1-4 describes how resting potentials vary as sodium and potassium ion concentrations are changed. Because there is a passive leak of charged ions through the membrane, the resting potential would not exist at the level of −90 mV unless it were actively maintained. This is accomplished by two active transport mechanisms that exchange Na+ ions for K+ and calcium (Ca2+) ions. One might ask: with all the potassium ions and other positively charged ions in the cell, and with a relatively small amount of negatively charged chloride ions in the cell, how can the interior of the cell be negative with respect to the outside? The answer is that an array of intracellular organic and inorganic anions inside the cell—molecules that do not cross the membrane—carry net negative charges sufficient to make the overall balance of charge negative.

K+

INa

 [K + ]  or, in log base 10 terms, Vm (K + ) = −61.5 log  + i  .  [K ]o 

o

Na+

[1-2]

where z here is the valence (the number of positive or negative electrical charges) of the ion, F is the Faraday constant (96,500 coulombs/ equivalent), and Vm is the transmembrane potential difference (transmembrane voltage, measured in millivolts). During equilibrium, the total of the potential energies from diffusion and electric forces is zero, and no net ionic movement occurs. Therefore, the sum of Equation 1-1 and Equation 1-2 may be set to zero. This yields the Nernst equation, which describes (in measurable electrical units) the potential that must exist for a single ionic species, here K+, to be in equilibrium across the membrane of a resting cardiac cell:  [K + ]  Vm (K + ) = −26.7 ln  + i  [1-3]  [K ]o 

o

Ca2+

IK1

KCNJ2

Ito1

KCND

Ito2

?

IKr

KCHNH2IKCNE2

IKs

KCHQ1IKCNE1

IKp

KCNK?

B Figure 1-2  Ion channels underlie cardiac excitability. A, Key ion channels (and electrogenic transporter) in cardiac cells. K+ channels (green) mediate K+ efflux from the cell; Na+ channels (purple) and Ca2+ channels (orange) mediate Na+ and Ca2+ influx, respectively. The Na+/ Ca2+ exchanger (red) is electrogenic, because it transports three Na+ ions for each Ca2+ ion across the surface membrane. B, Ionic currents and genes underlying the cardiac action potential: top, depolarizing currents as functions of time, and their corresponding genes; middle, ventricular action potential; bottom, repolarizing currents and their corresponding genes. (From Marban E: Cardiac channelopathies. Nature 415:213-218, 2002. ©Nature Publishing Group, http://www.nature.com.)

cell, across the membrane. Voltage-gated channels for sodium, potassium, and calcium appear to operate in similar ways, sharing many of the same structural features. In addition, each type of channel can be subdivided into several subtypes with different conductance or gating properties (Fig. 1-2). VOLTAGE-GATED CHANNELS Voltage-gated channels open in response to an applied electric potential. The source of this voltage can be an action potential propagated from an adjacent cell or the electric field of an artificial pacemaker electrode. If depolarization of the membrane exceeds a threshold voltage, an action potential is triggered, resulting in a complex cascade of ionic currents flowing across the membrane into and out of the cell. As a result of this flow of charge across the membrane, the potential gradient across the membrane changes in a characteristic pattern of events that produce the cardiac action potential (Fig. 1-3).

1  Cardiac Electrical Stimulation

50 Phase 1 Phase 2

–50

Phase 0

0 Threshold

Transmembrane potential (mV)



Phase 3

Phase 4 –100 100

200

300

400

500

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

Selective membrane-bound proteins (ion channels) determine the passive transmembrane flux of an individual ion species. The transmembrane currents determine or influence cellular polarization at rest, action potential depolarization and repolarization, conduction, excitation-contraction coupling, and myofibril contraction. The channels that regulate transmembrane conductance of Na+ and Ca2+ are voltage gated. The sodium channel is a large protein molecule composed of approximately 1830 amino acids.10 It contains four internally homologous repeating domains, believed to be arranged around a central water-filled pore lined with hydrophilic (“water-loving”) amino acids. It is estimated that there are 5 to 10 Na+ channels per square micrometer (µm2) of cell membrane. When an alteration changes the membrane potential to about −70 to −60 mV (the threshold potential), four to six positively charged amino acids move across the membrane in response to the change in electric field. This causes a change in the conformations of the channel proteins, resulting in opening of the channel. After a single Na+ channel changes to the open conformation, about 104 Na+ ions enter the cell. On depolarization of the membrane, the Na+ channels remain open for less than 1 millisecond (msec). After rapid depolarization of the membrane, the Na+ channel again changes to the closed conformation. In addition to the Na+ channel, specialized proteins are suspended in the cell membrane that have differential selectivity for K+, Ca2+, and Cl− ions, with much different time constants for activation and inactivation. SINOATRIAL AND ATRIOVENTRICULAR NODAL CELLS In contrast to Purkinje fibers and working myocardial cells, sinoatrial (SA) and atrioventricular (AV) nodal cells are characterized by no true resting potential. Following repolarization, the transmembrane potential of the nodal cells reach a minimum of approximately −60 mV. The SA and AV nodal cells have a slow, inward, depolarizing, combined Na+/K+ current known as the “funny” current (If ). These If channels open and cause the transmembrane potential to rise slowly until Ca2+ channels are activated and an action potential is generated. The rate of

7

this rise in transmembrane potential is affected by factors such as autonomic nervous stimulation and hormones. In these nodal structures, depolarization is primarily mediated by inward Ca2+ conductance through specialized Ca2+ channels. There are two types of Ca2+ channels in the mammalian heart. The L-type channels are the major voltage-gated pathway for entry of Ca2+ into the myocyte, and they are heavily modulated by catecholamines.11 The T-type channels contribute to spontaneous depolarization of the cell associated with automaticity (pacemaker currents). The pore of the Ca2+ channel has a functional diameter of about 0.6 nanometers (nm), larger than that of the Na+ channels (0.3-0.5 nm).12 The selectivity for Ca2+ is high, up to 10,000-fold greater than that for Na+ or K+. The key elements are high-affinity binding sites for Ca2+, positioned along a single file pore. “Elution” of a Ca2+ ion occurs when another Ca2+ ion enters and is selectively bound. In normal situations, the SA node has the most pronounced automaticity and highest spontaneous firing rate in the heart. If the SA node is not functioning properly, however, or if block develops between the SA node and the AV node, the AV node will display pacemaker activity at a slightly slower rate than the SA node. In the absence of a pacemaker signal from the AV node, Purkinje fibers will activate at an even slower rate than either the SA or the AV node. MAINTENANCE OF RESTING MEMBRANE POTENTIAL The resting membrane potential is maintained by the pumping of Na+ ions out of the cell and K+ ions into the cell. The Na+,K+—adenosine triphosphatase (ATPase) pump moves three Na+ ions out of the cell in exchange for two K+ ions moved into the cell.13-15 The basic unit of the Na+,K+-ATPase protein (pump) consists of one alpha (α-) and one beta (β-) subunit. The α-subunit is large (1016 amino acids) and spans the entire membrane, whereas the β-subunit is a smaller glycoprotein. There appear to be about 1000 pump sites/µm2 of cardiac cell membrane. The fully activated pump cycles about 50 to 70 times per second (interval of 15-20 msec/cycle). Similarly, the Na+-Ca2+ pump moves three Na+ ions out of the cell in exchange for one Ca2+ ion.16,17 Therefore, both transport mechanisms result in the net movement of one positive charge out of the cell, polarizing the membrane and maintaining a negatively charged interior. The function of both exchange mechanisms depends on the expenditure of energy in the form of high-energy phosphates and is susceptible to interruptions in aerobic cellular metabolism (e.g., during ischemia). THE CARDIAC ACTION POTENTIAL When the voltage gradient across the membrane of a myocyte decreases so that the inside of the cell becomes less negatively charged with respect to the outside of the cell, a critical transmembrane voltage difference is reached, the threshold voltage. At threshold, the cell membrane suddenly undergoes a further depolarization that is out of proportion to the intensity of the applied stimulus. This abrupt change in the potential across the membrane is the start of a cascade of inward and outward currents that together are known as an action potential.18 The cardiac action potential is an enormously complex event and consists of five phases19 (see Fig. 1-3): phase 0, the upstroke phase of rapid depolarization; phase 1, the overshoot phase of initial rapid repolarization; phase 2, the plateau phase; phase 3, the rapid repolarization phase; and phase 4, characterized by a slow, spontaneous depolarization of the membrane in cells with spontaneous pacemaker activity, until the threshold potential is again reached and a new action potential is generated. Phase 0: Rapid Depolarization The upstroke of the action potential is triggered by a decrease in the potential gradient across the membrane to the threshold potential of −70 to −60 mV. On depolarization of the membrane to the less negative threshold voltage, the Na+ channels open, resulting in an influx of

8

SECTION 1  Basic Principles of Device Therapy

positively charged ions (the inward Na+ current) and rapid reversal of membrane polarity. The rate of depolarization in phase 0 ranges from 800 volts per second (V/sec) in Purkinje cells to 200 to 500 V/sec in atrial and ventricular myocytes. In these cells, the inward Na+ current is primarily responsible for phase 0 of the action potential. In SA and AV nodal cells, where the inward Ca2+ current predominates, the upstroke velocity of phase 0 is much lower (20-50 V/sec). Phase 1: Initial Repolarization After voltage-dependent activation of the Na+ current in phase 0, the membrane potential rapidly changes from negative to positive. The increased conductance of Na+ is rapidly followed by voltage-dependent inactivation. Phase 1 is characterized by the transient outward K+ current (IKto). The outward movement of K+ is a major contributor to the various repolarization phases. It is complex and has a number of discrete pathways.20,21 Most K+ currents demonstrate rectification, that is, decreased K+ conductance with depolarization. The K+ currents include the instantaneous inward rectifier K+ current, the outward (delayed) rectifier K+ current, the transient outward currents, and ATP-, Na+-, and acetylcholine-regulated K+ currents. The initial repolarization, however, is mainly the result of activation of a transient outward K+ current and inactivation of the fast inward Na+ current.22 The transient outward K+ current has two components, one voltage gated and the other activated by a local rise in Ca2+. Phase 2: Plateau The net current during the plateau phase is apparently small, although the individual currents (inward Na+ and Ca2+ and outward K+) are each about an order of magnitude larger.23 Among the inward currents are the slowly activating Na+ current, a Ca2+ current, and an Na+-Ca2+ exchange current. Outward currents include a slowly activating K+ current (IKs), a Cl− current, a more rapidly activating K+ current (IKr), an ultra-rapidly activating K+ current (IKur), and the Na+-K+ electrogenic pump. During phase 2 of the action potential, the absolute refractory period, the cardiac cell cannot be excited by an electrical stimulus, regardless of its intensity. Phase 3: Final Repolarization Deactivation of inward Na+ and Ca2+ currents occurs earlier than for the K+ currents, favoring net repolarization of the membrane. When the membrane is sufficiently repolarized, an inward K+ rectifier current is progressively activated, resulting in a regenerative increase in outward currents and an increasing rate of repolarization. Repolarization is also accomplished by the function of the Na+,K+-ATPase pump. The membrane potential eventually becomes stable, so that inward and outward currents are again in balance and the resting potential reestablished. Between the end of the plateau phase and full repolarization, the cell is partially refractory to electrical stimulation. During this period, the relative refractory period, a greater stimulus intensity is required to generate an action potential than required after full recovery of the resting membrane potential. For clinical measurement of myocardial refractoriness, the effective refractory period, stimulation of the heart is usually performed at twice the threshold current, as determined during late diastole. Phase 4: Automaticity and the Conduction System Automaticity is the property of certain cells by which they are able to initiate an action potential spontaneously. It has been known for centuries that the heart can exhibit spontaneous contraction even when completely denervated. Leonardo da Vinci observed that the heart could “move by itself.”24 William Harvey reported that pieces of the heart could “contract and relax” separately.25 Many cells within the specialized conduction system have the potential for automaticity. Not all parts of the heart, however, possess this property. In fact, cells in different areas of the heart have different transmembrane potentials, thresholds, and action potentials. Fast responses are characteristic of ordinary working ventricular muscle cells and His-Purkinje fibers, with resting membrane potentials of −70 to −90 mV and rapid

conduction velocities. Normal SA and AV nodal cells have slow responses, with resting potentials of −40 to −70 mV and slow conduction velocities. Cells, or group of cells, with the fastest rate of spontaneous membrane depolarization during phase 4 are the first to reach threshold potential and initiate a propagated impulse. Therefore, cells with the steepest slope in phase 4 become the heart’s natural pacemaker. Ordinary working myocardial cells usually are not automatic. Normally, depolarization is initiated at the SA node.26,27 Figure 1-4 shows action potentials from different types of cardiac cells.28 Rather than maintaining a stable resting membrane potential, the repolarization of the action potential is followed by a slow depolarization from about −71 to −54 mV, the threshold required to initiate another action potential. This slow, spontaneous depolarization drives cardiac automaticity and is related to a specialized current (If ). In the case of AV nodal cells, the fast upstroke is carried predominantly by an inward Ca2+ current. Repolarization is caused by delayed activation of the K+ current. The balance of inward and outward currents determines the net “pacemaker” current and is finely regulated by both adrenergic and cholinergic neurotransmitters. In the presence of AV block or abnormal SA nodal function, AV junctional cells in the region of the proximal penetrating bundle usually assume the role of pacemaker at rates slower than that of the sinus node. In the absence of disease in the AV junction, the escape rhythm occurs with a frequency that is about 67% of the sinus rate.29

Artificial Electrical Stimulation of Cardiac Tissue Artificial lipid membranes in their pure form are electrical insulators. The myocyte cell membrane (sarcolemma) is much more complicated. Specialized protein molecules in the membrane allow it to be conductive.30,31 These proteins, either singly or in certain groupings, form channels that open and close for transport of specific ions through the membrane in response to particular stimuli. The channel proteins are the end stages of processes that provide both active and passive transport of ions and molecules through the membrane. When application of a pacemaker or defibrillator pulse produces a local electric field gradient, ion drift in the extracellular fluid at that site, as well as ion flow within the cell and within the membrane, are affected by the field. The field-induced ion drift cannot be uniform within and outside the cell because of the different drift properties of different ion types, different ion and protein concentrations within and without the cell, and the barrier impedance effect of connections between cells. The effect of the stimulus is to change the transmembrane voltage of nearby myocytes sufficiently so that depolarization begins in and spreads from these myocyte membranes. Propagation of the stimulus to nearby myocytes occurs because the local transmembrane depolarization changes the voltage gradient across adjacent membranes sufficiently to trigger depolarization of those membranes. To cause an action potential to spread throughout the whole heart, an electrical pulse must stimulate a minimum of approximately 50 closely coupled cardiac cells.5 The result is a self-regenerating action potential that progresses in a wavelike, relatively slow manner beyond the local effect of the pacemaker stimulus. Away from immediate vicinity of the electrode, transmission of depolarization and its velocity depend in part on the resistance and capacitance properties of the membrane, on the opening and closing of ion channels, and on ion flows through the sarcolemma. MYOCARDIAL CELL ELECTRICAL PROPERTIES A single Purkinje fiber typically has an internal resistance that is two to three times greater than that of blood. The specific membrane resistance of a Purkinje fiber is on the order of 104 ohms–square centimeter (Ω-cm2). The time constant of the surface membrane is on the order of 10 msec, and the membrane capacity is about 1 microfaraday (µF)/cm2. Gap junctions are intercellular channels that provide a



9

1  Cardiac Electrical Stimulation

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

P

U

Q

S

Time (msec) 0 100 200 300 400 500 600 700 Figure 1-4  Action potentials from different cell types in the heart. Specialized pacemaker and conduction cell types have distinct action potential characteristics. Activation times for a normal heartbeat demonstrate the spread of activation throughout the whole heart. AV, Atrioventricular. (From Malmivuo J, Plonsey R: Bioelectromagnetism: principles and applications of bioelectric and biomagnetic fields. New York, 1995, Oxford University Press.)

SINGLE-CELL EXCITATION BY ELECTRICAL STIMULUS The basic principle underlying cardiac pacing is the ability of an externally applied voltage difference to generate a self-propagating wavefront of cardiac depolarization in the heart. This wavefront is the result of propagation of action potentials across the myocardium, a process that involves the opening and closing of ion channels in the membrane. In reality, an applied external electrical stimulus initiates cardiac excitation as a result of passive effects on the transmembrane potential. This is a result of a change in potential that is induced in the extracellular space. Because of the relatively high impedance between cells across the membrane and through the gap junctions compared with the lower impedance path through the extracellular space, relatively small amounts of current pass through the cell membrane. When an external stimulus is applied, the potential within the intracellular space does not vary appreciably over the length and width of a cardiac myocyte (Fig. 1-5). However, near the stimulating electrode, the extracellular stimulus causes a field gradient to develop in the extracellular space. The transmembrane voltage (Vm) is equal to the difference in potential inside the cell (Vi) minus the potential outside the cell in the extracellular space (Ve), so a change in extracellular potential with little change in intracellular potential across the length of the cell results in a net difference in Vm. This is represented as follows:

Vm = (Vi − Ve )

Thus, as considered along the length of a cardiac myocyte, one end has a more negative transmembrane potential than the other. This leads to a net hyperpolarization of the cell on one end and depolarization on the other end. The induced change in Vm has the capacity to open voltage-gated ion channels within the cardiac cell membrane. Voltagegated Na+ channels are opened by a decrease in Vm. If an external

Vi Vm+

Vm–

A

+

–

Ve Vm

Voltage

pathway for electrical communication between myocytes. Their diameter is about 2 nm and length about 12 nm. Gap conductance is voltage sensitive.32 Under pathologic conditions, such as severe hypoxia or ischemia, gap junctions will not function normally.

B

Vi Ve

Figure 1-5  Response of a cardiomyocyte to electric field. Relatively little current passes through the relatively high impedance cell membrane. A, Cardiomyocite exposed to extracellular electric field. The intracellular voltage (Vi) is relatively low impedance and thus does not vary from one side of the cell to the other while the extracellular voltage (Ve) drops from one end of the cell to the other. B, Transmembrane potential (Vm) is defined as Vm = Vi − Ve, so the left side of the cell is hyperpolarized by the external field while the right side of the cell is depolarized. If the extracellular field gradient is large enough, the rightsided portion of the cell may reach threshold and initiate an action potential that may activate the entire cell and spread to adjacent cells through gap junctions. (From Dosdall DJ, Fast VG, Ideker RE: Mechanisms of defibrillation. Annu Rev Biomed Eng. 12:233-258, 2010.)

10

SECTION 1  Basic Principles of Device Therapy

stimulus is strong enough to cause Vm on the depolarized side of the cell to rise above the threshold to activate the inward Na+ channels, the influx of Na+ ions rapidly leads to further depolarization of the cell and initiates an action potential. This action potential spreads to adjacent cells by passing ions through gap junctions and raising Vm in adjacent cells, and initiating an active response through the activation of Na+ channels. FACTORS DETERMINING CAPTURE BY AN ELECTRICAL STIMULUS In order for an electrical stimulus to stimulate (capture) myocardium, it must be applied with sufficient amplitude, for a sufficient duration (measured in milliseconds), at a time when the myocardium is electrically excitable. Many clinical factors determine whether a stimulus of given amplitude will result in capture, including proximity of the electrode to the myocardium, pathology of the underlying cardiac tissue, size and shape of the electrode, and effects of drugs and hormones, as well as electrolytes. For routine clinical practice, the most important factor is that the lead must be positioned close to well-functioning myocardium in a secure manner. (It is also important that the lead does not stimulate the diaphragm or the phrenic nerve, that it is at a site that results in good hemodynamic function, and that the location does not allow the electrode to bump other electrodes.) After adequate lead positioning, the strength-duration relation is the next most important consideration. CURRENT DENSITY, ELECTRIC FIELD GRADIENT, AND PROPAGATION OF DEPOLARIZATION There are two approaches to thinking about how an electrical stimulus induces a self-propagating wavefront of depolarization within myocardium. The current density approach considers the magnitude of current flowing through a given mass of myocardium between the stimulating electrodes and finds this to be the critical factor required to induce a regenerative wavefront of depolarization. From this point of view, the stimulation threshold is a function of current density (amperes/cm2) in the excitable myocardium underlying the electrode.33-36 The electric field approach holds that the critical factor affecting myocardial depolarization is the magnitude of the electric field gradient (volts/cm in viable tissue) that is induced in the myocardium beneath the stimulating electrode.37,38 These approaches are fundamentally the same in that there is a mathematical relationship between the electric field gradient and the current density at the stimulating electrode. Because reactance as well as resistance is present, Ohm’s law in this context must be stated in terms of impedance, as follows:

v = iz

[1-5a]

where v is the stimulus voltage, i is the current, and z is the impedance to current flow. Note that each of these is a vector. Also note that z varies with current density at the electrode (because of interface properties), with direction of current flow (anisotropy), and with domain (extracellular or intracellular). The total energy of the pacing stimulus is determined by the applied voltage, the current, and the duration of the stimulus, as follows:

t

J t = ∫ v t i t dt 0

[1-5b]

where Jt is the energy delivered (expressed in joules) from time zero to time t during the pulse, vt is the voltage, and it is the current at the electrode at instantaneous time t during the pulse. This equation indicates that the total energy delivered during the pulse is proportional to the area under a curve. The curve is formed in the vertical dimension by the instantaneous product of voltage and current applied to the electrode, and in the horizontal dimension by the time elapsing during the pulse. When the time t reaches the programmed pulse duration, the pulse ends. The area under the curve then represents the electrical energy transferred during the time span of the pulse. The equation can easily be solved in real time with a small digital computer, provided

that continuous phasic measurements of the voltage and current in the wire going to the electrode are available. (Note that this equation has blood flow and pressure energy analogs.39) For pacemakers in almost all clinical situations, it is more practical and very reasonable to calculate the energy delivered by a constantvoltage pulse through a pacemaker electrode by means of an approximation. The assumptions are that (1) the voltage during the pulse really is constant and (2) the current flowing is linearly related to the voltage. Because neither is quite true, the questions become (1) whether the information obtained proves useful and (2) to what degree it can be misleading. If the assumptions are close to being correct, the following equation can be approximately correct and useful: V V2 t= t, R R where J is the energy delivered, V is the “constant” voltage, I is the current, R is the resistance, and t is the stimulus duration. This equation indicates that the total energy delivered can be estimated by multiplying the voltage reading displayed on the pacing systems analyzer (PSA) by the current reading on the PSA and multiplying that product by the stimulus duration. Alternatively, the voltage reading can be squared, then divided by the resistance reading, and the quotient multiplied by the stimulus duration. The virtue in this calculation is that ordinary equipment can provide a quick, approximate determination of the total energy delivered per stimulus. In clinical practice, however, the usual range of good or acceptable values for pacing threshold current and voltage is known, and there is rarely a need to calculate the delivered energy. The stimulation threshold of isolated cardiac myocytes depends on their orientation within an electric field. The threshold is lowest when the myocytes are oriented parallel to the field and highest when the axis of the myocytes is perpendicular to the field.40 It is clear that myocardial stimulation may be induced with anodal or cathodal stimulation, or both, although with somewhat different characteristics. Various investigators have used several different parameters to express stimulation threshold, including current (mA), potential (V), energy (J), charge (Q), pulse width (t, msec), and voltage multiplied by stimulus duration (V-sec).41-47 For the purposes of this chapter, the myocardial stimulation threshold for pacing is defined as the minimum stimulus amplitude at any given pulse width required to consistently achieve myocardial depolarization outside the heart’s refractory period.48 Stimulation thresholds measured with a constant-voltage (CV) generator are stated in volts, and those of a constant-current (CI) generator are stated in milliamperes (mA). Note that a true constant-voltage generator may yield a slightly different stimulation threshold than the pseudoconstant-voltage generators in ordinary use. Why must an electrical pacing stimulus rely on propagation in myocardium rather than directly exciting the entire heart? With the magnitude of a stimulus generated by a pacemaker pulse generator, the electric field gradient and ion current density near the electrode are great enough to trigger an action potential only extremely close to the electrode. Unless propagation of the local depolarization occurs, no cardiac contraction will result. In contrast to cardiac pacing with direct production of only local depolarization and dependence on self-propagation of depolarization, defibrillation involves direct depolarization or hyperpolarization of a large portion of the heart. This is necessary to provide at least the minimum local voltage gradient required to produce depolarization over major portions of the myocardium. To do so requires a stimulus intensity that is much greater than that required for pacing49 (see Chapter 2 for a thorough discussion of defibrillation).

J = VIt = V

ELECTRIC POTENTIAL GRADIENTS AND CURRENTS FOR STIMULATION AND DEFIBRILLATION When the electric field exceeds approximately 1 V/cm in the extracellular space during diastole, myocardial stimulation (capture) results. The electric field must be achieved only in the area of several dozen



cells and may be achieved typically with less than 0.5 mA. When the electric field strength is increased to 6 V/cm, ventricular fibrillation may occur if the stimulus is applied during the vulnerable period (approximately the peak of the T wave). An area where the voltage gradient is at least 6 V/cm must be achieved at a distance of 1 cm from the electrode. To accomplish this, a current of approximately 20 mA is required to initiate ventricular fibrillation. This same field strength (6 V/cm) is also required to interrupt ventricular fibrillation (defibrillation). However, defibrillation requires the minimum field intensity to be approximately 6 V/cm at almost all points in the myocardium. In order to achieve this minimum electric field gradient at the same instant over the entire heart, a very large shock current, typically more than 10 A, must be applied. This current is thousands of times greater than the current required for pacing.49 When a pacemaker pulse is applied, electrically excitable myocytes respond with a wave of depolarization followed by repolarization in the myocardium. Depolarization of a small, local group of myocytes begins the self-propagating process. The initial depolarization adjacent to the electrode produces a potential gradient great enough at neighboring myocytes to result in their depolarization. The process is a self-regenerating mechanism that requires time to spread, and once established, it is largely independent of the distance from the stimulating electrode. Defibrillation is largely accomplished by providing an extremely large potential gradient between the defibrillation electrodes. This produces, at one instant within local myocardium everywhere in the heart, at least a 6 V/cm gradient. In contrast, pacing is accomplished by providing a gradient of approximately 1 V/cm or less at a local site and relies on self-propagation to spread the depolarization process throughout the myocardium. VIRTUAL ELECTRODE EFFECT Virtual electrodes are important because they can serve as sites that initiate or prevent depolarization of myocytes.50 Newton and Knisley51 defined virtual electrodes as “experimentally observed regions of large delta Vm that arise distant from the stimulating electrode.” “Delta Vm” (ΔVm) here represents the change in voltage difference across the myocyte membrane during application of the stimulation current. A virtual electrode can be described as a collection of charge predominately of one sign at a site away from a regular electrode. If, in an electrically neutral solution, ions of one charge sign are moved away by an electric field, a relative excess of ions of the other charge sign will be left or will move in the opposite direction. For example, if the regular electrode is paced with a negative-going stimulus, a site elsewhere in the tissue (i.e., not at this physical electrode) may become transiently positive and alter the transmembrane potentials of cells at that site.52-55 In cardiac tissue, a virtual electrode can occur as an effect of anisotropy after a defibrillation shock and can initiate refibrillation. A virtual electrode may also result from ion redistribution patterns in a nonanisotropic medium. In anisotropic tissue, charge redistribution produced by a physical electrode can have a “dog bone” shape.56 The virtual electrodes occur at sites where ions flow into or out of cells by crossing the cell membrane. An applied unipolar stimulation current flows into some cells while simultaneously flowing out of others, thus producing negative and positive virtual electrodes at different locations. In an anisotropic medium, a cathodal pulse produces a virtual electrode with a dog bone shape oriented perpendicular to cardiac fibers and containing a large ΔVm, and a pair of virtual electrodes containing negative ΔVm at locations along the fibers. Theoretical models have shown that this effect can be attributed to unequal tissue anisotropy in the intracellular and extracellular spaces; that is, intracellular current at a given location favors the longitudinal direction 10 to 1, whereas extracellular current favors the longitudinal direction only 3 to 1 (Knisley, personal communication, 2005). Knisley and Pollard57 studied the effects of electrode-myocardial separation on cardiac stimulation of rabbit hearts in conductive solution. The

1  Cardiac Electrical Stimulation

11

electrode-myocardial separation altered the spatial distribution of ΔVm and increased the pacing threshold. In regard to electrode-myocardial tissue separation, Shepard et al.,58 during 1980s transthoracic defibrillator implantation procedures, sewed the electrodes used for rate sensing onto the outer surface of the pericardium. Sensing voltages were normal. Pacing thresholds of these electrodes measured were high (3.7 ± 1.9 mA and 4.5 ± 2.19 V at 0.5msec stimulus duration), as would be expected both from current density considerations and from Knisley’s findings. The initial impedance was 1209 ± 383 Ω, and the chronic impedance was 1550 ± 358 Ω at a median follow-up time of 964 days. The thresholds had by then decreased to 3.8 ± 2.07 V and 2.7 ± 1.8 mA. The transpericardial distance from underlying myocardium did result in high initial pacing thresholds (and in the special distribution of ΔVm near the electrodes). However, the long-term tissue reaction of pericardium and underlying myocardium to the presence of these electrodes was not detrimental in terms of threshold evolution. Virtual electrodes normally exist near an electrode when a pacing pulse is applied. Nikolski and Sambelashvili,59 studying Langendorffperfused rabbit hearts, found that stimuli of magnitude five times threshold produced “make” or stimulus-onset excitation from virtual cathodes, whereas near-threshold stimuli produced “break” or stimulus-termination excitation from virtual anodes. In studying how cardiac tissue damage at an electrode results in a pacing threshold increase, Sambelashvili et al.60 found that the virtual electrode effect was destroyed by very strong (40 mA, 4 msec, biphasic, 240/min) pacing pulses applied for 5 minutes. Fluorescent optical mapping showed that decrease or loss of the virtual electrode polarization was associated with pacing threshold increase. Propidium iodide staining showed tissue damage within an area about 1 mm in diameter surrounding the electrode. Another view of the effect of strong stimuli as just described is that local tissue damage at a pacing electrode increases the distance from the electrode to normal myocytes. Because electric field strength and current density decrease with distance from the electrode, an increase in pacing stimulus amplitude applied to the electrode is necessary to restore the stimulus current density to the pacing threshold level of myocytes at the outer edge of the damaged tissue. ANODAL AND CATHODAL STIMULATION The polarity of a stimulus can be cathodal or anodal. In determining how an artificial electrical stimulus interacts with cardiac cells to initiate this self-propagating wavefront, the mechanism of excitation of cardiac tissue by an electrical stimulus depends on the polarity, strength, and duration of the stimulus. Cardiac excitation may occur immediately adjacent to the electrode by direct stimulation or at a distance from the electrode through virtual electrode effects. The two mechanisms by which electric current stimulates cardiac excitation close to electrodes are cathodal make and anodal break. Virtual electrodes cause cardiac excitation by cathodal break and anodal make phenomenon. Cathode make excitation occurs adjacent to a cathode from the high current density immediately surrounding the electrode. Current flows in the extracellular space and causes hyperpolarization and depolarization of opposite ends of the cell, as shown in Figure 1-5. Current also passes through the cell membrane and causes an increase in Vm that exceeds the threshold for fast Na+ channels to activate. Excitation spreads outward from the central excitation site in an elliptical shape because of tissue anisotropy (Fig. 1-6). Cathode make occurs on the rising edge of the stimulation pulse (Fig. 1-7, A). An anode tends to hyperpolarize cells immediately adjacent to the electrode, but virtual cathodes form on both sides of the central anode. Anode make excitation causes activation to spread on both sides of the central hyperpolarized region until the excitation merges into a single wavefront and spreads elliptically (Fig. 1-7, B). The threshold for stimulation with cathode make is typically higher than for anode make because the virtual cathodes that form at a distance from the central

12

SECTION 1  Basic Principles of Device Therapy

Cathodal make 0 – mV –100 Anodal break

0 + mV –100 Figure 1-6  Action potentials are initiated by cathodal and anodal excitation. Excitation make and break phenomena occur under and near electrodes. Cathodal make occurs when the excitation current directly excites the tissue under the electrode. Anodal break occurs under the excitation electrode on termination of the stimulating pulse. (From Antoni H: Electrical properties of the heart. In Reilly JP, editor: Applied bioelectricity: from electrical stimulation to electropathology. New York, 1998, Springer, p 160.)

0

3

6

anode are weaker and have less depolarizing power than the direct effects of a central cathode. Cathode break excitation occurs when the stimulating pulse is turned off (Fig. 1-7, C). While the stimulation pulse is turned on, the tissue under the stimulating electrode is depolarized, and the fast Na+ channels open and then become unexcitable. Propagation of the excitation is repressed somewhat by the virtual anodes on either side. On termination of the stimulating pulse, the virtual anodes disappear, and excitation spreads from the active tissue under the electrode to the hyperpolarized regions under the virtual anodes. Because the tissue under the virtual anodes was hyperpolarized, the Na+ channels are fully excitable, and the excitation wavefront spreads rapidly through the area previously covered by virtual anodes. Cathode make occurs at a lower threshold than cathode break and is not observed unless tissue is refractory when the stimulation pulse commences but becomes excitable during the pulse. If the tissue has recovered by the time the pulse turns off, cathode break may occur. If the tissue is unexcitable surrounding an anode at the beginning of a stimulation pulse, anode break excitation may result from a mechanism similar to that of cathode break. Excitation may be initiated by the termination of the excitation pulse (Fig. 1-7). Hyperpolarization directly below the anode causes the Na+ channels to become excitable. Depolarization from the virtual cathodes then rapidly propagates into the hyperpolarized region under the anode (Fig. 1-7, D). The virtual cathodes are relatively weak compared to the hyperpolarizing effect of the anode, but the activation spreads from the depolarized region into the hyperpolarized region because the cell membrane has a nonlinear response by which hyperpolarization decays more rapidly than does depolarization. Because anode break relies on virtual cathodes to initiate excitation in recently recovered tissue, however, this mechanism of excitation has the highest threshold of all, the make or break excitation mechanisms.

9

12

15

CM

A

AM

B

15 mV

–85 mV

CB

–185 mV

C

AB

D Figure 1-7  Computer simulation of four types of excitation with stimulating electrode (black rectangle). Cathode make (A), anode make (B), cathode break (C), and anode break (D) excitation are shown at various times; columns are time (t) in milliseconds. In A and B, t = 0 when the stimulation pulse turns on, and in C and D, t = 0 when the stimulation pulse is turned off. (From Wikswo JP, Roth BJ: Virtual electrode theory and pacing. In Efimov IR, Kroll MW, Tchou PJ, editors: Cardiac bioelectric therapy: mechanisms and practical implications. New York, 2009, Springer Science, pp 283-330.)



1  Cardiac Electrical Stimulation

Anode and cathode make and break phenomenon have been detected in computer models and verified in animal experiments.54,61 Optical mapping techniques with voltage-sensitive dye have been particularly useful in illuminating these mechanisms of stimulation.

Strength-Duration Relationships CHRONAXIE, RHEOBASE, ENERGY, AND PULSE DURATION THRESHOLDS





Irh

I=

[1-5c]

−t

threshold in terms of pulse duration can be misleading if the stimulus amplitude setting is omitted or unknown. In 1892, Hoorweg66 used a voltage source, a galvanometer, and lowleakage capacitors to conduct quantitative stimulation studies. He found that the voltage at which a capacitor must be charged to cause depolarization of nerves and muscles is an inverse function of the capacitance of the capacitor, as follows: Vc = aR + b / c



The intensity of an electrical stimulus (measured in volts or milliamperes) that is required to capture the atrial or ventricular myocardium depends on the duration of the stimulus.62,63 The historical background and the relations between the various electrical factors have been reviewed, further studied, and clearly stated by Blair,64 and by Geddes65 for electrode-electrolyte interface function and models. The interaction of stimulus amplitude and stimulus duration (pulse width) defines the strength-duration curve (Fig. 1-8). Lapique described this exponential relationship of stimulus intensity (current), I, and pulse duration, t, as follows:

1− e τ

where Irh is the rheobase current, e is the mathematical constant 2.718, and τ is the membrane time constant. The voltage or current amplitude required for endocardial stimulation has an exponential relation to the pulse duration, with a relatively flat curve at durations of longer than 1 msec and a rapidly rising curve at durations of less than 0.25 msec. Because of this fundamental property, a stimulus of short pulse duration must be of much greater intensity to capture the myocardium than a longer-duration pulse. Conversely, increasing the pulse width to longer than 1 msec has minimal influence on the intensity of the stimulus that is required for capture. Therefore, if one defines the stimulation threshold in terms of pulse amplitude without also specifying the pulse duration, important information is neglected. Similarly, specifying the capture

Threshold (V, µJ, µC)

x

Energy

4 x x

3

x x

x

x

x

x

x

Charge

x

2 Potential

Rheobase

1 Chronaxie 0 0

.2

.4

.6

.8

1.0

1.5

Pulse width (msec) Figure 1-8  Relationships between chronic ventricular canine constantvoltage strength-duration curves expressed in terms of potential ( V ), charge (µC), and energy (µJ) for a tined unipolar lead with an 8-mm2, polished, ring-tip electrode. Thresholds are measured at the point of gain of capture. Rheobase is the current or voltage threshold to the right that is independent of pulse width. Chronaxie is the pulse width at twice the rheobase. (From Stokes K, Bornzin G: The electrodebiointerface [stimulation]. In Barold SS, editor: Modern cardiac pacing. Mt Kisco, NY, Futura, 1985, pp 33-78.)

[1-6]

In this experimentally determined equation, Vc is the threshold voltage to which a capacitor of capacitance C must be charged to produce stimulation on discharge. R is the resistance of the circuit through which the capacitor is discharged, and a and b are coefficients that vary with the specimen (tissue) tested. Here, the experimentally determined constant a has the dimension amperes and the constant b has the dimension ampere-seconds. (The capacitance can be derived from the equation

C=

Q V

where C is the capacitance, Q is the charge on either conductor, and V is the magnitude of the potential difference across the capacitor; the capacitance has the dimension ampere-seconds [or coulombs] per volt; 1 coulomb per volt = 1 farad). Hoorweg determined that there was only one specific capacitance value for which the threshold charge was a minimum. He also determined that the threshold charge was a linear function of the stimulus duration, “intersecting the y-axis above zero.” However, Hoorweg did not have the capability to measure thresholds at very short pulse durations. In reality, the threshold charge increases toward infinity as the pulse duration approaches zero. In 1901, Weiss67 reported that the threshold charge required for stimulation increases linearly with stimulus duration. He called this relationship the “formule fondamentale.” If I represents the magnitude of the current, t is the stimulus duration, and a and b are constants determined by analyzing the data, Weiss found that: t

t

5

13

∫ Idt = at + b 0

[1-7]

Or, because ∫ idt = Q, the charge in the capacitor at threshold stimu0 lus duration t, Q = at + b. For a constant-current stimulus, this can be stated in words as threshold charge requirement = a amperes times t seconds + b, where b is a constant with dimensional units of ampereseconds. The values of the constants a and b vary with the tissue tested. The left-hand side of Equation 1-7 indicates that the charge delivered into the electrode is, for a constant current, the magnitude of the current multiplied by the duration of the current. The right-hand side states that the charge required to stimulate at threshold is a minimum of b ampere-seconds plus the product of the current level a and the stimulus duration t. Weiss67 noted that, for various stimulus durations tested, the quantity of charge required to initiate depolarization remained constant. The statement that the threshold charge requirement does not change with stimulus duration is true only within a limited range. The limitation leads to the concept of rheobase, as stated in 1909 by Lapique;68 when pulse duration is increased beyond a limited range, the current requirement does not decrease further. The charge delivered continues to increase as pulse duration is increased. Lapique called this minimum current the rheobase, or the lowest stimulus current that continues to capture the heart when the stimulus duration is made extremely long. In this situation, further increases in stimulus duration no longer reduce the magnitude of current required to stimulate the heart. Note from Equation 1-7 that the ratio a/b has the dimensions/ amperes divided by (ampere-seconds); this expression reduces in dimensions to the reciprocal of the stimulus duration. Lapique called the time seen in the a/b ratio the chronaxie, specified in seconds. Chronaxie time experimentally turns out to be approximately the threshold pulse duration at twice rheobase amplitude, and it has become defined

SECTION 1  Basic Principles of Device Therapy

as such; this is illustrated in Figure 1-8. Note again that rheobase is specified in terms of current, and chronaxie is specified in terms of time. Lapique69 determined that stimulation at the chronaxie pulse width approaches the minimum threshold energy. Therefore, the two most important reference points on a current or voltage strength-duration curve are rheobase and chronaxie. Obtaining a true rheobase typically requires pulse widths of 10 msec or greater. For clinical purposes, rheobase can be approximately measured at pulse widths of 1.5 to 2 msec. The value obtained may be slightly greater than the true rheobase. Therefore, the chronaxie stimulus duration obtained by setting the stimulus current at twice the approximate rheobase current and then finding the minimum stimulus duration that will result in capture is slightly low. In a time-saving, useful, and reasonable clinical sense, one may empirically set the stimulus duration at a value determined from experience, such as 0.4 to 0.5 msec and then determine the current and voltage thresholds. Safety factor allowances of current and voltage are then added or subtracted based on the patient’s current and projected clinical status (see later discussion). The goal is to find the combination of pacing threshold stimulus current, voltage, and pulse width that results in minimal charge drain from the pulse generator battery at normal pulse rates. Figure 1-8 shows that, for capture to be accomplished, the least charge was required at the shortest pulse duration (0.1 msec in the figure), but the least energy was required at about 0.3-msec stimulus duration. A very short pulse duration puts thresholds close to the steeply ascending limb of the voltage or current curve, where slight fluctuations can risk loss of capture. For this reason, Irnich70,71 recommended that chronaxie, or the pulse duration slightly to the left of chronaxie, is the most efficient pulse duration for both pacing and defibrillation. In most cases, the pulse duration at chronaxie or slightly greater appears to represent the best overall compromise between adequate safety and generator longevity. One also considers the expected stability or instability of the patient’s status, patient compliance, and follow-up arrangements, both short and long term. Pacing thresholds can also be specified in terms of only energy (microjoules, µJ) or only pulse duration (milliseconds). However, reliance on specifications in either of these units alone can be misleading. Doing so discounts that, regardless of the pulse duration or energy used, the heart cannot be stimulated unless the stimulus amplitude exceeds the rheobase current.72 For example, if rheobase is achieved with a pulse amplitude of 0.5 V and 10 msec, the pulse at this point on the strength-duration curve will have 5 µJ of energy (current × voltage × time) with a 500-Ω lead (1 mA × 0.5 V × 10 msec = 5 µJ). A 0.4-V, 20-msec pulse on a 500-Ω lead has 6.4 µJ of energy (28% greater than the “threshold” energy at 10 msec) but will not result in cardiac capture. Increasing the pulse duration to 100 msec at 0.4 V results in 32 µJ of energy (540% greater than “threshold” energy at 10 msec) but will still not capture the heart, for the same reason. No matter how far the pulse duration is extended (with increased energy), the myocardium will not be captured unless the stimulus amplitude is at least as great as the rheobase value. Therefore, calculation of the energy threshold at very long pulse durations does not provide clinically useful information.73 Coates and Thwaites74 studied the strength-duration curve in 229 patients with 325 leads. The mean atrial chronaxie (n = 101) was 0.24 ± 0.07 msec, and the mean ventricular chronaxie (n = 224) was 0.25 ± 0.07 msec. Mean atrial and ventricular rheobase values were 0.51 ± 0.2 V and 0.35 ± 0.14 V, respectively. Because the pulse generators were set at factory-nominal pulse durations of 0.45 to 0.50 msec, the authors concluded that pacing was suboptimal from the efficiency point of view. They pointed out that battery drain would be reduced by programming pulse duration to the chronaxie value and then programming the voltage to double the chronaxie value. In a study of excitability of rat atria during postnatal development up to 120 days, Gurjao de Godoy et al.75 found that atrial rheobase decreased with animal age and was altered by electric field orientation. Atrial chronaxie increased only with age. The clinician might keep in mind that the chronaxie is influenced not only by electrode material, size, and

stimulation mode, but also by clinically varying biochemical factors. Certainly, marked variations in pacing threshold, either up or down, do occur months and years after implantation in some children and adults.76 As stated earlier, on the left side of the chronaxie point of the strengthduration curve, the current and voltage rise rapidly as stimulus duration decreases. At pulse durations greater than chronaxie, the slope of the strength-duration curve gradually flattens. Small changes in stimulus amplitude are then less likely to result in loss of capture, especially if the threshold has increased because of a pathophysiologic condition. This increased safety comes at the cost of decreased battery life. PRACTICAL APPLICATION TO THRESHOLD MEASUREMENT To determine the strength-duration relationship, the clinician must measure the stimulation threshold at specific amplitudes and pulse durations. However, the result obtained varies somewhat with the method used. Usually the threshold, as measured at a specific stimulus duration (e.g., 0.5 msec), will be slightly higher when the stimulus amplitude is gradually being increased than being decreased. The threshold can also be measured by holding the output voltage constant and changing the stimulus duration. The stimulation threshold then is defined as the lowest-amplitude pulse duration that results in consistent capture of the myocardium. Only the threshold measured in this way can be clinically useful, as noted earlier. It also can be deceptive, because the strength-duration curve is not linear. For example, if the pulse duration threshold is 0.5 msec at an amplitude of 2 V, reprogramming the pacemaker stimulus to a pulse duration of 1 or even 1.5 msec would provide a very small margin of safety 10 9 8

Volts

14

7

D

6

2×V E

5 4

C

3 2

A

2×V 2 × PW

B

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

Pulse duration (msec) Figure 1-9  Programming of pulse amplitude and pulse duration based on analysis of the strength-duration curve in a patient evaluated at the time of pulse generator replacement (6 years after lead implantation). The rheobase voltage was 1.4 V, and the chronaxie pulse width (PW) was 0.30 msec. Note that the stimulation threshold, determined by decreasing the stimulus amplitude at a constant pulse duration of 0.5 msec, was about 2 V (point A). Doubling of the PW on the relatively flat portion of the strength-duration curve (point B) provides little safety margin for ventricular capture. In contrast, consider a threshold value on the steeply ascending portion of the strength-duration curve (point C). Doubling of the pulse amplitude doubles the safety margin on this portion of the curve, but it lies very close to the curve (point D). An appropriate setting for the chronic pacing pulse might be achieved by doubling the pulse amplitude from point A to point E. Also note that a similar programmed setting would have been obtained had the pulse duration been tripled from point C. Thus, the shape of the strengthduration curve has an important influence on the choice of the amplitude and duration of the pacing pulse.



1  Cardiac Electrical Stimulation 1.0

Effect of Pacing Rate on Stimulation Threshold

0.8 0.7 0.6 0.5

Increment

0.4

Decrement

0.3 0.2 0.1 0.0 2250 2000 1750 1500 1250 1000

750

500

250

Cycle length (msec) Figure 1-10  Pacing thresholds determined by gradually incrementing and decrementing the pulse amplitude until gain and loss of capture, respectively, are demonstrated in a patient with complete atrioventricular block. The pacing threshold was determined at cycle lengths of 2000, 1500, 1000, 750, and 500 msec and with a constant pulse duration of 2.0 msec. To prevent variation in cycle length during incrementing and decrementing pulse amplitudes, a backup pulse was delivered at 25 msec. Note that the threshold values determined in this manner, with increments and decrements of the stimulus amplitude, are similar. Therefore, the Wedensky effect may be marginal when the pacing cycle length is maintained at a constant value.

(Fig. 1-9). Instead, doubling the stimulus amplitude to 4 V at a pulse duration of 0.5 msec would provide an adequate margin of safety. In contrast, consider that, in this same patient, the pulse duration threshold is 0.15 msec at a pulse amplitude of 3.5 V. Increasing the pulse width to 0.45 msec would also provide an adequate safety margin. The reason that a threefold increase in pulse width is not adequate in the first example but is acceptable in the second relates to the location of the threshold stimulus on the strength-duration curve. Consideration of programming the pulse generator to twice the voltage threshold found at the chronaxie pulse duration is only a starting point. Determination of the safety margin necessary for any particular patient depends on several physiologic factors, as discussed next. Capture Hysteresis (Wedensky Effect) The threshold stimulus amplitude measured by decreasing the voltage or current until loss of capture occurs is sometimes less than that determined by increasing the stimulus intensity from below threshold until gain of capture occurs. This hysteresis-like phenomenon is the Wedensky effect,77 the effect of subthreshold stimulation on the subsequent suprathreshold stimulation when the stimulus amplitude is being increased (Fig. 1-10). Langberg et al.78 observed no demonstrable capture hysteresis at pacing cycles longer than 400 msec. They concluded that the Wedensky effect was related to asynchronous pacing in the relative refractory period when incrementing the stimulus intensity, versus the synchronous late-diastolic stimulation when decrementing the stimulus amplitude until loss of capture. Swerdlow et al.79 showed in a study of 40 patients that cardiovascular collapse occurs at AC current levels less than the ventricular fibrillation current threshold. The continuous capture threshold for AC current is less than the capture threshold for a single ordinary pacing stimulus. The authors suggested that continuous capture at low levels of AC current requires a cumulative effect of subthreshold stimuli; this is a variation of the Wedensky effect. The safety standard for 60-hertz (Hz) current leakage longer than 5 seconds should be 20 microamperes (µA) or less to avoid intermittent capture.

Hook et al.80 reported a significant increase in ventricular pacing threshold in 10 of 16 patients at 400 msec and in 15 of 16 at 300 msec (relative to a pacing cycle length of 600 msec). The phenomenon was not observed at every trial (e.g., 12 of 72 trials at 400 msec). The patients were all candidates for ICDs, and 9 of 12 were receiving antiarrhythmic drugs. The leads were bipolar: a pair of epicardial corkscrews in 11 patients and endocardial screw-in lead in five patients. The atrial stimulation threshold has also been shown to vary as a function of pacing rate.81,82 Katsumoto et al.83 reported that 29 of 36 patients exhibited constant-current atrial pacing threshold energy and current variations as a function of pacing rate in the range of 60 to 120 beats per minute (bpm). The pacing was done with activated vitreous carbon electrodes. Kay et al.84 found significant human atrial threshold changes as a function of pacing rate (125-300 bpm) using constant-voltage stimulation. They found a significant increase in rheobase voltage (Fig. 1-11), chronaxie, and minimum threshold energy at pacing rates greater than 225 bpm using platinized (low-polarizing) unipolar electrodes. They also determined strength-interval curves and found no correlation between atrial effective refractory period and rheobase voltage, chronaxie, or rate-dependent changes in either of these values. They concluded that the phenomenon is probably related to the “opposing effects of decreasing cycle length on the action potential duration and the slope of the strength-interval curve. Thus, if the pacing interval shortens to a greater extent than the refractory period and pacing stimuli are delivered during the ascending limb of the strength-interval curve (the relative refractory period), the diastolic threshold will increase.”84 The increase in stimulation threshold with increasing stimulation rate probably has minimal implications for bradycardia pacing. It is important for antitachycardia pacing, however, because threshold must be measured at rates required to interrupt the arrhythmia. In addition, the safety factors used with antitachycardia pacemakers must be based on thresholds measured at the clinically appropriate rates, rather than during pacing at resting rates.

Strength-Interval Relationships Voltage and current stimulation thresholds vary as a function of the coupling interval of the stimulus to prior beats and to the stimulation frequency used for the basic drive train. Figure 1-12 shows a typical ventricular constant-current strength-interval curve. At relatively long extrastimulus coupling intervals (>270 msec), the intensity of the extrastimulus required for ventricular capture is relatively constant, approaching the rheobase value. At shorter extrastimulus coupling 4 Atrial pacing rheobase (V)

0.9

Rheobase (V)

15

3 2 1 0 125

150

175

200

225

250

275

300

Pacing rate (beats per min) Figure 1-11  Effect of pacing rate on atrial rheobase voltage in patients. The rheobase voltage increases significantly for pacing rates greater than 225 beats/min (values shown as mean ± SEM). (From Kay GN, Mulholland DH, Epstein AE, Plumb VJ: Effect of pacing rate on human atrial strength-duration curves. J Am Coll Cardiol 15:1618-1623, 1990.)

16

SECTION 1  Basic Principles of Device Therapy

9 8 7 Current (mA)

6 5

Distal unipolar cathodal

4

Bipolar

3

Proximal unipolar anodal

2 1 0

Because the purpose of a pacing stimulus is to produce high current density at the electrode-electrolyte interface in the heart, the ideal situation is to deliver a low current from a very small surface area of electrode. If high current density (current magnitude divided by the plane area of the myocardium touched by the electrode) can be obtained with a small, high-impedance electrode, the current drained from the battery will be less than with lower impedance. At pacing threshold, a certain current density exists. If the interface area is made smaller, the current can be reduced equivalently and the same current density maintained. Minimizing the current needed for stimulation requires an optimal combination of interface area, current density, surface material and structure of the electrode, and increased impedance from making the electrode interface area smaller. An increase in impedance made to reduce current drain from the battery should be accomplished by making changes in the electrode, not by increasing the resistance of the wires to and from the pacemaker. Opposition to Stimulus Current Flow

240

280

320

360

400

Delay (msec) Figure 1-12  Strength-interval curves. Unipolar distal cathodal, unipolar proximal anodal, and bipolar strength-interval curves during an acute study in a patient with a temporary bipolar lead (equal-sized cathode and anode). The bipolar and unipolar anodal refractory periods are equal and are shorter than the unipolar cathodal refractory period. (From Mehra R, Furman S: Comparison of cathodal, anodal, and bipolar strength-interval curves with temporary and permanent pacing electrodes. Br Heart J 41:468, 1979.)

intervals (<250 msec), the intensity of the extrastimulus must be increased for it to elicit myocardial capture. The exponential rise in amplitude required for capture at short coupling intervals is the result of encroachment into the refractory period of the myocardium. During the relative refractory period, corresponding to the repolarization phase of the action potential, the myocardium can be induced to generate a new action potential if the stimulus has sufficient intensity. During the absolute refractory period, corresponding to the plateau phase of the action potential, no depolarization can be achieved, regardless of the stimulus intensity. The strength-interval curve of atrial and ventricular myocardium is shifted to the left at shorter basic drive cycle lengths. Therefore, the effective refractory period, which is usually measured at a pulse duration of 2 msec and an amplitude that is twice that of the late-diastolic threshold, decreases as a function of the pacing cycle length. At relatively slow pacing rates (<200 bpm), minimal interaction occurs between the strength-duration and strength-interval curves. As discussed previously, however, at rapid pacing rates (>225 bpm), pacing stimuli during the basic drive may encounter the relative refractory period. If the stimulus amplitude becomes subthreshold, 2 : 1 capture of the myocardium results.

Factors Opposing Pacemaker Current Flow When an electrical stimulus is applied to a pacing lead in contact with the heart, factors opposing current flow include the electrical resistance of the pacemaker lead wires, the impedance of the electrode-tissue interface, and characteristics of ion movement in the myocardium. All these are important to both the efficiency of stimulation and the longevity of a pacemaker or ICD battery. CONDUCTOR RESISTANCE AND CONDUCTION LOSSES Energy dissipation in the lead from the pulse generator to the electrode is proportional to the square of the current. To minimize charge drain from the battery, low lead resistance and low current are desirable.

The resistance of the wire leading from the pulse generator to the stimulating electrode—the cathode—is typically in the range of 50 to 150 Ω. The resistance of the pathway leading to the anode may be 50 to 150 Ω, or it may be lower, as in some unipolar pacing systems. For some practical purposes, one can consider the electrode and the tissue to act as resistances, rather than having a combination of reactive and resistive elements (i.e., impedance). Measuring the ratio of peak or average voltage to peak or average current when a pulse is applied provides useful information about load on the pulse generator and expected battery life. An increase in resistance calculated from voltage/current ratios may indicate a lead fracture, poor lead contact in the connector block, or an electrode that is not in good contact with tissue or blood. Very low impedance values may indicate an insulation defect. The absence of gross change from normal lead resistance does not rule out a broken lead or an insulation defect. An insulation defect can act as a shunt to ground. Its effect depends on its magnitude and location. A broken wire can separate and produce an extremely high measured lead resistance (e.g., 1800 Ω). However, the broken ends may remain in contact or in intermittent contact. If the ends are in continuous contact, lead resistance may not increase significantly. Intermittent contact can cause erratic pacing, sensing, or lead resistance measurements. Impedance at Electrode-Tissue Interface The resistance (ohmic polarization) of the stimulating electrode is inversely and exponentially proportional to the size of the electrode and the temperature and conductivity of the tissue.85,86 For practical purposes, in vivo temperature and conductivity are essentially constant. Decreasing the geometric surface area of the electrode, as illustrated in Figure 1-13, is a way to increase the current density at the electrode for any given level of output current from the pulse generator. Size reduction alone, however, has the negative effect of increasing polarization impedance. In effect, the Helmholtz capacitance at the interface is decreased by reducing the area of the electrode. One remembers from the equation 1 t v t = ∫ i tdt C 0 that the voltage vt across the Helmholtz capacitance at the electrode is proportional to the accumulated charge and inversely proportional to the capacitance. If the electrode contact area with the tissue is made smaller to increase the current density and nothing else is changed, the capacitance will be decreased. For a given applied charge, that action will increase the Helmholtz capacitance voltage, because of the inverse relationship between the voltage and the capacitance. For a given electrode material (as a first approximation), capacitance (C) is a function of the interface surface area (Ai):

C = eA i /d

[1-8]



1  Cardiac Electrical Stimulation

Ohmic pacing impedance (Ω × 102)

100 (N = 48 models, 323 canines) 50

10 5

The total reactance in a simple series circuit is the scalar sum of the inductive and capacitive reactances, each of which varies with the frequency content of the applied signal. Impedance also varies with signal frequency content; it is the vector sum of reactance and resistance. Impedance has a magnitude and a phase angle, both dependent on the rates of change of the applied voltage. The phase angle represents the difference in timing of sinusoidal current flow peaks compared with sinusoidal voltage peaks when a sinusoidal voltage is applied to a circuit. A voltage or current pulse of any shape can be broken down mathematically into combinations of sinusoidal components. In a simple series circuit, where R is the resistance and X is the sum of the capacitive and inductive reactances, the magnitude of impedance Z is defined as follows:

1 .1

.5 1.0 5.0 10 Geometric surface area (mm2)

50 100

Figure 1-13  Leading-edge ohmic polarization or electrode resistance of chronic canine ventricular leads (12 weeks after implantation) measured with a Medtronic Model 5311 pacing systems analyzer as a function of the unipolar cathode’s geometric surface area. Each data point was obtained from a population of animals testing one lead design. The study included 323 animals and 48 lead models.

where e is the dielectric constant of the Helmholtz double layer and d is its thickness. Because e and d are essentially constants in vivo, the capacitance of the electrode varies as a function of the interface surface area. How does one obtain a very large surface area to contact the tissue and not use a large-diameter electrode? Initially, this was accomplished by making the electrodes porous.87-89 Microporous electrodes, such as activated carbon or platinized platinum, have a higher real surface area.90,91 The capacitance of microporous surfaces is greater than that of porous surfaces and still greater than that of polished surfaces. Fractal-surface electrodes have been best at meeting the requirements of very small geometric areas and large Helmholtz interface areas at the same time. Another approach to the problem of decreasing polarization is to use a material that supplies its own majority charge carriers. An example is the silver–silver chloride (Ag/AgCl) electrode in chloride solution. The majority carrier, in this case Cl−, cannot be depleted; Cl− evolves at the cathode and is formed at the anode. Therefore, voltage across the electrode interface has a waveshape (waveform) essentially identical to that of the applied constant current, and the electrode is “nonpolarizing.” Unfortunately, no clinically usable, nonpolarizing (charge-carrier supplying) electrode material seems available for permanent implantation. Ag/AgCl electrodes, for example, are not used for permanent pacing because the anode erodes and the AgCl eventually dissolves from the cathode.92,93 The ideal, practical stimulating electrode is made of a corrosionresistant, biocompatible material with small geometric size and very high interface surface area, which gives it the characteristics of very high capacitance and a lower polarization voltage. Impedance Impedance can be defined as the vector sum of all forces opposing the flow of current in an electric circuit. Pure inductors and pure capacitors have no energy losses. They store or release energy in or from an electric field (capacitor) or a magnetic field (inductor). They also change the time relationships between varying voltage and current. The effects can be quantified in terms of reactance. Capacitive and inductive reactances oppose each other. If a sine-wave voltage is applied to a pure capacitance, the current peaks occur 90 degrees earlier than the voltage peaks. If a sine-wave voltage is applied to a pure inductance, the current peaks occur 90 degrees later than the voltage peaks.

17

Z = R2 + X2

[1-9]

for a single R-X series combination. For various combinations of reactance and resistive elements in parallel and in parallel-series combi­ nations, each component of impedance (Z1, Z2, Z3, . . . ) used in determining the total impedance must be treated as a vector quantity. Serial and parallel connection calculations using such vector components cannot be accurately performed by ordinary scalar, algebraic, Ohm’s law manipulations.94 Right ventricular (RV) and coronary sinus (CS) catheters connected to a single pulse generator output socket represent impedances in parallel. Reactance is positive for inductance (i.e., when a sine wave voltage is applied, the voltage leads the current, as seen on an oscilloscope screen) and negative for capacitance (when a sine wave voltage is applied, the voltage lags the current). Reactance and Charge Movement at Electrode-Electrolyte Interface The electrode and the tissue act not as pure resistances but as reactances that change the time relationships between voltage and current when a pacemaker or defibrillator pulse is applied. The only elements of almost-pure resistance in the pacemaker circuit external to the pulse generator are the lead wires. When the pacing pulse begins, several events occur almost immediately. Electron motion in the lead wires and electrodes must result in ion motion in the interstitial fluid. The electric field between the electrodes results in (1) reversible or irreversible oxidation-reduction processes occurring at the interface between the electrode and the electrolyte, and/or (2) the electrode-tissue interface’s acceptance of current without charge crossing the interface. In the latter case, the interface acts as a capacitor. Figure 1-14 shows the result is the sequence of events. The leading-edge voltage/current ratio is referred to as ohmic polarization. The increasing charge on the Helmholtz capacitor results in increasing opposition to further current inflow. The figure shows this as polarization overvoltage. The rates of voltage and current change depend on rates of ion movement within the electrolyte. These in turn depend on the mass and charge of ion species, species interactions and concentrations, and the applied electric field. The result is that lead wire current and voltage are not linearly related. The circuit is reactive. The vector combination of this reactance and the resistive elements (e.g., lead wires) make up the impedance faced by the pulse generator output. CHARGE CONDUCTION AND TRANSMEMBRANE POTENTIAL CHANGES IN CARDIAC PACING AND DEFIBRILLATION The directed motion of electric charge is the result of electric field gradients. Currents at pacemaker or defibrillator electrodes and within the heart can be considered in several categories, including (1) ion movement currents external to cells, within cells, and across cell membranes, and (2) charge separation currents that bring opposing charges to the electrode-electrolyte interface and membrane intracellularextracellular interfaces, without charge actually crossing the interface.

18

SECTION 1  Basic Principles of Device Therapy

Pulse generator pulse

Polarization effect

ELECTRICAL CHARGE MOVEMENT IN WIRES AND IN TISSUE

Vp

Leading edge

Trailing edge

Vo

Whether contraction of a myocyte occurs when a pacemaker pulse is applied is a function of several factors: (1) the amplitude, waveform, and duration of ion flow in the extracellular tissue resulting from the electric field gradient produced by the pulse; (2) the state of the preexisting voltage difference from the outside to the inside of the cell (degree of polarization or depolarization of the cell membrane); and (3) the biochemical state of the cell. Irnich104,105 discussed fundamental laws of electrostimulation in 1990, then three myocardial stimulation theorems in 2002, introduced by George Weiss 100 years earlier: (1) the pacing threshold, as measured by the voltage-time product of the stimulus, is a linear function of the pulse duration; (2) there is a minimum of the delivered threshold energy requirement that depends on the pulse duration; and (3) pulse shape plays no role in electrostimulation. The third theorem is a concept with clinically apparent limitations and requires some interpretation,106,107 as discussed later. Electron Flow versus Ion Flow

Time (pulse width) Figure 1-14  The effect of polarization (upper diagram) has been separated from the capacitively coupled constant-voltage pacing pulse (lower diagram) to help clarify the electrical manifestation of electrode polarization. At the leading edge of the pulse, polarization is essentially zero. With time (pulse duration), the voltage resulting from polarization (sometimes called polarization overvoltage, Vp) increases. When the pacing pulse is shut off, polarization overvoltage decays exponentially as a result of diffusion. (From Stokes K, Bornzin G: The electrodebiointerface [stimulation]. In Barold SS, editor: Modern cardiac pacing. Mt Kisco, NY, Futura, 1985, pp 33-78.)

The latter currents are capacitive, charging the Helmholtz capacitor at the electrode and changing the voltage difference across the cell membrane before it depolarizes. In 1969, Otto Schmidt95 described biologic information processing using the concept of “interpenetrating domains.” In the heart, two domains of charge flow exist: extracellular and intracellular. The two domain pathways are different in their anisotropic characteristics, and they are interwoven. Current passes from one domain to the other through cell membranes. Bidomain theory and studies are based on the concept that at every point in the heart, there are two electric potential vectors, one intracellular and one extracellular. Mathematically, each domain is continuous and occupies the entire domain. Membrane current leaves one domain and enters the other at the same coordinates. The analytical and experimental work from these concepts has provided insight about depolarization wavefront passage over the myocardium, current flow directions in relation to wavefront direction, and virtual electrode formation.96-99 Roth99 showed that stimulation of myocytes by application of a defibrillation shock can result in transmembrane voltage (Vm) changes by four mechanisms: (1) direct polarization of the tissue, as described by cable theory (see later discussion); (2) production of virtual anodes and cathodes, as described by bidomain models with unequal anisotropy ratios of the intracellular and extracellular spaces; (3) polarization of the tissue secondary to change in orientation of cardiac fibers (curving); and possibly, (4) polarization of individual cells or groups of cells because of a sawtooth electric potential effect. Sawtooth alterations in potential may result from resistive discontinuities introduced by intercellular gap junctions. Plonsey and Barr100 predicted the occurrence of sawtooth effects in 1986. Under what circumstances and to what degree, if any, the sawtooth effect might become important in fibrillation or defibrillation is uncertain.101-103

An interface between ion drift current and electron drift current occurs at the electrode-electrolyte interface. In pacemaker and defibrillator leads and electrodes connected to the negative output socket of the pulse generator, current flow consists of a net drift of electrons through the negative lead toward the electrode surface. In contrast, in the extracellular fluid, the current flow is ion drift. The motion of an individual ion is random, with the drift imposed by the electric field superimposed on the random motion. Drift velocity is proportional to the magnitude of charge on the ion, the mean time between collisions with other ions divided by the mass of the ion, and the electric field strength and direction. In addition, ions interact with each other. For a pacemaker electrode in the heart, all these properties, together with the Helmholtz double-layer capacitance effect, mean that current flow in the lead wires going to the electrodes does not vary instantaneously in proportion to the applied voltage. CAPACITANCE AND POLARIZATION When an electric field is applied to an electrode, the interface with the surrounding tissue stores electric charge separated by charge sign, thereby acting as a capacitor. A capacitor requires a finite amount of time to charge or discharge to any specified voltage. The relationship between the voltage across a pure capacitor at any time t, when i represents the instantaneous current flow and the capacitance has magnitude C (measured in farads), is t 1 v t = = ∫ i t dt C 0 and the instantaneous current i entering or leaving the capacitor is dv it = C dt where dv/dt represents the rate of change of the voltage across t the capacitor. The amount of charge stored, ∫ i t dt , is directly propor0 tional to the net area under the curve of instantaneous current entering and leaving the capacitor plotted against time. Note also that, for a given amount of charge, the voltage across the capacitor is inversely proportional to the capacitance. A perfect capacitor with no energy loss in storage or discharge is described by these equations. Neither cell membranes nor man-made capacitors are perfect capacitors. In nonperfect capacitors, some energy is dissipated in the capacitor during storage or discharge. Pacemaker and defibrillator electrodes act as imperfect capacitors by means of the Helmholtz double-layer effect (described later). The capacitance of canine myocytes in isosmotic solution is about 175 picofarads (pF). Cell membrane capacitance is about 0.01 microfarad (µF) per square millimeter of membrane surface area.108 The



1  Cardiac Electrical Stimulation

Input current from pulse generator cathode Lead resistance

Helmholtz doublelayer capacitance

Electrode interface with electrolyte in adjacent myocardium

Tek Output current to pulse generator anode Myocardial bulk impedance

T Trig’d

19

M Pos: 1.980ms

2

1

Faradic charge Warburg transfer resistance diffusion impedance Figure 1-15  Equivalent circuit representation of an electrodeelectrolyte interface, including the Helmholtz capacitance and the Warburg impedance. (Modified from Rodman JE: Solution, Surface and solid state assembly of porphyrins [PhD thesis]. Cambridge, University of Cambridge, 2000, Fig 4.6.)

capacitance of pacemaker electrodes has been reported to range from 0.2 µF/mm2 (smooth surface) to 40 µF/mm2 (fractal-surface electrode). The area of pacemaker tip electrodes is usually in the range of 4 to 10 mm2. Therefore, the capacitance might be 200 µF or more. The time constant for a 200-µF capacitor C discharging through a 500-Ω resistor R is R × C, or 100 msec. This means that the voltage across the capacitor, or across an ideal electrode-electrolyte interface, if the capacitance were a truly constant 200 µF and the interface impedance 500 Ω, would on discharge decrease from its maximum level to 37% of maximum in 100 msec. Figure 1-15 shows the Helmholtz capacitance and Warburg impedance in a schematic representation of a typical pacing configuration. These effects are caused by the electrode/tissue interface and the conversion of electric charge from electron flow to ion flow. Figure 1-16 demonstrates the decay of the Helmholtz capacitance following a constant current pulse. After termination of the pacing pulse, the voltage decays slowly toward baseline, but does not reach it before the second pulse is delivered. If the impedance between the positive and negative sockets of the pulse generator were made to go to very high impedance when the pacemaker pulse ends, almost no current would flow in the catheter wires between pulses. The Helmholtz capacitor at the electrode would discharge mainly through resistive components in the electrode-electrolyte interface. In that case, when the constant-current stimulus stops, the voltage measured across the catheter pins would keep the same polarity and would decay slowly rather than go to zero precipitously. Figure 1-16 also illustrates this decay by an electronic circuit simulation during pacing of a fibroblast cell culture through a bipolar cardiac catheter. The top trace shows a biphasic constant-current pulse with 3 msec between the end of the initial, negative phase and the beginning of the positive phase. The duration of each phase is 1 msec. The pulse generator circuit here goes to a high impedance rather than a low impedance when each phase ends. The bottom trace shows the voltage response. At the end of each phase, the voltage measured at the pulse generator terminals initially goes rapidly toward zero but, partway toward zero, it begins to change only slowly. This is the effect of the Helmholtz capacitor discharging into the surrounding electrolyte, and it is a function of the rates of ion movement at that time during the discharge cycle. BIPHASIC STIMULUS EFFECT Of great interest is the effect of decreasing the delay time between the end of the first and the beginning of the second phase. Compare Figure 1-16 with Figure 1-17. When the stimulus is biphasic with almost no delay (here about 10 µsec) between phases, the slow decay of voltage

Ch1 500mV

Ch1

Ch2 500mVBw 1 msec

Figure 1-16  Slow decay of pacing catheter lead voltage between biphasic stimulus phases. The plot shows current (top trace) and induced voltage (lower trace) from a Bloom stimulator delivered between the pacing tip of an Endotak catheter and a “hot can” electrode in saline. At the end of each phase, the voltage measured at the pulse generator terminals initially goes rapidly toward zero but then, partway toward zero, it begins to change only slowly. This is the effect of discharge of the Helmholtz capacitor into the surrounding electrolyte. This voltage response between phases is a function of the rates of ion movement in the electrolyte.

at the end of the first phase is aborted by the beginning of the second, opposite-polarity phase. The amplitude and duration of the voltage decay after the second phase are each less when the second phase occurs immediately rather than being delayed by 3 msec. In practical terms, because of Helmholtz effects, electrical activity continues to occur at and near the electrode after the stimulating pulse stops. Details of the activity depend on the stimulus (uniphasic or biphasic), delay time between phases, characteristics of each phase, and the action of the pulse generator output circuit when the stimulus stops (high impedance, low impedance, or recharge pulse). The electrical activity also depends on the characteristics of the electrode (e.g., geometric vs. electrical surface area), the ionic solution, and the tissue. Tek

Stop

M Pos: 1.980ms

2

1

Ch1 500mV

Ch2 500mVBw 1 msec

Ch1

Figure 1-17  Effect of near-zero delay between phases on voltage response to biphasic constant-current pulse. This plot was made under the same conditions as Figure 1-16, except that the delay between phases was decreased from 3 msec to approximately 100 µsec. When the stimulus is biphasic with almost no delay between phases, the slow decay of voltage at the end of the first phase is aborted by the beginning of the second, opposite-polarity phase.

20

SECTION 1  Basic Principles of Device Therapy

Cc

CAPACITANCE, POLARIZATION, AND TRUE IMPEDANCE When using an ordinary pacing systems analyzer, one thinks of impedance as a voltage/current ratio. If reactance is present, as it is at a pacemaker or defibrillator electrode because of the Helmholtz capacitance effect, the voltage/current ratio as seen on the PSA can be deceptive (see Biventricular Pacing). Because of the charging and discharging of the Helmholtz capacitor at the electrode-electrolyte interface, the applied voltage and resultant current are not linearly related. Therefore, the circuit is reactive, and impedance rather than resistance is the correct measure of opposition to current flow (see earlier discussion). There are practical problems in easily measuring true impedance during implantation procedures. In most clinical circumstances, however, the voltage/current ratio provides a reasonable estimate of the load into which the pulse generator is working, which is a major factor in determining current drawn from the battery.

T–

Rw –

Rc Rt

IPG Ca i0 T+

Rw+ Ra

CHARGE GROUPING AND POLARIZATION When an electrical stimulus is applied to the heart, the negatively charged cathode in contact with the myocardium attracts positively charged ions in the extracellular space. Likewise, the positively charged anode attracts negatively charged ions. This grouping phenomenon is known as polarization. The word “polarization” in this chapter has the following meanings: 1. The grouping and subgrouping by charge sign of electrons in the metal, and of ions in the electrolyte, at and near pacemaker or defibrillator electrodes, resulting in the formation of two or more electric charge layers and voltage gradients during the application of a pacemaker or defibrillator stimulus, and their subsequent decline after cessation of the stimulus. 2. The charge separation across a cell membrane, measurable as a voltage gradient between the inside and the outside of the cell. 3. The opposition to current flow within a battery, which results, when current is being drawn from the battery, in a decrease in the voltage available at its terminals. CAPACITANCE EFFECT OF HELMHOLTZ DOUBLE LAYER About 1879, Helmholtz suggested that a layer of ions (inner Helmholtz layer) is attracted to the surface of a charged electrode, and that this layer is bounded on its nonelectrode side by a layer of oppositely charged ions (outer Helmholtz layer) in the solution.109 This interface acts as a capacitor. Helmholtz assumed for this model that no electron transfer reactions occur at the electrode and that the electrode environment is only an electrolyte. These (and more complicated charge redistributions) occur because of attraction and repulsion interactions between an electrode held at a given electric potential (e.g., 2 V negative during 0.5-msec pacemaker pulse) and the ions and charged molecules in the interstitial fluid. The charge placed on the electrode, by electrostatic attraction, forces the accumulation of a polarized water layer and a second layer of hydrated, oppositely charged ions adjacent to the electrode surface. The oppositely charged ions come from the electrolyte. The interface at the electrode becomes an electrical double layer. This is in effect a charged capacitor.110 However, the actual interface is not a simple double layer. The Helmholtz model does not take into account other factors, such as absorption on the surface, thermal buffeting, and interaction between solvent dipole moments and the electrode. Models more complicated than the Helmholtz include the Gouy-Chapman and Gouy-Chapman-Stern.111 These models have dissimilar shapes for the plot of electric potential variation with distance from the electrode. The layers have relatively high dielectric constants and form an interface that behaves electrically like a capacitor (Cc and Ca in Fig. 1-18). In a semiconductor or in localized regions of an electrolyte, an excess of positive or of negative

Figure 1-18  Simplified equivalent circuit for cardiac pacemaker. T− and T+ are, respectively, the negative (cathode) and positive (anode) terminals of the implanted pulse generator (IPG; in a unipolar device, T+ is inside the can). The direction of current flow is shown with i0. RW− is the resistance of the conductor wire leading to the distal tip of the lead (cathode), whereas RW+ is that of the wire leading to the bipolar lead’s proximal ring electrode or the unipolar pulse generator’s can. Rt is the resistance of the tissue between the bipolar lead’s two electrodes, or the lead tip to the pulse generator can in a unipolar system. Rc is the ohmic polarization of the cathode, whereas Ra is that of the bipolar anode or unipolar generator can. Cc is the capacitance of the cathode, and Ca is that of the anode (can). The equivalent circuit is the same for unipolar and bipolar systems, but the values for some of the components differ.

charge may be present. If the excess is of positive charge carriers, the positive carriers are the majority carriers and the negative charge carriers the minority carriers. An excess of negative carriers makes these the majority and the positive carriers the minority. The second layer of the double layer is formed of hydrated ions and more water (Fig. 1-19). In physiologic electrolytes, the ions include Na+ and Cl− in major concentrations (majority carriers). Other ions present in lower concentrations (minority carriers) include hydronium (H3O+), hydroxyl (OH−), and phosphate (HPO42−). The ions attracted to or repelled from the electrode during the electrical stimulation pulse make up a separation of charge in the tissue electrolyte. When the pacemaker pulse is applied as a negative voltage to the electrode, electrons accumulate in the electrode. Reversible reactions may form metal-oxide complexes on the surface of the electrode. A primary water layer forms on the electrode. Positive ions surrounded by water molecules—a water shell—make a secondary water layer. This accumulation of positive ions in the electrolyte near the electrode unbalances local electrolyte charge neutrality. Secondary ion rearrangements occur in great complexity, with several names for the various processes. When the pacemaker pulse stops, the ions, being no longer attracted to or repelled from the electrode (depending on the polarity of the electrode and of the ions), are forced by their charges to rearrange themselves back toward their original, more electrically neutral positions (determined by random motion) in the electrolyte near the electrode. Ion rearrangement is not as fast as the transmission of an electric potential in a wire. The result is that, after a pacemaker or defibrillator pulse stops, the Helmholtz capacitor acts as if it were a temporary battery of declining voltage discharging itself into the tissue. The decaying voltage gradient in the tissue persists long enough to be detected by a pacemaker or defibrillator and may be great enough to interfere with sensing in autocapture devices in some situations. Other factors being equal, the voltage decay duration is greater and the polarization voltage is less when the electrode surface area is great versus small.



1  Cardiac Electrical Stimulation

+ H3O + Na – Cl + Na

– OH

+ Na

+ Na

A

B

C

Figure 1-19  Hypothesized structure of the Helmholtz double layer. At the left is an uncharged cathodic electrode interface. When the electrode is charged, a layer of surface hydration develops (region A). Region B contains a second, more loosely held hydration layer with hydrated ions. Based on electrostatic considerations, layer B has a high dielectric constant, approaching that of pure water. It has a thickness of less than 10 Å. Region C represents bulk solution. (From Stokes K, Bornzin G: The electrode-biointerface [stimulation]. In Barold SS, editor: Modern cardiac pacing. Mt Kisco, NY, Futura, 1985, p 63.)

ELECTRODE-ELECTROLYTE INTERFACE PROCESSES At equilibrium, an electrode in a solution will have, relative to a reference electrode, a potential that depends on the composition of the electrode, the composition of the solution, and other factors such as temperature.112 When a pacemaker pulse is applied, the electrode is forced to a different electric potential. This new potential forces ion flow to occur in the vicinity of the electrode. Ion secondary rearrangements occur distally. The net ion drift near the electrode is in a direction that will produce a new equilibrium, a balance of charge across the interface. At least two types of processes involving ion flow occur at the interface between a pacemaker or defibrillator electrode and tissue. First, oxidation-reduction reactions, which can be reversible or irreversible, involve electron movement across the interface and constitute faradic current. The second process, nonfaradic current, occurs without transfer of electrons across the interface. It consists of electron flow in or out of the electrode itself and a flow of ions in various layers or “clouds” toward or away from the interface. This nonfaradic process is similar to charging or discharging an electrical capacitor, but at an electrode in electrolyte, there is directed electron drift in or out of the electrode and directed ion drift within the electrolyte. The ion flow and the electron flow each constitute an electric current, yet no charge crosses the interface. Away from the electrode in the body of the electrolyte, the electric potential gradient causes ions to move away from or toward the electrode region, depending on their charge. Ion mobility characteristics, concentration gradients, and temperature gradients also affect ion movement. In the heart the process is more complicated than in an electrolyte solution alone, because of the anisotropic properties of the extracellular and intracellular domains. Faradic Current Flow Faradic current is so named because Michael Faraday (1791-1867) quantitatively described the effects of electric current through an electrolyte in 1834, stating that “the chemical decomposing action of a current is constant for a constant quantity of electricity, notwithstanding the greatest variations in its sources, in its intensity, in the size of

21

the electrodes used, in the nature of the conductors (or nonconductors) through which it is passed, or in other circumstances.”113 If one thinks of an electrode-electrolyte interface as resembling a variable resistor and variable capacitor in parallel, faradic current flow would be flow—electron transfer—through the resistor. Whether electron transfer across the interface occurs depends on the properties of the electrode and the electrolyte and on the applied electric pulse characteristics. The current crossing the interface when one is charging a battery is faradic current produced by oxidation-reduction reactions. Capacitive Current Flow This mechanism is not a transfer of electrons in one direction or the other across the electrode-electrolyte interface. Capacitance current flow is the accumulation of charge on the electrode at the interface balanced by a corresponding accumulation of charge of net opposite sign in the electrolyte adjacent to the interface. Charge flows in the lead, and in the electrolyte, but not across the interface. The flow of charge into the capacitor is measurable current in a pacemaker lead, even though in a perfect capacitor, no charge crosses the interface. The resulting arrangement of ion groupings at and near the electrode-electrolyte interface can be very complicated.114 The accumulation of charge separated by charge sign at the interface defines the Helmholtz double-layer capacitance.115 Energy is stored as an electrostatic field that exists between the positive and negative charge layers. When a pacemaker pulse is applied to an electrode in the heart, the charge movement into or out of the Helmholtz capacitance—both electrons in the electrode and ions in the electrolyte—is an electric current, even though no charged particle necessarily crosses the interface. At a pacemaker electrode, the capacitance is not constant. It is dependent on current density, electrolyte composition, and the area and other surface characteristics (e.g., fractal surface) of the electrode material.116 Because of the minute distances separating positive and negative charge layers at an electrode in electrolyte, the capacitance is high, ranging from about 0.2 to 40 mF/mm2 of geometric area of the electrode.117 The fractal-surface electrode has a much higher capacitance than a smooth-surface electrode. dv states that the current i entering or leaving The equation it = C dt a capacitor at time t varies directly as the magnitude of the capacitance (C, measured in farads) multiplied by the rate of change of the voltage difference across the capacitor. Because a constant-voltage pulse is nominally a rectangular wave, its rate of change is small except when the pulse is being turned on or off. Therefore, current flow into the capacitor (i.e., into the pacemaker electrode-electrolyte interface, Helmholtz capacitance) occurs very rapidly during the leading edge of a nominal constant-voltage pacemaker pulse. Then, as charge accumulates at the electrode-electrolyte interface, the accumulating negative charge collection at the electrode surface slows the rate of further accumulation. It does so because the accumulating negative charge opposes the inflow of additional negative charge. Finally, unless the pacemaker output voltage were to be raised during the pulse, no further accumulation can occur. The voltage across the capacitor becomes equal and opposite to the constant-voltage pulse that had been driving current flow into the lead. Note that the Helmholtz capacitance is influenced by the current density at the electrode, the types and numbers of ions in the electrolyte, the material and surface of the electrode, the temperature, and other factors. Relation of Electrode-Tissue Interface Capacitance to Polarization Voltage 1 t Note that in the expression vt = ∫ i tdt , describing the voltage that C 0 occurs across a capacitor when current flows into it, the developed voltage varies inversely as the capacitance (C, measured in farads). If

22

SECTION 1  Basic Principles of Device Therapy

both sides of this equation are multiplied by C, the equation becomes t t Cvt = ∫ i t dt . The amount Qt of stored charge at any time t is ∫ i t dt . 0 0 Therefore, Cvt = Qt, and vt = Qt/C. That is, for a given amount of charge put into the defibrillator or pacemaker electrode, the voltage across the interface will be decreased if the surface area of the electrode (and therefore the capacitance) is increased. This voltage is the polarization voltage. During its decay time after the stimulation pulse has ended, the polarization voltage can interfere with autosensing of capture. For a given stimulus voltage, the greater the polarization voltage at the electrode-tissue interface, the lesser the voltage gradient elsewhere in the tissue between the cathode and the anode. A fractal-surface electrode, with its high capacitance,117 has, for the same charge flow in and out, a lesser potential difference across the electrode-electrolyte interface than a smooth-surface electrode. Therefore, the fractal surface is preferred. No electrode is completely free of charge transfer across the interface between its surface and the electrolyte. Some current flow across the interface does occur by means of electron transfer through oxidationreduction reactions. At a pacemaker electrode, these reactions, if reversible, may be desirable. Irreversible charge transfer results either in corrosion or in electrochemical reactions in the tissue. For present-day implantable pacemaker electrodes, if the pulse duration is short, or if the pulse is biphasic with a sufficiently short time between balanced phases, the charge transfer reactions do not normally result in gross corrosion or gross tissue changes. Therefore, those reactions that do occur must be largely reversible.118 Ideal and Non-Ideal Electrodes To avoid corrosion and deposition of toxic products in the adjacent tissue, electrodes that undergo irreversible reactions only minutely are highly desirable for pacemakers and defibrillators. Michael Faraday worked out the quantitative relationship between electrode-electrolyte chemical changes and the total amount of charge passed through the electrode into the electrolyte.119 Electrodes that under the conditions of use produce irreversible electrochemical reactions by means of these faradic currents entering or leaving the electrolyte are non-ideal electrodes. Equal back-and-forth transfer of electrons from an inert electrode to an electrolyte occurs naturally and is an equilibrium situation. When an external voltage is applied, electrodes that then do not produce irreversible chemical reactions are so-called ideal electrodes. Although metal oxide films form on the electrodes, occurrence of other oxidation-reduction reactions and corresponding electron transfer through the interface do not occur or are minimal in ideal electrodes. In this sense, no real electrode can be completely “ideal” in all circumstances. If the magnitude and duration of polarization secondary to occurrence of the electric pulse is brief, electrochemical reactions occurring at the electrode may reverse. If the charge redistribution time is long, the effects of charge redistribution on the electrode and on the tissue may become irreversible. When biphasic pulses are used, minimal postpulse polarization persists, provided the time between phases is in the microsecond range. If the time between phases is increased into the millisecond range, the duration and amplitude of persisting polarization increases (see Figs. 1-16 and 1-17). OPTIMIZING IMPEDANCE FOR MINIMUM BATTERY CONSUMPTION The effects of high impedance on the pacing system vary with the cause and location. The manner of resistance distribution in the system can produce different effects, increasing or decreasing the efficiency of the stimulating pulse. To clarify these issues, one needs to examine the leads and the electrode-tissue interface impedance and conduction losses. A clinically useful electrode does not waste energy in mini–heat production or in undesirable chemical reactions, and it also has a low pacing threshold. Therefore, the electrode should have minimal energy

loss in the lead wires, minimal energy loss in moving ions in the electrolyte during charge and discharge of the Helmholtz capacitance, and high current density at the electrode, allowing depolarization of the membranes of nearby myocytes with less lead current than would otherwise be needed. All these factors influence the impedance as seen at the pulse generator. Note that fibrous tissue formation around an electrode does not necessarily increase the impedance but does increase the pacing threshold. (For further discussion of these issues, see Clinical Aspects of Myocardial Stimulation by Pacemakers.)

Impedance in the Extracellular Electrolyte WARBURG IMPEDANCE In 1899, Warburg described an effect of ion motion caused by a voltage difference between two electrodes in an electrolyte as an “impedance.” The Warburg impedance is an effect of ion diffusion activity under the influence of the potential gradient at the interface between electrode and electrolyte. It is said that “a Warburg impedance can be difficult to recognize because it is nearly always associated with a charge transfer resistance and a double layer capacitance.”120 The Helmholtz capacitance concept is based on the separation of charges in the electrode from charges in the electrolyte at angstromdimension levels. Warburg developed his impedance concept to model the effects of diffusion. When a sine-wave voltage is applied across an electrolyte, this impedance manifests itself as a 45-degree phase shift between the voltage and the current if the current density is infinitely low. The Warburg impedance magnitude becomes small compared with other impedances at the electrode-electrolyte interface as the electrolyte concentration is made greater and as the stimulation frequency increases.121 Ovadia and Zavitz122 made an impedance spectroscopy study of the interface between a platinum electrode and a metabolically active, perfused living heart. Three impedance spectral components were found: the Warburg impedance, a thin-film impedance, and a single high-angle constant-phase impedance. The impedance spectrum was not single-valued and was not stable in time. Figure 1-15 shows one equivalent circuit representation of an electrode-electrolyte interface, including the Helmholtz capacitance and Warburg impedance (modified from Rodman123). The Helmholtz capacitance–Warburg impedance effects can become especially important in explaining some of the threshold measurements and seeming contradictions found during biventricular pacing procedures,124,125 as discussed later. CAPACITANCE AND ELECTRIC POTENTIAL GRADIENTS RELATED TO MEMBRANE DEPOLARIZATION IN PACING AND DEFIBRILLATION The cell membrane does charge and discharge similar to a nonlinear capacitor. Ions under both electric gradient and biochemical-signaling control traverse the membrane as if it were both a variable and a highly ion species–selective resistor connected in parallel with a variable capacitor. When a negative-going pacemaker stimulus reaches the membrane, the stimulus reverses the voltage gradient across the membrane, beginning processes that in effect discharge the capacitor. Krassowska and Neu126 pointed out that the response of the cell to an external field is a two-stage process. The initial polarization proceeds with the cellular time constant (magnitude <1 µsec). The actual change of physiologic state proceeds with the membrane time constant (several milliseconds). The membrane resistance changes as ion channels open and close (see Fig. 1-2). The membrane capacitance may vary with ion concentrations and with stimulating voltage and frequency.127 Membrane capacitance and resistance determine the membrane time constant. An ordinary pacemaker stimulus can reverse the membrane potential to threshold levels only locally. The voltage gradient produced in the tissue by a pacemaker electrode decreases to below-threshold



1  Cardiac Electrical Stimulation

values with distance from the electrode. In order to depolarize much more myocardium at once than a pacemaker does, a defibrillator electrode must have a much greater voltage gradient between its electrodes. To initiate fibrillation during the vulnerable period or to defibrillate, an extracellular electric field of approximately 6 V/cm in a large volume of tissue is required, whereas for pacing, a local electric field of only about 1 V/cm in a small volume is necessary. Ideker et al.49 noted that, because the potential gradient decreases rapidly with distance from the stimulating electrode, a current about 40 times greater than the diastolic pacing threshold is required to generate an electric field of 6 V/cm approximately 1 cm from the stimulating electrode. To generate the 6-V/cm field over the entire ventricular myocardium, a current thousands of times greater than the low milliampere level necessary for pacing is required. The difference in voltage and current requirements for pacing and for defibrillation can be described using a “lumped” and distributed systems analogy. A line of cars is stopped at a traffic light. When the light changes from red to green, the front car moves, which prompts the next car to move. The motion propagates back down the line of cars in the manner of a string of myocytes undergoing sequential depolarization, or a line of falling dominoes; this is an analogy for pacing. Defibrillation is analogous to all the cars being connected by an iron rod. Much force would be required to move the whole string of cars at once, even though the force required for each individual car is relatively little. Similarly, to depolarize most of the ventricles at once, a very large voltage is needed to obtain the small transmembrane voltage gradient reversal that is required at each distant myocyte.

Cathodal and Anodal Stimulation, Constant Voltage, Constant Current, Monophasic and Biphasic Stimulation ANODAL AND CATHODAL STIMULATION Generally speaking, anodal stimulation is associated with properties that are less desirable than those of cathodal stimulation. If the stimulating electrode is a unipolar cathode, the strength-interval curve has a shape that is similar to a strength-duration curve. The shape of the typical anodal strength-interval curve is somewhat different from that of a cathodal curve (Fig. 1-20). Given equal-sized electrodes, the 8.0 7.0 6.0

mA

5.0 4.0 3.0 2.0 1.0

Anodal Cathodal

0.0 190 210 230 250 270 290 310 330 350 370 390 410 S1-S2 coupling interval Figure 1-20  Anodal and cathodal ventricular strength-interval curves are demonstrated in a patient with atrioventricular block, using unipolar constant-current stimuli. The anodal threshold is slightly higher than the cathodal threshold at coupling intervals greater than 250 msec. At coupling intervals of less than 210 msec, the anodal threshold is about equal to the cathodal threshold. In some patients, the anodal threshold may not decrease (“dip”) to a value that is lower than the cathodal threshold.

23

anodal stimulation threshold is generally somewhat greater than that for cathodal stimulation at long coupling intervals. With progressively more premature coupling intervals, there is a dip in the anodal curve to threshold values lower than those of the cathodal curve. At still more premature coupling intervals, the anodal curve rises steeply, as observed with cathodal stimulation.128 The “dip” phenomenon is not always seen.129 If the electrodes of a bipolar lead are of equal size, or if the anode is smaller than the cathode, it is possible to stimulate earlier in the cardiac cycle, in the relative refractory period, than with unipolar cathodal stimulation. This may occur because the threshold for anodal stimulation with electrodes of equal size is lower than for cathodal stimulation in the strength-interval curve at shorter pulse widths. Anodal stimulation generating tachyarrhythmias in ischemic or electrolyte-unbalanced hearts is well documented.130,131 This is one of several reasons that implantable bipolar leads are designed with considerably larger anodes than cathodes. In fact, it is commonly believed that the bipolar anode needs to be large enough that anodal stimulation is prevented. It is unlikely that this actually prevents anodal stimulation, however, because most pulse generators are set at outputs much higher than threshold. Even a 50-mm2 anode probably can reach threshold at a stimulus intensity well below 5 V and 0.5 msec. It is likely that delivery of the stimulation pulse with many bipolar leads at times actually results in capture of the myocardium at both the cathode and the anode. Therefore, the size relationship between cathode and anode probably has greater significance than currently known. Many newer bipolar leads have smaller anodes because the cathodes are smaller. The anode/cathode ratio should be maintained, and the designs should be supported with strength-interval tests. Bipolar pacing with equal-sized electrodes has been common in temporary pacing and in implanted epicardial systems, in which two unipolar leads are often used for bipolar pacing. The clinician should keep in mind the arrhythmogenic possibilities of such combinations in addition to their occasional advantages. For example, if two epicardial leads have been used in the past for bipolar pacing, with one electrode on the right ventricle and the other at an easily available site on the left ventricle, the hemodynamic result compared with unipolar pacing could have been better or worse. The sequences of contraction of the left ventricle periphery, septum, and right ventricle are hemodynamically important. Another reason that anodes should be relatively large is to preclude unacceptable corrosion rates. Platinum, for example, does not corrode under the cathodic current densities used in pacing, but anodic potentials can cause corrosion. To preclude significant corrosion, the proximal electrode (anode) of a transvenous bipolar pacing lead is made large enough to ensure that current density is low in that electrode. Careful examinations of explanted platinum polished distal tips (cathodes) have often revealed rough surfaces that are the result of corrosion. One explanation for the occurrence of corrosion at the distal tip electrode is that it is a cathode only during the stimulation pulse. Some pulse generators also have a fast-recharge pulse immediately after the stimulus to neutralize the afterpotential. Because the polarity of the fast recharge is then reversed, this portion of the pulse is actually anodal. Therefore, as “cathode” size decreases, current density increases and corrosion can rise to eventually unacceptable levels. Nonetheless, excellent canine performance has been reported with electrodes as small as 0.6 mm2 in geometric surface area.132 If the electrode is porous or fractal, the surface area of the electrolyte interface is actually large despite the small volume encompassed by the geometric surface. This can reduce current density and interface voltage at the electrodeelectrolyte interface to a level below that required for corrosion by the fast-recharge pulse.

Unipolar and Bipolar Stimulation All pacing systems require both a cathodal and an anodal electrode to complete the electric circuit from the pulse generator to, through, and from the myocardium back to the pulse generator. Therefore, all

SECTION 1  Basic Principles of Device Therapy

pacemakers are bipolar with respect to the body, but not the heart. Ordinary terminology defines a bipolar system as one that has two electrodes in contact with myocardium. A unipolar system has only one electrode in contact with myocardium. The effects of unipolar and bipolar configurations on sensing are profound, as covered in Chapter 2, and the engineering aspects are equally important, as discussed in Chapter 3. In an ordinary unipolar pacing system, one electrode—the cathode—is in contact with myocardium. The metal case of the pulse generator is the anode. Bipolar leads may have higher stimulation thresholds than unipolar leads if the two electrodes are of equal size. In this case, the output voltage delivered by the pulse generator is divided equally between the two electrodes, causing the measured threshold to be higher. The voltage required to force threshold current through two small electrodes is greater than that required for one small electrode. By making the cathode relatively small and the anode relatively large, the overall impedance can be kept at about the same value as for the cathode alone. The threshold current can be delivered with little increase in the driving voltage. Making the anode large makes it a relatively poor stimulating electrode, thereby reducing the arrhythmogenic potential of anodal stimulation, compared with a bipolar lead with equal-sized electrodes. To produce bipolar thresholds that are statistically equivalent to unipolar leads, the anode must be at least three times larger than the cathode.133 It has become standard practice in the design of pacing leads to ensure that the anode area is at least 4.5 times larger than the geometric surface area of the cathode. Bipolar electrode spacing can also be important, especially for sensing (see Chapter 2). It is generally accepted, for example, that as the electrodes are moved closer together, the signal-to-noise ratio, resistance to crosstalk, and median frequency of the sensed signal increase, whereas signal amplitude may decrease. If the spacing is too close, transvenous electrode pairs can have high stimulation thresholds, because current can be shunted between the two electrodes, producing a partial short circuit. In his classic thesis, De Caprio134 determined that the optimal bipolar spacing (for sensing) was about 8 mm.

Constant-Current versus Constant-Voltage Stimulation A constant-current generator delivers a current pulse that is rectangular in shape. That is, the trailing edge (Ite) of the pulse equals the leading-edge (Ile) current in amplitude and the pulse is flat on top. The current delivered by a constant-current pulse generator is independent of pacing impedance to the limits of the power source (Fig. 1-21). If impedance decreases, the pulse generator output voltage automatically decreases to keep the current constant. In the presence of a highresistance circuit, such as one with a lead partial fracture, the limits of the power source can be reached. In this case, the pulse generator does not generate enough voltage to maintain a preset level of current. The result is delivery of a lower current than programmed. Although under normal operating conditions the current waveform of a constant-current pulse generator is rectangular, the resultant voltage waveform is not. It would be rectangular if the circuit beyond the pulse generator were purely resistive. However, because of the reactance in the circuit, capacitive in nature at the electrode-electrolyte interface, the initial leading-edge voltage (Vle) rises rapidly, then more slowly to a larger, trailing-edge voltage (Vte), as shown in Figure 1-22. At the end of the pulse, an immediate drop in voltage equal to Vle is observed, followed by an exponential decay back to baseline. The voltage rise during the pulse, called the overvoltage, is caused by an increase in pacing impedance due to electrode polarization, and the afterpotential is caused by the gradual dissipation of that polarization. At present, probably all implantable pulse generators are capacitively coupled devices of approximately constant voltage. Because the pulse generator output comes from a capacitor, the voltage delivered

Voltage at 10 mA

18 16

Voltage at 5 mA

14 Volts and milliamperes

24

12 10

Current at 10 mA Voltage at 2 mA

8 6

Current at 5 mA

4

Current at 2 mA

Voltage at 1 mA

2 Current at 1 mA 100

1000

10,000

Resistance (ohms) Figure 1-21  The effects of pacing impedance on constant-current leading-edge waveform amplitudes using a Medtronic Model 5880A external pulse generator. Current is independent of impedance until the battery is “saturated.” This occurs when the voltage cannot increase any more to keep the current constant. (From Barold SS, Winner JA: Techniques and significance of threshold measurement for cardiac pacing. Chest 70:760, 1976.)

decreases as charge is delivered from the capacitor to the electrode in the heart. The fully charged capacitor in the output circuit of the pulse generator delivers a set leading-edge voltage (Fig. 1-23). This results in continuing charge transfer from the capacitor to the heart during the duration of the pulse. The difference in voltage at the beginning and end of the pulse is directly proportional to the amount of charge transferred out of the capacitor: Vle − Vte = Q*/C pg



in which Vle is the leading-edge voltage of the pulse, Vte is the trailingedge voltage, Q* is the amount of charge removed from the capacitor

Delivered voltage

Resultant voltage

5V

Resultant current 20mA

5V

Delivered current 20mA Constant voltage

Constant current

Figure 1-22  Voltage and current waveforms from a capacitively coupled, constant-voltage implantable pulse generator, the Medtronic Model 7000A (left), and the constant-current mode output of an external research stimulator, Medtronic Model 1356 (right). A unipolar lead with a polished platinum, 8-mm2 ring tip is used with a 900-mm2 titanium anode in 0.18% NaCl solution. The voltage across the lead is measured against a 100-mm2 Ag/AgCl electrode to eliminate the anode’s polarization from the waveform. (From Stokes K, Bornzin G: The electrodebiointerface [stimulation]. In Barold SS, editor: Modern cardiac pacing. Mt Kisco, NY, Futura, 1985, pp 33-78.)



25

1  Cardiac Electrical Stimulation 40

32

Volts and milliamperes

28 24

2 1 .8 .6 .4

.1

16

8

x x

2 1 .8 .6 .4

x x

.3

.5 .8 1.0 2.0

x x

x

Chronic

x x

x x

Peak (3 weeks)

x x

x

x

Implant

.2

.2

Leading edge

20

Threshold (volts)

Threshold (mA)

36

12

10 8 6 4

10 8 6 4

.1

.3

.5 .8 1.0 2.0

Pulse width (msec) Trailing edge Leading edge

Voltage

4

Current

Trailing edge

0

Figure 1-24  Logarithmic plots of canine ventricular constant-current (left) and constant-voltage (right) strength-duration curves for a passively fixed, atraumatic, unipolar lead with a polished platinum, 8-mm2 ring tip at various times after implementation (three different animals). (From Stokes K, Bornzin G: The electrode-biointerface [stimulation]. In Barold SS, editor. Modern cardiac pacing. Mt Kisco, NY, Futura, 1985, p 41.)

100

1000

high-impedance lead than for a low-impedance lead. Also, the rate of charge delivery decreases as the pulse generator capacitor output voltage decreases. This means that the voltage droop line is not a straight line, although it may appear so when the impedance is very high. The output voltage pulse becomes almost rectangular (i.e., Vle − Vte is very small) when the impedance of the lead-heart-lead circuit into which the pulse generator capacitor is discharging is very high. If the impedance is low, the pulse waveform droop is steeper. If it is very low (<200 Ω), the capacitor may not be able to maintain even an approximate constant voltage. Constant-current thresholds are related to constant-voltage thresholds by impedance. However, impedance is not constant during the pulse. It increases with polarization. Therefore, constant-current strength-duration curves can have slightly different shapes than constant-voltage curves. Most modern endocardial leads (e.g., iridium oxide, fractal, platinized platinum, or “activated” carbon surfaces) have relatively low-polarizing electrodes. These tend to have a lower-voltage rheobase and a somewhat higher chronaxie than highly-polarizing electrodes135 (Figs. 1-24 and 1-25). In fact, a nonpolarizing electrode made from Ag/AgCl has the same shape strength-duration curve for constant-voltage stimulation as for constant-current stimulation

10,000

Resistance (ohms) Figure 1-23  Effect of pacing impedance on capacitively coupled, constant-voltage leading-edge and trailing-edge waveform amplitudes. The leading-edge voltage remains constant as a function of pacing impedance at values of 200 Ω or greater (Medtronic Model 5950 pulse generator). The trailing edge of the voltage waveform changes slightly with impedance up to about 1000 Ω. Current falls significantly with increasing impedance. Any constant-voltage source may no longer be constant at very low pacing impedances, because the battery becomes unable to supply enough current to maintain a steady voltage. These very low impedance values are not likely to be encountered clinically in a properly functioning pacing system. (From Barold SS, Winner JA: Techniques and significance of threshold measurements for cardiac pacing. Chest 70:760, 1976.)

(measured in coulombs), and Cpg is the capacitance of the capacitor (measured in farads). Obviously, what is delivered through the lead wire to the pulse generator is not truly a constant-voltage pulse. The change in voltage over the pulse duration (“droop”) is directly proportional to the total charge delivered (Q*) during the pulse. For a given leading-edge voltage, the charge delivered will be less for a

2.0

2.0

1.5 Polished platinum

1.0

0.5

Cl 0.0

0.5

1.0

Pulse width (msec)

1.5

Ag/AgCl 1.0

0.5

CV

0.0

A

Threshold (V)

Threshold (V)

1.5

CV Cl

0.0 2.0

0.0

B

0.5

1.0

1.5

2.0

Pulse width (msec)

Figure 1-25  Constant-current (CI) compared with constant-voltage (CV) canine ventricular strength-duration curves for passively fixed, atraumatic electrodes. A, Polarizing, 8-mm2, polished ring-tip electrode. B, Nonpolarizing, Ag/AgCl, 8-mm2 electrode.

26

SECTION 1  Basic Principles of Device Therapy

1.4 *

Threshold (V/cm)

1.2

M B

2.52.5

1.0

2.5 *

0.8

55

5

0.6

10

20

0.4

1010

electrode, together with the present and future states of the myocardium, determine the pacing margin of safety. This margin is the difference between the stimulus magnitude delivered by the pulse generator and the magnitude, now and in the future, required for myocardial stimulation. The electrodes normally are the major factor in determining the impedance into which the pulse generator delivers its stimuli. In addition, the electrodes, along with the pulse generator circuitry, determine the sensing characteristics of the pacing system.140 Size, material, surface structure, and biologic response of the tissue to the particular electrode material are critical factors in electrode design. These factors are reviewed in the following sections, along with some effects of shape, spacing, fixation methods, and material.

0.2

SIZE OF THE STIMULATING ELECTRODE (GEOMETRIC SURFACE AREA)

0.0 0

5

10

15

20

Duration (msec) Figure 1-26  Monophasic (M) and biphasic (B) strength-duration curves obtained in strips of frog ventricular myocardium superfused with a solution containing 3 mmol/L of potassium. At pulse duration of 20 msec, there is no significant difference in the rheobase threshold. Note that the M waveform produces a lower threshold than the B waveform at pulse durations of 5 msec or less. Therefore, although the M and B waveforms have a similar rheobase, chronaxie is less with M than with B stimuli. (From Knisley SB, Smith WM, Ideker RE: Effect of intrastimulus polarity reversal on electric field stimulation thresholds in frog and rabbit myocardium. J Cardiovasc Electrophysiol 3:239, 1992.)

(see Fig. 1-25). Therefore, polarization affects the value of the constantvoltage stimulus chronaxie. This accounts for most of the differences in shape between constant-voltage and constant-current strengthduration curves.

Monophasic and Biphasic Waveforms A biphasic defibrillator stimulus, compared with a monophasic stimulus, has a lower voltage threshold for atrial and ventricular defibrillation.136 Knisley et al.137 studied the effect of biphasic and monophasic stimuli on excitation thresholds in rabbit and frog ventricles (Fig. 1-26). At very long pulse durations (20 msec), there was no difference in the stimulation threshold for monophasic versus biphasic waveforms. Therefore, reversal of waveform polarity during a long stimulus did not affect rheobase voltage. At shorter pulse durations (2.5 msec), the threshold voltage was significantly greater for biphasic than for monophasic waveforms. Therefore, the chronaxie pulse duration was significantly increased by an intrastimulus polarity reversal during shorter stimuli. The energy requirements for pacing and defibrillation are different for monophasic stimuli compared with biphasic stimuli. The energy requirement for pacing is greater with biphasic stimuli. For defibrillation, however, biphasic stimuli decrease the energy requirements required. Overlapping biphasic waveforms for pacing the atria in single-lead DDD pacing have been studied. One object was the reduction of atrial pacing thresholds and, consequently, reduction of the incidence of diaphragmatic stimulation compared with monophasic pacing under the same conditions.138 This configuration uses unipolar anodal and cathodal pulses delivered to two closely spaced, floating atrial electrodes so that the pulses of opposite polarity overlap. Phrenic stimulation remains a problem.139

Design Features of Pacing Electrodes That Affect Performance The performance of any pacing electrode is one of the major deter­ minants of the pacing threshold. Stimulation characteristics of the

Stimulation threshold varies as an inverse function of the size or geometric surface area of the stimulating electrode.37,141-143 Irnich38 stated that, in theory, the chronic stimulation threshold of a spherical electrode should decrease as its radius (geometric surface area) decreases to a certain value. With further decreases below that value, the chronic threshold should begin to increase. Therefore, the electric field strength E necessary for stimulation is a nonlinear function of the electrode radius r and of the potential V applied to a spherical electrode. Stimulation threshold increases during the first several weeks after lead implantation. This increase in threshold is related to the development of a conductive but nonexcitable fibrotic capsule that forms around the electrode and separates it from normal, excitable myocardium37,38,144,145 (Fig. 1-27). The fibrous capsule is a biologic response to the presence of a foreign body, in this case the electrode. With perhaps rare exceptions, the fibrous capsule is not a response to the electrical stimulus.146 In reality, if r is the electrode radius and d is the approximate mean thickness of the fibrotic capsule, r + d describes the dimensions of the effective or “virtual electrode,” as defined by Furman et al.147 If the stimulus voltage is held constant, the field strength varies as a function of the thickness of the nonexcitable capsule. Higher voltages must be applied to maintain the intensity of the field at the interface of the virtual electrode and the excitable myocardium. Therefore, chronic thresholds (for steroid-free electrodes) must be higher than the values at implantation. This theory was modeled by Irinch,37 as follows140: 2 V r  E=  [1-10] r  (r + d)  Equation 1-10 is lacking a crucial parameter required by the strength-duration relationship—stimulus duration. According to Irnich, the equation is based on measurements taken at a pulse duration of 1 msec. Holding everything else constant, the capsule thickness is the same for a large electrode radius as for a small one. If, for

Figure 1-27  Schematic representation of a simple collagenous capsule around an electrode, which separates it from excitable tissue. The length of the arrow in the left panel indicates the radius (r) of the spherical electrode. In effect, the fibrous tissue becomes a “virtual electrode” with radius r + d (arrow in right panel), where d is the thickness of the fibrous tissue (stippled area). (From Stokes K, Bornzin G: The electrode-biointerface [stimulation]. In Barold SS, editor: Modern cardiac pacing. Mt Kisco, NY, Futura, 1985, pp 33-78.)



1  Cardiac Electrical Stimulation

example, two electrodes of different radii (0.5 and 5 mm) develop a 1-mm-thick capsule (d), the electrode with the smaller actual radius will be associated with a greater percentage increase (r + d vs. r) with the formation of the virtual electrode (300% vs. 20%, respectively). This helps to explain the observation that smaller electrodes have lower acute thresholds but develop a greater rise in threshold over time.86,89 There is a point (r = d) at which chronic threshold reaches a minimum. According to Equation 1-10, when d is greater than r, thresholds increase. Therefore, there is a theoretical minimum spherical electrode radius at which thresholds must reach their minimum value. Irnich found this to be 0.72 mm for polished electrodes. In reality, because currently used electrodes are not polished and are not spherical, the relationship between stimulation threshold and electrode radius is more complex. Electrode size is more meaningful when it is discussed in terms of not only geometric surface area but also electrical surface area, which can be vastly increased by having a fractal surface. EFFECTS OF MAXIMIZING ELECTRODE SURFACE AREA Maximizing the electrical surface area of the electrode has the effect of increasing the Helmholtz capacitance. For a given stimulus amplitude expressed in terms of charge delivered, the voltage across the capacitance will decrease as the magnitude of the capacitance increases (V = Q/C; see earlier discussion). This means that increasing the Helmholtz capacitance decreases the polarization voltage, provided the delivered charge is kept the same. A very small-diameter electrode

A

27

with a large surface area (due to its fractal structure) can have a high Helmholtz capacitance and a high current density at the electrodetissue interface. The result is an electrode with high impedance and a lower threshold current requirement. Sensing performance may be good or not so good, depending on the electrode and pulse generator design (see later discussion). ELECTRODE SURFACE STRUCTURE (INTERFACE SURFACE AREA) Based on animal studies, electrodes with pores that allow tissue ingrowth have thinner fibrous capsules and somewhat lower chronic thresholds than electrodes with solid or polished surfaces.148-151 These claims have been somewhat controversial when applied to humans.152-155 Pore structures of certain dimensions are known to promote rapid tissue ingrowth156 (Fig. 1-28). Activated-carbon electrodes are reported to have porosity on the order of 10 nm.157 It has been argued that microporosity provides rapid stability of the electrode at its interface with the myocardium and prevents electrode motion that otherwise may result in tissue irritation and high chronic stimulation thresholds. The importance of microporosity among the many other factors involved in electrode mechanical stability is uncertain. RELATIONSHIP OF ELECTRODE SIZE AND SENSING PERFORMANCE The amplitude of the electrogram sensed by the amplifier of the pulse generator is less than the amplitude of the available signal in the

B

C Figure 1-28  Electron micrographs of microporous electrode surfaces. A, Activated carbon surface of Siemans Model 412S/60 electrode at about 8000× magnification. B, Medtronic Model 4011 platinized surface at 6900× magnification. C, Polished platinum surface of Medtronic Model 6971 at 7000× magnification. The polished platinum surface has little microstructure. Its actual microscopic surface area is similar to its apparent or geometric surface area (8 mm2). The platinized platinum surface (B) is composed of particles so small that they absorb visible light, and the surface appears black. The true surface area of the interface is many orders of magnitude greater than the geometric surface area. (From Seeger W: A scanning electron microscopic study on explanted electrode tips. In Aubert AE, Ector H, editors: Pacemaker leads. Amsterdam, 1985, Elsevier Science, pp 417-432.)

SECTION 1  Basic Principles of Device Therapy

myocardium. The attenuation of the signal is related to the ratio of the source impedance (of the electrode-myocardial interface) and the input impedance of the sensing amplifier. The higher the source impedance, the more the signal amplitude is attenuated for a given input impedance. Signal attenuation can be greatly minimized by increasing the input impedance of the sense amplifier, so that the ratio of input impedance to source impedance is very large.

Clinical Aspects of Myocardial Stimulation by Pacemakers STABILITY OF STRENGTH-DURATION CURVES AND THE DISTORTING EFFECT OF POLARIZATION None of the relationships published in the last 100 years has exactly duplicated the capacitively coupled, constant-voltage strength-duration curves that have been measured empirically with various types of electrodes. This may be caused, in part, by the confounding effect of polarization, which varies significantly with electrode size, shape, surface finish, and material. Another problem, however, is that the voltage and current strength-duration relationships, typically plotted on linear coordinates, are really logarithmic158 (see Fig. 1-24). When plotted in this manner, the strength-duration curve for constantcurrent stimuli is now seen to be a straight line up to rheobase, whereas the constant-voltage line turns upward, as seen at stimulus duration 2 msec in Figure 1-24. The effect of polarization on reducing current flow in constant-voltage pacing does not occur with constant-current pacing, because with the latter the pulse generator develops whatever voltage is necessary to force the set level of current into the circuit. The slope of the strength-duration curves for passively fixed, atraumatic leads does not change with time after implantation (see Fig. 1-24). The acute, peak, and chronic mean thresholds of five ventricular and three atrial, passively fixed, atraumatic leads in 77 canines had a (log10) slope of about 0.60 ± 0.07 V/msec.158 The correlation coefficient for these relationships was typically more than 0.99 at pulse durations of less than 0.5 msec for polished electrodes and at durations of less than 1.0 msec for low-polarizing designs. Traumatic electrodes, such as those with active fixation, typically had a lower slope immediately after implantation, which shifted to about the same 0.60-V/msec value within days as the acute trauma to the electrode-tissue interface healed. Because strength-duration curves are typically plotted on linear rather than logarithmic coordinates, changes in voltage and current threshold after implantation are usually seen as a shift upward and to the right. In most cases, because it is related to the slope of the curve (the tissue’s time constant), chronaxie does not change significantly with time. However, Cornacchia 159 showed that the administration of propafenone shifted both the strength-duration curve and the chronaxie to the right. The pacing threshold also increased. Some older pulse generator designs attempted to compensate for the gradual decline in stimulus voltage that results from battery depletion by automatic “stretching” of the pulse duration to maintain the stimulus energy at an almost constant value. Although this feature was designed to prevent loss of capture, increasing the pulse duration from a programmed value of 0.5 msec adds little to the pacing safety margin, because it approaches an essentially flat portion of the strengthduration curve. In fact, although this automatic increase in pulse duration was added in the name of safety, it is a wasteful use of battery energy. Preventing the decrease in output voltage of the pulse generator, while giving adequate warning of the depleted state of the battery, is a better solution. The pulse-duration increase method did temporarily prolong pacing capture, although inefficiently. This was valuable for patients who were far away from medical care for a short time. It also did provide a warning in that both transtelephonic checking devices and clinical devices measure the actual pulse width. Newer pulse generators have abandoned this approach to compensating for battery depletion in favor of automatic regulation of stimulus voltage. They maintain constant stimulus amplitude despite declining battery voltage.

THRESHOLD CHANGES AS A FUNCTION OF TIME AFTER LEAD IMPLANTATION It is well known that stimulation thresholds change as a function of time after lead implantation, typically rising to a peak value after several weeks.160-163 With older polished electrodes, some patients had thresholds that evolved over longer periods, up to 6 months.164 Luceri et al.165 observed that, after the acute rise in stimulation threshold, 43% of 120 patients had stable chronic thresholds for up to 8 years; 17% had chronic thresholds that decreased at a rate of 5% per year, and 19% had thresholds that rose at a rate of 14% per year; 20% had thresholds that varied widely around a stable mean. Modern pacing electrodes have a porous or microporous surface structure. Most also elute a glucocorticosteroid. As shown in Figure 1-29, the technological progression from polished to porous to microporous to steroid-eluting electrodes has significantly reduced the evolution of stimulation threshold as a function of time after implantation.166 In fact, as can be determined from the follow-up data available, the thresholds of porous and microporous steroid-eluting leads do not change significantly with increasing time after implantation.167-170 The reasons for threshold changes as a function of time and electrode design are to be found in the foreign body response to the electrodes. ELECTRODE-TISSUE INTERFACE AND FOREIGN BODY RESPONSE The body’s response to an implanted device has been relatively well characterized.171 Figure 1-30 shows schematically what typically occurs at an implantation site as a function of time. If one analyzes the tissues adjacent to canine ventricular endocardial atraumatic electrodes as a function of time after implantation, the evolution depicted in Figure 1-30 is typically observed (with one exception). In the first 1 to 3 days, there is no evidence of cellular inflammation, and thresholds do not change significantly. This lack of a cellular response in the first few days is an exception to the classic view of surgical wound inflammation. In a surgical wound, polymorphonuclear leukocytes (PMNs) normally are present during this period to scavenge necrotic debris and bacteria.

2

Threshold (V)

28

1

x x 3

x

1

x

x

xx

x

4

2 x 3 x 4 5 6

xx

0 0

2

4

8

12

Implantation time (wk) Figure 1-29  Canine ventricular voltage thresholds at 0.5 msec for 8-mm2 unipolar, transvenous, tined leads as a function of time after implantation. 1, Polished platinum ring-tip (manufacturer A); 2, polished platinum ring-tip (manufacturer B); 3, porous-surface platinum hemisphere (manufacturer C); 4, porous-surface titanium hemisphere (manufacturer B); 5, platinized Target Tip (manufacturer B); 6, steroid-eluting titanium porous-surface electrode (manufacturer B).



1  Cardiac Electrical Stimulation

Implantation of lead Myocardial membrane changes Increased local perfusion Increased endothelial permeability

Edema/fibrinolysis PMN cellular infiltration Macrophages and granulation tissue Exocytosis

Chemotaxis

H2O2 O2 radicals Enzymes

Fibroblasts Collagen capsule

Necrosis Myofibril disarray

More macrophages High stimulation threshold

Figure 1-30  Schematic representation of the inflammatory process and foreign body response to lead implantation. This has been correlated approximately with time after implantation of a cardiac pacemaker lead, based on histologic studies of canine electrode-tissue sites. PMN, Polymorphonuclear neutrophil; H2O2, hydrogen peroxide.

The classic view of inflammation holds that the initial event is dilation of blood vessels and alteration in the vascular endothelium, resulting in increased permeability. This increased perfusion and permeability of the blood vessel walls allows plasma to leak into the surrounding tissue, producing edema. Because plasma and serum are more conductive than myocardium, pacing impedance decreases almost immediately after lead implantation. After 2 to 3 days, a mixed cellular inflammation begins to appear in the tissues surrounding the electrode, and the stimulation threshold begins to rise. Tissue inflammation reaches its peak in about 1 week, with clearly evident interstitial edema and cellular necrosis. Pacing impedance typically reaches its minimum value at the inflammatory peak (about 1 week after implantation), after which it increases as the edema resolves. The stimulation threshold may or may not reach its peak at the same time as the nadir of pacing impedance. After the early, mixed cellular reaction, the inflammatory response is characterized by a gradual accumulation of macrophages that reach the electrode surface, adhere, and become activated. The major function of the macrophage is to dispose of dead or foreign cells and particles by the process of endocytosis116,172 (Fig. 1-31, A). With a large foreign object, such as a pacing electrode, the process of endocytosis is impossible. In this case, the macrophage tries to destroy the large foreign body by the process of exocytosis, the extracellular release of enzymes and oxidants (Fig. 1-31, B). Macrophages spread over the electrode surface, often differentiating into foreign body giant cells. Their lysosomes migrate to the membrane surface, releasing hydrolytic enzymes and oxidants into the surrounding tissue and onto the electrode surface. Additional macrophages settle on the foreign body giant cells and are activated. Myocytes adjacent to the electrode-tissue interface are bathed in a “soup” of inflammatory mediators resulting from exocytosis. These mediators of the inflammatory response dissolve the subcellular collagen beams, struts, and nets that hold the myocytes in the normal orderly array of myocardium.173 After 3 to 4 weeks (for a stable, biocompatible device), the more global inflammatory response has essentially resolved, with the

EXOCYTOSIS

ENDOCYTOSIS Phagocyte

29

Phagocyte

Particle

Lysosome

Lysosome

Device

Activation

Membrane invagination

Adhesion activation

Lysosomal release extra membrane

Vacuole

Enzymes H2O2

Lysosomal release within membrane

O2

A

Etc.

B

Figure 1-31  Endocytosis and exocytosis. A, Schematic representation of phagocytosis by endocytosis. A small particle attaches to the macrophage’s membrane. The membrane invaginates, encapsulating the particle in a membrane-lined vacuole. The vacuole and lysosome migrate toward each other, and their membranes fuse. The particle is then destroyed by lysosomal enzymes and oxidants. B, Schematic representation of the process of “frustrated phagocytosis” by exocytosis, or the acute foreign body response on the device surface, which results in lysosomal release of inflammatory mediators on the device surface and into the adjacent tissue.

30

SECTION 1  Basic Principles of Device Therapy

development of a collagenous capsule surrounding the electrode. If the electrode is biocompatible and stable relative to the myocardium, few significant visible histologic changes occur in the adjacent tissues after this period. Thresholds may subsequently decrease, remain stable, or increase, depending on the intensity of the chronic foreign body response at the electrode-tissue interface. Chronic atraumatic polished electrodes typically have a layer of foreign body giant cells on their surface (Fig. 1-32, A). These are covered by a layer (or layers) of macrophages that can be surprisingly thick. This cellular component of the capsule is not an acute, transient phenomenon but has been observed on electrodes up to 13 years after implantation in Stoke’s canine studies. These cells are covered by a layer of collagen that is oriented along the surface of the electrode. The myocytes adjacent to the collagen layers are disarrayed and are interspersed with collagen fibers that are radially oriented with respect to the surface of the electrode. There may also be circular “holes” in this layer of disoriented myocytes, which appear to be infiltrated with fatty material. Groups of macrophages with fused membranes (foreign body giant cells) enveloping bundles of myocardial fibers adjacent to the electrode were observed 1 to 4 weeks after implantation. The presence of this fatty infiltration, its location, and its severity depend on the stability, myocardial location, and shape of the electrode relative to the vector of myocardial contraction. Outside the disarrayed myofiber zone, one finds normally oriented myocardium. Therefore, it is clear that the concept of a simple fibrous capsule (Irnich’s “d”) separating the electrode surface from viable myocardium does not completely describe this complex biologic response and its effect on pacing. PASSIVELY FIXED LEADS Porous and Microporous Electrodes Electrodes with pores 10 to 100 µm in diameter allow the formation of collagen because of chronic inflammation. Therefore, in a sense,

macroporous electrodes promote inflammation. In addition, the chronic capsule surrounding porous electrodes can be associated with significant myofibrillar disarray (Fig. 1-32, B). Despite these factors, porous electrodes are associated with a relatively thin collagenous capsule. Histologic study of the chronic tissue reaction surrounding porous electrodes has shown that inflammatory cells are usually not found on the surface of the electrode. Rather, the active cellular component of the fibrous capsule is located within the pores. Similarly, microporous electrodes tend to have few cells on their surfaces with well-healed interfaces (Fig. 1-32, C). Target Tip design electrodes show an active cellular component deep in the grooves of the electrode (Fig. 1-32, D). The smooth outer surface interfaces directly with collagen. Steroid-Eluting Electrodes It has long been recognized that systemically administered glucocorticosteroids decrease both acute and chronic stimulation thresholds. In the early days of cardiac pacing, sometimes one or more of the several mercury cells in the pulse generators would internally short circuit and fail suddenly, causing loss of capture. In these cases, intravenous hydrocortisone was an effective emergency measure for temporarily regaining and maintaining capture on the way to the operating room. Initially, the threshold-lowering effect of systemic glucocorticoid therapy was thought to be caused by effects on the myocyte membrane and sodium and potassium “retention.”174 Steroid-eluting leads, in which dexamethasone sodium phosphate or acetate is gradually released from a reservoir within or around the electrode, have become standard for permanent pacing. It has been conclusively demonstrated that steroid elution decreases chronic thresholds. Despite the clear clinical benefits of steroid elution on the evolution of stimulation thresholds, however, the mechanism of this effect is not well understood.175 Although steroid-eluting, porous electrodes tend to have thinner capsules than polished-platinum electrodes, these capsules surrounding steroid-eluting leads can be difficult

A

B

C

D

Figure 1-32  Optical micrographs of chronic electrode-tissue interfaces. A, Capsule surrounding a polished electrode (upper left) has a relatively thick layer of activated phagocytic cells, including foreign body giant cells on the surface and macrophages further out. Collagenous material encapsulates these cells, with fibrous stringers extending outward into the myocardium. B, Myofibrillar disarray is seen between the collagenous capsule and normal myocardium. The capsule surrounding a porous electrode (upper right) appears to be thinner. However, the active cellular component of the capsule and some of the collagen have been removed with the electrode to facilitate trimming of the tissues. C, Microporous (Target Tip) electrode-tissue interface is shown at lower left. Note the thin, fibrous capsule over the external surfaces and at the ridges. D, Some macrophages are seen deep in the Target Tip electrode’s grooves at higher magnification.



1  Cardiac Electrical Stimulation

to differentiate from those steroid-free electrodes in blind analyses. The capsule surrounding a steroid-eluting lead is characterized by minimal cellularity between the collagen and the electrode surface and minimal myofibrillar disarray. EFFECT OF GLUCOCORTICOIDS ON EVOLUTION OF PACING THRESHOLDS Glucocorticosteroids are known to stabilize the membranes of phagocytes, such as the macrophage, through interaction with surface receptors, inhibiting release of lysosomal contents.176 It has been shown, however, that dexamethasone and its derivatives have no significant electrical effects on myocyte membranes.177,178 The salutary effect of systemic glucocorticoids on stimulation threshold probably results from inhibition of the release of inflammatory mediators from the cellular components of the fibrous capsule. When steroid therapy is discontinued, the release of inflammatory mediators resumes, and stimulation thresholds once again increase. The same phenomenon minimizes inflammation (foreign body response) on and adjacent to the surface of a steroid-eluting electrode. Acute stabilization of macrophage membranes on the electrode surface reduces or minimizes lysosomal release, thereby minimizing myofibrillar disarray and myocyte membrane damage. Chronic steroid elution suppresses the slow, insidious leakage of inflammatory mediators, thereby preventing threshold increase, without the risk of systemic side effects. MECHANICAL INSTABILITY AT MYOCARDIUMELECTRODE INTERFACE Mechanically unstable electrodes can provoke a significant pathologic response in the myocardium. Consider, for example, a transvenous corkscrew electrode that “rocks” at the interface with the endocardium because the lead behind it is too stiff. In some cases, large pockets of activated macrophages may develop on either side of the helix, resulting in the formation of a “preabscess.” In other cases, thick collagenous capsules form and then differentiate into cartilage and, if the instability is severe enough, into bone. Because active-fixation myocardial electrodes interfere with the contractile motion of the adjacent myofibers, myocardial degeneration with fatty infiltration may result in patterns that are clearly related to this mechanical interference. For example, the center of a myocardial helix may fill in with fatty “cells.” A myocardial pin may also provoke myocardial degeneration and fatty infiltration. Therefore, mechanical instability at the electrode-myocardium interface can produce marked histologic and physiologic effects that result in deterioration of the stimulation threshold. A particularly notable example of a lead with a high potential for exit block was an active-fixation lead that used an electrically inactive helix and a platinum-iridium electrode. The presence of a stiff, metal, J-shaped retention wire increased the stiffness effect of this lead. It had an exit block rate of about 10%.

TABLE

1-1 

31

SECOND-ORDER EFFECTS OF ELECTRODE DESIGN Spatial and Size Relationships between Electrode Pairs As previously established, within clinically relevant limits, porous and microporous electrode size and shape per se have no significant effects on the chronic stimulation thresholds (at least with steroid-free electrodes). Nonetheless, the size and shape of the electrode can still be important factors under certain circumstances. For example, a small displacement can affect the performance of a smaller electrode because its high-density field influences relatively few cells.179,180 A larger electrode is not as threshold efficient, but perforation is less likely, and small displacements have less effect on lead performance. The ring-tip electrode design is essentially a small electrode made into a large shape.181 This allows high electric field strength coverage of a larger number of myocardial cells, while affording a lower probability of perforation than a lead with a smaller tip. Although myocardial screws, barbs, and hooks are traumatic to the underlying myocardium, these electrodes all have points that serve as sources of high field strength. Therefore, myocardial electrodes of this design can still have good chronic performance, assuming that the lead is not so stiff as to apply undue force on the electrode-myocardium interface. Electrode Fixation The most efficient electrode design has poor long-term performance if it is mechanically unstable, resulting in both high stimulation thresholds and histopathologic evidence of excessive trauma. For chronically stable, steroid-free electrodes in canines, however, there was no correlation between the mechanism of fixation and the chronic stimulation threshold (Table 1-1). Given similar geometric surface areas, there were comparable chronic canine ventricular thresholds with tined, flanged, screw-in, and hook-in polished electrodes, whether endocardial or myocardial. These results were confirmed by human studies in both the atrium and ventricle.182-184 Therefore, the hypothesis that fixation remote from the electrode stimulation site is necessary for low chronic thresholds does not appear to be valid.185,186 It has not been established whether this is also true for steroid-eluting electrodes. However, a variety of leads are available that combine active fixation and steroid elution. The long-term pacing thresholds of these leads are much lower than those for similar models without steroid elution. Electrode Materials To minimize inflammation and its subsequent effects on stimulation threshold, electrodes for permanent pacing should be biocompatible and resistant to chemical degradation. Many materials, such as platinum, titanium, titanium oxide, titanium nitride, carbon, tantalum pentoxide, iridium, and iridium oxide, have been shown to be acceptable for use as pacing electrodes.187-189 Even silver, which is toxic in neurologic tissue and is certainly not corrosion resistant, has been used successfully in the heart as an electrode.190 Titanium and tantalum have been known to acquire a surface coating of oxides, which some argue may impede charge transfer at the electrode interface. To prevent this,

Chronic (12-Week) Unipolar Canine (Constant-Voltage) Thresholds of Polished Platinum Electrodes at 0.5 msec vs. Fixation Mechanism

Model No. 6901 6950 6961 6971 4016 6959 6955 4951 6917

Insertion Fixation Endocardial flange Endocardial long tines Endocardial short tines Endocardial medium tines Endocardial helix Endocardial helix Endocardial helix Epicardial hook Epicardial helix

Electrode Location and Shape Endocardial cylinder Endocardial cylinder Endocardial ring Endocardial ring Myocardial helix Endocardial ring Myocardial helix Myocardial hook Myocardial helix

Surface Area (mm2) 11 11 8 8 10 13 10 10 12

Stimulation Threshold (V) 1.9 ± 0.77 1.6 ± 0.48 1.8 ± 0.69 1.5 ± 0.57 1.1 ± 0.30 1.3 ± 0.42 1.6 ± 0.51 1.5 ± 0.44 1.4 ± 0.53

From Stokes K, Bornzin G: The electrode-biointerface (stimulation). In Barold SS, editor: Modern cardiac pacing. Mt Kisco, NY, 1985, Futura, p 40. *P <.05 considered statistically significant.

n 11 4 9 8 6 8 10 10 17

P vs. Model 6971* 0.2 >0.5 0.35 — 0.1 0.4 >0.5 >0.5 >0.5

32

SECTION 1  Basic Principles of Device Therapy

titanium has been coated with platinized platinum or vitreous carbon. Others accept the concept that the oxide layer acts as the dielectric of a capacitor, so that no charge is transferred across the dielectric during stimulation. In that situation, current continues to enter and leave each side of the capacitor (see earlier discussion of Helmholtz capacitance). In any case, titanium oxide electrodes are highly corrosion resistant, and these combinations have been found to have excellent long-term performance as pacing electrodes.190-194 Some theories have held that the foreign body reaction is a response to the electrode material.195,196 As a result, many attempts have been made to improve thresholds through the use of more biocompatible materials. For example, some claim that pyrolytic carbons are more biocompatible and produce lower chronic thresholds than platinum.94,197-199 Besides having a different composition, these electrodes are also microporous. When similar electrodes are polished, we have found that the chronic canine thresholds are not significantly different from those observed with polished platinum electrodes. Proper electrode material selection certainly is necessary to prevent unacceptable toxic responses or corrosion. At this point, however, apparently no biocompatible electrode composition significantly improves thresholds. Not all conductive materials are suitable for use as electrodes. Certain materials, such as zinc, copper, mercury, nickel, and lead, are associated with toxic reactions in the myocardium and are unsuitable for use in permanent pacing leads.182 Stainless steel materials are highly varying in composition and microstructure, with great variation between production lots. Electrodes made from one lot may be acceptable, whereas those made from another lot may corrode unacceptably. Because some of the corrosion products may be inflammatory or toxic, high thresholds occasionally occur with these materials. Therefore, stainless steel materials are no longer used for implantable electrodes. The polarity of the electrode may also have an important influence on its chemical stability. For example, Elgiloy, a non-noble and highly polarizing metal alloy, was used in the past as a cathode. It could not be used as an anode because it is susceptible to significant corrosion. Although Elgiloy had other excellent qualities, it is no longer used as a pacemaker electrode. The materials presently used for permanent pacing electrodes include platinum-iridium, titanium (oxide), platinum- or carboncoated titanium, platinized platinum, vitreous carbon, and iridium oxide. The platinized platinum, “activated” carbon, and iridium oxide electrodes are associated with a reduced degree of polarization. A negligible degree of corrosion occurs with these materials. The vitreous carbon electrodes have been improved by roughening the surface, a process known as activation that increases the surface area of the interface, thereby reducing polarization. Fractal coating of the distal electrode also has been introduced as a method for reducing electrode polarization. This technique involves coating the distal electrode with a microscopic granular structure that produces a complex surface. Electrode Location (Epicardial, Endocardial, Intramyocardial) As shown previously in canine ventricles for electrodes without steroid elution, the type of fixation (as opposed to the location of the electrode) has no effect on stimulation thresholds, in that epicardially applied corkscrews, transvenous corkscrews, and passively fixed endocardial electrodes all tend to have about the same chronic thresholds. Most epicardial or myocardial pacing leads are used in pediatric patients, in whom threshold complications are relatively common.200,201 It is generally perceived that epicardial and myocardial electrodes without steroid elution have less desirable threshold performance than endocardial leads. One reason that the performance of epicardial and myocardial leads has lagged behind that of endocardial passively fixed leads is the mechanical limitations involved. Epicardial and transvenous active-fixation leads are significantly more complex and are more difficult to design with all the attributes necessary for low, chronic stimulation thresholds. Nonetheless, there is no inherent reason that transvenous active-fixation or myocardial electrodes should not perform as well as or better than endocardial systems.

The first transvenous steroid-eluting active-fixation lead, the Medtronic Model 5078 (Medtronic, Minneapolis), was bipolar, with an electrically inactive, fixed helix surrounding a 5.8-mm2 porous, platinized cathode. This lead had excellent electrical performance in both the atrium and the ventricle.202-205 However, difficulties with helix distortion and entrapment in some cases resulted in the abandonment of the design. Active-fixation steroid-eluting leads provide significantly reduced pacing thresholds compared with standard active-fixation leads. Active-fixation electrode thresholds tend to be slightly higher than those of passive-fixation steroid-eluting electrodes.206 Steroid elution has also been applied to epicardial leads; with a significant reduction in long-term pacing threshold compared with steroid-free leads.207 PHARMACOLOGIC EFFECTS ON CARDIAC STIMULATION Several types of antiarrhythmic drugs have been demonstrated to increase stimulation thresholds in both humans and animals. Type 1 drugs decrease Na+ conductance and decrease the rate of rise of the action potential. Therefore, it should be no surprise that these agents can increase the threshold for pacing. Type 1A drugs, such as quinidine208 and procainamide,209 may result in increased thresholds, especially when administered in high doses.210 Type 1C drugs, such as encainide, flecainide, and propafenone, have been associated with increased pacing thresholds.211-215 The increase in stimulation threshold with these drugs has been demonstrated to correlate with the change in QRS duration. In addition to the type 1 agents, propranolol has been shown to result in an increase in pacing threshold when administered intravenously.216 Amiodarone, lidocaine, tocainide, and verapamil reportedly have minimal effects on pacing threshold, although we have seen one patient with congenital heart disease in whom amiodarone reproducibly increased the pacing threshold to the point of exit block.217 METABOLIC EFFECTS ON CARDIAC STIMULATION Stimulation Thresholds Stimulation thresholds may rise during sleeping or eating, factors associated with withdrawal of sympathetic tone and increased vagal tone.218,219 In contrast, factors associated with increased sympathetic tone, such as exercise or assumption of the upright posture, are associated with a decrease in threshold. The myocardial stimulation threshold is increased by several metabolic abnormalities, including hypoxemia, hypercarbia, metabolic alkalosis, and metabolic acidosis. In the presence of respiratory or cardiac arrest, the pacing threshold may increase by well over 100%, resulting in loss of capture despite the use of a conventional safety margin. Because of this observation, antitachycardia pacing devices, such as pacemakers and ICDs, are often designed to deliver high-intensity stimuli during anti-tachycardia pacing or after a high-energy shock. In patients undergoing implantation of an ICD, Khastgir et al.220 were unable to demonstrate an increase in pacing threshold 10 and 60 seconds after defibrillation. Ventricular fibrillation was electrically induced, however, and defibrillation was promptly performed under controlled circumstances in this study. PHYSIOLOGIC AND PHARMACOLOGIC EFFECTS ON PACING In the presence of a primary respiratory arrest, pacing stimuli may not capture the myocardium until adequate ventilation and pH balance are restored. Therefore, careful attention to respiration and pH must be maintained during anesthesia in patients with implanted pacemakers to ensure continued myocardial capture. Ischemia produces variable effects on stimulation threshold, depending on the location of the pacing electrode relative to the ischemic myocardium. In the presence of acute myocardial ischemia, the resting membrane potential decreases



1  Cardiac Electrical Stimulation

(cells become partially depolarized), the action potential upstroke velocity decreases, and the action potential duration dramatically shortens. In the presence of metabolic blockade with 2,4-dinitrophenol, Delmar221 noted an upward shift in the strength-duration curve, indicating an increase in the current required for capture at all pulse durations. Therefore, if the stimulating electrode is located in an ischemic region, the stimulation threshold would be expected to increase. With further ischemia and infarction, the myocardial threshold may rise dramatically. This may be seen clinically in patients who develop an acute inferior myocardial infarction (MI) with right ventricular (RV) infarction; in such patients, a previously implanted pacemaker may suddenly lose capture. However, if the stimulating electrode is located in a nonischemic region (e.g., right ventricle), activation of the sympathetic nervous system may reduce the pacing threshold. Hyperkalemia has been shown to increase the stimulation threshold when the serum K+ concentration exceeds 7 mEq/L.222,223 In contrast, in the presence of hypokalemia, intravenous (IV) K+ may decrease the pacing threshold and restore capture of a subthreshold pulse.224 In addition, the reduced excitability during hyperkalemia can be corrected by IV calcium.225 Hyperglycemia in the range of 600 mg/dL may increase stimulation thresholds by as much as 60%.226 Therefore, patients who have diabetes or renal failure—conditions associated with the potential for altered glucose metabolism and electrolyte abnormalities—may need a larger safety margin than other patients. Hypothyroidism has also been demonstrated to increase pacing thresholds, an effect that is reversible with thyroxine replacement.227,228 As stated previously, glucocorticosteroids decrease stimulation thresholds and have been used to treat exit block both acutely and chronically.229 Endogenous and synthetic catecholamines are effective in lowering pacing thresholds.230,231 The effect of IV epinephrine or isoproterenol is to decrease the stimulation threshold initially, followed by an increase. IV or sublingual isoproterenol has been demonstrated to reverse high pacing thresholds related to antiarrhythmic drug toxicity.232

Biventricular Pacing Optimal clinical outcomes with permanent pacing require that the hemodynamic effects of the pacing site and the timing of atrial and ventricular contraction be carefully considered. As a result, cardiac resynchronization pacemakers and ICDs typically use electrodes in the coronary venous system to allow left ventricular (LV) stimulation. Implantation of permanent pacing leads in the cardiac veins is the predominant method for chronic LV stimulation.233 Although epicardial LV pacing using standard myocardial leads remains an effective method for biventricular pacing for the treatment of congestive heart failure, it requires a more extensive operation than a transvenous approach. Transvenous stimulation of the left ventricle requires placement of the lead into a branch of the cardiac venous circulation, usually a posterolateral vein. The electrical properties of chronic stimulation of both the right and left ventricles are complex. The com­ plexity of dual ventricular stimulation has been reported in canine studies.234 The initial clinical approach to cardiac resynchronization therapy (CRT) involved the use of a pulse generator with an atrial output circuit and a single ventricular output circuit. Thus, a pacing stimulus could be delivered in a series configuration to the right ventricle as the anode and the left ventricle as the cathode (or vice versa). Such a configuration had the advantage of having very high impedance because both the anode and the cathode were small, tip electrodes of the RV and LV leads. This approach was limited by the inability to time RV and LV stimuli separately and the marked prolongation of the intracardiac electrogram, which resulted from recording both RV and LV activation during sensing. Such a configuration was not acceptable for ICDs because oversensing of both RV and LV activation occurred in the combined electrogram. The next configuration used clinically involved splitting the current between the RV and the LV electrodes in a shared-cathodal

33

configuration. This configuration was limited by reduced current density at both the RV and the LV tip electrode resulting from the parallel circuit design and inability to separate timing of RV and LV stimulation. In addition, the ventricular electrogram was a composite of RV and LV activation, leading to double counting when the interventricular conduction time was prolonged. Pacing impedance in the ventricular split, bipole configuration was much higher than for either electrode alone. The anode was programmable to be either the pulse generator casing (unipolar split-cathodal configuration) or the ring electrode on the RV lead (bipolar split-cathodal configuration). The magnitude of the current flowing to either ventricular electrode was determined by the impedance of the two leads. The apparent pacing threshold in the left ventricle was dramatically affected by this splitcathodal configuration. For example, Mayhew et al.125 found that, when the unipolar LV (coronary venous) threshold was measured using the coronary venous tip electrode as the cathode and the pulse generator casing as the anode, the mean threshold was 0.7 ± 0.5 V at 0.5-msec pulse duration. Splitting the cathode between the LV tip and the RV tip electrodes increased the apparent LV threshold to 1.0 ± 0.8 V. When the anode was changed from the pulse generator casing to the RV ring electrode, the apparent LV threshold further increased to 1.3 ± 0.9 V. Therefore, a split-cathodal configuration greatly increased the apparent pacing threshold, compared with pacing the coronary venous lead alone. A further observation with the split-cathodal pacing configuration is that the impedance measured with a PSA does not behave as predicted by the simple application of Ohm’s law. For example, when the bipolar pacing configuration was used with the RV ring electrode as the anode,125 the pacing impedance was measured to be 705 Ω in the right ventricle and 874 Ω in the left ventricle. Using a split-cathodal bipolar configuration, the measured impedance was 516 Ω, higher than would be predicted by Ohm’s law, with both ventricles stimulated in parallel (390 Ω). The higher-than-predicted impedance is explained by the size and shape of the combined cathodes being different from those for either electrode alone. The combined-cathodal configuration results in a Warburg resistance and capacitance. Combining electrodes of similar size essentially doubles the cathodal surface area. Because the electrode resistance of a hemispherical electrode is about proportional to the square root of the electrode surface area, doubling the size of the cathode by combining the RV and LV electrodes decreases the Warburg resistance by a factor of 1/ 2. This helps to explain why the measured impedance using a split-cathodal configuration is not halved, as would be predicted by doubling the size of electrodes joined in parallel. Combining electrodes in parallel also increases the voltage droop during a constant-voltage pulse. Coronary venous pacing leads tend to result in higher impedance compared with endocardial RV leads (based on a lower volume of blood surrounding the electrode), and this further tends to shunt current from the higher-impedance electrode to the lower-impedance electrode when both ventricles are stimulated in parallel. This factor serves to increase the apparent LV threshold with the split-cathodal configuration. Current CRT devices use separate RV and LV output circuits, with multiple options for programming the stimulation vector. Thus, the left ventricle can be stimulated by a single electrode in the cardiac venous system or in the bipolar configuration with two electrodes in the cardiac veins. In addition, the unipolar configuration can be used with a single electrode in the cardiac venous system and the oppositepolarity electrode being the RV ring, the pulse generator casing, or another, more proximal electrode of a quadripolar LV lead. The polarities of the LV electrodes can also be reversed. With these multiple options for pacing the left ventricle, the undesirable effects of the older, split-cathodal pacing configuration have been overcome in newer CRT devices. However, because of the additional output circuit, there is additional battery drain with these newer devices. Another feature of CRT devices involves the use of bipolar or even quadripolar cardiac venous leads. These leads differ from traditional right atrial or RV leads in that the surface area of the two cardiac venous electrodes is often similar. Therefore, when the leads are

34

SECTION 1  Basic Principles of Device Therapy

programmed to the bipolar pacing configuration, the impedance is much higher than with unipolar pacing, because current flows between two small electrodes rather than between one small and one large electrode. In addition, the bipolar pacing threshold (measured in volts) is often considerably higher than the unipolar configuration, because the current delivered with higher impedance is reduced. The important clinical advantage of multiple electrode LV pacing leads and independent output circuits with programmable polarity options is the much better chance of finding a cardiac venous site that offers a low stimulation threshold while minimizing the chances of phrenic nerve stimulation.

Automated Capture Features To ensure ventricular capture and allow programming of a low margin of safety, newer pacemakers use algorithms that automatically detect ventricular capture. Based on the measured stimulation threshold, the amplitude of the pacing stimulus is automatically adjusted to provide a programmed margin of safety. The Autocapture feature of St. Jude Medical (St. Paul, Minn.) pacemakers automatically adjusts the amplitude of the stimulation pulse by detecting capture in the ventricle from the evoked ventricular electrogram (the evoked response). The St. Jude pacemakers may use either a bipolar or a unipolar ventricular pacing lead with low polarization properties for the distal electrode. The presence or absence of ventricular capture is determined by sensing of the evoked response from the tip electrode. These devices automatically determine the evoked response gain and sensitivity levels by delivering five paired ventricular pulses of 4.5 V at a minimum pulse duration of 0.5 msec or the programmed value. The first of the paired pulses measures the evoked response, and the second pulse is delivered within 100 msec of the first (i.e., during physiologic refractory period of myocardium) to determine the level of polarization. If the evoked-response (ER) amplitude is greater than 2.5 mV, the measured lead polarization is less than 4.0 mV, and the ratio of the ER/ER sensitivity is greater than 1.8 : 1, the device will automatically determine that the safety margin is acceptable and recommend Autocapture as a programmed feature. The Autocapture feature uses unipolar pacing from the tip electrode and determines capture on a beat-to-beat basis. If a ventricular stimulus is not followed by a detectable evoked response, a second test pulse is given at a value 0.25 V greater than the last threshold measurement, known as the automatic pulse amplitude (APA). If a pulse is not followed by detectable capture, a backup pulse is delivered within 80 to 100 msec at an amplitude of 4.5 V. If two consecutive APA pulses are not followed by an evoked response, the threshold is measured to determine whether the APA needs adjustment. Specifically, the pulse is incremented 0.25 V above the last APA. If capture is not confirmed, the APA is repeated in 0.125-V increments until two consecutive captured events occur. For these devices, all lossof-capture pulses are immediately followed by a backup pulse. A potential complicating factor with detection of the evoked response is differentiating fusion from capture. In the DDD(R) pacing mode, precisely timed intrinsic conduction can result in false detection of loss of capture. To differentiate true loss of capture from fusion, the atrioventricular (AV) delay is incremented by 100 msec after two consecutive loss-of-capture events to search for intrinsic conduction. If intrinsic conduction is indeed present during this extension of the AV delay, the backup pulse is eliminated. On the other hand, if subsequent backup pulses or APA increments are required due to loss of capture, the paced or sensed AV delay is shortened to either 50 or 25 msec, respectively. This sequence can complicate the interpretation of electrocardiographic tracings with irregular AV delays. However, knowledge of the function of the Autocapture algorithm allows recognition that this is a normal phenomenon. Automatic capture algorithms have also been applied by most other manufacturers. The Sorin Group pacemakers (Milan, Italy) detect the evoked response on the tip electrode of a unipolar lead, provided that the polarization properties of the electrode are favorable. Boston Scientific pacemakers (Natick, Mass.) use a lower-capacitance output

capacitor to minimize the afterpotentials on the ventricular lead as a method for improving detection of the evoked response. The smalleroutput capacitor increases the droop of the pacing pulse, but the afterpotential is greatly reduced. The smaller-output capacitor may have the effect of slightly increasing the apparent stimulation threshold, an effect that is usually measurable only when the pacing threshold exceeds 2.5 V. Other manufacturers offer variants of the automatic capture algorithm that do not deliver backup pulses on a beat-to-beat basis but provide automatic determination of pacing threshold at programmed intervals during the day. These devices determine the pacing threshold and adjust the pacing amplitude to provide a programmed margin of safety. In general, automatic capture algorithms function quite effectively and may reduce the risk of loss of capture due to fluctuations in pacing threshold caused by drugs, metabolic derangement, or lead dislodgment. The capability for reducing the programmed margin of safety is effective for prolonging battery life and may reduce the frequency of clinic follow-up visits. The Medtronic Ventricular Capture Management feature determines a strength-duration threshold at a programmable interval (nominally, once per day). After the amplitude threshold is determined at a pulse duration of 0.4 msec, the pulse amplitude is doubled and a pulse duration threshold is measured. The permanent ventricular stimulation amplitude is then automatically reprogrammed using a programmable amplitude safety margin (usually twice the threshold) or a programmable minimum amplitude, whichever is higher. The nominal values for ventricular capture management are a safety margin of twice the threshold with a pulse duration of 0.4 msec and a minimum ventricular amplitude of 2.5 V. During measurement of the pacing threshold, each test pulse is followed by a backup pulse 110 msec later to ensure that a pacing pause does not occur. If the automatically measured ventricular stimulation threshold is greater than 2.5 V at 0.4 msec, the ventricular output is automatically programmed to 5.0 V and 1.0 msec. The Medtronic Atrial Capture Management feature is designed to periodically measure the atrial stimulation threshold and adapt the atrial output to a programmable amplitude safety margin. This feature does not use the evoked potential to determine the presence or absence of atrial capture. Rather, the pacemaker searches for evidence that atrial test pulses reset the sinus node (Atrial Chamber Reset Method) or observes the ventricular response to determine whether a captured atrial test pulse is conducted to the ventricles through the AV conduction system. The Atrial Capture Management feature performs an atrial amplitude threshold at 0.4-msec pulse duration and after loss of capture is detected (defined as two of three test pulses indicating loss of capture); the amplitude setting is increased until atrial capture is confirmed. Because this feature does not rely on detection of the evoked response, there is no restriction on the type of atrial lead that can be used. This feature will not measure thresholds if the sinus rate is consistently faster than 87 bpm.

Adequate Margin of Safety A pacemaker must be programmed with a safety factor that allows for changes in pacing threshold. The clinician needs to know how much thresholds actually change in the acute to chronic period, as well as chronically on an hour-to-hour and day-to-day basis. Settings as low as 2.5 V and 0.5 msec at implantation appear to ensure capture for most adult patients with modern microporous, steroid-eluting leads. Higher values should be used with older-technology leads. Some patients have large variations in threshold. The factors involved in estimating how great a margin of safety to allow in programming are the (1) presently measured threshold, (2) probability of a catastrophe if pacing ceases, (3) patient’s ability to recognize intermittent loss of capture if occurring, (4) pharmacologic milieu, and (5) frequency of pacing threshold monitoring. In most patients, a 1.5 : 1 or 2 : 1 ratio of output voltage (Vo) to threshold voltage (Vthr), or output current to threshold current, provides a reasonable safety factor. The balance over



LONG-TERM AND DIURNAL VARIATIONS IN PACING THRESHOLD McVenes et al.241 found no significant threshold changes in adult canines using then-modern chronic atrial or ventricular leads as a function of eating, sleeping, or exercise. This finding is supported by Kadish et al.,242 who found no changes in chronic human pacing thresholds during a 24-hour period in four of five patients studied. Although in one patient the threshold at 0.6 msec changed from 1 to 1.5 V between 3 and 6 pm, these investigators concluded that “ventricular pacing thresholds do not show substantial diurnal variability.”242 Grendahl and Schaanning243 also found minimal variation in pacing threshold during the day, after meals, or during sleeping or physical activity.184 Shepard et al.76 reviewed 4942 pacing threshold measurements in 257 patients with 312 non-steroid-eluting leads at up to 295 months after implantation. The median in-use time was 17 months. Of the measurements, 1053 were in children younger than 12 years. At stimulus durations of 0.5 ± 0.04 msec, for thresholds measured 1 month or more after implantation, the mean threshold was 1.2 ± 0.66 V for endocardial electrodes and 2.8 ± 1.39 V for epicardially applied electrodes. Highest mean thresholds were in the 6- to 12-year-old age group. In patients with five or more measurements after 3 months of use, an increase in pacing threshold occurred after 3 months in 24%. An additional 21% had at least one threshold that exceeded the post3-months individual patient mean by three standard deviations. Other clinical events possibly related to threshold increases included a doubling of the threshold in a child during two successive summers, when mild cold–like symptoms began. The effects of various drugs on thresholds have been reported, but neither the test protocols nor the results have been consistent.244 Therefore, minimal literature appears to support the statistical validity of any particular safety factor.

35

On the basis of the earlier report of Preston et al.,239 Barold et al.245 suggested that Vo/Vthr must be at least 1.75 to ensure an adequate safety margin, assuming a 50% increase in energy at threshold throughout the day. Ohm et al.177,246 studied threshold evolution as a function of implant duration for an 8-mm2 polished platinum ring electrode and found that this lead had a peak threshold of 2.2 ± 0.75 V at a pulse duration of 0.5 msec measured 2 weeks after implantation. Assuming an output setting of 5 V and 0.5-msec pulse duration, the average patient had a Vo/Vthr of 2.3 : 1 during the peak threshold time. The 98th percentile patient (mean ± 2 standard deviations) had a Vo of 5 V and a Vthr of 2.2 ± 0.75 V, or a safety margin of about 1.35 : 1 at peak threshold. Unpredictable or unusual situations (e.g., myxedema) may occur that justify greater safety factor ratios.247 The important clinical point is that, for patients who are always or intermittently dependent on pacing to stay alive, a much greater pacing safety factor reduces the risk of otherwise unexplained sudden death. PROGRAMMING VOLTAGE VERSUS PULSE WIDTH FOR MAXIMUM PULSE GENERATOR LONGEVITY A common clinical concern for programming of the pulse generator to optimize battery longevity relates to whether it is more useful to program the amplitude or the duration of the output pulse. Based on examination of the strength-duration relationship, it is more efficient to reduce the Vo of the pulse, because the current drain varies as the square of voltage. Figure 1-33 illustrates the effect of doubling the Vthr (at a constant pulse width) or tripling the threshold pulse width (without changing voltage) on current drain. In this example, the rheobase voltage was determined to be 1 V and the chronaxie duration was 0.3 msec. The stimulation threshold was 4 V at a pulse duration of 0.1 msec or 2 V at 0.3 msec. Tripling the pulse duration at 4 V to 0.3 msec provided an adequate (2 : 1) safety margin with a current of 8 mA per pulse (4.8 µA continuous current) and a stimulus energy of 9.6 µJ. Similarly, doubling the Vthr at a pulse duration of 0.3 msec from 2 to 4 V yielded an identical current drain (8 mA/pulse, or 4.8 µA) and safety factor. If the patient had a higher threshold, for example 2 V at 1 msec, doubling of the voltage or tripling of the pulse width would still give the same current drain (12 µA), but the safety factor would be significantly different. Tripling the pulse width would provide a marginal (at best) safety margin, because threshold is approaching rheobase on the flat portion of the strength-duration curve. It would be necessary to double the voltage in this case to ensure a 2 : 1 safety margin. The foremost consideration for programming voltage and pulse duration is patient safety.

22 20 18

Volts, µJ

years is that of safety factor beyond threshold versus how long the pulse generator can be used. Again, clinical factors and patient reliability in checking need to be considered. Children may experience a higher rate of exit block because of their active inflammatory responses. In all patients, but especially children, use of steroid-eluting electrodes is recommended whenever possible.235-237 In a study reporting on 4953 threshold measurements made up to 20 years after implantation, long-term thresholds and threshold variations of non-steroid-eluting electrodes were found to be greater in children than in adults.76 The safety factors provided during the peak threshold phase are highest for steroid-eluting porous electrodes and lower for (in descending order) microporous electrodes, porous electrodes, and polished-tip electrodes, based on data represented in Figure 1-29. There may also be situations in which the patient’s pacemaker must be set at the maximum output at implantation (e.g., if the patient will not be available for follow-up). In these cases, an acceptable balance of safety margin and battery longevity might be obtained with a setting of 2.5 V and 0.5 msec for a steroid-eluting electrode, or a setting of up to 5 V and 0.5 msec with other leads. These numbers are at best only general guides. It is much better to arrange, by whatever means possible, for actual follow-up threshold measurements, both acutely and chronically. Modern perception of the range of chronic circadian threshold variation is based mainly on the work of Sowton and Norman238 (mid1960s), Preston et al.239 (late 1960s), and Westerholm240 (1971). The first two studies used constant-current generators and reported thresholds in terms of energy. Because the computed energy “thresholds” could have been a reflection of changes in impedance, the actual variation in voltage threshold cannot be determined from these studies. Westerholm, who reported both voltage and energy data, noted substantial circadian variations in both parameters. In all these studies, the leads were primarily epicardial/myocardial with polished electrodes. These human data, however, may be of marginal relevance to modern constant-voltage generators and porous, steroid-eluting electrodes.

1  Cardiac Electrical Stimulation

Energy (2× voltage)

16 14 12

Energy (3× PW)

10 8

Energy (threshold)

6 4

Threshold (V)

2 0

0.0 0.2 0.4 0.6 0.8 1.0 1.2 1.4 1.6 1.8 2.0 2.2 PW (msec) Figure 1-33  Effect on current drain and safety margin of programming the stimulus voltage to twice the threshold (at a constant pulse duration) or programming the pulse width (PW) to three times the threshold value (at the threshold voltage). See text for discussion.

36

SECTION 1  Basic Principles of Device Therapy

Summary Myocardial stimulation is the fundamental principle underlying artificial cardiac pacing. Perhaps the most important concept for programming of an implantable pacing system is a thorough understanding of the strength-duration relationship. Pulse generators allow the clinician to program both the pulse amplitude (in volts) and the pulse duration (in milliseconds). The stimulation threshold is a function of both these parameters. The exponential shape of the strength-duration curve must always be considered when programming the output pulse to ensure an adequate margin of safety between the delivered stimulus and the capture threshold. For example, pulse durations of 1 msec and greater are located on the flat portion of the strength-duration curve, whereas pulse durations of less than 0.15 msec are on the steeply rising portion of the curve. The practical importance of these facts can be appreciated by considering two points on the strength-duration curve shown in Figure 1-9. If the clinician determines the threshold to occur at point A (2 V and 0.5 msec) by decrementing the stimulus voltage at a constant pulse duration, programming of the pulse duration to 1 msec (point B) would provide very little margin of safety. Similarly, if the threshold is measured to be at point C (3.5 V and 0.15 msec) by decrementing the pulse duration at a constant voltage, doubling of the stimulation voltage to 7 V (point D) also would provide a poor safety margin. When one considers the shape of the strength-duration curve, a more appropriate programmed setting would be provided by doubling the threshold voltage at a pulse duration of 0.5 msec (point E, 4 V and 0.5 msec). As a general rule, if the threshold is determined by decrementing the stimulus voltage, an adequate margin of safety can be assumed by doubling the voltage if the pulse duration used was greater than 0.3 msec. The two most important points on the strength-duration curve (rheobase and chronaxie) are easily estimated with modern pulse generators (see Fig. 1-8). Rheobase can be estimated by decrementing the output voltage at a pulse duration of 1.5 to 2 msec. Chronaxie can then be estimated by determining the threshold pulse duration at twice the rheobase voltage. If instead the threshold is determined by decrementing the pulse duration, an adequate safety margin can be assumed by tripling the pulse duration only if the threshold is 0.15 msec or less. If the rheobase and chronaxie are measured, doubling of the threshold voltage at the chronaxie pulse duration provides an excellent method for programming a pacing system. In most circumstances, however, experienced clinicians will measure the pacing threshold at an initial stimulus duration of 0.4 or 0.5 msec, then make decisions based on their knowledge of the patient’s problems and medications. What constitutes an adequate safety factor depends on knowledge of the patient’s status. How dependent on the pacemaker is the patient? How medically stable is the patient? What medicines that can influence the pacing threshold is this patient taking? Have important threshold variations been seen, or are they anticipated in this patient after the initial stabilization period? Are there problems with the leads? How old is the pulse generator, and what is the known history of this pulse generator and attached leads in other patients? These factors determine clinical judgments about voltage and current safety, frequency of pacing and medical status monitoring, and early or late replacement of the pulse generator and leads. When programming the pulse generator at implantation, the clinician must also consider the acute to chronic evolution of the stimulation threshold. Because an acute rise in threshold typically occurs during the first several weeks after lead implantation, voltage and pulse duration may need to be programmed to higher values than needed for chronic pacing. The physician should reevaluate the stimulation threshold after the acute rise (and sometimes subsequent fall). For

most patients, the pacing system can be programmed to chronic output settings at a follow-up evaluation about 6 weeks after lead implantation. Although these recommendations may not be as applicable to patients receiving a steroid-eluting lead, caution is still warranted. The importance of drug and electrolyte effects on the strengthduration curve should also be appreciated. For patients requiring antiarrhythmic therapy, the stimulation threshold should be measured a number of times after drug initiation to ensure an adequate margin of safety for pacing. Similarly, patients who are more likely to experience alterations in electrolyte concentration (e.g., patients with renal failure, patients taking potassium-wasting diuretics) may need their pace­ makers to be programmed with a greater margin of safety. Perhaps most important, the degree to which the heart depends on pacing to sustain life or to prevent severe symptoms must be factored into the choice of a programmed margin of safety. For pacemakerdependent patients, a pacing pulse at least 2.5 times the chronic capture threshold is usually recommended. In contrast, patients unlikely to experience severe symptoms should failure to capture occur may have their pacemaker programmed to a lower margin of safety, perhaps twice the threshold. The effect of pacing rate on the stimulation threshold should also be considered for patients who require antitachycardia pacing, with the pacing threshold measured at all rates likely to be used. In the presence of high impedance caused by lead fracture, the current output of a constant-voltage pulse generator decreases. Loss of capture can occur. If lead insulation failure occurs, the impedance as seen by the pulse generator may decrease because of current shunting to noncardiac tissue. This results in an increase in the current from the pulse generator without a change in the nominal output voltage. This change may not be detected early if threshold is determined only by the voltage required for pacing capture. Measuring the voltage/current ratio allows detection of the nominal impedance and alterations in lead insulation or in wire continuity. Because some wire fractures intermittently make and break contact, a normal impedance measurement does not always ensure that the lead is intact. Pacing impedance is determined by four factors: (1) resistance in the conductor wire pathways, (2) polarization at the electrode-tissue interfaces, (3) resistance (small geometric size for high resistance) at the electrode-tissue interface, and (4) impedance/resistance of the tissues between the electrodes. The first two factors are energy inefficient, decreasing the current available for stimulation, whereas the third factor decreases current drain without decreasing the efficiency of stimulation. An ideal electrode would have, among other attributes, high resistance and high capacitance (low polarization voltage) at the electrode-tissue interface. Pacing with a monophasic stimulus is more energy efficient than pacing with a bipolar stimulus. The pacing threshold is greater at normal stimulus durations for biphasic stimuli than uniphasic stimuli with the same total duration. For successful defibrillation, biphasic stimuli are more energy efficient. A biphasic stimulus with proper characteristics reduces the postpulse ion rearrangements. Biphasic stimuli also may reverse continuing local and undesirable chemical processes at the electrode.

Conclusion Electrical stimulation, not only of the heart but also of other types of tissue (e.g., brain), is being used clinically more and more. Increasing knowledge of fundamental factors in electrical stimulation and practical developments for particular clinical circumstances is ongoing. Such knowledge will help cardiologists, as well as physicians and surgeons of other disciplines, make good decisions and expand the range of patient care.



1  Cardiac Electrical Stimulation

37

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Tse HF, Lau CP, Leung SK, et al: Single lead DDD system: a comparative evaluation of unipolar, bipolar and overlapping biphasic stimulation and the effects of right atrial floating electrode location on atrial pacing and sensing thresholds. Pacing Clin Electrophysiol 19:1758-1763, 1996. 139. Izquierdo R, Rodrigo G, Pelegrin J, et al: Single lead DDD pacing using electrodes with longitudinal and diagonal atrial floating dipoles. Pacing Clin Electrophysiol 25:1692-1698, 2002. 140. Ripart A, Fletcher R: Sensing. In Ellenbogen KA, Kay GN, Wilkoff B, editors: Clinical cardiac pacing, Philadelphia, 1992, Saunders. 141. Barold SS, Ong LS, Heinle RA: Stimulation and sensing thresholds for cardiac pacing: Electrophysiologic and technical aspects. Prog Cardiovasc Dis 24:1, 1981. 142. Barold SS, Ong LS, Heinle RA: Stimulation and sensing thresholds for cardiac pacing: electrophysiologic and technical aspects. Prog Cardiovasc Dis 24:1, 1981. 143. Smyth NPD, Tarjan PP, Chernoff E, et al: The significance of electrode surface area and stimulating thresholds in permanent cardiac pacing. J Thorac Cardiovasc Surg 71:559, 1976. 144. Parsonnet V, Zucker IR, Kannerstein ML: The fate of permanent intracardiac electrodes. J Surg Res 6:285, 1966. 145. Thalen HJT, Van den Berg JW: Threshold measurements and electrodes of the cardiac pacemaker. Acta Pharmacol Nederl 14:227, 1966. 146. Akyurekli Y, Taichman GC, White DL, et al: Myocardial responses to sutureless epicardial lead pacing. In Meere C, editor: Proceedings of the VI World Symposium on Cardiac Pacing, Montreal, 1979, PaceSymp. 147. Furman S, Hurzler P, Parker B: Clinical thresholds of endocardial cardiac stimulation: a long-term study. J Surg Res 19:149, 1975. 148. Wilson GJ, MacGregor DC, Bobyn JD, et al: Tissue response to porous-surface electrodes: basis for a new atrial lead design. In Moore C, editor: Proceedings of the VI World Symposium on Cardiac Pacing, Montreal, 1979, PaceSymp. 149. Amundson D, McArthur W, MacCarter D, et al: Porous electrode-tissue interface. In Moore C, editor: Proceedings of the VI World Symposium on Cardiac Pacing, Monstreal, 1979, PaceSymp. 150. Amundson DC, McArthur W, Moshaffafa M: The porous endocardial electrode. Pacing Clin Electrophysiol 2:40, 1979. 151. MacGregor DC, Wilson GJ, Lixfeld W, et al: The porous surface electrode: a new concept in pacemaker lead design. J Thorac Cardiovasc Surg 78:281, 1979. 152. Breivik K, Ohm O-J, Engedahl H: Acute and chronic pulse-width thresholds in solid versus porous tip electrodes. Pacing Clin Electrophysiol 5:650, 1982. 153. Berman ND, Dickson SE, Lipton IM: Acute and chronic clinical performance comparison of porous and solid electrode design. Pacing Clin Electrophysiol 5:67, 1982. 154. Freud GE, Chinaglia B: Sintered platinum for cardiac pacing. Int J Artif Organs 4:238, 1981. 155. MacCarter DM, Lundberg KM, Corstjens JP: Porous electrodes: concept, technology and results. Pacing Clin Electrophysiol 6:427, 1983. 156. MacGregor DC, Pilliar RM, Wilson GJ, et al: Porous metal surfaces: a radical new concept in prosthetic heart valve design. Trans Am Soc Artif Intern Organs 22:646, 1976. 157. Elmqvist H, Schuller H, Richter G: The carbon tip electrode. Pacing Clin Electrophysiol 6:436, 1983. 158. Stokes KB, Bornzin G: The electrode-biointerface (stimulation). In Barold SS, editor: Modern cardiac pacing, Mt Kisco, NY, 1985, Futura, pp 33-78. 159. Cornacchia O, Maresta A, Nigro P, et al: Effect of propafenone on chronic ventricular pacing threshold in patients with steroideluting (capture) and conventional leads. Eur J Card Pacing Electrophysiol 2:A88, 1992. 160. Pearce JA, Bourland JD, Neilsen W, et al: Myocardial stimulation with ultrashort duration current pulses. Pacing Clin Electrophysiol 5:52-58, 1982. 161. Meyers GH, Parsonnet V: Engineering in the heart and blood vessels. New York, 1989, Wiley-Interscience. 162. Hurzeler P, Furman S, Escher DJW: Cardiac pacemaker current thresholds versus pulse duration. In Silverman HT, Miller IF, Salkind AJ, editors: Electrochemical bioscience and bioengineering, Princeton, NJ, 1973, Electrochemical Society, pp 124. 163. Barold SS, Winner JA: Techniques and significance of threshold measurement for cardiac pacing. Chest 70:760, 1976. 164. Brownlee WC, Hirst R: Six years experience with atrial leads. Pacing Clin Electrophysiol 9(6 pt 2):1239-1242, 1989.

165. Luceri RM, Furman S, Hurzeler P, et al: Threshold behavior of electrodes in long-term ventricular pacing. Am J Cardiol 40:184, 1977. 166. Bornzin GA, Stokes KB, Wiebusch WA: A low-threshold, lowpolarization platinized endocardial electrode [abstract]. Pacing Clin Electrophysiol 6:A-70, 1983. 167. Elmqvist H, Schuller H, Richter G: The carbon tip electrode. Pacing Clin Electrophysiol 6:436, 1983. 168. Mond H, Stokes KB: The electrode-tissue interface: the revolutionary role of steroid elution. Pacing Clin Electrophysiol 15:95107, 1992. 169. Ohm O-J, Breivik K: Pacing leads. In Gomez FP, editor: Cardiac pacing, electrophysiology, tachyarrhythmias, Madrid, 1985, Editorial Group, pp 971-985. 170. Hoff PI, Breivik K, Tronstad A, et al: A new steroid-eluting electrode for low-threshold pacing. In Gomez FP, editor: Cardiac pacing, electrophysiology, tachyarrhythmias, Mt Kisco, NY, 1985, Futura, pp 1014-1019. 171. Anderson JM: Inflammatory response to implants. ASAIO Trans 34:101-107, 1988. 172. Henson PM: Mechanisms of exocytosis in phagocytic inflammatory cells. Am J Pathol 101:494-514, 1980. 173. Robinson TF, Cohen-Gould L, Factor SM: Skeletal framework of mammalian heart muscle: arrangement of inter- and pericellular connective tissue structures. Lab Invest 29:482-498, 1983. 174. Preston TA, Judge RD: Alteration of pacemaker threshold by drug and physiologic factors. Ann NY Acad Sci 167:686-692, 1969. 175. Stokes KB, Anderson J: Low threshold leads: the effect of steroid elution. In Antonioli GE, editor: Pacemaker leads, Amsterdam, 1991, Elsevier, pp 537-542. 176. Sibille Y, Reynolds HY: Macrophages and polymorphonuclear neutrophils in lung defense and injury. Am Rev Respir Dis 141:471-502, 1990. 177. Benditt DG, Kriett JS, Ryberg C, et al: Cellular electrophysiologic effects of dexamethasone sodium phosphate: implications for cardiac stimulation with steroid-eluting electrodes. Int J Cardiol 22:67-73, 1989. 178. Stokes KB, Kriett JM, Gornick CA, et al: Low-threshold cardiac pacing electrodes. In Frontiers of Engineering in Health Care: Proceedings of the Fifth Annual Conference IEEE Engineering in Medicine and Biology Society, 1983. 179. Irnich W: Considerations in electrode design for permanent pacing. In Thalen HJT, editor: Proceedings of the IVth International Symposium on Cardiac Pacing, Assen, The Netherlands, 1973, Van Gorcum, pp 268. 180. Irnich W: Engineering concepts of pacemaker electrodes. In Schaldach M, Furman S, editors: Advances in pacemaker technology, New York, 1975, Springer-Verlag, pp 241. 181. Mond H, Sloman JG, Cowling R, et al: The small tined pacemaker lead: absence of displacement. In Meere C, editor: Proceedings of the VI World Symposium on Cardiac Pacing, Montreal, 1979, PaceSymp, pp 29-35. 182. Baker JH, Shepard RB, Plimb VJ, Kay GN: Effects of fixation mechanism and electrode material on atrial stimulation threshold: long-term evaluation in 338 patients [abstract]. Pacing Clin Electrophysiol 15:54, 1992. 183. Cornacchia D, Jacopi F, Fabbri M, et al: Comparison between active screw-in and passive leads for permanent transvenous ventricular pacing [abstract]. Pacing Clin Electrophysiol 6:A56, 1983. 184. EI Gamal M, Van Gelder L, Bonnier J, et al: Comparison of transvenous atrial electrodes employing active (helicoidal) and passive (tined J-lead) fixation in 116 patients [abstract]. Pacing Clin Electrophysiol E 6:205, 1983. 185. Kay GN, Anderson K, Epstein AE, Plumb VJ: Active fixation atrial leads: randomized comparison of two lead designs. Pacing Clin Electrophysiol 12:1355-1361, 1989. 186. Rasor NS, Spickler JW, Clabaugh JW: Comparison of power sources for advanced pacemaker applications. In Proceedings of the 7th Intersociety Energy Conversion Engineering Conference, Washington, DC, 1972, American Chemical Society, p 752. 187. Hirshorn MS, Holley LK, Hales JR, et al: Screening of solid and porous materials for pacemaker electrodes. Pacing Clin Electrophysiol 4:380, 1981. 188. Schaldah M: New pacemaker electrodes. Trans Am Soc Artif Intern Organs 17:29, 1971. 189. Helland J, Stokes KB: Nonfibrosing cardiac pacing electrode. US Patent No 4033357. February 17, 1976. 190. Elmqvist H, Schuller H, Richter G: The carbon tip electrode. Pacing Clin Electrophysiol 6:436, 1983. 191. Thuesen L, Jensen PJ, Vejby-Christensen H, et al: Lower chronic stimulation threshold in the carbon-tip than in the platinum-tip endocardial electrode: a randomized study. Pacing Clin Electrophysiol 12:1592-1599, 1989. 192. Bornzin GA, Stokes KB, Wiebush WA: A low threshold, low polarization, platonized endocardial electrode. Pacing Clin Electrophysiol 6:A70, 1983. 193. Mugica J, Duconge B, Henry L, et al: Clinical experience with new leads. Pacing Clin Electrophysiol 11:1745-1752, 1988. 194. Djordjevic M, Stojanov P, Velimirovic D, et al: Target lead-low threshold electrode. Pacing Clin Electrophysiol 9:1206-1210, 1986. 195. Amundson DC, McArthur W, Moshaffafa M: The porous endocardial electrode. Pacing Clin Electrophysiol 2:40, 1979. 196. Timmis GC, Helland J, Westveer DC, et al: The evolution of low threshold leads. Clin Prog Pacing Electrophysiol 1:313, 1983.



197. Berman ND, Dickson SE, Lipton IM: Acute and chronic clinical performance comparison of porous and solid electrode design. Pacing Clin Electrophysiol 5:67, 1982. 198. Freud GE, Chinaglia B: Sintered platinum for cardiac pacing. Int J Artif Organs 4:238, 1981. 199. Kay GN, Anderson K, Epstein AE, Plumb VJ: Active fixation atrial leads: randomized comparison of two lead designs. Pacing Clin Electrophysiol 12:1355-1361, 1989. 200. Stokes KB: Preliminary studies on a new steroid eluting epicardial electrode. Pacing Clin Electrophysiol 11:1797-1803, 1988. 201. Hamilton R, Gow R, Bahoric B, et al: Steroid-eluting epicardial leads in pediatrics: improve epicardial thresholds in the first year. Pacing Clin Electrophysiol 14:2066, 1991. 202. Stokes KB, Frohling G, Bird T, et al: A new bipolar low threshold steroid eluting screw-in lead. Eur J Card Pacing Electrophysiol 2:A89, 1992. 203. Schwaab B, Frohling G, Schwerdt H, et al: Long-term follow-up of a bipolar steroid eluting pacing lead with active and passive fixation. In Antoniolo GE, editor: Pacemaker leads 1997, Bologna, 1997, Monduzzi Editore, pp 361-364. 204. Schwaab B, Frohling G, Schwerdt H, et al: Long-term follow-up of three microporous active fixation leads in atrial position. In Antoniolo GE, editor: Pacemaker leads 1997, Bologna, 1997, Monduzzi Editore, pp 365-368. 205. Schwaab B, Frohling G, Schwerdt H, et al: Atrial and ventricular pacing characteristics of a steroid eluting screw-in lead. In Antoniolo GE, editor. Pacemaker leads 1997, Bologna, 1997, Monduzzi Editore, pp 383-388. 206. Menozzi C: Comparison between latest generation steroideluting screw-in and tined leads: Long term follow-up. In Antoniolo GE, editor: Pacemaker leads 1997, Bologna, 1997, Monduzzi Editore, pp 389-394. 207. Nurnberg JH, Schopper H, Busscher U, et al: Retrospective comparison of epicardial steroid-eluting and conventional leads for pacing after corrective surgery in congenital heart disease. Pacing Clin Electrophysiol 20:1193, 1997. 208. Wallace AG, Cline RE, Sealy WC, et al: Electrophysiologic effects of quinidine. Circ Res 19:960-969, 1966. 209. Gay RJ, Brown DF: Pacemaker failure due to procainamide toxicity. Am J Cardiol 34:728-731, 1974. 210. Moss AJ, Goldstein S: Clinical and pharmacological factors associated with pacemaker latency and incomplete pacemaker capture. Br Heart J 31:112, 1969. 211. Hellestrand KJ, Burnett PJ, Milne JR, et al: Effect of the anti­ arrhythmic agent flecainide acetate on acute and chronic pacing thresholds. Pacing Clin Electrophysiol 6:892, 1983. 212. Salel AF, Seagren SC, Pool PE: Effects on encainide on the function of implanted pacemakers. Pacing Clin Electrophysiol 12:1439, 1989. 213. Montefoschi N, Boccadamo R: Propafenone induced acute variation of chronic atrial pacing threshold: a case report. Pacing Clin Electrophysiol 13:480-483, 1990. 214. Huang SK, Hedberg PS, Marcus FI: Effects of antiarrhythmic drugs on the chronic pacing threshold and the endocardial R

1  Cardiac Electrical Stimulation wave amplitude in the conscious dog. Pacing Clin Electrophysiol 9:660, 1986. 215. Bianconi L, Boccadamo R, Toscano S, et al: Effects of oral propafenone therapy on chronic myocardial pacing threshold. Pacing Clin Electrophysiol 15:148-154, 1992. 216. Kubler W, Sowton E: Influence of beta-blockade on myocardial threshold in patients with pacemakers. Lancet 2:67, 1970. 217. Irnich W, Gebhardt U: The pacemaker-electrode combination and its relationship to service life. In Thalen HJT, editor: To pace or not to pace: controversial subjects in cardiac pacing, The Hague, 1978, Martin Nijhoff, pp 209. 218. Gay RJ, Brown DF: Pacemaker failure due to procainamide toxicity. Am J Cardiol 34:728-731, 1974. 219. Preston TA, Fletcher RD, Lucchesi BR, Judge RD: Changes in myocardial threshold: physiologic and pharmacologic factors in patients with implanted pacemakers. Am Heart J 74:235, 1967. 220. Khastgir T, Lattuca J, Aarons D, et al: Ventricular pacing threshold and time to capture postdefibrillation in patients undergoing implantable cardioverter-defibrillator implantation. Pacing Clin Electrophysiol 14:768-772, 1991. 221. Delmar M: Role of potassium currents on cell excitability in cardiac ventricular myocytes. J Cardiovasc Electrophysiol 3:474486, 1992. 222. Gettes LS, Shabetai R, Downs TA, et al: Effect of changes in potassium and calcium concentrations on diastolic threshold and strength-interval relationships of the human heart. Ann NY Acad Sci 167:693-705, 1969. 223. Lee D, Greenspan K, Edmands RE, et al: The effect of electrolyte alteration on stimulus requirement of cardiac pacemakers. Circulation 38:124, 1968. 224. Walker WJ, Elkins JT, Wood LW, et al: Effect of potassium in restoring myocardial response to a subthreshold cardiac pacemaker. N Engl J Med 271:597, 1964. 225. Surawicz B, Chelbus H, Reeves JT, et al: Increase of ventricular excitability threshold by hyperpotassemia. JAMA 191:71-76, 1965. 226. Westerholm CJ: Threshold studies in transvenous cardiac pacemaker treatment. Scand J Thorac Cardiovasc Surg 8(suppl):1, 1971. 227. Schlesinger Z, Rosenberg T, Stryjer D, et al: Exit block in myxedema, treated effectively by thyroid hormone replacement. Pacing Clin Electrophysiol 3:737-739, 1980. 228. Basu D, Chatterjee K: Unusually high pacemaker threshold in severe myedema: decrease with thyroid hormone therapy. Chest 70:677-679, 1976. 229. Nagatomo Y, Ogawa T, Kumagae H, et al: Pacing failure due to markedly increased stimulation threshold 2 years after implantation: successful management with oral prednisolone—a case report. Pacing Clin Electrophysiol 12:1034-1037, 1989. 230. Haywood J, Wyman MG: Effects of isoproterenol, ephedrine, and potassium on artificial pacemaker failure. Circulation 32(suppl II):110, 1965. 231. Katz A, Knilans TK, Evans JJ, Prystowsky EN: The effects of isoproterenol on excitability, supranormal excitability and

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conduction in the human ventricle [abstract]. Pacing Clin Electrophysiol 14:710, 1991. 232. Levick CE, Mizgala HF, Kerr CR: Failure to pace following high dose anti-arrhythmic therapy: reversal with isoproterenol. Pacing Clin Electrophysiol 7:252-256, 1984. 233. Daubert C, Ritter P, Cazeau S, et al: Permanent biventricular pacing in dilated cardiomyopathy: Is a totally endocardial approach technically feasible? Pacing Clin Electrophysiol 19:699, 1996. 234. McVenes R, Stokes K: Alternative pacing sites: how the modern technology deals with this new challenge. In Antonioli GE, editor. Pacemaker leads 1997, Bologna, 1997, Monduzi Editore, pp 223-228. 235. Hurzeler P, Furman S, Escher DJW: Cardiac pacemaker current thresholds versus pulse duration. In Silverman HT, Miller IF, Salkind AJ, editors: Electrochemical bioscience and bioengineering, Princeton, NJ, 1973, Electrochemical Society, pp 124. 236. Stokes KB, Church T: The elimination of exit block as a pacing complication using a transvenous steroid-eluting lead. Pacing Clin Electrophysiol 10:748, 1987. 237. Till JA, Jones S, Rowland E, et al: Clinical experience with a steroid eluting lead in children [abstract]. Circulation 80:389, 1989. 238. Sowton E, Norman J: Variations in cardiac stimulation thresholds in patients with pacing electrodes. In Digest of the 7th International Conference on Medical and Biological Engineering, 1967, Stockholm. 239. Preston TA, Fletcher RD, Luchesi BR, et al: Changes in myocardial threshold: physiologic and pharmacologic factors in patients with implanted pacemakers. Am Heart J 74:235-242, 1967. 240. Westerholm C-J: Threshold studies in transvenous cardiac pacemaker treatment. Scand J Thorac Surg Suppl 8(suppl):1-35, 1971. 241. McVenes R, Lahtinen S, Hansen N, Stokes K: Physiologic and drug induced changes in cardiac pacing and sensing parameters [abstract 324]. Eur J Card Pacing Electrophysiol 2:A86, 1992. 242. Kadish A, Kong T, Goldberger J: Diurnal variability in ventricular stimulation threshold and electrogram amplitude [abstract]. Eur J Card Pacing Electrophysiol 2:A86, 1992. 243. Grendahl H, Schaanning CG: Variations in pacing threshold. Acta Med Scand 187:75-78, 1970. 244. Barold S: Effect of drugs on pacing thresholds. In Antonioli GE, Aubert AE, Ector H, editors: Pacemaker leads 1991, New York, 1991, Elsevier, pp 73-86. 245. Barold SS, Ong LS, Heinle RA. Stimulation and sensing thresholds for cardiac pacing: Electrophysiologic and technical aspects. Prog Cardiovasc Dis 24:1, 1981. 246. Ohm O-J, Breivik K: Pacing leads. In Gomez FP, editor: Cardiac pacing, electrophysiology, tachyarrhythmias, Madrid, 1985, Editorial Group, pp 971-985. 247. Basu D, Chatterjee K: Unusually high pacemaker threshold in severe myedema: decrease with thyroid hormone therapy. Chest 70:677-679, 1976.

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2 

SECTION 1  Basic Principles of Device Therapy

Principles of Defibrillation: From Cellular Physiology to Fields and Waveforms GREGORY P. WALCOTT  |  STEVEN M. POGWIZD  |  RAYMOND E. IDEKER

E

lectrical defibrillation is the only practical means for halting ventricular fibrillation (VF). Although it has been known for more than a century that application of an electric shock directly to the myocardium causes VF, and that the heart can be returned to normal rhythm by subsequent application of a shock of greater magnitude,1-3 knowledge of the mechanisms underlying the process of defibrillation was slow in developing. Only the introduction of novel techniques for the analysis of action potentials and activation sequences has allowed greater insight into the physiology of both fibrillation and defibrillation.4-8 It is hoped that this insight will result in a higher success rate for external defibrillation and improved design of implantable cardioverter-defibrillators (ICDs). A large part of current research is dedicated to determining the underlying reason for the success or failure of a defibrillating shock. VF is maintained by multiple activation fronts that are constantly moving in a pattern of reentry. Characteristics of the activation pattern and action potential are believed to be important determinants of whether a shock will successfully defibrillate the heart. A successful defibrillating shock is believed to extinguish most of these activation fronts, permitting the resumption of coordinated responsiveness.9-13 For the defibrillating shock to be completely successful, this must be accomplished without creating an environment that promotes susceptibility to reinitiation of fibrillation.9,10 It has been established that successful defibrillation requires that the shock results in adequate distribution of the potential gradient (a surrogate for local current flow) throughout the ventricular myocardium.14-16 Fundamentally, defibrillation is believed to be realized through an electrical pulse that causes an alteration in the transmembrane potential of myocytes. It most likely requires a rapid induction of changes in the transmembrane potential of the myocytes in a critical mass of myocardium (75%-90% of myocardium in dogs).13,14,16 Because this represents a large mass of tissue, depolarization must be achieved at a considerable distance from the stimulating electrode. To gain an understanding of this complex far-field process, various mathematical models have been generated and predictions of computer simulations compared with physiologic findings. Both discontinuities in the anisotropic properties of the extracellular and intracellular domains, as described by the bidomain model,17,18 and highly resistive discontinuities in the intracellular space (e.g., collagenous septa), as described in the secondary source model,19-21 may contribute to the far-field changes in the potential gradient that halt the activation fronts of fibrillation; these are discussed later. The mechanisms underlying degeneration into fibrillation in failed shocks remain incompletely understood. Residual wandering wavelets,22 nonuniform refractoriness,23 and areas of low potential gradient, in which critical points (centers of reentrant circuits) form,15 may be the sources of propagating wavefronts that can result in fibrillation through reentry. Centrifugal propagation from ectopic foci induced by the defibrillation shock may also play a role, especially in the atrium.10 This chapter expands on the subjects just mentioned, discussing characteristics of VF important to understanding defibrillation and characteristics of shock that lead to successful defibrillation, such as waveform shape and electrode configuration. A shock is traced from its origin at the defibrillation electrodes to its distribution through the

40

heart, with a discussion of its effect on the transmembrane potential and how it leads to the successful cessation of fibrillation.

Fibrillation Understanding defibrillation requires an understanding of fibrillation. Knowing the basic characteristics of VF and whether it is maintained by reentrant or focal activity, as well as the characteristics of the action potential and the excitability of the fibrillating tissue, helps to define the therapy that will be successful in stopping the arrhythmia. Ventricular fibrillation has been characterized as progressing through four stages, based on high-speed cinematography of electrically induced fibrillation in dog hearts24 (Fig. 2-1). A brief undulatory, or tachysystolic, stage lasting only 1 to 2 seconds occurs first. It is characterized by three to six undulatory contractions that resemble a series of closely occurring systoles and involve the sequential contraction of large areas of the myocardium. This is followed by a second stage of convulsive incoordination (15-45 seconds) during which more frequent waves of contraction sweep over smaller regions of the myocardium. Because the contractions in each region are not in phase, the ventricles are pulled in a convulsive manner. It is during this stage of fibrillation that the ICD shocks are given, about 10 to 20 seconds after the onset of fibrillation. In the third stage of tremulous incoordination, the independently contracting areas of the ventricular surface become even smaller, giving the heart a “trembling” appearance. Tremulous incoordination lasts for 2 to 4 minutes, before the fourth and final stage of atonic fibrillation occurs. Atonic fibrillation develops within 3 to 5 minutes after the onset of fibrillation and is characterized by the slow passage of feeble contraction wavelets over short distances. With time, the number of quiescent areas increases. Ischemia plays a role in the development of the third and fourth stages, because the fibrillating heart remains in the second stage if the coronary arteries are perfused with oxygenated blood.25,26 Driving the mechanical activity of the heart during fibrillation is the electrical activity of the myocardium. The electrical activity of the heart during fibrillation has been studied using both extracellular and optical recordings. Several groups have suggested that fibrillation is maintained by reentry. In most cases, reentry appears to be caused by “wandering wavelets” of activation, activation fronts that follow continually changing pathways from cycle to cycle. In some studies, the activation sequence appears moderately repeatable from cycle to cycle, following approximately the same pathway.27-29 Occasionally, a spiraling pattern of functional reentry emanates from the same region for several cycles. Sometimes, the central core of these spiral waves meanders across the heart.5 At other times, new reentrant activation fronts are generated when one front interacts with another during its vulnerable period. Study of VF suggests that there is a level of organization to the seemingly random patterns of wandering wavelets. Two competing hypotheses have been proposed to explain this organization. The “mother rotor” hypothesis was first proposed by Lewis30 in 1925 and revived in 1996 by Gray and Jalife.31 This hypothesis proposes that a single, stationary reentrant circuit (or mother rotor), located in the fastestactivating region of the heart, “drives” VF by giving rise to activation fronts that propagate throughout the remainder of the myocardium.29



2  Principles of Defibrillation: From Cellular Physiology to Fields and Waveforms

41

CELLULAR ACTION POTENTIAL AND EXCITABLE GAP DURING FIBRILLATION y y

W

y

y

X

B

A

C

D

E

F

Figure 2-1  Spread of waves in analysis of motion pictures obtained during four stages of fibrillation described by Wiggers. A, Spread of wavefront during initial, undulatory stage. B, Theoretical passage of impulses from point x to form a wavefront at y. C through F show the appearance of contraction waves in the small rectangular area W from A, magnified. C, Undulatory stage; D, convulsive stage; E, tremulous stage; F, atonic stage. (From Wiggers CJ: Studies of ventricular fibrillation caused by electric shock: cinematographic and electrocardiographic observations of the natural process in the dog’s heart—its inhibition by potassium and the revival of coordinated beats by calcium. Am Heart J 5:351-365, 1930.)

These wavefronts propagate away from this fast-activating region, encounter areas of unidirectional block, and break up into smaller, slower-moving waveforms that resemble Wiggers’ wandering wavelets. Experiments best demonstrating the “mother” rotor have been performed in small hearts, from guinea pigs or rabbits. Studies of larger hearts have shown areas of faster and slower activation across the heart, but the existence of a single reentrant rotor that drives fibrillation has not been clearly demonstrated.32,33 In contrast to the mother rotor hypothesis is the “restitution” hypothesis. Restitution properties of the heart have been recognized for many years. Restitution in the heart refers to the relationship between the duration of an action potential in a particular cell and the duration of the previous diastolic or resting interval. If the previous diastolic interval (DI) is short, the current action potential duration (APD) will also be short. If the previous DI is long, the current APD will be long. For a regular rhythm, the preceding DI is constant and, therefore, so is the duration of the subsequent APD. The relationship between an APD and the previous DI is often described graphically as a plot of APDn versus DIn-1 (Fig. 2-2). The steepness of the restitution curve is an important characteristic of this curve, especially at short DIs. If this slope is greater than 1, then, at a constant cycle length, a single perturbation in DI will cause the ensuing APDs and DIs to oscillate, with the oscillations progressively increasing until the site is refractory at the time of the next cycle, causing conduction block and VF initiation. During VF, it is hypothesized that, when the slope of the restitution curve is greater than 1, oscillations in DI and APD increase until block occurs and wavefronts break up. Figure 2-2 shows an example of a restitution curve recorded from the right ventricle of a pig. The relationship between a DI and the subsequent APD is well defined during paced rhythm but less well defined during VF. Understanding how VF is maintained may help develop new therapies that will make fibrillation easier to stop. If VF is maintained by a mother rotor, targeting of electrical therapy, either shocks or pacing, to the region that contains the dominant reentrant circuit may be successful in halting VF. If VF is started and maintained primarily by the DI restitution properties of the heart, drugs that decrease the slope of the restitution curve, especially at short coupling intervals, may be successful in halting VF.

In the past few years, knowledge of the characteristics of the action potentials during fibrillation has increased greatly. This is a direct result of the introduction of techniques for recording action potentials in whole hearts, either in vivo31 or in perfused, isolated hearts.4-6,10,31 During fibrillation, the action potentials are altered: the APD is decreased; the action potential upstroke is slowed (decreased firstorder derivative, dV/dt) and of decreased magnitude; the plateau phase is abbreviated; and DIs are abbreviated or absent (Fig. 2-3). During the first few seconds of VF (or atrial fibrillation), the activation rate is quite rapid; the mean cycle length of VF in patients undergoing defibrillator implantation was 213 ± 27 msec.6 DIs are rarely seen during early fibrillation, and the upstroke of most action potentials occurs before the transmembrane potential has returned to baseline from the previous action potential. The demonstration of an excitable gap in fibrillating atrial tissue,34 as well as evidence of an excitable gap in fibrillating ventricular tissue,35 suggests that there are periods late in the action potential in the fibrillating myocardium during which an electrical stimulus can capture a portion of the fibrillating myocardium. Knowledge of an excitable gap provides an opportunity to stimulate the tissue just in front of a fibrillating wavefront, to cause wavefront block. As described in Chapter 1, the electrical activity of the heart is controlled ultimately by ion channels located in the cell membrane of the myocyte. It has been established that both the voltage-gated fast channels (sodium [Na+]) and slow channels (Na+ and calcium [Ca2+]) are active during the first few seconds of VF.8 The fast-channel activity is indicated by the rapidity of the upstroke of the action potential (phase

200 180 160 APD60 (msec)

y

140 120 100 80 60 40 0

20

40

60

80

100

DI (msec) Figure 2-2  Restitution curve recorded from right ventricle of pig. Dynamic action potential duration APD60 restitution relationship during pacing and ventricular fibrillation (VF) in one animal. (APD60 is the action potential duration at 60% of return to resting membrane voltage.) Open circles represent data from pacing, and solid squares represent data throughout 60 seconds of VF. Data were recorded from the anterior right ventricle using a floating microelectrode. The heart was stimulated using decremental pacing at an initial pacing rate of 1 pulse every 450 msec. The stimulus-to-stimulus interval was progressively shortened until either the heart was refractory to the stimulation or VF was induced. Note that the open circles form an exponential relationship between the diastolic interval (DI) and APD60, whereas the relationship between DI and APD60 is not well defined during VF. (From Huang J, Zhou X, Smith WM, Ideker RE: Restitution properties during ventricular fibrillation in the in situ swine heart. Circulation 110:3161-3167, 2004.)

42

SECTION 1  Basic Principles of Device Therapy

I II III

MAP 250 msec ECM Figure 2-3  Recording during ventricular fibrillation in a human. Leads I, II, and III are body surface electrocardiograms. Note that there is no period of diastole between action potentials. MAP, Right ventricular monophasic action potentials; ECM, local bipolar electrogram. (From Swartz JF, Jones JL, Fletcher RD: Characterization of ventricular fibrillation based on monophasic action potential morphology in the human heart. Circulation 87:1907-1914, 1993.)

0) during early fibrillation and its sensitivity to administration of the Na+ channel blocker, tetrodotoxin. As fibrillation proceeds, the upstrokes of the action potentials become increasingly slower, with a decreased dV/dtmax, but the action potentials remain sensitive to tetrodotoxin until 1 to 5 minutes after initiation of fibrillation. A transition then occurs in which the action potential upstrokes become insensitive to tetrodotoxin. This suggests that the propagation of the action potential is no longer mediated primarily by the fast voltage-gated Na+ channels and may be mediated by slow voltage-gated Ca2+ channel activity in the later stages of fibrillation.4 The activation complexes recorded from the ventricular myocardium remain active only as long as the coronary arteries are perfused with oxygenated blood, suggesting that ischemia may be responsible for the loss of fast-channel activation during prolonged VF.26

Defibrillation As previously mentioned, application of a powerful electrical shock to the heart is the only reliable means of stopping fibrillation. Successful defibrillation can reflect either the immediate cessation of all activation fronts or the cessation of activation fronts after two to three cycles,11,36 followed by coordinated beating of the heart. Unsuccessful defibrillation can reflect a failure to inhibit the fibrillating activation fronts or the resumption of fibrillating activation fronts after their initial inhibition. WAVEFORMS, CURRENT STRENGTH, AND DISTRIBUTION DURING DEFIBRILLATION The two most common waveform shapes used clinically are the monophasic and biphasic waveforms. In monophasic waveforms, the polarity of the shock is unchanged at each electrode for the entire duration of the electrical shock. In biphasic waveforms, the polarity of the shock reverses at each electrode partway through the defibrillation waveform. Many studies, in both animals and humans, have shown that biphasic waveforms can defibrillate with less current and energy than monophasic waveforms, in both internal and transthoracic defibrillation configurations.37-40 Within each type, waveforms can be described as truncated exponential or damped sinusoidal shapes. ICDs use truncated exponential biphasic waveforms. Most external defibrillators have used damped sinusoidal monophasic waveforms, but because of the inductor necessary to shape the waveform, these defibrillators tend to be large and heavy. More recently, smaller, lighter external defibrillators have been developed that use truncated exponential biphasic waveforms similar to those used in ICDs. Damped sinusoidal biphasic waveforms are used in external defibrillators in Russia; similar to truncated exponential biphasic waveforms, these show improved efficacy over monophasic waveforms.41,42

However, not all biphasic waveforms are superior to monophasic waveforms, For example, if the second phase of the biphasic waveform becomes much longer than the first phase, the energy required for defibrillation increases and can eventually rise to a level greater than the energy required to defibrillate with a monophasic waveform (with duration equal to the first phase of the biphasic waveform).40,43,44 The optimum duration of the two phases of the biphasic waveform depends on the electrode impedance and the defibrillator capacitance.45-48 Several groups have shown that defibrillation efficacy for square waveforms follows a strength-duration relationship similar to that for cardiac stimulation;49,50 as the waveform becomes longer, the average current at the 50% success point (the current when one half of delivered shocks will succeed) becomes progressively less, approaching an asymptote called the rheobase.51 On the basis of this observation, some suggest that cardiac defibrillation can be mathematically modeled using a resistor-capacitor (RC) network to represent the heart45,51-53 (Fig. 2-4). As empirically determined, the time constant for the parallel RC network is 2.5 to 5 msec.45,47,53 In one version of the model,53 a current waveform is applied to the RC network. The voltage across the network is then calculated for each time point during the defibrillation pulse. The relative efficacies of different waveform shapes and durations can be compared by determining the current strength that is necessary to make the voltage across the RC network reach a particular value, called the defibrillation threshold. Several observations can be made from this model. First, for square waves, as the waveform duration becomes longer, the voltage across the network increases, approaching an asymptote, or rheobase. For truncated exponential waveforms, however, the model voltage rises, reaches a peak, and then, if the waveform is long enough, begins to decrease (see Fig. 2-4). Therefore, the model predicts that monophasic exponential waveforms should be truncated at a time when the peak voltage across the RC network is reached. Current or energy delivered after that point is wasted. In support of this prediction, strength-duration relationships measured in both animals53 and humans54 do not approach an asymptote but rather reach a minimum and remain constant as the waveform lengthens. This minimum occurs over a range of waveform durations and does not extend indefinitely. Schuder et al.55 showed that, if the duration of a waveform becomes too long, defibrillation efficacy decreases. Second, the model predicts that the heart acts as a low-pass filter.52 Therefore, waveforms that rise gradually should have better efficacy than waveforms that turn on immediately. This prediction has been shown to hold true for external defibrillation,56 internal atrial defibrillation,57 and internal ventricular defibrillation.58 Ascending ramps defibrillate with a greater efficacy than do descending ramps.58,59 Third, several groups have suggested that the optimal first phase of a biphasic waveform is the optimal monophasic waveform.46,53 If so, what does the model predict as the “best” second phase of a biphasic waveform? Empirically, the role of the second phase apparently is to return the model voltage response back to zero as quickly as possible, thereby maximizing the increased efficacy of the biphasic waveform over that of the monophasic waveform with the same duration as phase 1 of the biphasic waveform. If the network voltage does not reach zero, or if it overshoots zero, efficacy is lost. Swerdlow et al.47 showed in humans that the “best” phase 2 of a biphasic waveform returns the model response close to zero. Together, these models allow the clinician to choose optimal capacitor sizes for truncated exponential biphasic waveforms, the waveforms most often used in ICDs. The capacitor must be large enough to be able to raise the network voltage to its threshold value and still hold enough charge to drive the network voltage back to zero. For a 40-Ω interelectrode impedance and a network time constant of 2.8 msec, the minimum capacitor that can accomplish this has a capacitance of 75 microfarads (µF). These models can also help in choosing shock parameters, including phase duration for biphasic waveforms. Shock duration should be varied depending on defibrillation lead impedance; patients with low impedance should receive short-duration shocks, whereas patients

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Figure 2-4  Parallel resistor-capacitor (RC) network representation of heart. Responses of RC network to monophasic and biphasic truncated exponential waveforms with a time constant of 7 msec. The parallel RC network has a time constant of 2.8 msec. A, Input monophasic waveforms. The leading-edge current of the input waveform was 10 A. The waveforms were truncated at 1, 2, 3, 4, 5, 6, 8, and 10 msec.  B, Model response, V(t). Initially, as the waveform lengthens, V(t) increases until it reaches a maximum at about 4 msec, after which it begins to decrease. C, Input biphasic waveforms. The leading-edge current was 10 A. Phase 1 was truncated at 6 msec. Phase 2 was truncated after 1, 2, 3, 4, 5, 6, 7, and 8 msec. D, Model response does not change polarity until the phase 2 duration is longer than 2 msec. (From Walcott GP, Walker RG, Cates AW, et al: Choosing the optimal monophasic and biphasic waveforms for ventricular defibrillation. J Cardiovasc Electrophysiol 6:737-750, 1995.)

with high impedance should receive longer-duration shocks. Initially, the concept of “tilt” was used to vary shock phase durations with impedance. Tilt is defined as the quotient of the leading-edge voltage minus the trailing voltage at the end of the shock divided by the voltage at the beginning of the shock, expressed as a percentage. For example, a shock with starting voltage of 500 V and ending voltage of 100 V has a tilt of 80%. If tilt is held constant, shock duration varies linearly with shock impedance. Compared with model predictions (Fig. 2-5), fixedtilt waveforms are too short for patients with low impedance and too

long for patients with higher impedance. One defibrillator manufacturer has developed a reference table that gives waveform phase durations as a function of patient impedance using a model similar to the one previously presented.60,61 Their model uses “model time constants” of 2.5, 3.5, and 4.5 msec. It is recommended that the 3.5-msec constant be used first. If the defibrillation threshold for the 3.5-msec waveform is unsatisfactory, using the 4.5-msec constant is likely to be more effective than 2.5 msec. Waveforms that are too short tend to fail; waveforms that are too long tend to waste shock energy yet still defibrillate. The location of the defibrillation electrodes affects the magnitude of the shock necessary to defibrillate the heart. Typically, 200 to 360 J of energy are necessary for successful defibrillation, with the defibrillation electrodes located on the body surface, during transthoracic defibrillation with a damped sinusoidal monophasic waveform. Although less energy is required for a truncated exponential biphasic waveform,62 only about 4% to 20% of the current that is delivered to transthoracic defibrillation electrodes ever reaches the heart.63,64 Indeed, when the defibrillation electrodes are placed in the heart itself, usually only 20 to 34 J of energy is required, and the requirement may be as low as only a few joules when large, contoured epicardial electrodes are used.40,65 The strength of the shock also varies for different locations on or in the heart; epicardial patches defibrillate with a lower shock energy than transvenous electrode configurations.66 Although defibrillation efficacy is usually described by some measure of defibrillation shock “strength” (energy, voltage, or current), little insight into the mechanisms of defibrillation can be obtained from these measures. Knowing how the current (or voltage) of a defibrillation shock is distributed over the heart allows greater understanding of how defibrillation occurs. Several studies have measured the potential gradient distribution throughout the heart during a defibrillation shock.14,67,68 The potential gradient is a measure of the spatial variation of shock voltage across the heart. The potential gradient is measured in volts per centimeter (V/cm) of tissue. In a region with a high potential gradient, the difference in voltage between a given point and an area 1 cm distant from that point is high. Regions of low potential gradient have a measured voltage that is similar to that of nearby points. These studies show an uneven potential distribution for most electrode configurations, with areas of high potential gradient near the defibrillation electrodes and areas of low potential gradient in regions distant from the defibrillating electrodes. It has been hypothesized that a minimum potential gradient must be attained for successful defibrillation to occur, and that this requirement is independent of the current applied or the electrode configuration.15,16 After a shock that fails to defibrillate VF, the site of earliest activation immediately after the 65%

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2  Principles of Defibrillation: From Cellular Physiology to Fields and Waveforms

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44

SECTION 1  Basic Principles of Device Therapy

shock can be mapped and related to the electric field that was produced by the shock. For shocks near the defibrillation threshold, the sites of earliest activation after a failed shock occur in the areas of lowest potential gradient. The minimum potential gradient required for defibrillation is lower for biphasic than for monophasic waveforms (4 vs. 6 V/cm). A minimum potential gradient of 6 V/cm was required for successful defibrillation using a 10-msec truncated exponential monophasic waveform in the open-chest dog model.16 Similar findings were observed using a 14-msec truncated exponential monophasic waveform and multiple electrode configurations.15 In contrast, a minimum potential gradient of 4 V/cm was required for successful defibrillation using a truncated exponential biphasic waveform. Because higher shock strengths are required to induce a higher potential gradient, biphasic shocks successfully defibrillate with lower energy than monophasic shocks (i.e., a lower-voltage gradient is required). The requirement for a minimum potential gradient may reflect the need for a shock to prevent the generation of new activation fronts that can result in reinitiation of fibrillation.69 Examination of activation patterns after failed defibrillation for progressively larger shock strengths indicate that postshock activation occurs at numerous sites throughout the ventricle, and that reentry is common when the shock strength is much lower than that needed for defibrillation70 (Fig. 2-6). At shock strengths just lower than those required for defibrillation, postshock activation arises in a limited number of myocardial regions. The activation fronts then propagate to activate other regions of the myocardium for a few cycles before reentry occurs; activation becomes disorganized; and fibrillation is reinitiated. Although postshock activation sites can still arise in regions of lowest potential gradient after a shock slightly greater than that required for defibrillation, the cycles of activation that originate from these sites are slower. These activations terminate after a few cycles without reinitiating fibrillation.69-71 At least two inferences can be drawn from the model that the lowest potential gradient region determines whether or not a shock will halt fibrillation. First, different shock electrode configurations with the same defibrillation threshold (DFT) have the same lowest potential gradient value, but in different regions of the heart. Single-coil versus dual-coil defibrillator systems have similar DFTs. Mapping of postshock activations shows that the earliest recorded activity after the shock is the anterior left ventricle for dual-coil shocks but the posterior area for single-coil shocks.72 Thus the low gradient region presumably is anterior for dual-coil shocks and posterior for single-coil shocks, although this inference needs to be confirmed experimentally. Crossley et al.73 measured DFTs for shocks from either an electrode in the right ventricular (RV) apex or RV outflow tract to the pulse generator can and showed that DFTs were not significantly different for the two configurations. We can infer that the lowest potential gradient level was the same for each configuration, but likely in a different location. Second, high DFTs occur in some patients because either (1) the geometry of their heart is such that the low potential gradient region has a lower potential gradient than in most patients, or (2) these patients require a higher gradient level in the low gradient region to defibrillate. In either case, unusual lead configurations may help in these patients. A subcutaneous array can lower high DFTs,74 most likely by directing current to the left ventricular free wall. Likewise, an azygos vein electrode can be useful in lowering high thresholds,75 likely by directing current to the posterior portion of the heart. The previous discussion shows how defibrillation can fail because a shock is of insufficient strength. Another issue involves the defibrillation shock that becomes extremely large. At high shock strengths, the probability of defibrillation success again begins to decrease. It is believed that, at large strengths, defibrillation shocks can have detrimental effects on the heart. Increasing the shock strength to very high levels (>1000 V with transvenous electrodes) can result in activation fronts arising from regions of high potential gradient that reinduce VF.76 Cates et al.77 showed that, for both monophasic and biphasic shocks, increasing shock strength does not always improve the probability of successful defibrillation and may even increase the incidence

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Figure 2-6  Phase maps of single rabbit heart. Maps show the last cycle before a shock (left side) and the first cycle after a shock (right side) of 100 V (A), 200 V (B), 600 V (C), and 800 V (D) that failed to defibrillate. The defibrillation threshold was 800 ± 200 V in this series. Colors represent phase, and the symbols + and − indicate phase singularities or centers of reentrant circuits of opposite direction. Phase singularities as a marker of reentry were observed during ventricular fibrillation just before the shock in all cases. A and B, Postshock phase singularities were observed after failed 100-V and 200-V shocks. Visual analysis of animations of the optical recordings indicated that many of the phase singularities represented reentrant activations occurring immediately after the shock, so that the postshock interval was 0 msec. C and D, No phase singularities were observed after the 600-V and 800-V shocks. For the 600-V shock, activation propagated away in all directions from two early sites, at the apex and at the lateral base of the left ventricle, both of which appeared after a postshock interval of 42 msec. For the 800-V shock, a single wavefront of activation appeared at the apex and propagated away in all directions in a focal pattern after a postshock interval of 72 msec. (From Chattipakorn N, Banville I, Gray RA, Ideker RE: Effects of shock strengths on ventricular defibrillation failure. Cardiovasc Res 61:39-44, 2004.)

of postshock arrhythmias. Chapman et al.78 showed in dogs that the time required for the heart to recover hemodynamically after a defibrillation episode was shorter for biphasic than for monophasic shocks. Further, they showed that hemodynamic recovery took longer after high-energy shocks than after low-energy shocks. Reddy et al.79 showed that transthoracic defibrillation with biphasic shocks resulted in less postshock electrocardiographic evidence of myocardial dysfunction (injury or ischemia) than standard monophasic damped sinusoidal waveforms, and without compromise of defibrillation efficacy. One mechanism implicated in the means by which shocks cause damage to the myocardium is electroporation, the formation of holes or pores in the cell membrane. Electroporation may occur in regions where the shock potential gradient is high (>50 to 70 V/cm) and may even occur in regions where the potential gradient is much less than 50 V/cm.80 The very high voltage can result in disruption of the phospholipid membrane bilayer and in the formation of pores that permit



2  Principles of Defibrillation: From Cellular Physiology to Fields and Waveforms

the free influx and efflux of ions and micromolecules. Electroporation can cause the transmembrane potential to change temporarily to a value almost equal that of the plateau of the action potential. At this transmembrane potential, the cell is paralyzed electrically, being both unresponsive and unable to conduct an action potential. Exposure of the myocardium to yet higher potential gradients, probably greater than 150 V/cm, results in arrhythmic beating, and at extremely high potential gradients, necrosis may occur.81 The shape of the waveform alters the strength of the shock at which these detrimental effects occur. Use of a 10-msec, truncated exponential monophasic waveform for VF in dogs resulted in conduction block in regions where the potential gradient was greater than 64 ± 4 V/cm.82 Shocks that created even higher potential gradients in the myocardium (71 ± 6 V/cm) were required for conduction block when a 5-msec/5msec, truncated exponential biphasic shock was used. Adding a second phase to a monophasic waveform, thereby making it a biphasic waveform, reduced the damage sustained by cultured chick myocytes compared with that induced by the monophasic waveform alone.83 Therefore, biphasic waveforms are less likely to cause damage or dysfunction in high-gradient regions than monophasic waveforms. MODELS PROPOSED TO EXPLAIN INDUCTION OF CHANGES IN TRANSMEMBRANE POTENTIAL THROUGHOUT HEART DURING DEFIBRILLATION SHOCK A shock in the form of a square wave given across the defibrillation electrodes appears almost immediately as a square wave in the extracellular space of the heart. There is no significant distortion, because the extracellular space throughout the body is primarily resistive, with little reactive component. Phase delays and alterations of the appearance of the shock wave occur in the transmembrane potential, however, because of the capacitance and ion channels of the myocyte membrane.84 Consequently, a square-wave shock can elicit an exponential change with time in the transmembrane potential (Fig. 2-7). The nonlinear behavior of the membrane caused by the ion channels also affects the outcome of reversing the polarity of the defibrillation shock. Reversing the polarity may reverse the sign of the change in the transmembrane potential in some regions of the myocardium, and

Extracellular space

10 msec

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B Figure 2-7  Effect of square-wave shock on extracellular potential and transmembrane potential. The square-wave shock appears immediately as a relatively undistorted square wave in the extracellular space. It appears as an exponentially increasing change in the transmembrane potential. When a shock of a given polarity is delivered during the action potential plateau (A), the depolarization obtained is different in magnitude and time course from the hyperpolarization obtained with a shock of the opposite polarity (B). (From Walcott GP, Knisley SB, Zhou X, et al: On the mechanism of ventricular defibrillation. Pacing Clin Electrophysiol 20:422-431, 1997.)

45

the nonlinear behavior of the membrane can alter the magnitude and the time course of this change. As discussed previously, reversing the polarity of the shock may not reverse the sign of the change in the transmembrane potential in all regions of the heart; some areas may be hyperpolarized with both shock polarities.85 This may reflect the nonlinear behavior of the membrane ion channels. Several models have been formulated in an attempt to explain the mechanisms by which the defibrillation shock is distributed throughout the myocardium to restore coordinated, effective action potentials. As yet, none of the models adequately describes all the experimental findings on the action potential changes during defibrillation. It is well established experimentally that changes occur many centimeters from the defibrillating shock electrodes. These changes in transmembrane potential can result in new action potentials or prolongation of the action potential as described previously.23,86 Direct excitation can be observed,23 even far from the electrode (>30 mm),87 as well as across the entire heart.88 Although the one-dimensional cable model described in Chapter 1 adequately describes the generation of self-propagating action potentials close to an electrode as required for pacing, it fails to account for the far-field changes observed during defibrillation. During stimulation or defibrillation, this model predicts that the tissue near the anode should be hyperpolarized, whereas the tissue near the cathode should be depolarized.89 The magnitude of the hyperpolarization or depolarization decreases exponentially with the distance from the electrodes according to the membrane space constant, the distance at which the hyperpolarization or depolarization has decreased by 63%. For cardiac tissue, the space constant is only 0.5 to 1 mm.89,90 Therefore, the onedimensional cable equations predict that tissue more than 10 space constants (~1 cm) distant from the defibrillation electrodes should not directly undergo changes in transmembrane potential because of the shock field. That is, new action potentials should not arise by direct excitation at distances greater than 1 cm from the electrodes. This model fails to describe the experimentally observed global distribution of action potentials during defibrillation. Therefore, several additional mathematical formulations have been proposed, including the sawtooth model,21,91-93 the formation of secondary sources at barriers in the myocardium,20,94 and the bidomain model,95 to explain how a defibrillation shock affects the transmembrane potential a long distance from the shocking electrodes. In the simplest formulation of these models, the extracellular and intracellular spaces are considered to be low-resistance media and the membrane to be a high-resistance medium in parallel with capacitance. The simple case models incorporate only passive myocardial properties. The models have been rendered more realistic by the addition of active components to represent the ion channels in the membrane, gap junctions, and membrane discontinuities.96,97 By convention, the current is defined as the flow of positive ions from the anode to the cathode. Sawtooth Model The one-dimensional cable model positions two low-resistance continuous spaces that conduct current from the shock, the intracellular space and the extracellular space, separated by a high-resistance cell membrane. In the sawtooth model the intracellular space is divided by a series of high-resistance barriers, the gap junctions. Because of these high-resistance barriers, current moving in the intracellular space is forced to exit into the extracellular space and reenter the cell on the other side of the barrier. Exit and reentry of the current from the intracellular domain results in hyperpolarization near the end of the cell closest to the anode and depolarization near the end of the cell closest to the cathode. A tracing of the changes in transmembrane potential along a fiber during the shock should therefore resemble the teeth of a saw, with each tooth corresponding to an individual cell21,91-93 (Fig. 2-8). Increases in the junctional resistances are predicted to increase the magnitude of the potential changes at the ends of the cell.21 Although gap junctions are of low resistivity, they can present significant junctional resistance under certain conditions, such as hypoxia98,99 and calcium depletion.100 As the resistance of the gap junctions

SECTION 1  Basic Principles of Device Therapy

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Figure 2-8  Sawtooth model of transmembrane potential during a shock. A, Transmembrane potential shown is a summation of the membrane potential profile of the cable model and the periodicity arising from the periodic changes in intracellular resistance. The anode is to the left and the cathode is to the right in this one-dimensional model. The fiber is divided into 31 cells of equal length separated by junctions of high resistance. In the figure, the junctional resistance is shown much greater than is believed to occur in cardiac fibers, to allow the sawtooth pattern to be seen. B, Two parts of the summation in panel A are shown; Vm0, transmembrane potential profile of cable model; Vm1, periodicity arising from periodic changes in intracellular resistance. (Modified from Krassowska W, Pilkington TC, Ideker RE: Periodic conductivity as a mechanism for cardiac stimulation and defibrillation. IEEE Trans Biomed Eng 34:555-560, 1987.)

increases, it is predicted that a greater fraction of current passes preferentially across the cell membrane rather than along the cell. The sawtooth model adequately describes the requirement for a minimum potential gradient, because the magnitudes of the hyperpolarization and polarization at the ends of the cells are directly proportional to the strength of the stimulus. It also adequately describes the generation of action potentials at a distance from the electrodes and the differences in threshold stimuli between cathodal and anodal stimulation. Sawtooth changes in transmembrane potential have been observed in preparations of isolated cardiomyocytes;101,102 however, such a pattern in isolated cells would be consistent with the cable model. This pattern has not been observed in a syncytium of cardiac cells.20,103,104 Secondary Source Model Although the resistivity of the gap junctions at the boundaries between the cells may not adequately explain the physiologic effects of defibrillation, the resistivity of other intracellular discontinuities and

interruptions may well play a role. Most theories concerning the generation of action potentials and their propagation across the ventricle, such as the bidomain theory described later, consider the myocardium to be a uniform electrical continuum. This assumption does not take into account the discontinuities of the intracellular domain, where the myocardium is interrupted by barriers such as connective tissue septa, blood vessels, and scar tissue. As described previously for the sawtooth model, the intracellular current, on encountering such a barrier, must leave the intracellular space, cross the barrier, and reenter the intracellular domain on the other side. Depolarization should occur on one side of the barrier and hyperpolarization on the other side. Therefore, the barrier acts as a set of electrodes during the shock, becoming a secondary source of action potentials (Fig. 2-9). These secondary sources are important causes of depolarization and hyperpolarization throughout the myocardial tissues during a shock.20 The resistive barriers act in a manner similar to that described for the sawtooth model. In this case, however, the resistive barriers represent larger discontinuities, which tend to increase with age and cardiac hypertrophy.105 Computer simulations have shown that the cathodal stimulation delivered to the myocardium near an oval scar results in three distinct activation fronts: the primary activation front and secondary fronts at the distal and proximal edges, which are generated by the exit and reentry of current from the intracellular and extracellular spaces106 (see Fig. 2-8). Optical techniques have directly recorded changes in transmembrane potentials throughout a monolayer of ventricular myocytes.20 Localized regions of depolarization and hyperpolarization coincided with discontinuities in the monolayer, resulting in slow conductance. Microscopic regions of depolarization and hyperpolarization have also been observed in isolated slabs of left ventricle, although their correlation with anatomic structures was not possible using optical mapping techniques because of the light-scattering properties of myocardium.107 The significance of secondary sources was demonstrated in whole hearts by mapping of action potentials and determination of shock thresholds before and after the generation of a transmural lesion in the myocardial walls of dogs.94 Generation of the lesion resulted in the development of a region of direct activation in the area of the lesion, in addition to the region of direct activation resulting from the stimulating electrode observed before the lesion. Furthermore, the strength of the shock required to cause direct

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Figure 2-9  Computer model of secondary sources adjacent to scar elicited by single, large, pacing pulse. Area to the right represents myocardium that contains a rectangular scar (stippled region). On the left is the blood pool with a pacing catheter in it. Note that a pacing pulse depolarizes not only tissue near the cathode but also tissue near the scar. (From Street AM: Effects of connective tissue embedded in viable cardiac tissue on propagation and pacing: implications for arrhythmias. Durham, NC, 1996, Department of Biomedical Engineering, Duke University, p 134.)



2  Principles of Defibrillation: From Cellular Physiology to Fields and Waveforms

10

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Figure 2-10  Isochronal activation maps. Maps obtained after cathodal stimulation before and after creation of a transmural incision that caused secondary sources adjacent to the lesion. Isochrones are drawn at 5-msec intervals, timed from the onset of the S1 or S2 stimulus. Arrows represent direction of activation. Darkened regions represent areas directly activated by the stimulus. Black vertical bars represent the approximate location of the transmural incision. A, S1 stimulus delivered before incision. B, 75-mA S2 stimulus delivered before incision. C, 250-mA S2 stimulus delivered before incision. D, S1 stimulus delivered after incision. E, 75-mA S2 stimulus delivered after incision. F, Orientation of the long axis of myocardial fibers. (From White JB, Walcott GP, Pollard AE, et al: Myocardial discontinuities: a substrate for producing virtual electrodes to increase directly excited areas of the myocardium by shocks. Circulation 97:1738-1745, 1998.)

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activation in the area of the lesion was less than one half of that required before generation of the lesion (Fig. 2-10). The effects of secondary sources have major implications for the probability of successful defibrillation at different shock strengths in individual patients, particularly elderly patients, as well as the potential for reentry. Furthermore, the size and placement of operative lesions may play significant roles in the success of subsequent defibrillation. Bidomain Model The bidomain model is an extension of the one-dimensional cable model into two or three dimensions. That is, the extracellular and intracellular spaces are represented as single, continuous domains that extend in two or three dimensions and are separated by the highly resistive cell membrane95 (Fig. 2-11). If the conductivities of the intracellular and extracellular spaces are constant in all directions, the

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Figure 2-11  Circuit diagram of two-dimensional bidomain model. Top of the resistor network represents the extracellular space; bottom of network represents the intracellular space. The symbol reΔx represents the extracellular resistivity in the x direction; reΔy represents the extracellular resistivity in the y direction; riΔx and riΔy represent the intracellular resistivities in the x and y directions. The rectangles represent the cell membrane. For a passive model, the rectangle would be replaced by a parallel resistor-capacitor network. For an active model, the rectangle would be replaced by a membrane ion model.

F

model collapses to the one-dimensional cable model. Anisotropy refers to the manner in which conductivities change with the direction of myocardial fiber orientation. Clerc108 showed that conductivity is higher in the direction parallel to the long axis of myocardial fibers (longitudinal) than in the direction perpendicular to the fibers (transverse) for both the intracellular and extracellular spaces. If conductivities change with direction but change the same for the intracellular and extracellular spaces, the bidomain model collapses to the onedimensional cable model. Studies have shown that the anisotropy ratio in ventricle is about 3 : 1 in the extracellular space and 10 : 1 in the intracellular space. When anisotropy ratios are used, the bidomain model begins to give new insights into how shocks change the transmembrane potential. Similar to the one-dimensional cable model, the bidomain model predicts that hyperpolarization occurs in tissue that lies under the extracellular anodal electrode. Likewise, depolarization occurs in tissue under the extracellular cathodal electrode. Unlike the one-dimensional cable model, the bidomain model also predicts that depolarization occurs along the long axis of the myocardial fibers at distances just a few millimeters from the anode. A similar effect is predicted to occur at the cathode, with hyperpolarization at distances of a few millimeters.109 Therefore, the effect on the transmembrane potential near the shocking electrode is predicted to be much more complicated by the bidomain model than by the one-dimensional cable model. The power of the bidomain model, however, is that it hypothesizes that there should be changes in the transmembrane potential, either hyperpolarization or depolarization, across the entire heart. In this model, the change in transmembrane potential elicited by the shock depends on the distribution of intracellular and extracellular current, which is affected by the change in potential gradient with distance, the distance from the electrode, and the orientation of the myocardial fibers. Experimental studies show a complex pattern of transmembrane potential changes during the delivery of a defibrillation shock, similar to those predicted by the bidomain model.85,110,111 The transmembrane potential changes that occur during the delivery of a defibrillation shock can lead to the initiation of reentry and subsequent reinitiation of fibrillation after the shock. Reentrant circuits can be described by the mathematical concept of a phase singularity.112 Phase can be used to describe the cardiac action potential, with 0 phase assigned to the upstroke of the action potential and 2Π phase assigned to the end of the action potential. A reentrant circuit can be thought of as a circle of phase starting at 0 (excitation) and continuing

48

SECTION 1  Basic Principles of Device Therapy

Postshock transmembrane voltage (mV´)

Postshock activation isochronal maps (5 ms) 1st beat 2nd beat 125 11.5  11.5 mm

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Shock Figure 2-12  Creation of shock-induced phase singularity. Left upper panel shows the change in transmembrane potential at the end of a +100/−200 V biphasic shock (i.e., at 15th msec of 16-msec shock), which resulted in a single extra beat. Scale is shown in millivolts, calibrated to a control 100-mV action potential. Point of phase singularity is indicated by the black circle. Middle upper panel shows a 5-msec isochronal map that depicts the initiation of the postshock spread of activation. The map starts at the onset of the 8-msec phase 2 of the shock (polarity reversal). Lower panels show optical recordings from several recording sites used to reconstruct the activation maps. Right upper panel shows a continuation of the reentrant activation shown in the middle panel. Reentrant activity self-terminates after encountering refractory tissue in the lower right corner of the field of view (see traces in right lower panel). (From Efimov IR, Cheng Y, Van Wagoner DR, et al: Virtual electrode-induced phase singularity: a basic mechanism of defibrillation failure. Circ Res 82:918-925, 1998.)

to 2Π (recovery). The reentrant circuit moves around a central point, called a phase singularity. Efimov et al.110 showed that a defibrillation shock can impose changes on the transmembrane potential extending from 0 phase through 2Π phase (Fig. 2-12). Thus, a reentrant circuit is generated and fibrillation induced. The authors suggest that induced reentrant circuits may be one way that defibrillation shocks can fail. The secondary source and bidomain models may not be mutually exclusive; rather, both may contribute to the changes in transmembrane potential. The exact mechanism by which an electrical pulse results in defibrillation remains incompletely understood at the level of the cell membrane and the ion channels. As the transmembrane potential attains values closer to the typical resting transmembrane potential than the usual minimum of −65 mV observed in fibrillating myocytes, this may allow the voltage-gated Na+ channels to recover sufficiently and the myocytes to regain full excitability. EFFECT OF DEFIBRILLATING SHOCK FIELD ON CELLULAR ACTION POTENTIAL The final common pathway of changes in the transmembrane potential caused by a defibrillation shock involves effects on the shape and duration of the cellular action potential. The shock can have one of three effects on the myocardium, depending on the local strength of the shock and its timing with respect to the local action potential. If the shock is delivered during the early plateau, there will be little or no prolongation of the action potential. If the shock is strong enough and is delivered relatively late during the action potential, it will initiate a new action potential. A shock that is strong enough but is delivered during early phase 3 of the action potential will modify and prolong an ongoing action potential without initiating an entirely new action potential (Fig. 2-13).

To defibrillate the heart successfully, a defibrillating shock (1) must stop most or all activation wavefronts on the heart and (2) must not reinitiate fibrillation. The extension of refractoriness hypothesis helps to explain how a shock can stop fibrillation. A shock can prolong the refractory period of an action potential without triggering a new potential if it is of sufficient strength and is delivered at an appropriate interval with respect to the upstroke.23,113 If the first activation front that forms after a defibrillation shock encounters tissue with an extended refractory period, the front will be stopped because it cannot propagate into the region of refractory tissue. Also, a defibrillation shock must not restart fibrillation. If only part of the front encounters tissue with an extended refractory period, only part of the front will be halted. The rest of the activation front will propagate forward and will eventually move into the area that could not be stimulated or that would not allow propagation (unidirectional block). This process of stimulating some tissue and creating unidirectional block in adjacent regions creates a reentrant circuit that eventually breaks down into fibrillation. The critical point is that point at which a critical shock strength intersects a critical level of refractoriness, leading to the formation of a reentrant circuit.114-116 Optical mapping has revealed another type of critical point about which reentry can occur in response to the shock electric field.110 This type of critical point is formed when the shock electric field causes a large region of hyperpolarization immediately adjacent to a large region of depolarization. According to Efimov et al.,117 a biphasic shock waveform is more efficacious because the second phase of the shock obliterates the regions of hyperpolarization and depolarization so that the critical-point reentrant wavefront does not arise. Both types of critical points share two attributes; first, they produce a wavefront that appears immediately after the shock, and second, the wavefront forms a reentrant circuit.



2  Principles of Defibrillation: From Cellular Physiology to Fields and Waveforms 1.6 V/cm, 2 msec

S1-S2 (ms)

0 (mV)

222 225

80

A

100 msec 8.4 V/cm, 2 msec S1-S2 (ms) 90 110 130, 150, 170 180 200 210 220 230

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0

80

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230 90 S1-S2 (msec) 100 msec

Figure 2-13  Transmembrane recordings from guinea pig papillary muscle. Recordings show an all-or-none response to a weak field stimulus and action potential prolongation in response to a larger stimulus. A, Recordings that illustrate the response to an S2 stimulus of 1.6 V/cm oriented along the fibers. The S1-S2 stimulus interval for each response (milliseconds) is indicated to the right of the recordings. The responses are much different, even though the change in S2 timing was only 3 msec. An S1-S2 interval of 222 msec caused almost no response, whereas an interval of 225 msec produced a new action potential.  B, Range of action potential extensions produced by an S2 stimulus generating a potential gradient of 8.4 V/cm oriented along long axis of myofibers; recordings from same cell as in A. The action potential recordings, obtained from one cellular impairment, are aligned with the S2 time. An S1 stimulus was applied 3 msec before phase 0 of each recording. The longest (230 msec) and shortest (90 msec) S1-S2 intervals tested are indicated beneath their respective phase 0 depolarizations. The S1-S2 interval for each response is indicated to the right. (From Knisley SB, Smith WM, Ideker RE: Effect of field stimulation on cellular repolarization in rabbit myocardium: Implications for reentry induction. Circ Res 70:707-715, 1992.)

REFIBRILLATION CAUSED BY DEFIBRILLATION SHOCK Although activation wavefronts consistent with the two types of critical points are observed following shocks given during the vulnerable period of paced rhythm and after shocks much weaker than the DFT during VF, these wavefronts are usually not observed after shocks near DFT strength during VF.118-123 Instead, earliest activation is not recorded until 50 to 60 msec after the shock, an interval called the isoelectric window, and wavefronts spread centrifugally away from this early site in a focal activation pattern. This pattern is present not only in electrical epicardial recordings (Fig. 2-14) but also in optical epicardial recordings (Fig. 2-15), which are not subject to saturation of the amplifiers by the shock as are electrical recordings. This activation pattern may represent epicardial breakthrough of an intramural reentrant circuit caused by a critical point. However, the same focal activation pattern has been observed in three-dimensional transmural electrical

49

recordings with plunge needles, during which an isoelectric window was also observed intramurally.122 This window has also been observed in optical recordings of a thin layer of spared epicardial tissue in which the intramural myocardium was frozen.123 Following a failed shock slightly weaker than the DFT, the focal activation pattern is seen for several cycles before the less organized pattern of VF reappears (see Fig. 2-14). However, this same focal activation pattern can be observed for one or more cycles after a successful defibrillation shock (see Fig. 2-15, B). These findings suggest that the failed shock reinitiates VF by causing a trigger (the focus) that interacts with the substrate (the ventricular tissue). If the trigger focus lasts long enough for reentry to develop in the ventricular myocardium substrate, the shock fails. If the trigger focus halts before reentry is induced in the myocardial substrate, the shock succeeds. This concept raises the possibility that cardiac disease or drugs can alter DFT by altering the trigger, the substrate, or both. For example, conditions that increase the number of cycles of the focal trigger or that decrease the cycle length of the trigger might increase DFT. Similarly, conditions that allow more easily induced reentry in the substrate, such as an increased dispersion of refractoriness, could also increase DFT. Although several mechanisms have been hypothesized, the cause of the focal activations after the shock is not definitively known, nor is it known why increasing the shock electric field decreases the number of focal postshock activations. One hypothesis is that the focus is not a true focus but arises from “tunnel propagation.”124 According to this hypothesis, a small part of a VF activation wavefront is not extinguished but propagates intramurally during the postshock isoelectric window, where it is undetected by three-dimensional mapping because it is so small. When sufficient tissue recovers, it is able to break out as a large wavefront. This wavefront appears to arise focally because the small, slowly propagating tunnel of activation is not detected by mapping. According to this hypothesis, a biphasic waveform has a lower DFT than a monophasic waveform because it decreases the number and size of the tunnels for the same-strength shock. However, the finding at multiple focal cycles is difficult to explain with the tunnel propagation hypothesis; multiple focal cycles, not just one, are present after the shock (see Fig. 2-14). Another hypothesis is that the focal activation arises where an intracellular calcium (Cai) “sinkhole” is present after the shock.123 This hypothesis is based on optical recordings demonstrating that Cai decreases during the isoelectric period in the region where the focus will later appear (Fig. 2-16). After the formation of the sinkhole, the action potential in this region is immediately preceded by an increase in Cai, suggesting the potential arises from triggered activity. According to this hypothesis, a biphasic waveform has a lower DFT than a monophasic waveform because it causes less postshock heterogeneity in Cai.125 In one study the DFT after VF of 10 seconds was significantly decreased by flunarizine, a drug that prevents activity triggered by delayed afterdepolarization (DAD).126 However, another study did not find that DFT after 10 seconds of VF was altered either by flunarizine or by pinacidil, a drug that prevents activity triggered by early afterdepolarization (EAD).127 A third study found that DFT after 20 seconds of VF was not altered by flunarizine, but that DFT was reduced by flunarizine after 7 minutes of VF by 82%.128 This suggests that after long-duration VF, but not short-duration VF, the focal postshock activation may arise from a DAD. Another finding indicates that the mechanism of defibrillation may differ depending on the duration of VF. The earliest postshock activation after the isoelectric period arises primarily in the working myocardium following VF of about 30 seconds, but arises primarily in Purkinje fibers after VF of 180 seconds.129 These differences in defibrillation after short- versus long-duration VF are probably related to the different characteristics of the activation wavefronts maintaining these short or long VFs.130-132 It has even been reported recently that endocardial ventricular activation is not reentrant but arises focally from Purkinje fibers in up to 30% of cases during long-duration VF133 (Fig. 2-17). This finding raises the startling possibility that to defibrillate

50

SECTION 1  Basic Principles of Device Therapy

64

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Figure 2-14  Example of postshock cycles after failed shock in pig. Defibrillation shock was of a strength near the defibrillation threshold (DFT), delivered from electrodes in the right ventricular apex and the superior vena cava. Recordings were made simultaneously from 504 electrodes distributed globally over the porcine ventricular epicardium. Electrode sites are indicated in gray on a polar projection, with the atrioventricular groove at the periphery and the left ventricular apex in the center. Anterior is at the bottom of the projection, and the left ventricle is to the right. Each panel shows in black the electrode sites at which an activation occurred at any time during each 10-msec interval. Numbers above the panels indicate the start of each 10-msec interval relative to the shock onset. Red arrows indicate the site of earliest recorded activation for each cycle. The first cycle appeared on the epicardium 64 msec after the shock at the anteroapical left ventricle. The second cycle (154 msec) arose on the epicardium in the same region as the first cycle and also propagated away in a focal pattern. The third (235 msec) and fourth (315 msec) cycles arose before the activation front from the previous cycle disappeared. (From Chattipakorn N, Fotuhi PC, Ideker RE: Prediction of defibrillation outcome by epicardial activation patterns following shocks near the defibrillation threshold. J Cardiovasc Electrophysiol 11:1014-1021, 2000.)



2  Principles of Defibrillation: From Cellular Physiology to Fields and Waveforms

51

RV Posterior coronary vein

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B Figure 2-15  Similar activation patterns of postshock cycle 1 after failed (A) and successful (B) shocks. Maps from a pig heart are oriented as shown at top of figure; RV, right ventricle; LV, left ventricle; LAD, left anterior descending (coronary artery). Numbers above maps are times in milliseconds relative to defibrillation shock onset. Recording sites where action potential upstroke reached 50% of the action potential amplitude at any time during each 8-msec interval are black. Activation arose focally at the left ventricular apex where the shock electric field is weak and appeared 84 msec after the shock in A and 72 msec after the shock in B. Activation wavefronts propagated focally away from the apex toward the base in an organized pattern for both the failed and the successful shock episodes. (From Chattipakorn N, Banville I, Gray RA, Ideker RE: Mechanism of ventricular defibrillation for near-defibrillation threshold shocks: a whole-heart optical mapping study in swine. Circulation 104:1313-1319, 2001.)

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SECTION 1  Basic Principles of Device Therapy

Cai

A

100 msec

Vm Vm

Activation isochrone

Cai

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–8 –4

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Figure 2-16  Intracellular calcium (Cai) “sinkhole” and earliest postshock activation during unsuccessful shock in isolated rabbit heart. A, Optical signals from the site marked by an asterisk in B. The isoelectric window is the time between the time of the defibrillation shock (red line) and the time when the postshock Vm tracing crosses F (fluorescence strength, black line). B, Optical maps after failed shock (200 V), with time of the shock as time 0. C, Consecutive isochronal activation maps (left), isochronal repolarization map (middle), and Cai maps (right). Cai sinkholes (white arrows) are seen at the same site before the onset of the repetitive activations; Act, activation; Rep, repolarization. D, Optical tracings from the center (top two tracings) and edge (bottom two tracings) of the Cai sinkhole. Horizontal line segments on the tracings represent the levels of F. (From Hwang GS, Hayashi H, Tang L, et al: Intracellular calcium and vulnerability to fibrillation and defibrillation in Langendorff-perfused rabbit ventricles. Circulation 114:2595-2603, 2006.)



2  Principles of Defibrillation: From Cellular Physiology to Fields and Waveforms

2

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RV apex

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Figure 2-17  Recordings and activation times for 15 cycles after ventricular fibrillation in dog. A, Soon after onset of VF; B, after 5 minutes of VF. Left ventricular endocardial basket recordings from 64 unipolar electrodes are shown at the top of each panel, with a right ventricular (RV) unipolar recording and ECG lead (aVR) below them. Below the ECG are the ventricular activation times recorded at the 64 basket electrodes during the same period, with each short vertical line representing an activation. At the bottom of each panel, the activations for all 64 electrodes are shown on a single line. The 15 cycles are numbered at the top of B. In A, endocardial activations are present throughout the cardiac cycles, consistent with reentry. In B, however, activation is highly synchronized with activation present during only a small part of the cardiac cycle, so that 14 large temporal gaps are present between the 15 cycles of activation in which no activations are present. This finding is inconsistent with reentry near the left ventricular endocardium. In both A and B, RV activation and ECG appear disorganized, as expected for VF. (From Robichaux RP et al: Periods of highly synchronous, non-reentrant endocardial activation cycles occur during long-duration ventricular fibrillation. J Cardiovasc Electrophysiol 21:1266-1273, 2010.)

long-duration VF, a shock must not only halt reentrant activation wavefronts, but also halt focal, possibly triggered, activity. Thus, although significant advances have been made in our understanding of the mechanisms of defibrillation, it is clear that there is still much more to be learned.

Acknowledgment This work was supported in part by National Institutes of Health grant HL-42760.

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Walcott GP, Walker RG, Krassowska W, et al: Choosing the optimum monophasic and biphasic waveforms for defibrillation. Pacing Clin Electrophysiol 17:789, 1994. 49. Blair HA: On the intensity-time relations for stimulation by electric currents. II. J Gen Physiol 15:731-755, 1932. 50. Lapicque L: L’Excitabilite en fonction du temps. Paris, 1926, Libraire J Gilbert, p 371. 51. Irnich W: The fundamental law of electrostimulation and its application to defibrillation. Pacing Clin Electrophysiol 13:14331447, 1990. 52. Sweeney RJ, Gill GM, Jones JL, Reid PR: Defibrillation using a high-frequency series of monophasic rectangular pulses: observations and model predictions. J Cardiovasc Electrophysiol 7:134143, 1996. 53. Walcott GP, Walker RG, Cates AW, et al: Choosing the optimal monophasic and biphasic waveforms for ventricular defibrillation. J Cardiovasc Electrophysiol 6:737-750, 1995. 54. Gold MR, Shorofsky SR: Strength-duration relationship for human transvenous defibrillation. Circulation 96:3517-3520, 1997. 55. Schuder JC, Stoeckle H, West JA,Kesar PY: Transthoracic ventricular defibrillation in the dog with truncated and untruncated exponential stimuli. IEEE Trans Biomed Eng 18:410-415, 1971. 56. Walcott GP, Melnick SB, Chapman FW, et al: Comparison of damped sinusoidal and truncated exponential waveforms for external defibrillation. J Am Coll Cardiol 27:237A, 1996. 57. Harbinson MT, Allen JD, Imam Z, et al: Rounded biphasic waveform reduces energy requirements for transvenous catheter cardioversion of atrial fibrillation and flutter. Pacing Clin Electrophysiol 20:226-229, 1997.

58. Hillsley RE, Walker RG, Swanson DK, et al: Is the second phase of a biphasic defibrillation waveform the defibrillating phase? Pacing Clin Electrophysiol 16:1401-1411, 1993. 59. Schuder JC, Rahmoeller GA, Stoeckle H: Transthoracic ventricular defibrillation with triangular and trapezoidal waveforms. Circ Res 19:689-694, 1966. 60. Denman RA, Umesan C, Martin PT, et al: Benefit of millisecond waveform durations for patients with high defibrillation thresholds. Heart Rhythm 3:536-541, 2006. 61. Keane D, Aweh N, Hynes B, et al: Achieving sufficient safety margins with fixed duration waveforms and the use of multiple time constants. Pacing Clin Electrophysiol 30:596-602, 2007. 62. Bardy GH, Marchlinski FE, Sharma AD, et al: Multicenter comparison of truncated biphasic shocks and standard damped sine wave monophasic shocks for transthoracic ventricular defibrillation. Transthoracic Investigators. Circulation 94:2507-2514, 1996. 63. Camacho MA, Lehr JL, Eisenberg SR: A three-dimensional finite element model of human transthoracic defibrillation: paddle placement and size. IEEE Trans Biomed Eng 42:572-578, 1995. 64. Lerman BB, Deale OC: Relation between transcardiac and transthoracic current during defibrillation in humans. Circ Res 67:1420-1426, 1990. 65. Karlon WJ, Eisenberg SR, Lehr JL: Effects of paddle placement and size on defibrillation current distribution: a threedimensional finite element model. IEEE Trans Biomed Eng 40:246-255, 1993. 66. Block M, Hammel D, Isburch F, et al: Results and realistic expectations with transvenous lead systems. Pacing Clin Electrophysiol 15:665-670, 1992. 67. Tang ASL, Wolf PD, Afework Y, et al: Three-dimensional potential gradient fields generated by intracardiac catheter and cutaneous patch electrodes. Circulation 85:1857-1864, 1992. 68. Tang ASL, Wolf PD, Claydown FJ III, et al: Measurement of defibrillation shock potential distributions and activation sequences of the heart in three-dimensions. Proc IEEE 76:11761186, 1988. 69. Chen PS, Wolf PD, Melnick SD, et al: Comparison of activation during ventricular fibrillation and following unsuccessful defibrillation shocks in open-chest dogs. Circ Res 66:1544-1560, 1990. 70. Chattipakorn N, Banville I, Gray RA, Ideker RE: Effects of shock strengths on ventricular defibrillation failure. Cardiovasc Res 61:39-44, 2004. 71. Shibata N, Chen PS, Dixon EG, et al: Epicardial activation after unsuccessful defibrillation shocks in dogs. Am J Physiol 255: 902-909, 1988. 72. Walcott GP, Chau WC, Ideker RE, et al: Early activation following shocks delivered in humans: the difference between transvenous and “active can” ICD systems. J Am Coll Cardiol 29:427A, 1997. 73. Crossley GH, Boyce K, Roelke M, et al: A prospective randomized trial of defibrillation thresholds from the right ventricular outflow tract and the right ventricular apex. Pacing Clin Electrophysiol 32:166-171, 2009. 74. Verma A, Kaplan AJ, Sarak B, et al: Incidence of very high defibrillation thresholds (DFT) and efficacy of subcutaneous (SQ) array insertion during implantable cardioverter defibrillator (ICD) implantation. J Interv Card Electrophysiol 29:127-133, 2010. 75. Kommuri NV, Kollepara SL, Saulitis E, et al: Azygos vein lead implantation for high defibrillation thresholds in implantable cardioverter defibrillator placement. Indian Pacing Electrophysiol J 10:49-54, 2010. 76. Walker RG, Walcott GP, Smith WM, Ideker IE: Sites of earliest activation following transvenous defibrillation. Circulation 90(pt 2):I-447, 1994. 77. Cates AW, Wolf PD, Hillsley RE, et al: The probability of defibrillation success and the incidence of postshock arrhythmia as a function of shock strength. Pacing Clin Electrophysiol 17:12081217, 1994. 78. Chapman FW, El-Abbaday TZ, Walcott GP, et al: Dysfunction following transthoracic defibrillation shocks in dogs. Pacing Clin Electrophysiol 20:1128, 1997. 79. Reddy RK, Gleva MJ, Gliner BE, et al: Biphasic transthoracic defibrillation causes fewer ECG ST-segment changes after shock. Ann Emerg Med 30:127-134, 1997. 80. DeBruin KA, Krassowska W: Electroporation and shockinduced transmembrane potential in a cardiac fiber during defibrillation strength shocks. Ann Biomed Eng 26:584-596, 1998. 81. Schuder JC, Gold JH, Stoeckle H, et al: Transthoracic ventricular defibrillation in the 100 kg calf with symmetrical one-cycle bidirectional rectangular wave stimuli. IEEE Trans Biomed Eng 30:415-422, 1983. 82. Yabe S, Smith DM, Duabert JP, et al: Conduction disturbances caused by high current density electric fields. Circ Res 66:11901203, 1990. 83. Jones JL, Jones RE: Decreased defibrillator-induced dysfunction with biphasic rectangular waveforms. Am J Physiol 247: H792H796, 1984. 84. Walcott GP, Knisley SB, Zhou X, et al: On the mechanism of ventricular defibrillation. Pacing Clin Electrophysiol 20:422-431, 1997. 85. Clark DM, Rogers JM, Ideker RE, Knisley SB: Intracardiac defibrillation-strength shocks produce large regions of hyperpolarization and depolarization. J Am Coll Cardiol 27:147A, 1996.

86. Zhou X, Wolf PD, Rollins DL, et al: Effects of monophasic and biphasic shocks on action potentials during ventricular fibrillation in dogs. Circ Res 73:325-334, 1993. 87. Daubert JP, Frazier DW, Wolf PD, et al: Response of relatively refractory canine myocardium to monophasic and biphasic shocks. Circulation 84:2522-2538, 1991. 88. Colavita PG, Wolf PD, Smith WM, et al: Determination of effects of internal countershock by direct cardiac recordings during normal rhythm. Am J Physiol 250: H736-H740, 1986. 89. Weidmann S: Electrical constants of trabecular muscle from mammalian heart. J Physiol 210:1041-1054, 1970. 90. Kléber AG, Riegger CB: Electrical constants of arterially perfused rabbit papillary muscle. J Physiol 385:307-324, 1987. 91. Krassowska W, Frazier DW, Pilkington TC, Ideker RE: Potential distribution in three-dimensional periodic myocardium. Part II. Application to extracellular stimulation. IEEE Trans Biomed Eng 37:267-284, 1990. 92. Krassowska W, Pilkington TC, Ideker RE: Potential distribution in three-dimensional periodic myocardium. Part I. Solution with two-scale asymptotic analysis. IEEE Trans Biomed Eng 37:252-266, 1990. 93. Plonsey R, Barr RC: Effect of microscopic and macroscopic discontinuities on the response of cardiac tissue to defibrillating (stimulating) currents. Med Biol Eng Comput 24:130-136, 1986. 94. White JB, Walcott GP, Pollard AE, Ideker RE: Myocardial discontinuities: a substrate for producing virtual electrodes that directly excite the myocardium by shocks. Circulation 97:17381745, 1998. 95. Tung L: A bidomain model for describing ischemic myocardial DC potentials, Cambridge, 1978, Massachusetts Institute of Technology. 96. Trayanova N: Discrete versus syncytial tissue behavior in a model of cardiac stimulation. II. Results of simulation. IEEE Trans Biomed Eng 43:1141-1150, 1996. 97. Trayanova N: Discrete versus syncytial tissue behavior in a model of cardiac stimulation. I. Mathematical formulation. IEEE Trans Biomed Eng 43:1129-1140, 1996. 98. Kieval RS, Spear JF, Moore EN: Gap junctional conductance in ventricular myocyte pairs isolated from postischemic rabbit myocardium. Circ Res 71:127-136, 1992. 99. Shaw RM, Rudy Y: Electrophysiologic effects of acute myocardial ischemia: a mechanistic investigation of action potential conduction and conduction failure. Circ Res 80:124-138, 1997. 100. Shaw RM, Rudy Y: Ionic mechanisms of propagation in cardiac tissue: roles of the sodium and L-type calcium currents during reduced excitability and decreased gap junction coupling. Circ Res 81:727-741, 1997. 101. Knisley SB, Blitchington TF, Hill BC, et al: Optical measurements of transmembrane potential changes during electric field stimulation of ventricular cells. Circ Res 72:255-270, 1993. 102. Windisch H, Ahammer H, Schaffer P, et al: Optical multisite monitoring of cell excitation phenomena in isolated cardiomyocytes. Pflugers Arch 430:508-518, 1995. 103. Wikswo JP, Lin SF, Abbas RA: Virtual electrodes in cardiac tissue: a common mechanism for anodal and cathodal stimulation. Biophys J 69:2195-2210, 1995. 104. Zhou, X., Ideker RE, Blitchington TF, et al: Optical transmembrane potential measurements during defibrillation-strength shocks in perfused rabbit hearts. Circ Res 77:593-602, 1995. 105. Sommer JR, Scherer B: Geometry of cell and bundle appositions in cardiac muscle: light microscopy. Am J Physiol 248:H792H803, 1985. 106. Street AM, Plonsey R: Activation fronts elicited remote to the pacing site due to the presence of scar tissue. In Proc 18th Annu Int Conf IEEE Eng Med Biol Soc, Amsterdam, 1996, Piscataway, NJ, Institute of Electrical and Electronics Engineers. 107. Sharifov OF, Ideker RE, Fast VG: High-resolution optical mapping of intramural virtual electrodes in porcine left ventricular wall. Cardiovasc Res 64:448-456, 2004. 108. Clerc L: Directional differences of impulse spread in trabecular muscle from mammalian heart. J Physiol 255:335-346, 1976. 109. Knisley SB: Transmembrane voltage changes during unipolar stimulation of rabbit ventricle. Circ Res 77:1229-1239, 1995. 110. Efimov IR, Cheng YN, Van Wagoner DR, et al: Virtual electrodeinduced phase singularity: a basic mechanism of defibrillation failure. Circ Res 82:918-925, 1998. 111. Efimov IR, Cheng YN, Bierman M, et al: Transmembrane voltage changes produced by real and virtual electrodes during monophasic defibrillation shock delivered by an implantable electrode. J Cardiovasc Electrophysiol 8:1031-1045, 1997. 112. Iyer AN, Gray RA: An experimentalist’s approach to accurate localization of phase singularities during reentry. Ann Biomed Eng 29:47-59, 2001. 113. Knisley SB, Smith WM, Ideker RE: Effect of field stimulation on cellular repolarization in rabbit myocardium. Implications for reentry induction. Circ Res 70:707-715, 1992. 114. Chen PS, Wolf PD, Dixon EG, et al: Mechanism of ventricular vulnerability to single premature stimuli in open-chest dogs. Circ Res 62:1191-1209, 1988. 115. Frazier DW, Wolf PD, Wharton JM, et al: Stimulus-induced critical point: mechanism for electrical initiation of reentry in normal canine myocardium. J Clin Invest 83:1039-1052, 1989. 116. Winfree AT: When time breaks down: the three-dimensional dynamics of electrochemical waves and cardiac arrhythmias, Princeton, NJ, 1987, Princeton University Press. 117. Efimov IR, Cheng Y, Yamanouchi Y, et al: Direct evidence of the role of virtual electrode-induced phase singularity in success and



2  Principles of Defibrillation: From Cellular Physiology to Fields and Waveforms

failure of defibrillation. J Cardiovasc Electrophysiol 11:861-868, 2000. 118. Chen PS, Shibata N, Dixon EG, et al: Activation during ventricular defibrillation in open-chest dogs: evidence of complete cessation and regeneration of ventricular fibrillation after unsuccessful shocks. J Clin Invest 77:810-823, 1986. 119. Chattipakorn N, Fotuhi PC, Ideker RE: Prediction of defibrillation outcome by epicardial activation patterns following shocks near the defibrillation threshold. J Cardiovasc Electrophysiol 11:1014-1021, 2000. 120. Chattipakorn N, Banville I, Gray RA, Ideker RE: Mechanism of ventricular defibrillation for near-defibrillation threshold shocks: a whole-heart optical mapping study in swine. Circulation 104:1313-1319, 2001. 121. Wang NC, Lee MH, Ohara T, et al: Optical mapping of ventricular defibrillation in isolated swine right ventricles: demonstration of a postshock isoelectric window after near-threshold defibrillation shocks. Circulation 104:227-233, 2001. 122. Chattipakorn N, Fotuhi PC, Chattipakorn SC, et al: Threedimensional mapping of earliest activation after near-threshold

ventricular defibrillation shocks. J Cardiovasc Electrophysiol 14:65-69, 2003. 123. Hwang GS, Hayashi H, Tang L, et al: Intracellular calcium and vulnerability to fibrillation and defibrillation in Langendorffperfused rabbit ventricles. Circulation 114:2595-2603, 2006. 124. Constantino J, Long Y, Ashihara T, Trayanova NA: Tunnel propagation following defibrillation with ICD shocks: hidden postshock activations in the left ventricular wall underlie isoelectric window. Heart Rhythm 7:953-961, 2010. 125. Hwang GS, Tang L, Joung B, et al: Superiority of biphasic over monophasic defibrillation shocks is attributable to less intracellular calcium transient heterogeneity. J Am Coll Cardiol 52:828835, 2008. 126. Chattipakorn N, Ideker RE: Delayed afterdepolarization inhibitor: a potential pharmacologic intervention to improve defibrillation efficacy. J Cardiovasc Electrophysiol 14:72-75, 2003. 127. Zheng X, Walcott GP, Smith WM, et al: Evidence that activation following failed defibrillation is not caused by triggered activity. J Cardiovasc Electrophysiol 16:1200-1205, 2005.

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128. Zheng X, Li L, Dosdall DJ, et al: Following long duration ventricular fibrillation, flunarizine greatly reduces the defibrillation threshold and the incidence of refibrillation. Circulation 120(Suppl 2):1481, 2009. 129. Dosdall DJ, Osorio J, Robichaux RP, et al: Purkinje activation precedes myocardial activation following defibrillation after long-duration ventricular fibrillation. Heart Rhythm 7:405-412, 2010. 130. Wiggers CJ: The mechanism and nature of ventricular fibrillation. Am Heart J 20:399-412, 1940. 131. Huang J, Rogers JM, Killingsworth CR, et al: Evolution of activation patterns during long-duration ventricular fibrillation in dogs. Am J Physiol Heart Circ Physiol 286:H1193-H1200, 2004. 132. Huizar JF, Warren MD, Shvedko AG, et al: Three distinct phases of VF during global ischemia in the isolated blood-perfused pig heart. Am J Physiol Heart Circ Physiol 293:H1617-H1628, 2007. 133. Robichaux RP, Dosdall DJ, Osorio J, et al: Periods of highly synchronous, non-reentrant endocardial activation cycles occur during long-duration ventricular fibrillation. J Cardiovasc Electrophysiol 21:1266-1273, 2010.

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3 

SECTION 1  Basic Principles of Device Therapy

Sensing and Detection CHARLES D. SWERDLOW  |  JEFFREY M. GILLBERG  |  PAUL KHAIRY

Sensing of cardiac depolarizations and detection of arrhythmias

control the electrical therapies of pacemakers and implantable cardioverter-defibrillators (ICDs). When a wavefront of depolarization passes the tip electrode of an intracardiac lead, a deflection in the continuous electrogram signal travels instantaneously up the lead wire to the pacemaker or ICD, where sensing electronics amplifies, filters, digitizes, and processes the signal. A sensed event occurs at the instant when the sensing system determines that an atrial or ventricular depolarization has occurred. Dual-chamber pacemakers and ICDs have separate sensing systems for the atrium and ventricle. Appropriate sensing results in one sensed event for each activation wavefront in the corresponding chamber. Failure to sense activation wavefronts results in undersensing, which can cause inappropriate pacing, failure to switch modes, or failure to detect a tachyarrhythmia. Undersensing occurs if the depolarization signal has insufficient amplitude or frequency content to be recognized as a sensed event, or if a blanking period disables the sensing amplifier at the time of the event. Oversensing occurs when nonphysiologic signals or physiologic signals that do not reflect local myocardial depolarization are sensed. Oversensing can cause inappropriate pacing inhibition, pacemaker tracking, or ICD therapy. Detection algorithms process sensed events to classify the atrial or ventricular rhythm. This classification is used to control beat-by-beat paced events, to change the pacing mode, to store data regarding tachyarrhythmias, and to terminate sustained tachyarrhythmias with antitachycardia pacing or shocks.

Electrograms SURFACE ELECTROCARDIOGRAM VS. INTRACARDIAC ELECTROGRAM An electrogram (EGM) is a graphic display of the potential difference between two points in space over time. During the upstroke of a myocardial action potential, the inside of the cell abruptly changes from its resting negative potential (with respect to the outside of the cell) to a neutral or slightly positive potential. After about 250 to 400 milliseconds (msec), the cell membrane is then repolarized, with the inside of the cell returning to its resting, negatively charged state. Figure 3-1 illustrates how an EGM is recorded between two electrodes in contact with the myocardium. The electrocardiogram (ECG) is recorded from two electrodes on the surface of the body at some distance from the heart. The typical amplitude of its QRS complex is about 1 millivolt (mV). The locations of the two electrodes determine the vectorial “viewpoint” from which the electrical activity of the entire heart is observed from the body surface. In contrast, the ventricular endocardial unipolar EGM typically is 5 to 20 mV in amplitude when recorded from a small electrode on the tip of a lead placed in direct contact with the apex of the right ventricle (Fig. 3-2). The second electrode needed to record this unipolar EGM is the pacemaker or ICD metal “can,” which is located some distance from the heart. The location of this distant second electrode, sometimes called the “indifferent electrode,” has a much smaller effect on the signal’s properties, although it may record noncardiac electric potentials (e.g., from pectoral muscle). The ECG records electrical activity from the entire heart, whereas the EGM records only local wavefronts of depolarization and repolarization. The EGM depends on

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the viability of approximately 1 or 2 cm3 of myocardium immediately under the tip electrode,1,2 as depicted in Figure 3-2. ELECTRODE SYSTEMS Figure 3-3 contrasts endocardial unipolar (tip-to-can), bipolar (tip-toring), and integrated bipolar (tip-to-coil) electrode systems, and Figure 3-4 shows representative examples. Epicardial electrode systems may be either unipolar (tip-to-can) or bipolar (tip-to-tip). These different electrode configurations have EGMs with similar R-wave amplitudes and slew rates, provided that the interelectrode spacing is at least 10 mm, as is true of almost all commercial pacemaker and defibrillator leads. Because they are more likely to oversense than bipolar EGMs, unipolar electrode systems are contraindicated for ICDs and are used infrequently for modern pacemakers. ICD integrated bipolar electrodes sense between the right ventricular tip electrode and right ventricular high-voltage coil, with sensing characteristics closer to the bipolar than the unipolar configuration. Compared with true bipolar electrodes, integrated ICD bipolar electrodes are more likely to oversense myopotentials and electromagnetic interference (EMI).3,4 In one study, oversensing occurred in 40% of patients with integrated bipolar sensing, compared with 8% of patients with true bipolar systems.4 AMPLITUDE, SLEW RATE, AND WAVESHAPE The largest and steepest deflection on the local EGM, called the intrinsic deflection, occurs when the wavefront of depolarization passes the small-tip electrode. The EGM amplitude traditionally is defined as the peak-to-peak amplitude of the intrinsic deflection (measured in mV), as shown in Figure 3-5. The duration of a ventricular EGM usually is less than that of the QRS of the surface ECG, because the EGM is a local signal. The amplitude of an atrial electrogram (AEGM) or ventricular electrogram (VEGM) is determined primarily by the excitable tissue near the tip electrode and therefore is usually similar for unipolar and bipolar signals. Typical amplitudes are 5 to 30 mV for VEGMs and 1.5 to 6 mV for AEGMs.1,2,5 The maximum slope of the intrinsic deflection is the slew rate, measured in volts per second, which represents the maximum rate of change of EGM voltage. Mathematically, it is the first derivative of the voltage, dV/dt, so it depends on both the amplitude and the duration of the EGM, and it provides a crude representation of the frequency content. The frequency content of ventricular and atrial EGMs is similar and in the range of 5 to 50 Hz. T waves and far-field R waves have lower frequency content, whereas most myopotentials and EMI have higher frequency content (Fig. 3-6). Typical values for slew rates are 2 to 3 V/sec for VEGMs and 1 to 2 V/sec for AEGMs.3,4 Usually, an EGM with acceptable amplitude also has an acceptable minimum slew rate (>1 V/sec for VEGMs, >0.3 V/sec for AEGMs). EGMs with very low amplitude will not be sensed, regardless of the slew rate. Increasing the size of the tip electrode in the range of 2 to 10 millimeters (mm) has minimal effect on atrial EGM amplitude but increases EGM duration (Fig. 3-7). For short ventricular bipolar interelectrode spacing of 5 mm or less, the R-wave amplitude decreases, because the difference between the two unipolar EGMs from each electrode causes cancellation in the net bipolar signal. The slew rate increases, because the time between arrival of the wavefront at the two electrodes decreases more than the EGM amplitude. When two electrodes are widely separated, as in early Y-adapted cardiac



3  Sensing and Detection

1

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resynchronization electrode systems, two distinct intrinsic deflections may be recorded on the EGM—one representing right ventricular (RV) activation and the other left ventricular (LV) activation. The interval between these deflections is determined by the conduction delay between the ventricles near the two electrodes. The waveshapes of EGMs are quite variable (Fig. 3-8), probably because geometry of the trabecular endocardium adjacent to the tip electrode is complex. In one study at pacemaker lead implantation, 58% of unipolar EGMs were biphasic, with an initial upstroke followed by a roughly equal downstroke; 30% were predominantly monophasic negative, and 12% were predominantly monophasic positive.1 The ventricular depolarization recorded on the atrial electrode is referred to as the far-field R wave (FFRW). Oversensing of the FFRW confounds interpretation of the atrial rhythm. The amplitude of the FFRW depends on the location of the atrial electrode. It is greatest near the septum, intermediate in the right atrial appendage, and least on

     

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E Figure 3-1  An electrogram (EGM) is recorded between two electrodes in contact with the myocardium. A, At rest, both electrodes record a similar charge, with no potential difference between them.  B and C, As a wavefront of depolarization moves under electrode 1, a difference in electrical charge is generated such that electrode  1 becomes electrically negative with respect to electrode 2. D, As the wavefront propagates under electrode 2, no potential difference between the two electrodes is recorded. E, The depolarization wavefront is followed by a wavefront of repolarization, during which a potential difference of opposite polarity is recorded. Because the EGM is determined by the instantaneous potential difference between the electrodes, the amplitude and shape of the recorded signal are determined by the direction from which the wavefront approaches the electrodes. For example, if a wavefront of depolarization reached both electrodes at the same time, there would be no potential difference between the electrodes, and an EGM would not be inscribed.

Time Figure 3-2  Concept drawing of spatial and temporal relationships for unipolar endocardial EGM. Upper panel, Anatomic drawing. Lower panel, EGM recorded from a small-surface-area electrode at the tip of a pacemaker or defibrillator lead that makes direct contact with the endocardium in the right ventricular (RV) apex. The second electrode required to record an EGM is not shown, because it is a distant and “indifferent” electrode, usually the metal can of the pulse generator, and its location is not important provided that it is a substantial distance from the tip electrode. During a ventricular depolarization, the depolarization wavefront propagates from the septum, around the RV apex, and up the RV free wall (arrows). When the wavefront of depolarization arrives at location 1, just as it approaches the electrode, the initial positive deflection of the EGM occurs, at time 1. When the wavefront passes closest to the tip electrode at location 2, the major negative deflection on the EGM occurs, labeled as time 2. As the wavefront recedes from the electrode at location 3, the final portion of the EGM is inscribed at time 3. This local EGM is not affected by the depolarization wavefront as it travels farther from the electrode. Therefore, the local EGM is shorter in duration than the surface electrocardiographic QRS complex.

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SECTION 1  Basic Principles of Device Therapy

pulse generator. Greater sensing safety margins are preferred for activefixation leads. The method of lead tip stabilization, active screw-in or passive tines, has had no significant effect on sensing characteristics in most studies.8,9 Steroid-eluting electrodes reduce chronic pacing thresholds substantially, but also have no significant effects on sensing.10-13 Metabolic, Ischemic, Aging, and Drug Effects RV-Coil

Ring Tip

Figure 3-3  The three practical endocardial electrode configurations used by most pacemakers and ICDs. The distant and indifferent or “can” electrode is not shown because it is out of the field of view. The unipolar configuration used in Figure 3-1 to explain EGM formation simply records the signal between the tip electrode and the can. The tip electrode can be an active-fixation screw or a small-surface-area tip electrode with various geometries. This unipolar configuration is subject to considerable noise and interference signals and is not suitable for ICDs. The bipolar configuration uses the Tip and Ring electrodes shown in this figure. The interelectrode spacing is typically 12 to 15 mm, and the ring electrode may or may not make contact with the endocardium. The integrated bipolar configuration uses the Tip and RV-Coil electrodes shown in the figure. EGMs recorded from bipolar and integrated bipolar configurations are very similar, and one less conductor is needed for the integrated bipolar configuration. The main disadvantages of the integrated bipolar configuration are susceptibility to diaphragmatic myopotentials, undesired atrial EGMs in small hearts, and slower postshock recovery times caused by electrode polarization. RV, Right ventricular.

the right atrial free wall. Even if the FFRW has comparable amplitude to the P wave, its slew rate usually is much lower. In one series, the mean slew rate was 1.2 V/sec for AEGMs and 0.13 V/sec for FFRWs.1 If an active-fixation, screw-in tip electrode is successfully attached to the myocardium, the acute VEGM has a current of injury, with an elevated ST segment (Fig. 3-9) that is usually greatly reduced within 10 minutes after electrode fixation. During this 10-minute period, the EGM amplitude and slew rate usually do not change despite changes in waveshape, but the pacing threshold decreases by an average of 40%.2 Acute to Chronic Changes and Fixation The amplitude and slew rate of intracardiac EGMs typically decline during the first several days to weeks after lead implantation and then increase to chronic values that are slightly lower than those measured at implantation.6 The initial decrease in EGM amplitude is caused by the inflammatory response and edema at the electrode-tissue interface. This gradually resolves and is followed by development of a small, inexcitable fibrotic zone surrounding the electrode tip (Fig. 3-10). This inflammation and fibrotic tissue effectively increases the distance between the surface of the electrode and excitable myocardium that generates the EGM signal. Although chronic EGM amplitudes usually are reduced by less than 10% compared with acute amplitudes, chronic slew rates are reduced by 30% to 40%.7 The acute reduction in EGM amplitude is often greater with activefixation leads than with passive fixation leads. Atrial undersensing can occur during the acute phase despite adequate EGM amplitudes at implantation. To account for these time-related changes in EGM amplitude, the filtered EGM recorded at lead implantation should be at least twice the sensitivity threshold that will be programmed in the

The effects of metabolic abnormalities and drugs on pacing thresholds are well described. Much less information is available concerning their effects on EGMs and sensing. Factors that reduce EGM amplitude, slow conduction velocity, or diminish slew rate may produce either oversensing or undersensing. By prolonging the intracardiac EGM duration beyond blanking periods, ischemia or antiarrhythmic drugs can produce double-counting of the QRS complex.14 Similarly, drugs that prolong the PR or QT interval beyond the refractory period may result in oversensing.15,16 Undersensing may result from reduction in EGM amplitude or slew rate after myocardial infarction at the electrode-tissue interface, from drug and electrolyte effects,15,16 or from progression of conduction system disease. Acute ischemia causes ST-segment changes that can be detected on VEGMs. Monitoring of EGM ST-segment shifts has been proposed as a method for monitoring ischemia for pacemakers and ICDs.17 The likelihood of recording abnormal AEGMs (defined as ≥100 msec in duration or having ≥8 fragmented deflections) correlates with age of the patient (r = 0.34; P < .0005).18 Exercise, Respiratory, and Postural Effects The effect of exercise on the AEGM amplitude and slew rate is variable. Some studies have reported statistically significant decreases in amplitude that average 10% to 20% but may reach 40% in some patients.19,20 Other studies did not find significant changes between rest and exercise.21,22 Decreases in AEGM amplitude were not caused by atrial rate alone or by beta blockade.23 VDD/R lead studies with “floating” atrial electrodes showed particularly large decreases with exercise.24,25 Decreases in AEGM amplitude with lead maturation support the programming of a large safety margin for sensing at implantation to offset effects of lead maturation. P-wave amplitude increases significantly during full inspiration, during full expiration, and with erect posture.22 Respiratory variation averaged 9.7% for unipolar AEGMs and 11.5% for bipolar AEGMs.25,26 The effect of respiration on VEGMs was less, especially with the unipolar configuration.26 Ventricular Electrograms during Premature Ventricular Complexes, Ventricular Tachycardia, and Ventricular Fibrillation Premature ventricular complexes (PVCs) may have lower-amplitude R waves than sinus-rhythm R waves, as shown in Figure 3-11, but the reverse may also be true. For monomorphic ventricular tachycardia (VT), mean amplitude decreased only slightly from values in sinus rhythm—14% for epicardial EGMs and 5% for endocardial EGMs.27 In contrast, EGM amplitudes during ventricular fibrillation (VF) decreased by 25% for epicardial and 41% for endocardial EGMs. More importantly, EGMs in VF often have low, highly variable, and rapidly changing amplitudes and slew rates. Figure 3-12 shows endocardial spontaneous VF EGMs from different patients, illustrating variability in intrinsic deflections, amplitudes, slew rates, and morphologies. In a study of induced VF reproducibility, 50% of the variability was caused by interpatient differences and the other 50% occurred among repeated episodes in the same patient.28 In another study, the VEGM amplitude in VF was 1 mV or less in at least one VF episode in 29% of patients.27 If VF lasts for minutes, the amplitude and slew rate of the EGMs decrease. Atrial Electrograms during Rhythms Other Than Sinus Atrial activation from ectopic sites or atrial arrhythmias can alter the amplitude, frequency content, slew rate, and morphology of the



3  Sensing and Detection

59

Figure 3-4  Ventricular electrocardiograms (ECGs) recorded from different electrode configurations in the same patient. The central panel shows a left anterior oblique radiograph of a cardiac resynchronization ICD system. Each of the four tracings shows surface ECG lead II, EGM markers, and one ventricular EGM during atrial pacing at a rate of 75 bpm. Top left, Far-field EGM recorded between the right ventricular (RV) coil electrode and the electrically active ICD housing (CAN). Lower left, Integrated bipolar EGM recorded between RV tip and RV coil electrodes. Lower right, True bipolar EGM recorded between the RV tip and ring electrodes. Top right, Left ventricular (LV) unipolar EGM recorded between the LV tip electrode and can. EGM scale is 0.5 mV/mm, except for the LV unipolar EGM, which has a scale of 2 mV/mm. The downward EGM ventricular sense (VS) markers correspond to the time at which the true bipolar RV tip-ring EGM crosses the sensing threshold. Because the “field of view” of this EGM is local, its duration is short. It occurs early in the QRS complex of this patient with left bundle branch block. The integrated bipolar tipcoil EGM has a peak-to-peak amplitude and slew rate similar to those of the true bipolar EGM. However, its field of view is larger due to the size of the RV coil, and therefore the T wave is larger. Low-amplitude atrial EGMs are visible because of the proximity of the coil to the tricuspid anulus. Both the RV coil-CAN and the LV unipolar EGM are widely spaced, between an intracardiac electrode and the extracardiac can. Their duration is closer to that of the QRS complex. The intrinsic deflection of the LV unipolar electrode is late in the QRS complex, corresponding to late activation of the lateral left ventricle. The greater amplitude of the LV unipolar EGM reflects the greater muscle mass of the left ventricle. EGMs recorded from the superior vena cava (SVC) coil and from the atrial bipole (right atrium, RA) are not shown. Radiograph and EGMs are from different patients. Radiograph is for illustrative purposes only.

AEGM. Retrograde atrial activation during ventricular pacing reduces AEGM amplitude and slew rate by up to 50%.29 These EGM changes are more pronounced in the high right atrium than in the right atrial appendage or low right atrium.30 The frequency content of the AEGM is not significantly altered by retrograde atrial activation.31 Analysis of EGM turning-point morphology or the first-differential coefficient of slew rate has been used to discriminate sinus EGMs from those recorded during retrograde and ectopic atrial activation in small groups of patients.32 Atrial EGMs during atrial fibrillation (AF) are characterized by extreme temporal and spatial variability. EGMs tend to be most organized in the trabeculated right atrial appendage and more disorganized in the smooth right atrium or coronary sinus.33-35 Thus electrode spacing and positioning of atrial leads influence EGM characteristics during AF36-38 and may cause inconsistent diagnosis of AF based on rate criteria. The amplitude of chronic, unipolar pacemaker EGMs was 40% less in AF than in sinus rhythm.39 A comparison of acute AEGM amplitudes recorded with temporary pacing catheters showed that the mean sinus-rhythm EGM amplitude decreased only slightly in atrial flutter but decreased by about 50% in AF.35 Antiarrhythmic drugs may also

interfere with sensing during AF by reducing atrial rate, median frequency, and EGM amplitude.40 SUBCUTANEOUS ELECTROCARDIOGRAPHY The subcutaneous ECG is similar to the surface ECG because the two subcutaneous electrodes are sufficiently distant from the heart that they record electrical activity from the entire heart. As with the surface ECG, the amplitude of subcutaneous ECG signals usually is 1 mV or less. Simultaneous recordings of subcutaneous ECG signals and surface ECG signals from electrodes placed directly over the subcutaneous locations have similar amplitude and signal-to-noise ratio.41 Practical implantation considerations usually limit the subcutaneous electrode separation distance to 4 to 8 cm, compared with the typical surface ECG limb lead electrode separation of 40 to 60 cm. The orientation of the two subcutaneous electrodes relative to the heart can affect the amplitude of the signal recorded.42 Mapping studies on the chest skin with 4-cm electrode spacing in the range used by implantable loop recorders (ILRs) show larger intrinsic QRS amplitudes of 0.5 ± 0.1 mV for vertical orientation in the left parasternal zone and for horizontal orientation near the apex of the heart.

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SECTION 1  Basic Principles of Device Therapy

Time (∆T)

chamber is turned off or “blanked” for a short blanking period (20-250 msec) after each spontaneous depolarization or pacing stimulus, to prevent a single depolarization resulting in multiple sensed events. In the refractory period that follows the blanking period, the sense amplifier remains enabled. Sensed events occurring in refractory periods do not alter pacemaker timing cycles but may be sensed for tachyarrhythmia detection algorithms. BLANKING AND REFRACTORY PERIODS

Amplitude Slew rate =

Voltage (mV) (∆V) in volts/sec (∆T)

SENSING

ATRIUM

VENTRICLE

Electrogram

>1.5-2.0 mV

>5-6 mV

Slew rate (V/sec)

>0.3

>1.0

Figure 3-5  Major clinical descriptors of electrogram. The peak-topeak amplitude of the EGM is the difference in voltage recorded between two electrodes and is measured in millivolts (mV). The slew rate is equal to the first derivative of the EGM (dV/dt) and is a measure of the sharpness of the EGM and therefore its frequency content. Slew rate is measured in volts per second (V/sec). Usually, the amplitude of the EGM should be greater than 1.5 to 2.0 mV in the atrium and at least 5 to 6 mV in the ventricle at the time of lead implantation, to ensure adequate sensing. The slew rate should be at least 0.3 V/sec in the atrium and at least 1 V/sec in the ventricle.

Subcutaneous ECGs are used to detect arrhythmias in ILRs, to obviate the need for surface ECG electrodes during follow-up of pacemakers and ICDs, and to detect VT/VF in an ICD without intravascular electrodes.

Sensing The methods and technology of sensing and detection in ICDs and pacemakers share many features, but there are two major differences. First, ICDs need reliable sensing and detection during VF, but pacemakers do not. Second, pacemakers may use unipolar or bipolar sensing, whereas ICDs always use bipolar sensing. GENERAL CONCEPTS Figure 3-13 shows the primary functional operations of sensing systems used by pacemakers and ICDs. The raw signal passes from the leads to the connector, through hermetic feedthroughs with highfrequency filters and high-voltage protection circuitry, before reaching the sensing amplifier. After the signal is amplified, a band-pass filter processes it to reduce T waves, myopotentials, and EMI (filtering). Then, it is rectified to nullify effects of signal polarity (rectification). Finally, it is compared with the sensing-threshold voltage. At the instant the processed signal exceeds the sensing-threshold voltage, a sensed event is declared to the timing circuits and indicated by a marker pulse on the programmer marker channel. The sense amplifier in the same

Signal amplitude (mV)

Voltage (∆V)

Blanking periods and refractory periods are used to prevent undesirable behavior caused by oversensing or double-counting of cardiac activity (Figs. 3-14 and 3-15). The specifications of blanking/refractory periods have substantial impact on ICD sensing and pacing functions (Fig. 3-16). Same-chamber blanking/refractory periods after sensed events reduce double-counting of intrinsic cardiac depolarizations that may result in escape pacing at a rate slower than the programmed lower rate in pacemakers or inappropriate detection of VF in ICDs. After paced events, the same-chamber blanking/refractory periods are typically longer and prevent oversensing of the pacing artifact and evoked response. The blanking/refractory periods in the ventricle after atrial sensed or paced events and in the atrium after ventricular sensed or paced events are called cross-chamber blanking/refractory periods. Cross-chamber blanking periods help to prevent oversensing of the pacing artifact after a paced event in the opposite chamber. The atrial blanking period after ventricular events, postventricular atrial blanking (PVAB), is designed to avoid oversensing of ventricular pacing stimuli and FFRWs. Longer postventricular atrial refractory periods (PVARPs) prevent retrogradely conducted atrial activation from resetting atrial timing cycles for dual-chamber pacing. Crosschamber blanking in the atrium after a ventricular event must be minimized in ICDs with tachyarrhythmia detection (ICDs or atrial therapy ICDs) to avoid undersensing the atrial rhythm, particularly during high ventricular rates. Long PVAB periods prevent reliable sensing of AF and atrial flutter/atrial tachycardia (AT). However, short PVAB periods may result in atrial sensing of FFRWs. NOT SENSED

SENSED

70 50 40 30 20 10

Amplifier threshold R Waves

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4

EMI

T Waves Far-field R wave

3 2

Muscle potentials

1 0.7 1

2

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20

50

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Frequency (Hz) Figure 3-6  Signal amplitude versus frequency. This plot shows the approximate characteristics of the P and R waves that pacemakers and ICDs are intended to sense and the approximate characteristics of  the electromagnetic interference (EMI, muscle potentials), T waves, and far-field R waves that they are intended not to sense. The sense amplifier’s filters are designed to sense signals that are above the U-shaped amplifier threshold curve and to reject signals that are below the curve. P waves and R waves have similar frequency characteristics, but usually R waves have higher dominant frequency than P waves. Muscle potentials usually have higher-frequency components than intracardiac signals. T waves and far-field R waves have lower frequencies. As shown, there are some overlaps in these amplitude-frequency characteristics that cause oversensing or undersensing in particular situations. The ellipses representing the amplitude-frequency characteristics in  this figure are conceptual and are not based on quantitative measurements.

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3  Sensing and Detection

Unipolar (mV)



200 ms

1.7 mV

2 mm Bipolar contact

32 msec 1.6 mV

10 mm Bipolar contact

Bipolar (mV)

1 mV

5 0 5

1

5

1.3 mV

Orthogonal contact

20 msec 0.8 mV

Orthogonal floating

20 msec Figure 3-7  Effects of electrode configurations on atrial endocardial electrogram. The EGMs were obtained from a single patient with two catheters placed simultaneously in the right atrial appendage. One catheter had 2-mm ring electrodes (top three tracings), and the other catheter had 1-mm orthogonal electrodes. The surface electrocardiogram tracing is shown at the top of the figure. Time and voltage amplitude scales are shown. For each electrode configuration (right), the corresponding EGM is shown (left), with the peak-to-peak amplitude and EGM duration labeled. “Contact” refers to electrodes in contact with the atrial tissue. “Floating” refers to noncontact electrodes in the atrial chamber. Note that greater ring electrode spacing, from 2 to 10 mm, prolongs EGM duration without altering the amplitude. The unipolar EGM shows a wider and diminished atrial EGM and a prominent far-field ventricular EGM as well. The orthogonal electrode configurations provide EGMs of lesser amplitude and shorter duration, compared with the ring electrodes.

Implantable cardioverter-defibrillators that require tachyarrhythmia detection typically have shorter blanking and refractory periods than standard pacemakers, so that short cardiac cycles can be sensed reliably. As shown in the bottom marker diagrams of Figure 3-16, blanking periods may be adaptively extended based on noise-sampling windows (30-60 msec) if suprathreshold activity (due to cardiac or extracardiac sources such as EMI) is identified on the EGM immediately after a sensed event. If noise is seen in consecutive windows after a sensed event, the blanking period is “retriggered” for that beat to avoid double-counting or continuous oversensing. This operation may result in paradoxical undersensing of the cardiac rhythm when more sensitive sensing levels are programmed if noise is oversensed.43,44 The duration of the total atrial refractory period (TARP), equal to the atrioventricular (AV) delay plus the PVARP, in DDD pacing modes limits atrial tracking of the atrium at high sinus rates without affecting atrial sensing, as shown in Figure 3-17. Because the AV delay of most dual-chamber pacemakers shortens in response to increasing atrial rates or sensor input, the TARP also shortens. Several manufacturers now offer dual-chamber pacemakers that shorten the PVARP with increasing atrial or sensor-indicated rates, further reducing the TARP during exercise. The result of these newer algorithms is that the programmed upper tracking rate can be safely increased while providing protection at lower heart rates from initiation of pacemaker-mediated tachycardia caused by retrograde conduction.

Unipolar (mV)

55 msec

Bipolar (mV)

1.0 mV

3

4

7

8

9

5

0 5

0

50 msec Unipolar contact

2

100 (ms)

5 0 5

6

5

10

0 5

0

100 (ms)

Figure 3-8  Similarities of unipolar and bipolar electrograms. Examples recorded from a lead placed in the right atrial appendage in 10 patients. Note that the amplitudes of unipolar and bipolar EGMs are similar for each patient. The waveshapes of unipolar and bipolar EGMs for a given patient may be quite similar (patient 3) or quite different (patient 8), although these differences can be attributed to the relative size of the major inflections. Some of these differences may depend on whether the ring electrode for the bipolar recording makes contact with the myocardium. On the whole, intrapatient differences between unipolar and bipolar recordings appear to be less than interpatient differences.

120 mV 110 mV

Acute

18.4 mV

10 minutes

14.4 mV

1 hour

13.1 mV

210 mV 220 mV 120 mV 110 mV

210 mV 220 mV 120 mV 110 mV

10 mV 20 mV Figure 3-9  Acute current of injury at implantation. Top panel, Highresolution recording shows marked ST-segment elevation, indicating the current of injury when an active fixation screw-in tip electrode is extended into the endocardium. Middle panel, After only 10 minutes, most of the ST-segment elevation in the signal has disappeared. Bottom panel, The EGM is not appreciably different at 1 hour after implantation.

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SECTION 1  Basic Principles of Device Therapy

SENSING THRESHOLDS IN PACEMAKERS

Approximate outline of lead body and helix Figure 3-10  Gross microscopic slide specimen shows myocardium remaining after removal of endocardial active-fixation screw lead. The large dotted line shows where the lead body was located, and the colored staining shows a thin, fibrotic sheath around the lead body. The approximate location of the helical screw-tip electrode is shown by  the solid coiled line. The oval shape (dotted line) shows the size of the fibrotic capsule that formed around the helical extended-tip electrode. Most of the tissue outside the dotted lines stained red, indicating that it was active myocardium capable of conducting depolarizations. The tip region of this electrode is similar to that of the tip electrode in Figure 3-1, so propagation of depolarization wavefronts must travel around the tip electrode, in tissue largely out of the field of view on the right side of this figure.

Sensing thresholds in most pacemakers are programmable to a constant value. Ventricular sensing channels in conventional pacemakers typically operate at sensing thresholds of 2.5 to 3.5 mV, about 10 times less sensitive than those in ICDs. Therefore, pacemakers may undersense VF. Atrial sensitivity thresholds are typically 0.3 to 0.6 mV, to allow sensing of small-amplitude atrial EGMs during AF and to improve the accuracy of AF diagnostics. Unipolar sensing thresholds typically are set higher (less sensitive) than bipolar sensing thresholds to reduce oversensing of far-field cardiac and extracardiac signals that can lead to inappropriate pacemaker inhibition or tracking. Newer pacemakers automatically adjust the sensitivity setting to adapt to changes in EGM amplitude over time. Typically, these functions operate to modify sensing thresholds based on a series of 10 to 20 ventricular beats. One such algorithm employs two simultaneous sensing levels: the programmed sensitivity (inner target) and a value twice the programmed value (outer target) (Fig. 3-18).45 Sensed EGMs exceeding both target values decrease the sensitivity. Signals exceeding only the inner target increase the sensitivity. In this manner, a 2 : 1 sensing margin is maintained. Rapid, automated sensitivity adjustments may be desired when EGM amplitudes can be expected to change over a brief period, such as beat-to-beat variations from respiration, body position changes, or fluctuating EGM morphologies during AF.46 Far-field R-wave oversensing can be minimized by (1) selecting an atrial lead with a closely spaced bipolar electrode pair (≤10 mm), (2) choosing an implantation location that yields an FFRW/P-wave ratio of less than 0.5,47 (3) titrating programmed sensitivity to reject FFRWs without undersensing P waves and low-amplitude AF, and (4) using PVAB.

CHART SPEED 25.0 mm/s ECG

0.2 mV/mm LEAD II

V EGM

1.0 mV/mm

Slew rate=dV/dt ∆V

P-P Amplitude

∆t

Marker channel

V S

V S

V S

V S

V S

V S

V S

V S

Figure 3-11  Surface electrocardiogram (ECG) lead II, bipolar right ventricular electrogram, and event markers with downward pulses that show when sensing occurred. The QRS amplitude on the ECG is about 1 mV, which is typical. The peak-to-peak amplitude of the sinus R waves is about 10 to 12 mV, which is also typical. The slew rate is the maximum slope (dV/dt) of the EGM intrinsic deflection; it is difficult to measure with a paper speed this slow. The two premature ventricular complexes (PVCs) (fourth and sixth) have different amplitudes and shapes on both the ECG and the EGM. The sinus beat in the center, between the two PVCs, has its main intrinsic deflection during the last part of the ECG QRS complex. The left edge of the sense marker indicates the instant that sensing by the ICD occurred. Therefore, EGM morphology and timing of the sense marker may not correspond to the start of the QRS on the ECG as electrocardiographers expect. Each R wave was sensed only once because sensing is blanked by the ICD for 120 msec after each ventricular sense (VS).



3  Sensing and Detection

63

1 mV

1 mV

1 mV

1 mV

1 mV

1 mV

1 mV

1 mV

1 sec

1 mV Figure 3-12  Variability in electrograms during spontaneous ventricular fibrillation (VF). Bipolar EGMs recorded during VF detection and charging by Medtronic Gem ICDs are shown for nine patients. The 1-mV calibration markers are the same except for the bottom tracing, where the calibration pulse is about three times larger, meaning that the ventricular EGM for this patient was three times smaller than for the other eight patients. Also note that rapid intrinsic deflections are visible on all tracings, although some substantial beat-to-beat variations in amplitude occurred. The abbreviated sense markers are shown at the bottom edge of each stored EGM strip. The sense amplifier blanking period was 120 msec, and the programmed sensitivity in each case was 0.3 mV, except in the sixth tracing from the top, which had 0.6 mV sensitivity. Note that there was slight VF undersensing on the bottom tracing when large EGMs were followed by smaller EGMs.

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SECTION 1  Basic Principles of Device Therapy

Time of sensed event Electrograms

Filter

Output

Voltage

100

0

Input

1

Rectifier

0

Frequency

To timing circuits

Threshold Voltage

Amplifier

Output

Electrodes and lead

Blanking period

Threshold

0

Input

Variable for Auto-gain control

Time

Variable for Auto-gain threshold

Figure 3-13  Functional block diagram for pacemaker or implantable cardioverter-defibrillator (ICD) sense amplifier. The EGM signal from the two implanted electrodes is first amplified for subsequent processing. Band-pass filtering reduces the amplitude of lower-frequency signals such as T waves and far-field R waves and higher-frequency signals such as myopotentials and electromagnetic interference. After band-pass filtering, the signal is rectified to make polarity unimportant. The thresholding operation compares the amplified, filtered, and rectified signals with the sensing threshold voltage. At the instant the processed signal exceeds the sensing threshold voltage, the sense amplifier is blanked (turned off) for 20 to 120 msec, so each depolarization is sensed only once, and a sensed event is declared to pacemaker or ICD timing circuits. For pacemakers, the programmed sensitivity controls the constant sensing threshold voltage. For ICDs, the amplifier gain may be controlled by the input EGM amplitude. The programmable sensing threshold for ICDs controls the high and low limits on the sensing threshold, which automatically adjusts on a beat-by-beat basis (see text discussion). In actual circuits, some functions (e.g., amplification, filtering) may be integrated.

VENTRICULAR SENSING IN VENTRICULAR ICDS The guiding design principle is that sensing of VF and polymorphic VT should be sufficiently reliable that clinically significant delays in detection do not occur. Although high sensitivity is required to ensure reliable sensing during VF, continuous high sensitivity results in oversensing of cardiac or extracardiac signals during regular rhythm, which may cause inappropriate detection of VT or VF. To minimize both undersensing during VF and oversensing during regular rhythms, ICDs use feedback mechanisms based on R-wave amplitude that adjust the sensing threshold dynamically. To maximize the likelihood of detecting VF, blanking periods are kept short. Automatic Adjustment of Sensitivity Adjustment of Sensitivity in Normal Rhythm.  All ICDs automatically adjust sensitivity in relation to the amplitude of each sensed R wave (Fig. 3-19). At the end of the blanking period after each sensed ventricular event, the sensing threshold is set to a high value. It then decreases with time until a minimum value is reached. Compared with a fixed sensing threshold, automatic adjustment of sensitivity increases the likelihood of sensing low-amplitude and varying EGMs, while minimizing the likelihood of T-wave oversensing. The methods of the different manufacturers for automatic adjustment of sensitivity perform similarly after small R waves but differently

after large R waves. Figure 3-20 shows that after large R waves, the Boston Scientific ICDs increase the sensing floor. In the Cognis/Teligen family, the sensing floor is set to one-eighth the amplitude of the measured R wave if that value is greater than the programmed sensitivity. This prevents T-wave oversensing in the setting of large R waves and reduces oversensing of low-amplitude noncardiac signals (e.g., diaphragmatic myopotentials, EMI). However, it may increase the risk of undersensing during rare episodes of VF with highly variable EGM amplitude.48 Postpacing Automatic Adjustment of Sensitivity.  After ventricular pacing, all ICDs set ventricular sensitivity to a highly sensitive value to prevent pacing during VF. The sensitivity threshold then decays to the programmed sensitivity level (Fig. 3-21). Thus ICDs are especially vulnerable to oversensing of low-amplitude signals late in diastole during pacing, when the amplifier sensitivity or gain is maximal. Clinically, the most important manifestation is the oversensing of diaphragmatic myopotentials.3

V P Figure 3-14  Loss of atrial sensing with apparent ventricular undersensing and ventricular pacing. In the second complex, an atrial pacing stimulus follows the P wave because of undersensing in the atrium. The atrial pacing stimulus occurs at the start of the ventricular QRS complex. The ventricular EGM is not sensed because of blanking in the ventricular sensing amplifier immediately after the atrial pacing stimulus. This sequence is repeated in the fourth, sixth, and eighth complexes. Atrial undersensing may lead to apparent ventricular undersensing because the ventricular blanking period does not permit sensing of electrical signals for 10 to 40 msec after an atrial output pulse.

V P

V P

V S

V P

V S

V P

Figure 3-15  Oversensing of T waves causing interruptions in VVI pacemaker timing. The surface electrocardiogram (top tracing) and pacemaker marker channel (bottom tracing) show ventricular oversensing of T waves. The first three complexes show ventricular pacing (VP) with capture. The pauses after the third and fourth complexes result from T-wave oversensing (VS), demonstrated by the marker channel.



3  Sensing and Detection

65

ECG AEGM VEGM

ICD

Marker channel AS 0 100

AS 30 100

120 VS AS

30 AP 200

200 VP

30

Adaptive (retriggerable) blanking periods

AS

AS 30 100

0 AP 200

200 VP

30 120 VS

200 VP

AP

AS

AP

Pacemaker VS

VP

VP

ICD same chamber: 200 ms postpace 120 ms post V sense 100 ms post A sense Cross-chamber: 30 ms postpace blank 0 ms postsense blank 310 ms postventricular refractory

VP

VS

Pacemaker same chamber: 63 ms postpace or postsense (retriggerable) 230 ms ventricular refractory period Cross-chamber: 180 ms postventricular atrial blanking 250-400 ms postventricular atrial refractory (auto) 20-45 ms postatrial pace ventricular blanking

Figure 3-16  Basic blanking and refractory periods for DDDR mode (Medtronic Marquis DR ICD and EnPulse DR pacemakers). The top tracings show the surface electrocardiogram (ECG), atrial electrogram (AEGM), and ventricular electrogram (VEGM) signals. The bottom two marker diagrams illustrate atrial or ventricular pacing (AP, VP), atrial or ventricular sensing (AS, VS), blanking periods (purple), and refractory periods (blue). Blanking periods in ICDs are usually of fixed duration, whereas same-chamber blanking in pacemakers is generally adaptive, with short (30-50 msec) blanking periods that “retrigger” when suprathreshold signals are present during the blanking period. Adaptive blanking periods can extend indefinitely, resulting in activation of noise-reversion asynchronous pacing. Note the considerably shorter blanking periods on the atrial channel in the ICD; this allows more accurate sensing of atrial depolarizations during high ventricular rates, which is critical for tachyarrhythmia detection and discrimination algorithms. Also note the lack of ventricular refractory periods in ICDs; this allows inhibition of pacing when high rates are sensed. Ventricular blanking periods in pacemakers may be shorter than in ICDs because of the approximately 10-fold less sensitive sensing threshold in pacemakers compared with ICDs.

Automatic Gain Control.  Early Ventritex ICDs (St. Jude Medical) used automatic step adjustments of gain as a primary means for avoiding T-wave oversensing and ensuring detection of low-amplitude VF EGMs. This resulted in sensing errors when EGM amplitude changed abruptly.49-51 In Boston Scientific ICDs before the Cognis/Teligen models, “slow” Automatic Gain Control adjusts the dynamic gain of the sensing amplifier slowly in response to temporal changes in R-wave amplitude. This ensures that the peak of the sensed R wave reaches about 75% of the amplifier gain, and that the sensing EGM is not truncated or “clipped.”

The minimum value of amplifier range is one eighth of the maximum value, so the minimum sensitivity decreases as R-wave amplitude increases. In Cognis/Teligen models, the range sense amplifier is sufficient (0 to 32 mV) that “slow” Automatic Gain Control is not used. In these Boston Scientific ICDs, “Automatic Gain Control” refers to automatic adjustment of sensitivity. In older Boston Scientific ICDs, “fast” Automatic Gain Control refers to automatic adjustment of sensitivity. Thus, whatever the name applied, automatic control of sensitivity, rather than gain, has become the primary method of beat-to-beat sensing adjustment in all ICDs. 2 seconds

AS

VP

AS

AS

VP

AS

AS AS

VP

AS

AS AS

VP

AS

AS

VP

AS

AS

VP

AS

AS AS

VP

AS

AS

VP

AS

AS

VP

AS

AS

VP

AS

AS AS

VP

AS

AS AS

VP

AS

AS

VP

AS

AS

VP

AS

AS AS

VP

AS

AS AS

VP

AS

AS

VP

Figure 3-17  Intracardiac marker channels. In a patient with a DDD pacemaker, channels identify atrial events within postventricular atrial refractory periods (PVARPs) that are sensed but not tracked. The surface electrocardiogram (top tracing) demonstrates ventricular pacing at the upper rate limit during atrial flutter. The atrial marker channel demonstrates atrial events that were sensed within PVARP (AS) and atrial events sensed outside PVARP (AS). After a ventricular pacing stimulus (VP), some atrial EGMs are not sensed, resulting in apparent EGM dropout because the atrium is blanked for a period after delivery of a VP. Only atrial signals recorded outside PVARP (AS) are tracked.

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SECTION 1  Basic Principles of Device Therapy

mV

mV

10 8 6 4 2 0 2 4 6 8 10 Increase Increase Increase Increase Increase Decrease Increase Decrease sensitivity sensitivity sensitivity sensitivity sensitivity sensitivity sensitivity sensitivity 2:1 safety Barely Converging on 2:1 safety Optimal sensing margin sensing margin for sensing (equilibrium around a achieved 2:1 safety margin) Outer target (1/2 sensitivity level) Inner target (programmed sensitivity level)

Figure 3-18  Autosensing algorithm to maintain a 2 : 1 sensing safety margin. See text for details. (From Castro A, Liebold A, Vincente J, et al: Evaluation of autosensing as an automatic means of maintaining a 2:1 sensing safety margin in an implanted pacemaker. Autosensing Investigation Team. Pacing Clin Electrophysiol 19:1708-1713, 1996.)

Ventricular Blanking Periods Ventricular blanking periods prevent ventricular oversensing of samechamber signals (R-wave double-counting) and cross-chamber signals (atrial pacing pulses and P waves) in regular rhythms. Because reliable sensing in VF requires minimizing blanking periods, blanking periods in ICDs are short and may occasionally be insufficient to prevent oversensing. Short Same-Chamber Blanking Periods and R-wave DoubleCounting.  In adults who are not taking antiarrhythmic drugs, interEGM intervals for filtered EGMs in VF vary from about 130 to 300 msec, with a peak near 200 msec (Fig. 3-22). Therefore, fixed or nominal blanking ICD periods after ventricular sensed events range from 120 to 135 msec. R-wave double-counting occurs if the duration of the sensing EGM exceeds the ventricular blanking period. Cross-Chamber Blanking Periods and Undersensing of VT/VF.  Under most conditions, ICDs apply only the minimum cross-chamber ventricular blanking required to prevent “crosstalk” resulting from an

Unfiltered R ventricular EGM

T

atrial pacing stimulus. During high-rate atrial or dual-chamber pacing, ventricular sensing may be restricted to short periods of the cardiac cycle because of the combined effects of ventricular blanking after ventricular events and cross-chamber ventricular blanking after atrial pacing. If a sufficient fraction of the cardiac cycle is blanked, systematic undersensing of VT or VF may occur. When pacing and blanking events occur at intervals that are multiples of a VT cycle length, ventricular complexes may be repeatedly undersensed, delaying or preventing detection.52-54 This occurs most often with rate-smoothing algorithms. “Sensing” Other Ventricular Electrograms Implantable cardioverter-defibrillator may use information derived from other VEGMs (shock-channel EGM, LV EGM in resynchronization systems) to enhance their functionality. The EGM from the shock channel (far-field EGM) is not used by transvenous ICDs for rate counting because bipolar or integrated bipolar EGMs are less susceptible to extracardiac signals. However, some manufacturers (Boston Scientific, Medtronic) use morphologic characteristics of far-field EGMs for discrimination of supraventricular tachycardia (SVT) from VT. Ventricular fibrillation

1 sec

Automatically adjusting sensing threshold

Filtered and rectified EGM 0.3 mV

Programmed sensitivity

Figure 3-19  Automatic adjustment of sensitivity. ICDs adjust sensitivity in relation to the amplitude of each sensed R wave. The goal of this feature is to permit sensing of low-amplitude and varying-amplitude EGMs while minimizing T-wave oversensing. The figure shows two sinus beats followed by the onset of ventricular fibrillation (VF). The upper panel shows the unfiltered ventricular EGM. The lower panel shows the corresponding filtered and rectified EGM. After each sensed ventricular event, the sensing threshold is set to a predetermined fraction of the R-wave amplitude. For large R waves, the initial sensing threshold may have a maximum value. The threshold then decreases with time until it reaches a minimum value equal to the programmed sensitivity, which is nominally about 0.3 mV. For sinus beats, the threshold is larger than the T waves, preventing oversensing. When VF begins, the smaller R waves keep the threshold at a lower value, which allows sensing of R waves that are even smaller than the T waves in sinus rhythm.



3  Sensing and Detection

Automatically adjusted sensitivity (mV)

10.0 Boston Scientific Medtronic St. Jude Threshold

9.0 8.0 7.0 6.0 5.0 4.0 3.0 2.0 1.0 0.0 0

200

400

600

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Time (msec) Figure 3-20  Comparison of automatically adjusted sensitivity after sensed ventricular events for three manufacturers of ICDs after large (10 mV) R wave. The programmed sensing threshold is approximately 0.3 mV. After sensed ventricular events, Medtronic ICDs reset the sensing threshold to 8 to 10 times the programmed sensitivity, up to a maximum of 75% of the sensed R wave. The value of Auto-Adjusting Sensitivity then decays exponentially from the end of the (sense) blanking period, with a time constant of 450 msec, until it reaches the programmed (maximum) sensitivity. At the nominal sensitivity of 0.3 mV, there is little difference between the sensitivity curves of Medtronic ICDs after large and small spontaneous R waves. If the R wave is large, the entire Auto-Adjusting Sensitivity curve can be altered substantially by changing the programmed value of maximum sensitivity (see Fig. 3-19). At nominal settings, the St. Jude Threshold Start begins at 62.5% of the measured R wave for values between 3 and 6 mV. If the R-wave amplitude is greater than 6 mV or less than 3 mV, the Threshold Start is set to 62.5% of these values (3.75 mV and 1.875 mV, respectively). The sensing threshold remains constant for a Decay Delay period of 60 msec and then decays linearly with a slope of 3 mV/sec. Both the Threshold Start percent and the Decay Delay are programmable, over the range of 50% to 75% and 0 to 220 msec, respectively (see Fig. 3-21). Boston Scientific Cognis-Telegin ICDs set the starting threshold to 75% of sensed R waves with a maximum limit of 3/2 · Peak Running Average. Sensitivity then decays using digital steps, each seven-eighths the amplitude of the previous step. For sensed events, the duration of the first step is 65 msec, and the duration of subsequent steps is 35 msec. This results in a sensitivity of one-half the peak R wave in about 170 msec. (See text for further details.) After a paced ventricular event, all ICDs also adjust sensitivity dynamically, starting at the end of the (pace) blanking period, but the threshold starts at a more sensitive setting. (Modified from Swerdlow C, Friedman P: Advanced ICD troubleshooting. Part I. Pacing Clin Electrophysiol 28:1322-1346, 2005.)

Automatic analysis of the far-field EGM has also been proposed as a method to identify oversensing in ICDs resulting from lead fracture or sensing-lead connection problem. In one approach, the peak-peak far-field EGM amplitude is measured in a small window centered around each sensed event (on near-field channel) to discriminate rapid oversensing from true VT/VF. Oversensing is identified when sensed events on the near-field channel correspond to isoelectric periods on the far-field channel.55,56 In true VF, isoelectric periods are rare on the far-field channel (Fig. 3-23). Evaluating Sensing of Ventricular Fibrillation at Implantation Increasing interest in implanting ICDs without assessing defibrillation efficacy has focused attention on the extent to which adequacy of VF sensing can be determined from EGMs recorded in baseline rhythm.

67

Although the statistical correlation between R-wave amplitude in VF and baseline rhythm is weak,57,58 two studies reported that sensing of VF is adequate with nominal sensitivities near 0.3 mV if the baseline R wave is sufficiently large (≥5 mV or ≥7 mV).59 Rarely, clinically significant undersensing of VF or polymorphic VT may occur despite adequate sinus-rhythm R waves.48,60 In these cases, undersensing occurs because auto-adjusting sensitivity criteria respond inadequately to variations in R-wave amplitude, rather than consistently low-amplitude R waves. The reproducibility of this phenomenon is unknown, as is its predicted extent at implantation. Therefore, it is uncertain whether clinically appropriate testing at implantation can detect this infrequent cause of undersensing. During ICD implant with true bipolar sensing and current digital sensing amplifiers, clinically significant undersensing of VF is rare and unrelated to sinus-rhythm R-wave amplitude.61 Undersensing of spontaneous VT/VF in the VF zone is similarly rare. Reliable sensing of VF cannot be predicted from baseline EGMs if the baseline ventricular rhythm is paced. Sensitivity is programmed to a less sensitive value than nominal (e.g., to avoid T-wave oversensing), or patients have other implanted electronic ICDs, such as pacemakers, cardiac contractility modulation devices, or transcutaneous electrical nerve stimulation (TENS) units, that could cause device-device interactions. Postshock Sensing Postshock sensing is critical for redetection of VF after unsuccessful shocks and for accurate detection of episode termination. Electroporation, the process by which strong electric fields create microscopic holes in the cardiac cell membranes, has been proposed as the mechanism for postshock distortion of EGMs recorded from high-voltage electrodes.62 Because EGMs of dedicated bipolar sensing electrodes are minimally affected by shocks,63 they became standard for early epicardial ICDs. For transvenous ICDs, postshock sensing recovers more rapidly with true bipolar sensing configurations than with integrated bipolar sensing.64,65 This is a minor issue for current integrated bipolar leads with a pacing tip electrode–to–distal coil spacing of approximately 12 mm.66 ATRIAL SENSING IN DUAL-CHAMBER ICDS AND ATRIAL ICDS Accurate sensing of atrial EGMs is essential for accurate discrimination between VT/VF and rapidly conducted SVTs that satisfy ventricular rate criteria in dual- or triple-chamber ICDs. Rapid discrimination is essential to ensure prompt delivery of ventricular therapy while minimizing inappropriate shocks. Historically, some inappropriate detection of AT/AF has been considered an acceptable consequence of maintaining high sensitivity for detecting VT/VF. The atrial lead should be positioned at implantation to minimize FFRWs. Leads with an interelectrode spacing of 10 mm or less reduce oversensing of FFRWs. Atrial lead dislodgement, oversensing of FFRWs, or undersensing from low-amplitude AEGMs or atrial blanking periods can cause inaccurate identification of AEGMs. These errors in sensing may result in misclassification of VT as SVT, or vice versa. Postventricular Atrial Blanking and Rejection of Far-Field R Waves To prevent oversensing of FFRWs, older dual-chamber ICDs had fixed PVAB periods, similar to those in pacemakers (Fig. 3-24). With a fixed blanking period, the blanked proportion of the cardiac cycle increases with the ventricular rate. Atrial undersensing caused by PVAB causes underestimation of the atrial rate during rapidly conducted atrial flutter or AF, resulting in inappropriate detection of VT67 (Fig. 3-24, lower panel). Without PVAB, however, atrial oversensing of FFRWs could cause overestimation of the atrial rate during tachycardias with a 1 : 1 AV relationship.68 This may result in either inappropriate rejection of VT as SVT, if FFRWs are counted consistently as atrial EGMs, or inappropriate detection of SVT as VT, if FFRWs are counted inconsistently.69

68

SECTION 1  Basic Principles of Device Therapy

Postpace

Postsense

Filtered and rectified EGM Postsensed maximum

8x

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0.3mV

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Figure 3-21  Comparison of automatic adjustment of sensitivity after paced and sensed ventricular events. The filtered and rectified ventricular EGM, the corresponding sensing threshold, and ventricular (V) event markers are shown. Horizontal bars denote postpacing blanking periods. The blanking periods after paced events are longer than those after sensed events (250-350 vs. 120-140 msec), and the initial values of sensitivity are less. In this Medtronic ICD example, the initial sensing threshold after sensed events is 8 times the minimum programmed sensitivity of 0.3 mV, whereas the initial threshold after paced events is 4.5 times that value. The goal of a lower initial postpacing threshold is to prevent pacing into ventricular fibrillation. It compensates in part for the longer postpacing blanking period.

Medtronic ICDs also reject FFRWs algorithmically by identifying a specific pattern of atrial and ventricular events that fulfill specific criteria (Fig. 3-25). Intermittent sensing of FFRWs or frequent premature atrial events may disrupt this pattern, resulting in misclassification of a tachycardia. Therefore, it is preferable to reject FFRWs after sensed ventricular events by decreasing atrial sensitivity, if this can be done without undersensing of AF. Atrial sensitivity can be reduced to 0.45 mV with a low risk of undersensing AF. Less sensitive values should be programmed only if the likelihood of rapidly conducted AF is low. FFRW oversensing that occurs only after paced ventricular events (when auto-adjusting atrial sensitivity is maximal) does not cause inappropriate detection of SVT as VT, but it may cause inappropriate 100

Number of intervals

SENSING IN CARDIAC RESYNCHRONIZATION ICDS

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mode switching and can contribute to inappropriate detection of AF or atrial flutter. Medtronic ICDs (starting with Entrust) and Boston Scientific ICDs (starting with Vitality) may use brief atrial blanking or a period of reduced, automatically-adjusting sensitivity (or both) to reject FFRWs without preventing detection of AF (Fig. 3-26). St. Jude ICDs and Medtronic ICDs starting with Entrust provide programmable atrial blanking after sensed ventricular events to individualize the trade-off between oversensing of FFRWs and undersensing of AEGMs in AF. St. Jude ICDs also provide programmable atrial sensing Threshold Start and Decay Delay, corresponding to the same features in the ventricular channel.

600

Cycle length (msec) Figure 3-22  Sensed cycle lengths of human ventricular fibrillation (VF) EGMs by ICD sensing system compared with manual cyclelength measurements of same signals from many patients. Two histograms are superimposed. The red and orange vertical bars show the manually measured cycle lengths during VF (786 intervals); the blue bars show the intervals sensed by the ICD (772 intervals). Intervals of less than 120 msec, the blanking period, were not permitted. Note that there was some oversensing for intervals shorter than 180 msec, a slight amount of undersensing for intervals between 180 and 280 msec, and a small number of long intervals greater than 280 msec, which represent undersensing during VF. The peak in the histograms occurs at about 220 msec.

Left ventricular PVCs may activate the left ventricle but may not be sensed in the right ventricle before pacing is delivered. In this case, LV pacing could be delivered during the vulnerable period of the PVC. Boston Scientific cardiac resynchronization ICDs also sense in the left ventricle to reduce ventricular proarrhythmia by preventing pacing into the LV vulnerable period. The left ventricular protection period (LVPP) is a programmable interval (300-500 msec) following an LV event when the ICD will not pace in the left ventricle (Fig. 3-27). This prevents inadvertent delivery of an LV pacing pulse during the LV vulnerable period. The LVPP differs from other pacing refractory periods, which are designed to prevent inappropriate inhibition of pacing. In contrast, the left ventricular refractory period (LVRP), after a sensed or paced event on the LV lead, is a conventional refractory period. It prevents sensed events from causing inappropriate loss of cardiac resynchronization pacing following sensed events such as T-wave oversensing on the LV lead. The LVRP provides an interval following either an LV sense or an LV pace event (or leading ventricular pace event when LV offset is not programmed to zero), during which LV sensed events do not inhibit pacing. Use of a long LVRP shortens the LV sensing window. LVRP is available whenever LV sensing is enabled. Thus, LVRP minimizes unnecessary inhibition of resynchronization pacing while the LVPP minimizes the risk of LV pacing during the LV vulnerable periods.

Ventricular Oversensing: Recognition and Troubleshooting Oversensing is defined as sensing of unintended nonphysiologic signals or of physiologic signals that do not accurately reflect local depolarization. Nonphysiologic signals usually arise from extracardiac EMI (e.g.,



3  Sensing and Detection Lead noise

69

Ventricular fibrillation

Near-field EGM

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Marker F 2 S2 1 7 0 0

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Figure 3-23  Algorithm to discriminate pace-sense conductor failure of true bipolar ICD leads from ventricular fibrillation using far-field EGM (Medtronic Protecta ICDs, currently investigational). In pace-sense conductor failures, the far-field EGM has long isoelectric segments, which are not present in true VF. Stored EGM tracings during lead noise oversensing (left) and spontaneous VF (right) are shown; top tracings are the near-field (RVtip-RVring) EGM, middle tracings are the far-field EGM (RV coil-can), and the sensing markers are the lower tracing. Oversensing is identified by analyzing the peak-peak amplitudes of the far-field EGM in a 200-msec window centered at each of 12 sense markers (sensing derived from near-field EGM), as shown by the red arrows and boxes. Oversensing is identified when there are at least two far-field EGM analysis windows with isoelectric (<1 mV) amplitudes and at least one analysis window with peak-peak amplitude at least six times greater than the average of the two smallest amplitudes. Lead noise oversensing would have been properly discriminated from VF in these two examples.

ungrounded electrical equipment, antitheft detectors, surgical electrocautery) or mechanical problems in the sensing circuit, which may occur in the lead (insulation failure or fracture), connection between lead and header, or feedthrough wires from the header to the inside of the pulse generator. Rarely, electrical artifacts arise from retained fragments of abandoned intracardiac leads. Physiologic signals may be intracardiac (P, R, or T waves) or extracardiac myopotentials. Oversensing often results in characteristic patterns of stored EGMs and associated markers70,71 (Fig. 3-28). In pacemakers, oversensing may manifest as either failure to deliver an expected pacing stimulus or inappropriate tracking of nonphysiologic signals in the atrial channel of a dual-chamber ICD. In ICDs, oversensing usually manifests as inappropriate detection of VT or VF. This section focuses on oversensing in ICDs.

Figure 3-24  Effect of postventricular atrial blanking (PVAB) on atrial sensing. Upper panel shows the surface electrocardiogram (ECG), atrial electrogram (AEGM), and event markers during atrial sensed (As)–ventricular paced (Vp) rhythm. The first segment of each upper horizontal bar denotes the period of PVAB; the second segment denotes the postventricular atrial refractory period (PVARP); FFRW, far-field R wave on the atrial channel. The lower horizontal bars denote the postpacing ventricular blanking periods. With a short PVAB (left), the FFRWs are oversensed. A longer PVABP (right) prevents oversensing of FFRWs. Lower panel shows the ECG, AEGM, and ventricular electrogram (VEGM) tracings from a patient with atrial flutter with 2 : 1 atrioventricular (AV) conduction. The thin horizontal bars on the ventricular channel denote the periods of PVAB, which resulted in atrial undersensing of alternate atrial flutter EGMs (in boxes). The resultant incorrect calculation of atrial rate resulted in inappropriate ventricular therapy for atrial flutter (not shown), because the ICD did not classify the ventricular rate as less than the atrial rate.

PVAB

INTRACARDIAC SIGNALS Ventricular oversensing of physiologic intracardiac signals results in exactly two ventricular events for each cardiac cycle. This may result in inappropriate detection of VT or VF. T-wave Oversensing T-wave oversensing is the most common clinically-significant cause of intracardiac oversensing. Oversensing of spontaneous T waves may cause inappropriate detection of either VT or VF, depending on the sensed R wave–to–T wave interval and the programmed VF detection interval. T-wave oversensing is typically identified by alternating EGM morphologies.70 Device-measured intervals during T-wave oversensing may also alternate, but the degree of alternation often is small and

PVAB

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70

SECTION 1  Basic Principles of Device Therapy

P

Figure 3-25  Algorithmic rejection of far-field R waves (FFRW) by pattern analysis. ICDs with minimum cross-chamber blanking (Medtronic) reject FFRWs by the timing pattern of atrial and ventricular intervals. Atrial events are classified as FFRWs if all the following criteria are met: (1) there are exactly two atrial events for each V-V interval; (2) timing of one P wave is consistent with a FFRW (R-P interval <160 msec); (3) there is a stable interval between the FFRW and the ventricular electrogram (VEGM); (4) there is a short-long pattern of P-P intervals (to distinguish FFRW oversensing from atrial flutter); and (5) the pattern occurs frequently (4 of 12 intervals). A, Atrial electrogram (AEGM), VEGM, and dual-chamber event markers are shown. All five criteria for FFRWs are fulfilled. Horizontal double-ended arrows below event markers denote alternation of long (L) and short (S) atrial intervals. AR, Atrial refractory event; P, P wave; TS, ventricular event in ventricular tachycardia zone. B, Interval plot (left) and stored EGM (right) from episode of sinus tachycardia with consistent FFRW oversensing that is classified correctly. On the interval plot, open squares denote A-A interval and closed circles denote V-V interval. Horizontal lines denote ventricular tachycardia (VT) and ventricular fibrillation (VF) detection intervals of 400 and 320 msec, respectively. Alternating A-A intervals whose sum equals that of the VV intervals produce a characteristic “railroad track” appearance (arrow). The algorithm rejects FFRWs despite that FFRW oversensing does not occur for one V-V interval between seconds 9 and 10.  C, Interval plot (left) and stored EGM (right) from episode of sinus tachycardia with consistent FFRW oversensing that was detected inappropriately as VT. Intermittent oversensing of FFRWs occurs as sinus tachycardia accelerates gradually across the VT detection interval of 480 msec (arrow), resulting in inappropriate therapy (“Burst” antitachycardia pacing marker on interval plot, VT marker on EGM event markers).

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71

3  Sensing and Detection

Sensitivity threshold

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Rectified and filtered AEGM A P

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A As 75% of AS

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75% of 3/8 of AS As Blanking 15 ms

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Far-field V

Figure 3-26  Atrial sensing features intended to prevent far-field R-wave (FFRW) oversensing while still permitting detection of atrial fibrillation. These features, which combine limited atrial blanking with automatic adjustment of atrial sensitivity, are less strict than conventional blanking periods. A, Medtronic ICDs, starting with Entrust, have an optional “Partial+” postventricular atrial blanking period (PVAB) setting that increases the sensing threshold for a programmable duration (40-100 msec) after a ventricular event (2). Entrust operation after ventricular sensed events nominally has no absolute blanking or sensitivity changes. With Partial + PVAB enabled, the sensing threshold is raised to 75% of the prior sensed atrial electrogram (AEGM) amplitude to reduce the likelihood of sensing FFRWs. “Absolute” PVAB, another selectable option, blanks all atrial sensed events within the programmable PVAB interval. Because significant atrial undersensing can occur, this method is recommended only for addressing complications not addressed by the Partial + PVAB method. Partial + or Absolute PVAB does not affect atrial sensing operation after sensed events (1) or after atrial paced events (3). AP, Atrial paced event; AS, Atrial sensed event; mS, microseconds; VS, ventricular sensed event. B, Boston Scientific ICDs, starting with Vitality, introduce a brief (15 msec) blanking period after each sensed ventricular event, followed by a decrease in atrial sensitivity to three-eighths the mean P-wave amplitude. This is sufficient to prevent oversensing of FFRWs, provided that the amplitude of the AEGM is at least eight-thirds that of the FFRW.

2. St. Jude ICDs provide programmable Threshold Start and Decay Delay to reduce oversensing of spontaneous T waves. These features may be helpful even if the R wave is small (Fig. 3-30). 3. The apparent alternation of ventricular EGM morphologies caused by T-wave oversensing may be exploited to prevent inappropriate detection of VT. The SVT-VT morphology discriminator may be programmed “on” to classify alternate EGMs as “sinus” and thereby withhold inappropriate detection. The long-term reliability of this approach is not established. 4. As a temporary measure, some T-wave oversensing may be eliminated by forcing short-term ventricular pacing. This alters the sequence of repolarization and may reduce T-wave amplitude. However, unnecessary ventricular pacing may result in adverse, desynchronizing hemodynamic effects. 5. Medtronic ICD models (Secura VR/DR, Consulta) permit sensing from either the true bipolar EGM or the integrated bipolar EGM when a true bipolar lead is connected. Integrated-bipolar sensing may prevent T-wave oversensing, especially when true bipolar R-waves are small compared to the integrated bipolar R waves (Fig. 3-31).

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B difficult to recognize, especially during sinus tachycardia or if the QT interval is prolonged. T-wave oversensing may be divided into three classes: postpacing, large R wave (>3 mV) in spontaneous rhythm, and small R wave (<3 mV) in spontaneous rhythm (Fig. 3-29). Postpacing T-wave oversensing can inhibit bradycardia pacing49,72 or cause antitachycardia pacing to be delivered at the wrong cycle length.49 It does not typically cause inappropriate detection of VT but may increment VT or VF counters, making detection of nonsustained VT more likely. It is corrected by increasing the postpacing ventricular blanking period. Oversensing of spontaneous T waves often occurs in the setting of low-amplitude R waves because the sensing threshold decays from a fraction of the preceding low-amplitude R wave.73 To compound the problem, patients with low-amplitude R waves may require lower minimum sensing thresholds to ensure reliable sensing of VF. T-wave oversensing in these patients is a warning that detection of VF may be unreliable. The ventricular lead should be revised or a separate pace/ sense lead added if the safety margin for sensing VF is insufficient. Specific programming features may be used to reduce T-wave oversensing, provided that detection of VF is reliable, as follows: 1. The simplest approach is to reduce ventricular sensitivity to a less sensitive (higher) value. This may result in undersensing during VF if the R wave amplitude is low.

Figure 3-27  Event markers illustrate blanking periods in cardiac resynchronization ICDs. A, Blanking periods in Medtronic Marquis Insync III ICD. In condition 1, there is simultaneous right ventricular and left ventricular pacing (RV+LVp). There is a single cross-chamber blanking period (~30 msec in duration) on the atrial channel. In condition 2, RVp first with LVp delayed (up to 80 msec), and in condition 3, LVp first with RVp delayed, there are two cross-chamber blanking periods on the atrial channel, one for each ventricular paced event. In both cases, ventricular blanking is timed from the RVp event, resulting in a postpace ventricular blanking period that is longer by the V-V pacing delay (up to 80 msec). These additional blanking periods increase the risk of undersensing in both chambers. To date, there are no reports of undetected ventricular tachycardia (VT) caused by offsets in timing of V-V pacing. B, Left ventricular protection period (LVPP) of the Boston Scientific Renewal ICD. This ICD has RV timing and independent RV and LV stimulation channels without a programmable RV-LV delay. The LVPP allows LV sensing to occur with inhibition of the LV stimulus for a period that is programmable to prevent pacing into the vulnerable period of the left ventricle. Therefore, this feature allows biventricular sensing but RV-based timing. AS, Atrial sensed event; BiV, biventricular; LVP, left ventricular paced; LVS, left ventricular sensed; RVP, right ventricular sensed. (Courtesy Boston Scientific, Indianapolis; B from Kay G: Troubleshooting and programming of cardiac resynchronization therapy. In Ellenbogen K, Kay G, Wilkoff B, editors: ICD therapy for congestive heart failure. Philadelphia, 2004, Elsevier.)

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SECTION 1  Basic Principles of Device Therapy

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Figure 3-28  Types of oversensing resulting in inappropriate detection of ventricular tachycardia/fibrillation (VT/VF). A, B, and C, Oversensing of physiologic intracardiac signals. D, E, and F, Oversensing of extracardiac signals. A, P-wave oversensing in sinus rhythm from an integrated bipolar lead with distal coil near the tricuspid valve. B, R-wave double-counting during conducted atrial fibrillation (AF) in a biventricular sensing ICD. C, T-wave oversensing in a patient with low-amplitude R wave (note millivolt calibration marker). D, Electromagnetic interference from a power drill has higher amplitude on widely spaced, high-voltage EGM than on closely spaced, true bipolar sensing EGM. E, Diaphragmatic myopotential oversensing in a patient with an integrated bipolar lead at the right ventricular apex. Note that the noise level is constant, but oversensing does not occur until automatic gain control increases the gain sufficiently, about 600 msec after the sensed R waves. F, Lead fracture noise results in intermittent saturation of amplifier range, denoted by arrows. RA, Right atrium; RV, right ventricular sensing EGM; HV, high-voltage EGM. (From Swerdlow C, Shivkumar K: Implantable cardioverter defibrillators: clinical aspects. In Zipes DP, Jalife J, editors: Cardiac electrophysiology: from cell to bedside, ed 4, Philadelphia, 2004, Saunders.)

Spontaneous rhythm: Inappropriate therapy

Postpacing: Pauses in paced rhythm or ATP ECG

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Figure 3-29  Classification of T-wave oversensing. Left panel, Postpacing oversensing results in inhibition of pacing. Inhibition of antitachycardia pacing may also occur. Middle and right panels, T-wave oversensing in spontaneous rhythm can result in inappropriate detection of ventricular tachycardia or ventricular fibrillation. Management often depends on whether the R waves are small (middle) or large (right panel).



3  Sensing and Detection 8

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R

Threshold start Decay delay 60 msec R wave is sensed

Decay delay 0 msec

125 msec Figure 3-30  Specific programmable features to correct T-wave oversensing. Upper panel, Stored EGM from St. Jude ICD showing inappropriate shock for sinus tachycardia with T-wave oversensing. The lower strip was recorded from the programmer after reprogramming of the Decay Delay from 0 to 220 msec. Lower panel, Diagram of programmable features to avoid and correct T-wave oversensing in St. Jude ICDs. Automatic Sensitivity Control begins to adjust sensing at the end of the 125-msec ventricular blanking period. Both the initial sensitivity (Threshold Start as percentage of R-wave amplitude) and the time delay before onset of linear decrease in sensing threshold (increase in sensitivity) are programmable parameters. (From Swerdlow CD, Friedman PA: Advanced ICD troubleshooting. Part I. Pacing Clin Electrophysiol 28:1322-1346, 2005.)

EGM1: RVtip to RVring (1 mV)

EGM2: RVtip to RVcoil (1 mV)

Oversensed T wave Markers

↑ V S 5 4 0

↑ V S

↑ V S 5 4 0

↑ V S 5 4 0

↑ V S 5 4 0

↑ V S 5 5 0

↑ F S 2 8 0

↑ F S 2 3 0

↑ V S 5 3 0

Figure 3-31  True bipolar (RVtip to RVring) signal, integrated bipolar (RVtip to RVcoil) signal, and ventricular marker channel during true bipolar sensing shows intermittent T-wave oversensing (arrow) caused by low-amplitude true bipolar signal, despite raising sensing threshold from nominal 0.30 mV to 0.45mV. Programming sensing to the integrated bipolar signal, if available, would prevent T-wave oversensing.

74

SECTION 1  Basic Principles of Device Therapy

Vtip-Vring EGM

Filtered and rectified Vtip-Vring EGM

First differential R filtered and rectified Vtip-Vring EGM Vs

R T Vs

R T

Vs

Vs

R T

Vs

Vs

Vs

Figure 3-32  Algorithm to reject T-wave oversensing. Upper tracing, Vtip-Vring EGM with large T waves and T-wave oversensing. Middle tracing, Signal after standard sense amplifier filtering and rectification (purple line), automatic adjusting sensing threshold (red), and peak amplitude at each sensed event (blue; the peak amplitude for each sensed event is held until the next sensed event). Bottom tracing, First derivative of the filtered and rectified EGM (purple line) its peak amplitude at each sensed event (blue), and adaptive threshold used to separate possible R waves and T waves (red horizontal line). The algorithm recognizes oversensed T waves with lower frequency content than the preceding sensed R waves by automatically comparing signal amplitudes in the filtered/rectified sensing EGM and its first derivative (which further attenuates low-frequency signals) to identify patterns of alternating signal frequency content. Using an analysis window of six sensed events, the algorithm first automatically identifies possible R waves and T waves based on comparing peak amplitudes and patterns of the first derivative of the filtered/rectified sensing EGM. For each analysis window, the adaptive threshold is set to a fraction of the three largest first differential EGM peak amplitudes. The algorithm assumes that possible R waves are above threshold and possible T waves are below threshold. An alternating RT pattern with 3Rs and 3Ts or 4Rs and 2Ts in the six sensed events analysis window is required. RT pairs are confirmed when the R-wave/T-wave (R/T) ratio in the filtered/rectified sensing EGM is less than the R/T ratio measured in the first differential filtered/rectified EGM. This approach does not increase the risk of undersensing because the operation of the standard ICD sensing channel is not modified. It does not prevent oversensing of T waves, but rather allows the ICD to withhold therapy when T-wave oversensing pattern is recognized. Testing on more than 1000 VT/VF episodes showed no loss in sensitivity to true VT/VF. Vs, Ventricular sensed event.

6. An investigational ICD (Medtronic Protecta) has an algorithm that withholds ICD therapy when T-wave oversensing is identified as the source of device-detected tachycardia based on lower frequency content of the oversensed T waves (Fig. 3-32).74

between RV and LV electrodes.75 The composite ventricular EGM included deflections from the right ventricle and the left ventricle, both of which could be counted as separate R waves. Occasional R-wave double-counting may occur during PVCs and is more common with integrated bipolar than true bipolar leads.76 R-wave double-counting results in alternation of ventricular cycle lengths that produces a characteristic “railroad track” pattern on a plot of stored ventricular intervals. Because the second component of the R wave is sensed as soon as the blanking period terminates, the doublecounted RV-RV or RV-LV interval measures within 20 msec of the ventricular blanking period and is always classified in the VF zone (Fig. 3-33). Inappropriate detection of VF may occur despite a true ventricular rate below the VT detection interval. Most consistent R-wave double-counting now results from transient or reversible events such as hyperkalemia, drug effects, or lead failure in cardiac resynchronization ICDs; and most intermittent R-wave double-counting is not clinically significant. Some ICDs permit programming to increase the ventricular blanking period from the nominal value (120-125 msec) to higher values, up to 170 msec. Occasionally, reducing ventricular sensitivity can avoid R-wave doublecounting; but ventricular sensitivity should not be reduced unless reliable sensing of VF is confirmed at the reduced level of sensitivity. Rarely, lead revision is required. P-wave Oversensing P-wave oversensing may occur if the distal coil of an integrated bipolar lead is close to the tricuspid valve and if the sensed P-R interval exceeds the cross-chamber ventricular blanking period4 (Fig. 3-34). It is rare in adults with ventricular sensing electrodes near the RV apex, but it may occur in children or in adults if the RV electrode dislodges or is positioned in the proximal septum or inflow portion of the right ventricle. If P-wave oversensing occurs during a 1 : 1 rhythm, the sensed “R-R” pattern is similar to that of R-wave double-counting, provided that the sensed P-R or R-P interval is less than the VF detection interval. However, oversensing of P waves as R waves can cause inappropriate detection of VF during AF or atrial flutter, independent of the ventricular rate. Consistent oversensing of spontaneous P waves often requires lead revision. One amelioration strategy is to force atrial pacing using DDDR or Dynamic Overdrive modes. This shortens the ventricular cycle length (preventing ventricular sensitivity from reaching its minimum value) and introduces cross-chamber ventricular blanking after each atrial event (reducing the likelihood of oversensing P waves). Far-Field R-wave Oversensing The FFRW oversensing on the atrial channel, the inverse of P-wave oversensing on the ventricular channel, shows a pattern of alternating atrial cycle lengths with one sense marker timed close to the ventricular EGM. If rate and duration criteria for VT are fulfilled, FFRW oversensing may confound dual-chamber SVT-VT discrimination. FFRW oversensing does not cause inappropriate detection of VT if the ventricular rate is in the sinus zone, but it may cause inappropriate mode switching.

R-wave Double-Counting

EXTRACARDIAC SIGNALS

R-wave double-counting occurs if the duration of the sensed EGM exceeds the short ventricular blanking period in ICDs; currently, this is a rare clinical problem. Consistent R-wave double-counting may result from local ventricular delays in the baseline state in native rhythm or those resulting from use-dependent effects of sodium channel–blocking antiarrhythmic drugs or of hyperkalemia. Doublecounting may also occur with loss of RV capture in cardiac resynchronization ICDs: The ICD counts both the paced ventricular event and the conducted wavefront arriving at the RV bipole as a result of LV capture, if the interventricular conduction delay exceeded the ventricular blanking period. R-wave double-counting was a common problem in early ICDs that used Y-adapted or extended bipolar sensing

The distinctive feature of oversensing of extracardiac signals is replacement of the isoelectric baseline with high-frequency noise that does not have a constant relationship to the cardiac cycle.70,71,75,77 External Electromagnetic Interference With oversensing of external electromagnetic interference, signal amplitude is greater on the high-voltage EGM recorded from widely spaced electrodes than on the sensing EGM recorded from closely spaced electrodes. Although the interference signal may be continuous, oversensing is often intermittent because of auto-adjusting sensitivity. Clinical data may suggest a specific identifiable cause (Fig. 3-35). Some manufacturers (St. Jude, Sorin, Boston Scientific) have specific ICD



3  Sensing and Detection

75

ECG

A P

Marker

RVtip

LV-RV delay

RVringt

V blank A P

F S

RV Sensed event

VT interval = 400 msec FVT interval = 270 msec VF interval = 320 msec Placed event LV capture No RV capture Figure 3-33  Intermittent R-wave double-counting as a result of loss of RV capture in a cardiac resynchronization ICD patient with left bundle branch block. Left main panel shows surface ECG, atrial and ventricular marker channels, and true bipolar right ventricular (RV) EGM. The third, fifth, and sixth R waves are double-counted, as shown in insert on right. Simultaneous biventricular pacing activates the left ventricle (LV). After the LV-RV conduction delay, the wavefront arrives at the RV sensing bipole, corresponding to a sensed interval in the ventricular fibrillation (VF) zone (VS) and incrementing the VF counter. This condition alone will not result in VF being detected even if all beats are oversensed, but transient nonsustained ventricular tachycardia (VT) in the VF zone could result in inappropriate detection of VF, since the baseline VF counter may be as high as 50% of its maximum value.

algorithms to prevent inappropriate detection of VF in the presence of EMI (Fig. 3-36). Lead/Connector Problems Oversensing caused by lead insulation failure or conductor fracture, lead-lead mechanical interactions (“chatter”),76 or connector (header, adapter, set-screw) problems is intermittent and may occur only during a small fraction of the cardiac cycle. In ICDs, it is limited to the sensing EGM unless the problem relates to the RV coil in an integrated bipolar lead or to both sensing and high-voltage conductors/connectors. The rate of oversensed signals is always in the VF zone. Intervals below 200 msec are typical. Signals almost always display substantial variability in amplitude or frequency. High-frequency components are

common, and high-amplitude signals often saturate the amplifier (Figs. 3-37 to 3-39). In ICDs, extremely short R-R intervals near the ventricular blanking period (120-150 msec) do not represent successive cardiac depolarizations, except occasionally during VF. Medtronic ICDs count these very short intervals as a measure of nonphysiologic oversensing to provide early warning of lead fracture; they store this count as the sensing integrity counter (SIC). A high or rapidly accumulating Sensing Integrity Count (total >300 or 10/day for 3 consecutive days) is a sensitive indicator of pace-sense lead fracture,3,21,22 but in isolation is nonspecific. The most common cause of isolated, extremely short “R-R” intervals is a combination of an oversensed physiologic event and appropriately sensed R wave.

ECG

RV Tip-coil

Shock

Marker VT-1 VS 368 630

VS 695

VS 535

VS 538

VS 483

VS 560

VS 540

VS 520

VS 525

VS 435

VS 613

VS 535

VS 520

VS VS 425 510

VS 598

VS 508

VS 508

Figure 3-34  P-wave oversensing. The surface electrocardiographic (ECG) lead, integrated bipolar right ventricular electrogram (RV Tip-coil), and ventricular event markers are shown. P-wave oversensing caused by the proximity of the RV coil to the right atrium results in ICD sensing of alternating sinus (ventricular sensed, VS) and ventricular fibrillation (VF) intervals, the latter corresponding to the P-R interval. The sum of the VS and VF intervals equals the true R-R interval.

76

SECTION 1  Basic Principles of Device Therapy

Atrial

RV tip-coil Shock

--735

--

-783

-803

-783

-785

-803

-785

-705

-521 -193

-289 -501

-AS AS 129 143 205 -AS AS 455 473 121 -VF 201 209 --VP 403 160 501

AS 215 VS 380

AS 168

AS AS AS AS AS AS 117 127 137 88 156 246 AS AS AS AS AS AS AS 174 127 102 314 125 176 111 VS VF VF VF VF 429 137 234 176 180 VF VF VF VF VP 152 143 207 137 507

Figure 3-35  External electromagnetic interference caused by electric drill. Atrial EGM, integrated bipolar (RV tip-coil) EGM, and high-voltage (Shock) EGM are shown, together with a dual-chamber channel showing event markers. Electromagnetic interference (EMI) is continuous. Oversensing occurs first on the atrial channel, which has a higher sensitivity setting (and gain) than the ventricular sensing channel. Although the amplitude of external EMI is typically less on true bipolar sensing electrodes than on shocking electrodes, it may be similar on integrated bipolar and shocking electrodes.

In P-wave oversensing, the nonphysiologic short interval is the P-R interval. In R-wave double-counting, it is the interval between initial and terminal deflections of a true ventricular EGM. In T-wave oversensing, the interval is bounded by an oversensed T-wave and a PVC.76 However, these nonphysiologic combinations of physiologic signals rarely result in repetitive oversensing, which is common in lead or connector problems. Repetitive, transient oversensing may be identified by stored “nonsustained tachycardias.” The combination of isolated, extremely short RR intervals and repetitive rapid oversensing (defined as at least two nonsustained tachycardias with duration ≥5 intervals and mean cycle length ≤220 msec) has a positive predictive value for lead or connector problems of about 80%,76,78 even with a normal pace-sense impedance. Unusual cases of true VF76 and other unusual conditions, such as device-device interactions or header sealplug problems shortly after implant,79 can produce both isolated, very short intervals and rapid, repetitive oversensing. Role of lead impedance.  Modern pacemakers and ICDs perform daily, automated measurements of impedance in lead conductors. A large, abrupt change in impedance associated with rapid oversensing

VT-1 VS 368 630

VS 695

VS 535

VS 538

VS 483

VS 568

VS 540

VS 520

is, for practical purposes, diagnostic of a lead or header-connector problem. However, pace-sense conductor fractures typically present as rapid, repetitive oversensing without abnormal impedance.78 A fracture may generate sufficient oversensing to cause inappropriate shocks without an impedance increase or with an impedance increase in the interval between daily impedance measurements.78,80-82 Rarely, pacing precipitates lead-related oversensing.81 Lead fractures or headerconnector problems also present frequently as abrupt increases in impedance without oversensing. The Medtronic Lead Integrity Alert (LIA) is triggered by either abnormally high impedance relative to the patient’s baseline impedance or rapid oversensing unlikely to represent a physiologic event55,78 (Figs. 3-40 and 3-41). Once an alert is triggered, an audible tone sounds immediately and every 4 hours thereafter. LIA alerts also reprogram the number of intervals to detect VF (NID) from the value programmed at the physician’s discretion (nominally 18/24) to 30/40; EGMs are stored for any interval shorter than 200 msec. The purpose of the high NID is to reduce inappropriate shocks caused by transient, fracture-induced oversensing.78 The stored EGMs provide a diagnostic clue to determine the cause of rapid oversensing based on EGM

VS 525

VS 435

VS 613

VS 535

VS 520

VS VS 425 510

VS 598

VS 508

VS 508

Figure 3-36  Noise rejection algorithm prevents oversensing of external electromagnetic interference caused by electroconvulsive therapy, despite noise amplitude exceeding programmed sensitivity. Surface ECG, integrated bipolar sensing EGM, shock EGM, and ventricular marker channel are shown. The Dynamic Noise Algorithm in Boston Scientific Cognis-Telegin ICDs measures the energy in the upper half of the sensing bandwidth (20-85 Hz) and sets a noise-energy–based sensing level (“noise floor”), which is usually lower than the programmed maximum sensitivity. A higher-than-expected fraction of energy in the upper half of the bandwidth suggests the presence of noise, and the ICD responds by increasing the noise floor above the minimum programmed level of sensitivity while noise is present. (Courtesy Dr. Andrei Catanchin.)



77

3  Sensing and Detection

Atip-Aring

0.5mV

8mV

Vtip-Vring

5

5

5

0

0 A

0

0

S V S

5

0 A

5

0 A

0

R

0

R F

0

R F

F

4

0 A

0

R

F

F

S 2

S2

S1 S 2

F

0

1

4

6

3

5

6

7

0

6

3

6

0

0

0

0

0

0

0

0

0

0

0

0

S

A-A

Interval (ms) 1800

V 4

F

F

S1 S1 S

0

F

0

4

1 A

0

R

5

9 A

0

R

V F

V

F

F F

S1 S 3

S 2

S 2

S1 S1 S 2

F

6

3

3

0

9

7

2

5

0

0

0

0

0

0

0

0

5

1 A

0

R

3

S1 S

1 A

0

R F

0

R

V

F

F

4

S 2

S 2

D1

2

3

9

9

8

0

0

0

0

0 VF

S

VF = 300 ms 29.6 J

11.9 J

1500 1200

5

0 A

R V

3

5

0 A

R

V

S1 S1 S 2

5

9 A

5

V-V

F

5

0 A

30.1 J 30.1 J

20.2 J

900 600 400 200

−20

−10

0

10

20

30

40

50

60

70

80

Time (sec) [0=Detection] Figure 3-37  Oversensing caused by lead insulation failure. Top panel shows the atrial (Atip-Aring) EGM, the ventricular true bipolar (Vtip-Vring) EGM, and event markers. Bottom panel shows the interval plot. In the ventricular channel, there are intermittent, erratic signals that saturate the sensing amplifier. The interval plot highlights the wide range of sensed intervals, including many “short” intervals close to the ventricular blanking period of 120 msec. Arrows denote five inappropriate shocks.

characteristics.76 In newer ICDs with wireless telemetry, alerts in both groups may also initiate Internet notifications. Myopotential Oversensing Myopotential oversensing may persist for variable fractions of the cardiac cycle. Diaphragmatic myopotentials are most prominent on the sensing EGM. Oversensing usually occurs after long diastolic intervals or after ventricular paced events when amplifier sensitivity or gain is maximal. It often ends with a sensed R wave, which abruptly reduces sensitivity. In pacemaker-dependent patients, diaphragmatic oversensing causes inhibition of pacing, resulting in persistent oversensing and inappropriate detection of VF (Fig. 3-42). Clinically, this may manifest as syncope from inhibition of pacing followed by an inappropriate shock. This is an exception to the clinical rule that antecedent syncope usually indicates an appropriate shock. A short time constant for automatic adjustment of sensitivity increases the probability of this type of oversensing. It is most common in male patients who have integrated bipolar leads in the RV apex.3,4 Oversensing of diaphragmatic myopotentials may be corrected by reducing ventricular sensitivity, provided that VF sensing and detection are reliable at the reduced level of sensitivity. In pacemaker-dependent patients, oversensing may also be reduced by pacing at a faster rate. Occasionally, correction requires insertion of a new rate-sensing lead away from the diaphragm.

Pectoral myopotentials are more prominent on a far-field EGM that includes the ICD can rather than the near-field EGM. Because ICDs do not use this EGM for rate counting, oversensing of pectoral myopotentials does not cause inappropriate detection, but they may reduce the effectiveness of SVT-VT discrimination by EGM morphology algorithms during exercise-induced sinus tachycardia. Pectoral myopotentials are a major cause of oversensing in unipolar pacemakers and may result in either inhibition of ventricular pacing (if sensed by ventricular channel) or rapid ventricular pacing (if sensed and tracked by atrial channel). Myopotential oversensing may be suspected if symptoms occur during arm motion and can often be demonstrated in the clinic by having the patient forcefully press the hands together while the EGM is monitored. DETECTION ALGORITHMS FOR AUTOMATIC OPTIMIZATION OF PACEMAKER FUNCTION Pacemakers and ICDs incorporate special algorithms to prevent serious errors in pacemaker function such as inhibition during oversensing, inappropriately high pacing rates caused by tracking of atrial tachyarrhythmias, unnecessary ventricular pacing in patients with intact AV conduction, and loss of pacemaker capture. Ventricular safety pacing prevents inappropriate pacemaker inhibition caused by ventricular

78

SECTION 1  Basic Principles of Device Therapy

MR 81 PVC 0 [AG008VD] M

M

?

ECG Tip-RV Coil Tip-Ring AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS

1600 1400 1200 1000 800 600 400 200 0

VP

VP 750

VP 750

VP 750

VP 750

TS 360

VP 750

VP 750

VP 750

VP 750

ECG

LV pace RV pace

Ring-Can Tip-Ring AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS AS

4/ 13 4/ 14 4/ 15 4/ 16 4/ 17 4/ 18 4/ 19 4/ 20 4/ 21 4/ 22 4/ 23 4/ 24 4/ 25 4/ 26

Impedance

VP V-RRint

VP V-RRint

VP

VP TF 750 260

VP 750

VP 750

VS 740.30

VP 750

VP 750

VP 750

VP 750

Figure 3-38  Ventricular oversensing leads to inhibition of pacing in a pacemaker-dependent patient 2 weeks after replacement of an ICD pulse generator. Top left, In-hospital electrocardiogram (ECG) monitoring strip with 4.46-second pause in paced ventricular rhythm. The atrial rhythm is atrial fibrillation. Bottom left, Marked variability in right ventricular (RV) pacing impedance with constant impedance in left ventricular (LV) lead. The two right panels show telemetry Holter recordings, including one ECG channel, two ventricular EGMs, and event markers. Top right, Oversensing on tip-ring EGM (arrow, box) but not on tip-RV coil EGM. Oversensing is intermittent and saturates the amplifier. Bottom right, Oversensing on both tip-ring and ring-can EGMs, localizing the oversensing problem to connections related to the ring EGM (box). AS, Atrial sensed event; TS, ventricular interval in ventricular tachycardia (VT) zone; VP, ventricular paced event.

oversensing of atrial pacing stimuli (crosstalk). Noise reversion to fixedrate pacing prevents pacemaker inhibition during continuous ventricular oversensing. Automatic mode switching in dual-chamber pacing systems helps avoid inappropriate tracking of high atrial rates during atrial tachyarrhythmias. Ventricular Safety Pacing and Noise Reversion Ventricular safety pacing prevents inappropriate inhibition of ventricular pacing after atrial paced events. This feature augments the protection provided by cross-chamber blanking after atrial pacing stimuli. If a ventricular event is sensed after the delivery of an atrial pacing stimulus in a “nonphysiologic” period (10-100 msec) during the AV delay, a ventricular pacing stimulus is delivered at the end of the nonphysiologic interval (110-120 msec). This prevents pacemaker crosstalk from causing inhibition of ventricular pacing output, but it may also result in (noncaptured) ventricular pacing immediately after Before revision

After revision

Figure 3-39  Intraoperative fluoroscopic images corresponding to the data presented in Figure 3-38. The ICD has been removed from the pocket to enhance image quality. Before revision, the lead connector pin was not advanced completely into the header (arrow). The proximal connection between the ring electrode and the header was intermittent, accounting for the high impedance and oversensing illustrated in Figure 3-38. After revision, the ventricular electrode is advanced completely into the header.

true ventricular events that occur during the AV delay. Ventricular safety pacing is especially likely to occur during atrial undersensing, during AF, or during a junctional rhythm. Pacemakers also contain algorithms to protect against prolonged inhibition of ventricular pacing caused by oversensing. Noise-sampling windows (30-60 msec in duration) are used during the blanking periods to identify spurious signals and cause the ICD to change to an asynchronous “reversion” pacing mode, to ensure ventricular backup pacing.83 Noise-reversion operation is particularly important for unipolar sensing, which is more likely than bipolar sensing to exhibit oversensing of extracardiac signals84 (Fig. 3-43). Automatic Mode Switching to Avoid Ventricular Tracking of Atrial Tachyarrhythmias Pacemakers programmed to dual-chamber synchronous pacing modes may use automatic “mode switching” algorithms to initiate a temporary change to a nontracking pacing mode during paroxysmal atrial tachyarrhythmias, to avoid inappropriately high ventricular pacing rates. Methods of mode switching differ among pacemakers and ICDs and can be classified into five different categories85,86 (Fig. 3-44). Accurate atrial sensing is critical to appropriate mode switching, because all methods depend on measurement of atrial rate, A-V patterns, or both. The atrial sensing configuration (unipolar vs. bipolar), programmed atrial sensitivity, and atrial blanking periods influence the methods used by the various manufacturers and their performance for detection of atrial tachyarrhythmias. Algorithms to recognize repetitive blanking of atrial events during atrial flutter allow mode switching to occur more rapidly when this condition is confirmed (Fig. 3-45). Whatever method is employed, unreliable atrial sensing, whether from lowamplitude P waves or insensitive programming to avoid FFRW oversensing, will degrade performance of these algorithms.85-88 ICDs with long PVAB periods or low atrial sensitivity may fail to switch modes if a substantial fraction of atrial events are undersensed. Algorithms to Reduce Ventricular Pacing Unnecessary ventricular pacing contributes to adverse clinical outcomes (e.g., heart failure hospitalizations) and increased risk of AF.89,90 Algorithms that lengthen AV intervals adaptively to promote intrinsic conduction reduce ventricular pacing.91-94 Medtronic and Sorin



3  Sensing and Detection

79

Vtip−Vring

RVcoil−can

↑ F S

Marker

↑ ↑ ↑ V F F S 1 S1 S 7 6 0 0

8 2 0

↑ V S

5 7 0

↑ ↑ ↑ ↑ T F F F 3 S 1 S 1 S 1D 4 6 2 3 0 0 0 0

↑ ↑ V F S1S 2 0

7 3 0

↑ ↑ V F S 2S 0 0

↑ V S

7 7 0

5 1 0

↑ V S

↑ ↑ ↑ V F F S 1S 1S 3 3 0 0

6 8 0

Detected NID 18/24

↑ V S

7 5 0

↑ V S

7 6 0

4 4 0

↑ V S

↑ V S

7 6 0

↑ ↑ ↑ ↑ T T F F 3 S 3 S 2S 1S 0 0 3 2 0 0 0 0

↑ V S

4 1 0

3 3 0

↑ ↑ F V S 2 S 6 0

↑ ↑ ↑ T T F S 2 S 1S 4 3 0 0

7 1 0

↑ V S 1 8 0

5 6 0

↑ V S

6 2 0

Not detected NID 30/40

Figure 3-40  Stored electrogram shows intermittent oversensing caused by lead fracture. Pace-sense (Vtip-Vring), high-voltage (RVcoil-can), and marker channels are shown; VS, TS, and FS, intervals in sinus, ventricular tachycardia (VT), and ventricular fibrillation (VF) zones, respectively; FD, detection of VF. Upper and lower panels are continuous. Pace-sense channel shows intermittent, high-frequency, oversensing of nonphysiologic noise characteristic of lead fracture or header-connector problem. At the programmed number of intervals to detect VF (NID) 18/24, inappropriate detection of VF occurs near left of the bottom panel, resulting in an inappropriate shock (not shown). Coincident with detection of VF, oversensing decreases greatly. VF would not have been detected at higher values of NID. FS, Ventricular interval in VF zone; VS, ventricular sensed event; VP, ventricular paced event. (From Swerdlow CD, Gunderson BD, Ousdigian KT, et al: Downloadable algorithm to reduce inappropriate shocks caused by fractures of implantable cardioverter-defibrillator leads. Circulation 118:2122-2129, 2008.)

developed new pacing modes (MVP and SafeR) that operate in AAI(R) mode, with mode switching to DDD(R) to prevent excessive ventricular pauses or asystole.19,95,96 Single beats of AV conduction failure are allowed by both algorithms, with switching to DDD mode occurring only if higher levels of AV conduction block occurs (Medtronic) or if the PR interval is excessive. Once switched to DDD mode, both algorithms test for AV conduction periodically, to avoid unnecessary ventricular pacing for transient AV block. However, these pacing modes may be associated with pacemaker syndrome in patients with marked first-degree AV block, resulting in unnecessary pauses if complete heart block develops.97-99 Automatic Assessment of Pacemaker Capture (“Capture Management”) Automatic sensing of the evoked response (local myocardial depolarization) that results from a pacing stimulus allows for verification of myocardial capture. However, sensing of the evoked response is obscured by the decaying polarization afterpotential that follows the pacing stimulus.100 This polarization artifact can be attenuated by using high-capacitance, low-polarization leads or using multiphasic pacing impulses to neutralize the postpacing polarization. Figure 3-46 shows an example of the evoked response signal combined with polarization artifact after a pacing stimulus.100 To minimize the afterpotential from ventricular pacing stimuli and improve detection of the evoked response, some manufacturers use a smaller-output capacitor. This approach further enhances the accuracy of evoked response detection. Pacemakers use one of two different ventricular capture verification schemes: beat-to-beat approaches that monitor capture almost continuously and periodic approaches that test pacing thresholds intermittently (e.g., once every few hours). Algorithms that perform periodic capture threshold measurements typically set the pacing output at 1.5 to 2.0 times the threshold to avoid transient loss of capture until the next test. Beat-to-beat capture management algorithms adjust pacing

outputs to values that are just above pacing threshold but deliver higher-output backup pulses if loss of capture occurs.101 Ventricular capture verification has been incorporated into some ICDs and resynchronization ICDs. Algorithms for RV capture management use essentially the same methods as pacemakers, except that polarization artifact that challenges evoked response detection in bipolar pacing leads is eliminated or reduced by use of the RV coil-tocan EGM available in ICD systems.102 LV capture management may be achieved by pacing on the LV lead and analyzing the intraventricular conduction time sensed in the right ventricle. This method has been demonstrated to be safe, reliable, and clinically equivalent to operatorperformed LV threshold tests.103 Atrial capture management is more challenging than ventricular management because the atrial evoked response is much smaller than the ventricular response. Thus, reliable detection of the evoked response is difficult, and atrial capture detection using the evoked response method can be performed only using low-polarization electrodes.104 An alternative method relies on critical timing of the atrial and ventricular response to atrial pacing stimuli during normal sinus rhythm (Fig. 3-47). In ambulatory patients, this method has been reported to perform comparably with operator-performed atrial pacing threshold measurements without causing atrial proarrhythmia.105

Basics of Detection of Ventricular Tachycardia/Fibrillation Figure 3-48 provides an overview of the structure of VT/VF detection algorithms. Initial detection of tachycardia is based on rate and duration. The minimum duration of tachycardia required for detection is programmable, either directly (in seconds) or indirectly by setting the number of ventricular intervals required. A tachycardia episode begins when the minimum rate and duration criteria are satisfied. Computationally intensive features of detection algorithms, such

80

SECTION 1  Basic Principles of Device Therapy

Ohms >3000 2000 1500 1000 800 600 400 300 <200

1000 ohms

12/01/ 05

02/09/ 06

04/20/ 06

A

05/29/ 06

09/07/ 06

11/16/ 06

01/25/ 07

Last 80 weeks (min/max per week)

04/03/ 07

04/16/ 07

12 days

Ohms >3000 2000 1500 1000 800 600 400 300 <200

1000 ohms

02/05 06

B

04/17/ 06

06/26 06

09/04/ 06

11/13/ 06

Last 80 weeks (min/max per week)

01/18/ 01/31/ 07 07 Last 14 days

1st shock

LIA Oversensing Met

22:37:53 May 23

18:36:31 May 17

1st shock

LIA Impedance Met

LIA Oversensing Met

2nd shock

01:15:35 Aug 29

03:00:00

03:02:40

08:57:42

6 days Non-sustained tachycardia log

Non-sustained tachycardia log ID#

Date/time

V. cycle

Duration

ID#

Date/time

V. cycle

Duration

5 4

May 17 18:36:31 May 17 10:55:49

200 ms 190 ms

5 beats 5 beats

407 404

Aug 29 03:02:40 Aug 29 01:15:32

120 ms 200 ms

5 beats 5 beats

Ohms >2500 2000 1500 1000 800 600 400 300 <200

Impedance 1000 ohms

03/22/ 06

C

05/31/ 06

08/09/ 06

10/18/ 06

12/27/ 06

03/07/ 07

Last 80 weeks (min/max per week)

05/10/ 05/23/ 07 07 Last 14 days

Ohms

Impedance

At implant 496 ohms last >3000 ohms

1000 ohms

08/31/ 06

D

11/09/ 06

01/18/ 07

03/29/ 07

>3000 2000 1500 1000 800 600 400 300 <200

05/07/ 07

08/16/ 08/29/ 07 07 Last 14 days Last 80 weeks (min/max per week)

Figure 3-41  Individual patient examples of source data used in retrospective simulation of Lead Integrity Alert (LIA, Medtronic) designed to provide warning of lead failure triggered by either high pace-sense impedance or transient rapid oversensing. The goal of this feature is to provide patients and physicians with sufficient warning that patients will receive medical attention prior to delivery of inappropriate shocks.  A, True-positive impedance trigger: abrupt variations in impedance exceeding 1000 Ω, triggering a LIA alert 12 days before an inappropriate shock. B, False-positive impedance trigger shows gradual increase in impedance. New version of algorithm requires abrupt rise in impedance. C, Truepositive oversensing alert triggered 6 days before an inappropriate shock, despite constant impedance. D, LIA was not triggered prior to first inappropriate shock. However, because both impedance and oversensing triggers occurred 1 hour and 45 minutes after that shock, the automatic NID increase would have prevented the subsequent shocks 5 hours later. (From Swerdlow CD, Gunderson BD, Ousdigian KT, et al: Downloadable algorithm to reduce inappropriate shocks caused by fractures of implantable cardioverter-defibrillator leads. Circulation 118:2122-2129, 2008.)



3  Sensing and Detection

81

Atrial

RV Tip-Coil

Shock

AP-Sr -VP-Sr --

AP-Sr 985 VP-Sr 985

VP-Sr 985

AP-Sr 978 VP-Sr 978

AP-Sr 978 VP-Sr 978

AS 870 PVC 760

VP 870

(AS) 1153 VT 325

PVP

(AS) 1025 PVC 723

PVP

VF 160

VF 268

PVP

(AS) 1015 VF VF 153 198 VF VF 183 158 PVP

PVP

VS 550

VF 225 VF 150 Epsd PVP Suddn

Figure 3-42  Oversensing of diaphragmatic myopotentials. Atrial EGM, integrated right ventricular (RV Tip-Coil) bipolar sensing EGM, highvoltage (Shock) EGM, and marker channels from a Boston Scientific Prizm II ICD in a patient with complete heart block are shown. Low-amplitude, high-frequency signals are recorded on the integrated bipolar channel, inhibiting pacing and resulting in inappropriate detection of ventricular fibrillation (VF) at right (labeled Epsd, for episode start/end). Oversensing of diaphragmatic myopotentials usually occurs after long diastolic intervals or after ventricular paced events when amplifier sensitivity or gain is maximal. Repositioning of the lead eliminated the problem. AS, Atrial sensed event; AP, atrial paced event; PVC, premature ventricular complex; PVP→, extension of postventricular atrial refractory period (PVARP); Sr, sensed refractory event; VP, ventricular paced event.

VP

A

B

AB NS

VP

VP

VP

AB NS NS NS NS NS NS NS AB NS NS NS NS NS NS NS AB NS NS

90 bpm

70 bpm

Figure 3-43  A, Extension of refractory period resulting from detection of noise during the noise-sampling (NS) portion of the ventricular refractory period; VP, ventricular paced event. The first portion of the ventricular refractory period is absolute (AB). When suprathreshold signals are seen during the NS portion of the ventricular refractory period, a new NS window is initiated, “retriggering” the ventricular refractory period by a duration equal to the NS window. Continuous extension of the ventricular refractory period invokes noise reversion pacing at the programmed reversion rate. A, Normal VVI pacing at 60 bpm on the left. When electromagnetic interference begins, the ventricular refractory period is continuously retriggered, and noise reversion pacing starts at 90 bpm. B, Combination of ventricular inhibition and noise reversion. The noise reversion rate of this VVI pacemaker is 90 bpm, and the base rate is programmed at 70 bpm. Beats with arrows indicate noise reversion pacing caused by continuous retriggering of the ventricular refractory period, as in A. Some beats are slower than 70 bpm because of intermittent (not continuous) oversensing of noise such that ventricular refractory periods are not continuously retriggered. (B modified from Sweesy MW, Holland JL, Smith KW: Electromagnetic interference in cardiac rhythm management ICDs. AACN Clin Issues 15:391-403, 2004.)

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as morphology analysis, are activated only after initial detection occurs, to conserve power. Unbiased MAR

VENTRICULAR RATE AND COUNTING METHODS

Specificity

MAR biased to faster rates

The specific methods used to count ventricular intervals vary and influence the sensitivity and specificity of VT detection. Initial detection in the VF zone must be tolerant of undersensing. Detection usually occurs when a certain percentage (typically 70%-80%) of intervals in a sliding window (usually 10-40 intervals) fulfill the programmed rate criterion. Most ICD manufacturers also use this “X out of Y” counting in the slower (VT) zones, although Medtronic requires that all intervals exceed the rate criterion to detect VT. This consecutive-interval method diminishes inappropriate detection of AF without compromising sensitivity for detection of regular, monomorphic VT.106-108 St. Jude ICDs classify intervals based both on the last interval and on the average of the last four intervals (Figs. 3-49 and 3-50). A potential advantage of this method is that averaging minimizes the effects of intermittent undersensing by providing a “smoothing effect.”106 A disadvantage is that averaging exaggerates the effect of extremely short, oversensed intervals or extremely long, undersensed intervals. To satisfy tachycardia criteria, “binned” intervals do not need to be consecutive. As long as sinus rhythm is not redetected, the counter will continue to increment. This may be particularly problematic when the slowest VT zone is used as a monitoring zone: During sinus tachycardia in the monitoring zone, PVCs increment the VT counter, but the counter is not reset until the sinus rate decreases out of the monitoring zone, providing the potential for inappropriate therapy of sinus tachycardia109 (Fig. 3-51). Sorin ICDs first classify the rhythm on the basis of the ventricular cycle length. This algorithm relies on the “majority criterion,” which excludes the initial two, short ventricular cycles from analysis to withdraw cycles that may have unstable R-R intervals at the beginning of an arrhythmia.

RAC with high MS rate and cons counts x out of y with high x/y ratio x out of y with low x/y ratio RAC with low MS rate and non-cons counts Combined algorithms Beat-to-beat Sensitivity

Figure 3-44  Sensitivity and specificity of automatic mode-switching (MS) algorithms for atrial tachyarrhythmia detection. Relative performances of the five different classes of algorithms are shown; performance of the mean atrial rate (MAR), rate and count (RAC), and “x out of y” algorithms are shown twice to emphasize performance variability with specific parameter selections. Algorithms that use MAR calculation favor specificity over sensitivity but are generally slower-acting than algorithms that favor sensitivity (e.g., beat-to-beat). Atrial events with short and long atrial cycle lengths equally affect atrial rate estimation in “unbiased” MAR algorithms; long atrial cycle lengths have less impact on atrial rate estimation in “biased” MAR calculations, to reduce the impact of atrial undersensing. Cons, Consecutive; non-cons, nonconsecutive. (From Israel CW: Analysis of mode switching algorithms in dual chamber pacemakers. Pacing Clin Electrophysiol 25:380-393, 2002.)

A S

A S V P

A S V P

A S

1

V P

A S

V P

A S

V P

A S

V P

A S

V P

A S

V P

2

A S V P

A S

V P

3

6

A S V P

V P

A S

V P

7

A S

V P

1 V P

A S

2 V P

A S

3 V P

A S

4 V P

A S

5 V P

A S

6 V P

A S

7 V P

A S

8 V P

A S

A A A R S S

V P

V P

A A R R

A S

V P

Mode switch

A A A S S R

A S

A R

A R

V P DDI A A A S S R

A S

A R

V P

8

V P

A S

A R

A S

V P

V P

A A A A A A R R S R R S

V P

A R

V P

Mode switch

PVARP extension A S

A R V P

A S

A S

V P

5

A S V P

A S

V P

4

A S

V P

A A R R

V P

A A A A A A S R R S R R

V P

V P

A A S R

V P

Figure 3-45  Algorithm to prevent upper rate limit tracking of alternate atrial EGMs during atrial flutter (Blanked Flutter Search), even if EGMs recorded in postventricular atrial blanking period (PVAB), when sensing cannot occur. Upper panel, Simulated electrocardiogram (ECG) and ladder diagram. Black horizontal bars denote blanking periods, and white bars denote refractory periods for atrium (upper bars) and ventricle (lower bars). Lower panel, Stored ventricular EGM and dual-chamber EGM markers from a clinical event. Upper panel, The intention of this algorithm is to infer that alternate atrial EGMs are “hiding” in the PVAB. To determine whether this is occurring, the algorithm transiently interrupts atrial tracking to assess the atrial rhythm when it is likely that alternate atrial intervals are occurring in the PVAB. The specific criterion is that the measured A-A interval is both less than twice the sum of the sensed AV interval and the PVAB and less than the mode switch interval for 8 beats. When this occurs, the algorithm extends the postventricular atrial refractory period (PVARP extension). The next atrial EGM that would have been tracked is now in the PVARP (AR marker, first arrow) and is not tracked. The next atrial flutter EGM is sensed because it no longer times in the PVAB (AS marker, second arrow). The pacemaker then switches mode rapidly (DDI). Lower panel, In this stored clinical EGM, ventricular intervals are numbered to correspond to those in the simulation in the upper panel. As in the upper panel, the first arrow (AR) corresponds to PVARP extension, and the second arrow (AS) corresponds to a sensed atrial flutter EGM that would have been blanked in tracking mode. Mode switching follows immediately. VP, Ventricular paced event.



3  Sensing and Detection

Monitor-only Zones

Amplitude

Pace polarization artifact

Pacing pulse

−200 msec Time

Programmable monitor-only zones store EGMs of tachycardias based on rate and duration even if they are slower than the slowest zone in which VT therapy is programmed. If therapy is not programmed “on” for slow VT, the slowest rate zone may be programmed as a “monitor only” zone, with detection “on” and therapies “off.” However, interactions between the counters in the VT zone serving as a monitor zone and the next zone may restrict use of SVT-VT discriminators or decrease the number of intervals required for detection in the therapy zone. ICDs with dedicated monitor-only zones may avoid these potential interactions. However, as noted previously, in St. Jude ICDs the combination of counting method and criteria for resetting the VT counter during sinus tachycardia in the monitor-only zone may result in inappropriate therapy.109,110

Time

The SVT-VT discriminators are a programmable subset of an ICD’s detection algorithm for VT/VF. Discriminators withhold ventricular antitachycardia pacing (ATP) or shocks if SVT is diagnosed. These algorithms differ from the SVT detection algorithms used to switch modes during bradycardia pacing or deliver atrial therapy for AF or atrial flutter.

Amplitude

A

B

83

Evoked response

Discriminators

Amplitude

Detected signal (evoked response + pace polarization artifact)

Time

C Figure 3-46  Local myocardial depolarization (evoked response) is obscured by the postpacing polarization artifact. A, Pace polarization artifact is observed as a positive deflection that generally decreases toward the initial cell potential within about 200 msec after the pacing pulse. B, Evoked response is present immediately after the pacing pulse and shows an initial negative deflection. C, The signal sensed by the pacemaker is the summation of the pace polarization artifact and  the evoked response. (From de Voogt WG, Vonk BF, Albers BA, et al: Understanding capture detection. Europace 6:561-569, 2004.)

VENTRICULAR TACHYCARDIA/FIBRILLATION RATE DETECTION ZONES Tiered-therapy ICDs have up to three ventricular rate detection zones that permit programming of zone-specific therapies and SVT-VT discriminators (Fig. 3-52). The duration of a rapid ventricular rate required for detection is programmable independently in at least the slowest and fastest zone (see Zones and Zone Boundaries). The two slower zones are classified as VT zones, and the fastest one as a VF zone. All ICDs use strategies to minimize the risk of prolonged detection if successive VT intervals are classified in different zones. Except for Medtronic ICDs, sensed events in one zone increment counters in all slower zones. To avoid delays in detecting rhythms that straddle zone boundaries, Medtronic ICDs utilize a combined-count (summation) feature, which is satisfied if a sufficient fraction of intervals is in either zone.106

CONFIRMATION, REDETECTION, AND EPISODE TERMINATION If a tachycardia is classified as ventricular, ATP may be delivered immediately; but shocks require capacitor charging, which takes 6 to 15 seconds for maximum-energy shocks. The first shock in a shock sequence is “noncommitted,” meaning that it may be aborted during charging if the detection algorithm identifies termination of the arrhythmia during or immediately after charging of the high-energy capacitors (Fig. 3-53). In contrast, a “committed shock” is always delivered if the capacitor charges. Confirmation or reconfirmation is the brief process that occurs after charging by which ICDs determine whether to deliver or abort the first shock in a sequence. In all ICDs, the confirmation algorithm delivers the stored shock if a few intervals immediately after charge completion are shorter than the programmed VT interval (St. Jude, Boston Scientific, Sorin, Biotronik) or are within 60 msec of the VT interval (Medtronic ICDs before Protecta) (Fig. 3-54). Therefore, the first VF shock is effectively “committed” if the VT interval (or, in some ICDs, the monitor-only interval) is programmed to a long cycle length.111,112 A new shock confirmation algorithm in Medtronic ICDs (investigational, Protecta) aborts a defibrillation shock when ventricular cycle lengths increase by at least 60 msec relative to the detected tachycardia cycle length and exceed the VF detection interval, and the ventricular cycle length at the point of detection is regular. Tachycardias detected in the VF zone that do not have regular cycle length are judged against the VF interval plus 60 msec. Shocks for VT will abort on a 60-msec increase in cycle length relative to the point in detection, regardless of cycle length regularity. Once a shock is aborted, the device continues to apply tachycardia redetection and episode termination criteria. Depending on the manufacturer and programming, each shock after the first shock may be either noncommitted or committed. A shock is also committed if VT or VF is detected after a diverted shock and before episode termination.83 Redetection is the process by which ICDs determine whether VT or VF detection criteria remain satisfied after therapy is delivered. The duration for redetection is programmable independently of the duration for initial detection and typically is shorter than that for initial detection. ICD-defined arrhythmia episodes continue after each tiered therapy until either a tachyarrhythmia is redetected to initiate the next therapy or the rhythm is classified as normal (“sinus”), resulting in episode termination. In most ICDs, episode termination is based on (slow) rate and duration, with 3 to 8 beats classified as sinus. The

84

SECTION 1  Basic Principles of Device Therapy

APt AR

AS Atrial test pace not captured

Atrial backup pace VP

AP

VP

APt

AS Atrial test pace captured, no AV conduction

Capture verification windows VS

VS

VS

VP

VP AS

Atrial backup pace

APt

AP

Atrial test pace captured, AV conduction

A

Atrial overdrive rate

Atrial test pace captured

Atrial overdrive rate AP APt 270 ms Atrial overdrive rate

Atrial test pace not captured VP

VS

B

Atrial overdrive rate

Atrial overdrive rate AP APt 270 ms Atrial overdrive rate

Capture verification windows VS

VS

VS

Figure 3-47  Atrial capture verification method that relies on atrial or ventricular responses to atrial pacing. A high-level control algorithm determines which of two timing assessment methods should be used, based on the recent history of atrial pacing, atrial sensing, and atrioventricular (AV) conduction. At least two test paces are required to verify capture or loss of capture. Failed, inconclusive, or aborted threshold tests are repeated after a delay. This method cannot measure atrial capture during atrial paced–ventricular paced (VP) rhythms or rhythms with irregular atrial or ventricular cycle lengths. A, Atrial Chamber Reset method is used if recent history indicates primarily atrial sensed events (AS). A premature atrial test pace (APt) is delivered, and the resulting atrial timing is observed. An atrial refractory event (AR) within 10 bpm after the prior intrinsic atrial rate suggests loss of capture (top panel). The lack of an atrial refractory event in this period suggests resetting of the atrial timing and atrial capture (middle and bottom panels). B, The AV conduction check method is used if the recent history indicates that AV conduction is present. The average AV conduction time is measured before delivery of the APt. A backup atrial pace at the operating atrial output is also delivered 70 msec after the APt. The algorithm looks for ventricular sensed events (VS) in one of two capture verification windows (40 msec in duration) that are placed relative to the APt and the backup atrial paced events based on the average AV conduction time. A ventricular event sensed in the first capture verification window suggests capture by the APt (top panel). Lack of a sensed ventricular event in the first capture verification window and presence of a ventricular event in the second window is evidence of noncapture by the APt (bottom panel).

n

n

SVT-VT Discriminators

Initial detection y (V rate and duration)

Rate y zone VT

SVT-VT Discriminators

y ATP

Rx

Redetection y (V rate and duration)

n

Episode termination

y

VF

n

n

Shock

Confirmation after charging

y

n

Figure 3-48  Overview of ICD detection algorithm. Algorithms apply varying degrees of supraventricular tachycardia–ventricular tachycardia (SVT-VT) discrimination to redetection. ATP, antitachycardia pacing; n, no; Rx, antitachycardia therapy; V, ventricular; VF, ventricular fibrillation; y, yes.



3  Sensing and Detection

85

Current Interval

Figure 3-49  Interval classification based on current interval and average interval (St. Jude). The table shows classification of ventricular intervals by St. Jude ICDs based on both the value of the current interval and the average of the current interval with the preceding three intervals (average interval). Detection zones from slowest to fastest are Sinus, Tach A (slower ventricular tachycardia [VT] zone), Tach B (faster VT zone), and Fib (ventricular fibrillation [VF]). If both the current and the average interval are in the same zone, the current interval is classified in that zone. If they are in different VT or VF zones, the current interval is classified in the faster zone. If one interval is in the sinus zone and the other is in a VT or VF zone, the current interval is not classified (Not Binned).

Fib

Tach B

Tach A

Tach

Sinus

Fib

Fib

Fib

Fib

Not Binned

Tach B

Fib

Tach B

Tach B

N/A

Not Binned

Tach A

Fib

Tach B

Tach A

N/A

Not Binned

Tach

Fib

N/A

N/A

Tach

Not Binned

Sinus

Not Binned

Not Binned

Not Binned

Not Binned

Sinus

Average Interval Fib

number of intervals required for episode termination is programmable in St. Jude ICDs. Programming a short duration is advisable when ATP terminates VT, but it reinitiates promptly after termination. This permits repeating the first sequence of programmed ATP rather than escalating to subsequent therapies that include shocks.

SVT-VT Discrimination in Ventricular ICDs BUILDING BLOCKS FOR DISCRIMINATION The ICDs analyze information derived from a sequence of sensed atrial and ventricular EGMs to detect VT/VF and discriminate ventricular arrhythmias from SVT. These algorithms operate in a stepwise manner, using a series of physiologically relevant, logical “building blocks” (Table 3-1) based on the timing relationships and morphology of sensed EGMs. We focus on the “tool kit” provided by these building blocks rather than on the specific details of proprietary algorithms, reviewed elsewhere in detail.113 Detection algorithms combine complementary building blocks to form the final detection decision. Each building block has advantages and limitations. Some are redundant, and some interact. Interactions vary depending on the order in which blocks are applied. Clinicians and algorithm designers must consider these issues to understand trade-offs in clinical performance. Single-Chamber Ventricular Building Blocks The first five building blocks in Table 3-1 apply to single-chamber ICDs or the ventricular component of dual-chamber ICDs. The R-R interval + duration analysis building block classifies ventricular intervals into zones by programmable cycle length (or rate) thresholds. All singleand dual-chamber ICDs use cycle length or rate in combination with tachyarrhythmia duration as the most basic method of detection. When used alone, R-R interval + duration analysis provides the highest sensitivity for detection of VT/VF, but cannot discriminate VT/VF from SVT. The R-R regularity building block is used to discriminate between regular ventricular intervals caused by monomorphic VT and irregular ventricular intervals caused by rapidly conducted AF. The R-R onset building block is used to discriminate sudden-onset VT from gradual-onset sinus tachycardia. The VEGM morphology building block classifies tachyarrhythmias as SVT if morphologic characteristics of the ventricular EGM are similar to those of beats of known supraventricular origin. Otherwise, it classifies the rhythm as VT.

Dual-Chamber Building Block: Atrial vs. Ventricular Rate Direct or indirect comparison of atrial and ventricular rates is the cornerstone of most dual-chamber algorithms. In most studies, the ventricular rate exceeds the atrial rate in more than 80% of VTs in the VT zone of dual-chamber ICDs and a higher fraction of VTs in the VF zone.68,114 Thus, algorithms that compare atrial and ventricular rates as their first step (Boston Scientific Rhythm ID, St. Jude, Biotronik) apply SVT discriminators to fewer than 20% of VTs; this reduces the risk that they will misclassify VT as SVT, provided that the atrial rate is measured correctly. The Medtronic algorithm indirectly compares atrial and ventricular rates using dual-chamber pattern analysis as the first step to achieve a comparable result. The Sorin algorithm uses a conceptually different approach, by initially detecting tachyarrhythmia based on cycle length, then further classifying the arrhythmia as VT or SVT based on the stability of R-R intervals and an analysis of AV association and tachycardia onset. Dual-Chamber Building Blocks The next four building blocks in Table 3-1 are based on combinations of atrial and ventricular information. Their primary role is to discriminate VT with 1 : 1 ventricular-atrial (V-A, VA) conduction from SVT. Their secondary roles are to classify rhythms with stable N:1 atrialventricular (A-V, AV) relationships (stable ratio of N atrial events for each ventricular event) as SVT and to classify isorhythmic tachycardias (with AV dissociation and similar ventricular rate) as VT. The P-R dissociation building block can detect the presence of VT during SVT. The P-R patterns/relationships building block applies to stable associations of atrial and ventricular events. For tachycardias with 1 : 1 AV relationships, it analyzes the relative timing of P-R and R-P intervals to classify the rhythm as SVT or VT. It also identifies consistent N:1 AV patterns that occur primarily during atrial flutter. The chamber of origin building block discriminates between VT and SVT with 1 : 1 P-R association by identifying whether the tachycardia originates in the atrium or in the ventricle. In ICD patients, SVT usually begins with an intrinsic atrial event in the interval between the last ventricular event in the sinus rate zone and the first ventricular event in the VT zone. Conversely, at the start of spontaneous VT, there is usually no atrial event in this interval. The response of the opposite chamber to atrial or ventricular pacing has long been a critical tool in the electrophysiology laboratory for diagnosing tachycardias with 1 : 1 AV relationship. A novel ICD algorithm based on this premise has been shown to

86

SECTION 1  Basic Principles of Device Therapy

Atrial B

777

738

773

P

P

V

133

623

B

P

V

133

600

133

601

1

0

719

727

P

V

RV-LV

B

734

742

B

719

719

P

V

141

586

2

B

711

719

P

V

133

578

3

402 V

4

*

DD

S

S 367

T

B

695

680

125 515

R

R

P

R

141

F

344

371

V

R

676

F

R

645

R

F 648

125 520

R

R

R

157

340

348

V

R

652

R

R

9

R

R

8

F

125 527

R

652

547

7

F

125 523

R

8

P

F

6

F

25 527

664

128 523

5

4

T

B

656

R

F 652

125 531 R

F

R

R

10

652

125 527 R

125 R

11

12

Trigger F 660

527 R

F 664

125 535 R

676

125 539

R

R

12

F

723

570

P

R

187

R

R

** S

676

125 547

R

R

734

R

187

R

P

R

187

R

17

R

P

R

187

16

18

S 762

125 605 R

125 R

S 738

125 605 R

695

125 547

R

S 730

125 609

R

15

S

P

16

R

14

125 598 R

672

125 551

13

R

R

D

P

R

179

125 637 R

R

19

20

DDD

S

P 187

20

R

B

801

125

125 617

R

R

21

R

797 P 133

B 797

777 P

V

133

664

22

B 773

816 P

V

133

664

23

B 812 V 683

24

Figure 3-50  Illustration of interval classification based on current interval and average interval; T-wave oversensing initiates R-wave doublecounting in St. Jude ICD. The figure shows continuous recording of atrial EGM, event markers, and extended bipolar ventricular EGM (between Y-adapted left ventricular [LV] and right ventricular [RV] tip electrodes). The ICD is programmed with a single ventricular tachycardia (Tach) zone at 345 msec and a ventricular fibrillation (Fib) zone at 290 msec. The top panel shows atrial sensed, ventricular paced rhythm. In the second panel, the first ventricular EGM is an interpolated premature ventricular complex (PVC). Ventricular pacing is inhibited after the next P wave because the resultant PVC-V paced interval would have been less than the interval corresponding to the programmed upper rate limit (~450 msec). This results in a conducted R wave that is double-counted at the ventricular blanking period of 125 msec, but this interval is not classified in the VF zone because the interval average of (719 + 402 + 367 + 125)/4 = 403 msec exceeds the Tach detection interval. The third and fifth beats show postpacing T-wave oversensing (down arrows). The interval ending with the first oversensed T wave is classified in the Tach zone because the interval average is less than 345 msec: (367 + 125 + 516 + 344)/4 = 338 msec. The next ventricular interval, corresponding to the subsequent double-counted R wave, is classified in the Fib zone because the interval average is in the Tach zone ([516 + 344 + 371 + 125]/4 = 339 msec) but the current interval is in the Fib zone. Mode switching to the DDI pacing mode occurs after the counter of Tach + Fib beats equal 4 (DDI, asterisk). The third and fourth panels show conducted sinus rhythm after mode switching, with alternate intervals classified in the Fib zone. VF is detected when the Fib counter reaches 12. This counter increments for each interval classified in the Fib zone and is reset to zero by five consecutive sinus intervals. The count of Fib intervals reaches 12 after the third interval in the fourth panel, and Fib is detected (Trigger, D marker), resulting in capacitor charging. However, sinus slowing results in an increase in the average interval, so that the interval average is no longer in the Tach zone. This occurs at the second R wave in the fifth panel (S marker, double asterisk; interval average = [125 + 570 + 125 + 596]/4 = 354 msec). Once this occurs, alternate detected intervals are classified as “Sinus” or “Not Binned.” The shock is aborted when the count of consecutive binned Sinus intervals reaches 5 at the beginning of the bottom panel (RS = return to sinus, up arrow), and this is followed by mode switch to DDD pacing. Note that St. Jude ICDs have a “Bigeminal Avoidance” counter that prevents bigeminal rhythms from being classified as Tach. This counter does not apply in the Fib zone. R, Ventricular sensed event; V, ventricular paced event.

VT (monitor) T

A

347

T 347

T 345

T 347

17s

VS VS

B

383

347

VS 367

VS 365

18s

F

295

T

* 442

344

65s

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343

VS

363

263

* 343

66s

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VS

363

19s

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271

VS

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343

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343

20s

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363

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367

67s

253

VS

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340

359

21s

F 434

*

68s

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343

VS

343

C

344

VS

VS

357

106s

344

*

VS 263

VS

240

104s

345

VS 347

VS 342

107s

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F

*

367

441

363

108s

329

VS 343

100s

343

24s

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359

367

359

387

363

301

VS 367

359

*

70s

*

359

VS

VS

69s

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350

VS

23s

347

VS F 289

110s

426

75s

VF *

VS

22s

VS

358

VS

453

72s

*

*

*

*

*

F

F

*

343

375

111s

347

371

112s

Figure 3-51  Inappropriate detection of ventricular fibrillation (VF) caused by isolated premature ventricular complexes (PVCs) during sinus tachycardia in a monitoring zone. A, Sinus tachycardia gradually accelerated, triggering an episode recorded as ventricular tachycardia (VT) in the monitoring zone of a St. Jude ICD. B, Occasional PVCs are binned in the VF zone (F), adding to a cumulative counter. C, On the 12th such premature ventricular beat, VF is detected, leading to charging and an inappropriate shock (not shown) for sinus tachycardia. (From Mansour F, Khairy P: ICD monitoring zones: intricacies, pitfalls, and programming tips. J Cardiovasc Electrophysiol 19:568-574, 2008.)

88

SECTION 1  Basic Principles of Device Therapy

Tiered therapy: variable # ATP attempts followed by # shocks of increasing energy Bradycardia upper rate limit

Shock only; ATP not applied

Sinus

VT

FVT

VF

SVT discrimination zone

τToo fast for ATP (~240 ms) τExpect VF undersensing

Fastest SVT rejected

Different ATP or fewer trials (~320 ms)

Figure 3-52  Rate detection zones for implantable cardioverter-defibrillators. Some ICDs permit programming of an additional monitor-only zone; ATP, antitachycardia pacing. FVT, Fast VT; SVT, supraventricular tachycardia; VF, ventricular fibrillation; VT, ventricular tachycardia.

τ ~400 ms or 40 ms >slowest VT τ Consider SVT specificity

terminate or discriminate 1 : 1 tachycardias by using simultaneous atrial and ventricular ATP.115 Single-Chamber Atrial Building Blocks The final four detection building blocks are the atrial correlates of the single-chamber ventricular building blocks: P-P interval (atrial cycle length), P-P regularity, P-P onset, and AEGM morphology. Their primary role is to identify the presence (or absence) of atrial tachyarrhythmias and to invoke (or inhibit) the application of single chamber ventricular building blocks such as interval stability and/or EGM morphology. To date, atrial EGM morphology has not been applied in ICDs to discriminate between antegrade and retrograde conduction. SVT-VT DISCRIMINATION IN SINGLE-CHAMBER ICDS R-R Interval Regularity Measures of R-R interval regularity usually are referred to as measures of R-R interval “stability” or “stability algorithms.” Technical details and optimal programmed values differ among manufacturers. Specific measures of R-R regularity interact with the method of counting R-R

intervals and the duration required for detection. Requiring consecutive R-R intervals or a higher number of intervals to fulfill the rate and regularity criteria improves rejection of AF but increases the risk that detection of VT will be delayed.107,108,116-120 Regularity criteria can reject AF with ventricular rates slower than 170 beats per minute (bpm) in the absence of antiarrhythmic drugs. At faster rates, these criteria cannot discriminate AF from VT reliably, because R-R intervals in AF are more regular;116-120 and the criteria may prevent detection of VT in the presence of amiodarone or type IC antiarrhythmic drugs.117,119 These drugs may cause regular VT to become irregular, or they may cause rapid polymorphic VT to slow from the VF zone, where discriminators do not apply, to the VT zone, where regularity discriminators do apply. R-R regularity may also be used to select the first VT therapy: ATP for regular VT and shock for irregular and presumed VF or polymorphic VT. Sudden Onset Measures of abruptness of onset of a tachycardia have high specificity for rejecting sinus tachycardia,116,118-120 but may prevent detection of VT that originates during SVT or VT that starts abruptly with an initial

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Figure 3-53  Confirmation (“reconfirmation”) and aborted shock. Ventricular tachycardia (VT) in the ventricular fibrillation (VF) detection zone is detected and initiates capacitor charging (Chrg) before it terminates spontaneously. The noncommitted shock is diverted (Dvrt Chrg). AS, Atrial sensed event; VP, ventricular paced event.



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3  Sensing and Detection Coil-Can

RV tip-ring

2 4 0

F D

V V V V V V V V S S S S S S S S 2 2 2 2 2 2 2 2 1 1 1 2 2 2 3 3 0 0 0 0 0 0 0 0 VF VF Rx 1 Defib

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Figure 3-54  Confirmation failure. Two ventricular EGMs and event markers are shown from a single-chamber ICD. The programmed ventricular tachycardia (VT) and ventricular fibrillation (VF) detection intervals are 380 and 320 msec, respectively. Detection of rapid VT as VF results in capacitor charging (VF Rx 1 Defib). After the charging cycle ends (CE), Medtronic ICDs deliver the shock (CD) if the ventricular cycle length after charging is less than the programmed VT interval + 60 msec. In St. Jude and Boston Scientific ICDs, shocks for VF are confirmed if the ventricular cycle length is less than the VT interval. RV, Right ventricular. (From Swerdlow CD, Friedman PA: Advanced ICD troubleshooting. Part II. Pacing Clin Electrophysiol 29:70-96, 2006.)

TABLE

3-1 

Tachyarrhythmia Detection Building Blocks

Tachyarrhythmia Detection Building Blocks Purpose/Information Single-Chamber Ventricular Building Blocks Identifies sustained high ventricular rates R-R interval + duration R-R regularity Discrimination of monomorphic VT (regular cycle lengths) from rapid AF (irregular cycle lengths) R-R onset

Identifies sudden ventricular rate changes

VEGM morphology

Abnormal VEGM morphology may indicate ventricular tachyarrhythmias Intervals after entrainment of VT by burst pacing are less variable than intervals after burst pacing during SVT

Burst ventricular pacing

Key Dual-Chamber Building Block Comparison of atrial vs. VT diagnosed if atrial rate is less than ventricular ventricular rate rate Dual-Chamber Building Blocks P-R dissociation P-R dissociation usually indicates VT P-R patterns/relationships

Consistent P-R patterns/relationships usually indicate SVT Chamber of acceleration Identifies whether tachycardia initiates in atrium or ventricle Atrial or ventricular pacing, Discrimination of 1 : 1 rhythms using ventricular response in opposite chamber response to atrial extrastimuli Single-Chamber Atrial Building Blocks P-P intervals Identifies high atrial rates P-P regularity Regular atrial rate indicates organized atrial activity P-P onset Identifies sudden atrial rate changes AEGM morphology

Identifies atrial tachyarrhythmias and/or retrograde conduction

Potential Weaknesses SVT with high ventricular rates that overlap with VT/VF rates May lose effectiveness as ventricular rates during AF increase; 2 : 1 atrial flutter has regular R-R intervals; may cause underdetection of VT with irregular R-R intervals Not specific for atrial or ventricular tachyarrhythmias; may miss VT arising during sinus tachycardia Confounded by conduction aberrancy or changes in “normal” VEGM morphology Sensitive to single interval measurement, potential detection time delay, and potential proarrhythmia

Confounded by atrial undersensing or far-field R-wave oversensing

AV reentrant tachycardia; VT with 1 : 1 retrograde conduction; AF that conducts rapidly with apparent P-R dissociation AV reentrant tachycardia and VT with 1 : 1 retrograde conduction A single oversensed/undersensed event may result in misclassification. Primarily aids diagnosis for 1 : 1 rhythms; concerns for VT detection delay and proarrhythmia High atrial rates may be present during true VT/VF. Minimal benefit for ventricular tachyarrhythmia characterization Not specific for atrial or ventricular tachyarrhythmias (e.g., VT with 1 : 1 retrograde association) Confounded by far-field R waves and changes in “normal” AEGM morphology

AEGM, Atrial electrogram; AF, atrial fibrillation; AV, atrioventricular; SVT, supraventricular tachycardia; VEGM, ventricular electrogram VF, ventricular fibrillation; VT, ventricular tachycardia.

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SECTION 1  Basic Principles of Device Therapy

rate below the VT detection limit. In the latter case, the ICD misclassifies the “onset” of the arrhythmia as the gradual acceleration of the VT rate across the VT rate boundary. This criterion does not prevent SVT with sudden onset from being treated inappropriately. Ventricular EGM Morphology Analysis of ventricular EGM morphology,114,120-123 alone or in combination with stability, probably provides the best single-chamber SVT-VT discrimination for initial detection of VT. The morphology building block is the key element of single-chamber algorithms and the central single-chamber component of dual-chamber algorithms with this feature. For this reason, a more detailed analysis is provided. VENTRICULAR EGM MORPHOLOGY FOR SVT-VT DISCRIMINATION General Electrogram morphology algorithms are the most complex and effective building blocks in single-chamber detection algorithms. All morphology algorithms share common steps (Fig. 3-55): (1) record a template EGM of baseline rhythm; (2) construct and store a quantitative representation of this template; (3) record EGMs from an unknown tachycardia; (4) time-align the template and tachycardia EGMs; (5) construct a quantitative, normalized representation of each tachycardia EGM; (6) compare the representation of each tachycardia EGM with that of the template to determine its degree of morphologic similarity; (7) classify each tachycardia EGM as a morphology match or nonmatch with the template; and (8) classify the tachycardia rhythm as VT or SVT based on the fraction of EGMs that match the template. Steps 3 through 8 are performed in real time. Morphology algorithms differ according to EGM source or sources, methods of filtering and alignment, and details of quantitative representations. The features of specific algorithms are described in Figure 3-56. Limitations of Morphology Algorithms The START study (SAFARI trial) of arrhythmias recorded at ICD implant reported individual differences for the performance of singlechamber discrimination algorithms based primarily on EGM morphology.124 Morphology algorithms share common failure modes for inappropriate detection of SVT as VT: (1) inaccurate template, (2) EGM truncation, (3) alignment errors, (4) oversensing of pectoral

Extracted features stored in ICD memory

Template EGM (baseline rhythm)

(1)

(4)

(2)

Compare features

Time align EGMs

Match % = 1 – Area of Difference Area of Template

Tachycardia Template

(2)

Extracted features calculated in real time

(4)

(3)

Tachycardia EGM Figure 3-55  General steps in all morphology algorithms. See text for details.

myopotentials, (5) rate-related aberrancy, (6) SVT soon after shocks, and (7) inappropriate classification of VT as SVT. Inaccurate Template.  The template may be inaccurate because the baseline EGM changed (e.g., postimplantation lead maturation, intermittent bundle branch block) or because the template was recorded from an abnormal rhythm (e.g., idioventricular or bigeminal PVCs). Accurate SVT-VT discrimination requires periodic automatic or manual template updates. Most modern ICDs include an automatic template updating feature, except for cardiac resynchronization ICDs, which are intended to pace the ventricles continuously. If automatic updates are not available, the morphology algorithm should be updated when the EGM becomes chronic. However, the template cannot be updated without intrinsic AV conduction. If software permits, the template match should be verified initially and during follow-up. Electrogram Truncation.  EGM truncation (“clipping”) occurs when the recorded EGM signal amplitude exceeds the range of the EGM amplifier so that the maximum or minimum portion of the EGM is clipped. This removes EGM features for analysis and alters the timing of the tallest peak, which can affect alignment. The amplitude scale in Medtronic and St. Jude ICDs should be adjusted so that the EGM used for morphology analysis is 25% to 75% of the dynamic range (Fig. 3-57). Alignment Errors.  Alignment errors prevent match between a tachycardia EGM and a morphologically similar stored EGM. Mechanisms depend on the method used for EGM alignment (Fig. 3-58). Accurate alignment in the St. Jude algorithm is sensitive to the value of the sensing threshold at the onset of the ventricular EGM, as determined by Automatic Sensitivity Control. If a template EGM is acquired at the most sensitive setting of Automatic Sensitivity Control (either because of a slow sinus rate or after a ventricular paced beat), a low-amplitude peak at the onset of the ventricular EGM may be used for alignment. During tachycardia, the next R wave may occur before Automatic Sensitivity Control reaches its sensitive value. In this case, the lowamplitude peak at the onset of the ventricular EGM may not be used for alignment. If identical template and tachycardia EGMs are then compared, their representations in the morphology algorithm will not match. Usually, they are assigned morphology match scores of 0% (see Fig. 3-58). In patients who have dual-chamber ICDs and intact AV conduction, the most reliable template is acquired or verified during atrial pacing at a rate close to the VT rate. In single-chamber ICDs, options may include acquiring or verifying the template during exercise testing and altering the minimum sensitivity, Threshold Start, or Threshold Delay. If a template is acquired during atrial pacing or exercise testing, automatic template updating should be disabled to prevent overwriting the template with one acquired in baseline rhythm. Medtronic ICDs align EGMs based on their tallest (positive or negative) peaks. If an EGM has two peaks of near-equal amplitude, or if such peaks are caused artificially by truncation of large EGMs that exceed the programmed dynamic range, minor variations in their relative amplitudes may result in an alignment error.120 An alternative EGM source should be selected. The Boston Scientific morphology algorithm (Rhythm ID) simultaneously collects data from two different channels: the local bipolar rate-sense EGM and shock EGM.125,126 It considers the vector of depolarization wavefront (with shock EGM as the reference), the timing of the peak signal recorded on the rate-sense EGM, and alignment of eight characteristic points of the shock EGM. A correlation equation aligns peaks on the rate-sense EGM for each QRS complex of the unknown rhythm with the rate-sense EGM peak of the sinus template, then compares the eight captured points with the template. Oversensing of Pectoral Myopotentials.  In Medtronic ICDs, oversensing of pectoral myopotentials may prevent an SVT from matching the sinus template if the RV coil-to-can EGM is used for morphology analysis and the R-wave amplitude is small. The effect of myopotentials



3  Sensing and Detection

91

100 Shock

50 0 Template

50 B

A

C

Tachycardia EGM

A

–

B –

= A’

C’ B’

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B’ =

– A’

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= C’ 1

% Match = 

10



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60 40 20 0 −20 −40

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Wavelet Wavelet Amplitude Shape

20

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 ci*i[n]

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Wavelet transform

mV

Digitized electrogram (x[n]) 0

200 msec Smallest amplitude wavelets

C

− 0 +

0

200 msec 0

200 msec

Figure 3-56  Specific morphology algorithms. A, St. Jude MD algorithm. The positive and negative deflections in the sensing EGM are normalized and modeled as a series of three polygons (A, B, and C in the template EGM; A’, B’, and C’ in the tachycardia EGM). The normalized areas of these polygons are then computed. Each tachycardia EGM is compared with the template EGM in three steps. First, the difference in area of corresponding polygons is computed. Second, the absolute values of these differences are summed. Third, a match score is constructed to be inversely proportional to the sum of these differences. If a programmable number (= 4) of 8 complexes in a sliding window exceed the programmable threshold (nominally 60%), the rhythm is classified as supraventricular tachycardia (SVT). If not, it is classified as ventricular tachycardia (VT). B, Boston Scientific Vector Timing and Correlation (VTC) algorithm. This algorithm aligns shocking EGMs of the tachycardia and template based on the peak of the rate-sensing EGM. This method takes advantage of spatiotemporal differences between activation sequences in VT and baseline rhythm that cannot be detected from a single EGM. In this example, the peak of the shock EGM has similar timing to the peak of the rate-sensing EGM in sinus rhythm, but much later timing in VT. The amplitude of the shocking EGM (upper panel) is computed at each of 8 points selected on the basis of their timing relative to the peak of the sensing EGM (lower panel) and extracted as an eight-element “feature set,” corresponding to an eightdimensional vector. Feature sets of each tachycardia EGM are compared with the template EGM by computing the product-moment correlation coefficient between the two feature sets. This value, named the feature correlation coefficient (FCC), is a measure of mean spatiotemporal difference in ventricular activation between tachycardia and baseline rhythm. A threshold value for the feature correlation coefficient was selected to optimize SVT-VT discrimination on a test data set. This value is neither published nor programmable. If at least 3 of 10 complexes in a sliding window exceed the threshold, the rhythm is classified as SVT; if not, it is classified as VT. C, Medtronic Wavelet algorithm. The algorithm expresses the morphology of ventricular EGMs using the wavelet-transform. Wavelets are functions of constant shape and limited time duration.* They form the basis of a mathematical transformation that represents signals efficiently if they are both highly localized in time and preceded and followed by isoelectric intervals. Wavelets are therefore well suited for representing transient biomedical signals such as ventricular EGMs. The algorithm compares the morphology of ventricular EGMs during a tachycardia with a template recorded during baseline rhythm. This comparison is expressed as a percent-match score that describes the degree of morphologic similarity of the baseline and tachycardia EGMs using a programmable source. In general, the EGM recorded between right ventricular and superior vena cava coils is preferred as a default. This signal combines far-field sensitivity to morphology differences between VT and SVT with resistance to pectoral myopotentials. The algorithm begins processing when a tachycardia fulfills the programmed rate criterion for detection of VT and 8 beats remain to fulfill the programmed duration criterion for detection of VT. Percentmatch scores are calculated for each of the last 8 beats before detection. Beats with match scores less than a programmable threshold (nominally 70%) are classified as ventricular. Nominally, a tachycardia is classified as VT if 6 or more of the 8 analyzed EGMs are classified as ventricular. Otherwise, it is classified initially as SVT. If the tachycardia is classified as SVT, the algorithm is applied to each successive 8-beat sliding window until the rate criterion for VT is no longer fulfilled. (*Meyer Y: Wavelets: algorithms and applications. Philadelphia, 1993, Society for Industrial and Applied Mathematics.)

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SECTION 1  Basic Principles of Device Therapy

Range 9.8 mV

Range 14.4 mV

A

T S

5 9 0

V S

6 5 0

V S

4 5 0

No match 64%

V S

5 2 0

V S

3 4 0

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3 5 0

T S

T T 3 S 3 S 2 1 0 0

No match 64%

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T 3 S 3 0

3 2 0

No match 61%

T S 3 3 0

T S

3 4 0

T S

T 3 S 3 4 4 0 0

T S 3 3 0

T S

3 7 0

T S

Figure 3-57  Electrogram truncation as source of error for morphology algorithms. A, St. Jude MD template EGM is truncated (arrow) with an amplifier range of 9.8 mV. Truncation was corrected by increasing the range to 14.4 mV. Inconsistent truncation may prevent supraventricular tachycardia (SVT) morphology from matching the template. B, Medtronic Wavelet algorithm. EGMs in rapidly conducted atrial fibrillation exceed the maximum amplitude range of 8 mV, resulting in varying degrees of truncation (“clipping”) of the right ventricular (RV) coil-to-can + superior vena cava (SVC) signal compared with the template. Inappropriate detection of ventricular tachycardia (VT) occurs at right. Complexes below show the last eight EGMs before detection at higher resolution. These tachycardia EGMs are clipped (arrow) so that the peaks are cut off. The rhythm is classified as VT because six of the last eight EGMs have less than 70% match. This problem was corrected by expanding the EGM scale ±16 mV AS, Atrial sensed event; VP, ventricular paced event.

T 3 D 3 6 0 0 VT 0

No match 64%

4 mV 20 ms QRS Template

No match 67%

Match 76%

No match 61%

Match 76%

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SINUS

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386 397

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858 862 A 394

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862 A

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R 233

629 R

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229 R R Sensitivity

S

R 229

633

637

R

10

Figure 3-58  Alignment error in morphology algorithm. Interaction of Automatic Sensitivity Control and morphology analysis in St. Jude MD algorithm. Left panel, Stored EGM of supraventricular tachycardia (SVT) inappropriately detected as ventricular tachycardia (VT). Right panel, Programmer strip of validated template in sinus rhythm. Despite identical ventricular EGMs, morphology match scores are 0% in SVT and 100% in sinus rhythm. Slanted line denotes slope of Automatic Sensitivity Control. In sinus rhythm, Automatic Sensitivity Control reaches minimum value before the next ventricular EGM, so that the small peak at onset of EGM (arrows) is used for alignment. In SVT, Automatic Sensitivity Control does not reach minimum, and the small peak is not used for alignment.



3  Sensing and Detection

93

degrees of aberration often distort the terminal portion of the EGM sufficiently that the percent match is less than the nominal threshold. In St. Jude ICDs, reducing the fraction of EGMs required to exceed the match threshold from 5 of 8 beats to 3 or 4 of 8 beats may reduce this problem without compromising detection of monomorphic VT. Reducing the match percent required to exceed the match threshold may also prevent misclassification of aberrantly conducted SVT as VT, but it results in a greater chance of misclassifying monomorphic VT as SVT than does reducing the fraction of EGMs.

on match percent can be tested by pectoral muscle exercise. Select an alternative EGM source (e.g., distal coil to proximal coil) to prevent such oversensing. Pectoral myopotentials also pose a source of error in Boston Scientific ICDs, which necessarily incorporate the high-voltage EGM in morphology analysis. Oversensing of pectoral myopotentials is not a problem for St. Jude ICDs, which use near-field EGMs for morphology analysis. Rate-Related Aberrancy.  If complete bundle branch aberrancy occurs reproducibly, the template may be recorded during rapid atrial pacing (Fig. 3-59). However, subtle and varying aberrancy can confound templates acquired during baseline rhythm (Fig. 3-60). If a template is acquired at a fast rate, automatic template updating should be deactivated to prevent subsequent auto-acquisition of a slow, baseline template without aberrancy. During rapidly conducted AF, subtle

Supraventricular Tachycardia Soon after Shock.  After a shock, ICD detection algorithms reclassify the rhythm as “sinus” and revert to their initial detection mode within a few seconds, but postshock distortion of EGM morphology may persist for 30 seconds to several minutes.60,120 No SVT-VT discrimination algorithm relies on EGM morphology

V.C.6 TEMPLATE IN LBBB WITH ATRIAL PACING A. Capture confirmed: 1.0 V @ 0.5 ms 1.0 V

A. Capture lost: 0.75 V @ 0.5 ms 0.75 V

Return: 3.5 V

ECG

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100 750 754 A R 281 469

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100 750 750 A R 289 461

100 750 750 A R 289 461

100 750 746 A R 289 461

100 750 750 A R 889 461

RV

Baseline

Atrial pacing LBBB

Press here for Active Template Scoring

Now Scoring

Figure 3-59  Left bundle branch aberrancy error in morphology algorithm. Programmer strips show electrocardiogram (ECG), event markers, right atrial EGM, and right ventricular true bipolar EGM. Top panel, Measurement of atrial pacing threshold. Captured beats are conducted with rate-related left bundle branch aberrancy resulting in failure of morphology to match template collected in sinus rhythm. The last two ventricular EGMs at the slower sinus rate show 100% template match. A new template was acquired during atrial pacing. Middle panel, Near 100% morphology match to new template during conduction with left bundle branch aberrancy. Bottom panel, EGM morphology templates recorded during baseline sinus rhythm (left) and atrial pacing (right). LBBB, Left bundle branch block; RA, right atrium; RV, right ventricle.

94

SECTION 1  Basic Principles of Device Therapy

DD

RA

Trigger X S T T T T T T T T T T T T T T T T ATP ATP ATP ATP ATP 46 44 36 39 38 51 37 49 39 35 57 43 24 28 50 D= 355 348 355 352 348 352 355 348 348 352 348 348 352 348 348 343 343 343 359 343 355 352 352 352 352 352 352 352 348 352 352 348 348 348 348 348 X

R

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X

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ATP ATP ATP ATP ATP

R

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R R R D 332 340 332 336 336 332 340 336 336 336 336 340 336 336 340 340 336 336 336 332 336 336 336 336 336 336 336 340 336 336 344 340 R

R

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703 543 P

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527 P

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S 100

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R 156

673

S 0

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828

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309 301 250 367 297

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61 62 0 21 38 51 34 0 0 52 398 398 398 398 398 398 398 398 398 398 402 406 402 402 402 406 402 402 395 398 A A A A A A A A A A R

R 243

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R 273

R 273

R 281

R 297

R 289

303

R R R 297 297

528

RA

RV

Figure 3-60  Error in morphology algorithm caused by subtle rate-related aberrancy. Top panel and middle panel, Dual-chamber stored EGM from inappropriately treated sinus tachycardia. Right atrial (RA) EGMs, event markers, and true bipolar right ventricular (RV) EGMs are shown. Values immediately below event markers, corresponding to EGMs in ventricular tachycardia (VT) zone (labeled “T”), indicate the percent match between EGM morphology in VT and sinus morphology (seen in lower left panel). Corresponding “X” labels above event markers indicate that the morphology algorithm classifies the beat as VT because the match is less than 60%. Burst antitachycardia pacing (ATP) is delivered at the right of the first panel. Dotted arrow indicates that panels are not continuous. After several trials of ATP (not shown), a shock is delivered toward the end of the second panel (HV). “Trigger” at right of upper panel indicates (inappropriate) detection of VT. “D =” at right of upper panel indicates that the atrial rate is equal to the ventricular rate. S denotes intervals in the sinus zone above the VT detection interval of 360 msec. Time line is in seconds. Bottom panels show real-time programmer strips. Bottom left panel, 100% template match in sinus rhythm (arrows) and 0% match on premature ventricular complexes (PVCs). However, the bottom right panel, recorded during atrial pacing at a cycle length of 400 msec, shows only two EGMs with adequate match (check marks, arrows). Note that the surface electrocardiogram (ECG) does not show identifiable aberrancy despite sufficient changes in EGM to prevent template match.



3  Sensing and Detection

after shocks until the arrhythmia episode terminates. However, algorithms revert to their initial detection mode within a few seconds after a shock. If postshock SVT starts after the rhythm has been classified as sinus but before postshock EGM distortion dissipates, any morphology algorithm may misclassify SVT as VT.120 Therefore, distortion of the EGM after a shock could result in a repetitive sequence of inappropriate classification of SVT as VT, inappropriate shocks for SVT, and perpetuation of postshock EGM changes in SVT by each successive shock.

95

about whether comparable features are required for Medtronic and Boston Scientific algorithms. SVT-VT DISCRIMINATION IN DUAL-CHAMBER AND CARDIAC RESYNCHRONIZATION ICDS The integration of dual-chamber building blocks into detection algorithms may be considered in terms of the relative rates of the atrium and ventricle (Fig. 3-62). Operation for Atrial Rate Less Than Ventricular Rate

Inappropriate Classification of Ventricular Tachycardia as Supraventricular Tachycardia.  The St. Jude morphology algorithm, which analyzes only the rate-sensing electrode, continuously misclassifies up to 6% of VTs as SVT (Fig. 3-61). However, if it is restricted to tachycardias (with ventricular rate ≤ atrial rate in dual-chamber ICDs), only 2% of VTs are misclassified.114 The Medtronic morphology algorithm, which usually analyzes high-voltage EGMs, misclassifies about 1% of VTs as SVT.127 If misclassification occurs, an alternative EGM source may provide adequate discrimination. Limited data indicate that the corresponding error rate for the Boston Scientific algorithm is comparable or lower. When the St. Jude morphology algorithm is used in a single-chamber ICD, the sustained-duration time-out (Maximum Time to Diagnosis, Section V.F) should be programmed unless the algorithm is known to classify all clinical VTs correctly. Opinions differ

Dual-chamber detection algorithms revert to single-chamber operation when the atrial rate is slower than the ventricular rate and generally do not apply any single-chamber discriminators. Operation for Atrial Rate Equal to Ventricular Rate The vast majority of tachycardias with 1 : 1 AV relationship are SVT, primarily sinus tachycardia. VT with 1 : 1 VA conduction accounts for less than 10% of VTs detected by ICDs. The building blocks used by each manufacturer for distinguishing 1 : 1 AV conduction of SVT from 1 : 1 VA conduction are summarized in Table 3-2. A dual-chamber onset rule that evaluates both P-R and R-R interval onset improves specificity of R-R interval onset-type building blocks for sinus tachycardia with minimal loss of sensitivity for VT.128 Sorin ICDs classify

V.C.7 VT 3 DAYS POSTDISCHARGE ECG

T

T

T 100

T 100

T 100

T 100

T 100

T 100

T 100

T 100

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438

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379

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383

R

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0

11

Figure 3-61  Failure to detect ventricular tachycardia (VT) because of inappropriate classification of VT as supraventricular tachycardia (SVT) by morphology algorithm (St. Jude MD). Upper panel, Programmer strip from symptomatic tachycardia classified as SVT 3 days after implantation. Electrocardiogram (ECG), event markers, and right ventricular (RV) true bipolar EGM are displayed. Morphology match is 100% on most beats. The 12-lead ECG showed a left bundle branch block–type pattern, a broad R wave in V1 (distinctly different from conducted sinus QRS complexes), and atrioventricular (AV) dissociation. Lower panel, Event markers and RV EGM during tachycardia (left), sinus rhythm after manually delivered antitachycardia pacing (center), and induced VT at predischarge electrophysiologic study (EPS) (right). EGM morphology during spontaneous VT and sinus rhythm are similar and distinct from induced VT at predischarge EPS. True bipolar EGMs fail to distinguish VT from SVT in 5% to 10% of VTs.

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SECTION 1  Basic Principles of Device Therapy

nonconducted, it is considered ventricular in origin. Individual examples are shown in Figures 3-63 to 3-66. As in the electrophysiology laboratory, pacing methods may also be used to discriminate 1 : 1 tachycardias (see Active Discrimination).

300 VF

AF/VF

Ventricular rate (bpm)

250 (A-rate, V-rate)

(A=V) ST/VT/VF

200 VT 150

VF Interval (V>A)

100

AT/VT

VT Interval

(A>V) Sinus

50

AF/VT

AT

AF

Operation for Atrial Rate Greater Than Ventricular Rate The building blocks used by each manufacturer for distinguishing rapidly conducted AF and atrial flutter from VT during atrial arrhythmias are summarized in Table 3-3. Although various combinations of these building blocks successfully discriminate VT from SVT in VT zones, most VT during AF is sufficiently rapid to be detected in the VF zone,129 where some discriminators lose specificity or may not be applied at all. For example, R-R regularity cannot be applied in the VF zone because polymorphic VT has irregular R-R intervals, and undersensing of VF exaggerates the irregularity of measured R-R intervals. AV dissociation occurs uniformly during rapidly conducted AF,68 and morphology templates acquired during sinus rhythm often misclassify aberrantly conducted beats. Individual examples are shown in Figures 3-67 to 3-69. Specific Algorithms

0 0

100

200

300

Atrial rate (bpm) Figure 3-62  Dual-chamber “rate plane” highlights power of atrial versus ventricular rate. Dual-chamber rhythm classification can be visualized as a graph that plots atrial rate (A-rate) on the abscissa and ventricular rate (V-rate) on the ordinate. All points above the horizontal line representing the ventricular tachycardia (VT) intervals will be detected as VT or ventricular fibrillation (VF) on the basis of ventricular rate alone. Points that are both above this line and above and to the left of the line of identity are VT, and no additional discriminators need to be applied in this region, provided that atrial sensing is reliable. These discriminators can be restricted to rhythms with 1 : 1 atrioventricular (AV) relationship on the line of identity and those in which the atrial rate exceeds the ventricular rate in the region below and to the right of the line of identity. AF, Atrial fibrillation; AT, atrial tachycardia; ST, sinus tachycardia. (Modified from Morris M, Marcovecchio A, KenKnight B, et al: Retrospective evaluation of detection enhancements in a dualchamber implantable cardioverter defibrillator: implications for ICD programming [abstract]. Pacing Clin Electrophysiol 22:849, 1999.)

the 1 : 1 tachycardias based on “chamber of acceleration”: the first accelerated cycle is classified as “supraventricular” if two ventricular events are classified as “conducted,” if the R-R interval is greater than 75% of the preceding P-P interval. If the two ventricular events defining the first accelerated interval are ‘ ‘conducted,’ ’ the origin of the acceleration is declared “atrial.” Conversely, if the first accelerated beat is

TABLE

3-2 

Figures 3-70 to 3-74 show block diagrams of the principal dualchamber detection algorithms in clinical use today. Legends identify unique features of each algorithm. Single-Chamber vs. Dual-Chamber Discriminators Nominal programming of dual-chamber algorithms is safe.130,131 Dualchamber stored EGMs provide higher diagnostic accuracy for troubleshooting than single-chamber stored EGMs. However, dual-chamber discriminators cannot be implemented without the complications inherent in atrial leads. Also, dual-chamber ICDs introduce unique risks for underdetection of VT caused by cross-chamber ventricular blanking after atrial pacing (see Intradevice Interactions). In addition, with early dual-chamber algorithms, optimal values of programmable parameters were not known; initial approaches to the problem of atrial blanking versus FFRW oversensing had limited success; and atrial sensing problems and specific design flaws degraded performance. Not surprisingly, clinical studies of early dual-chamber algorithms reported no benefit over single-chamber algorithms.67,68,132,133 More recent, randomized prospective studies demonstrate moderate superiority in SVT-VT discrimination for dual-chamber over single-chamber algorithms.134,135 However, no agreement exists as to when the incremental implant complexity, price, risk of atrial lead complications, and reduced longevity of a dual-chamber ICD are justified for SVT-VT discrimination alone. Generally, dual-chamber ICDs should be considered in patients who are likely to have rapidly conducted SVT, those in whom monomorphic VT and sinus tachycardia rates are likely to overlap, and those who will benefit from additional diagnostics for atrial arrhythmias. Text continued on page 101

Discrimination of Tachycardias with Atrial Rate Equal to Ventricular Rate

Building Blocks Single chamber R-R stability P-P stability Sudden onset VEGM morphology Dual Chamber A rate vs. V rate P-R patterns AV association Chamber of acceleration

GDT Rhythm ID

Medtronic PR Logic

X

X [X] {X}

St. Jude Rate Branch

Sorin Parad+

Biotronik SMART

X

X

(X) X (X) [X]

X X

X X X X

X X X

X

GDT, Boston Scientific; VEGM, ventricular electrogram; A, atrial; V, ventricular; AV, atrioventricular. X, Algorithm uses this building block; [X], new PR Logic ST rule (Entrust) uses sudden R-R and P-R onsets; (X), “implicit” use of these building blocks via pattern analysis; {X}, Protecta ICDs (investigational) integrate Wavelet VEGM morphology into dual-chamber interval analysis.



3  Sensing and Detection

97

Date Time Type 17 - MAR - 2003 12:18 Spontaneous tial Detection Parameters 20 VF: 00 VT: SVT Inhibit = On, SRD 3:00 m:s 50 VT-1: SVT Inhibit = On Nonsutained Rhythm Average V Rate

SVT 158 bpm

End of Episode

25 cm/s

RA RV Tip–RV Coil

Shock AS ––

VS ––

AS 503

VS 543

VS 430

AS 508 VS 510

AS 415

AS 348 VS 440

VS 420

AS 375 VT-1 380

AS 363

AS 320

VT-1 365

(AS) 350

(AS) 378

VT-1 VS-Hy 335 405

(AS) 375

AS 398

AS 355

AS 338

(AS) 355 (AS) 318

VT-1 VT-1 VT-1 353 VT-1 365 VT-1 385 358 370

VT-1 383 VT-1 328 Epsd

Figure 3-63  Appropriate rejection of supraventricular tachycardia (SVT) with 1 : 1 atrioventricular (AV) conduction by AV relationship and morphology algorithm. Right atrial (RA), right ventricular (RV) rate-sensing, and Shock EGMs are shown in order with dual-chamber event markers. SVT begins with gradual warm-up, accelerating from a cycle length of approximately 540 msec to 380 msec in 3 seconds. The Boston Scientific Rhythm ID algorithm classifies morphology as SVT based on 1 : 1 AV association combined with constant shock morphology and constant timing relationship between peaks of two ventricular EGMs. The label ATR at bottom of event markers indicates that the rhythm is classified as SVT. Upper panel shows episode summary classifying rhythm as SVT.

TABLE

3-3 

Discrimination of Tachycardias with Atrial Rate Greater than Ventricular Rate

Building Blocks Single Chamber R-R stability P-P stability VEGM morphology Dual Chamber A rate vs. V rate P-R patterns AV association

GDT Rhythm ID

Medtronic PR Logic

St. Jude Rate Branch

Sorin Parad+

Biotronik SMART

X

X X {X}

X

X

X X

X X X

X

X X

X

X

X X

X

A, Atrial; AV, atrioventricular. GDT, Boston Scientific; X, Algorithm uses this building block; {X}, Protecta ICDs (investigational) integrate Wavelet VEGM morphology into dual-chamber interval analysis; V, ventricular; VEGM, ventricular electrogram.

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SECTION 1  Basic Principles of Device Therapy

SINUS TACHYCARDIA

VT WITH 1:1 RETROGRADE CONDUCTION

P R

P R A Programmable 1:1 VT-ST boundary (nominally 50% R-R interval)

N Programmable 1:1 VT-ST boundary (nominally 50% R-R interval) Junctional Retrograde Antegrade

A Sinus tachycardia

VT with 1:1 retrograde conduction

1 1 1 1 1 1 1 1 1 1 3 2 2 3 2 2 3 3 3 3 3 3 3 3 7 6 6 7 7 7 7 7 7 7 5 2 9 8 1 8 8 3 1 1 0 0 0 0 0 A 0 A 0 A 0 A0 A 0 A 0 A 0 A 0 A 0 A 0 A 0 A 0 A 0 A0 A 0 A 0 A 0 A 0 A 0 A 0 A 0 A 0 A 0 A 0 A 0 S S S S S S S S S S S S S S S S S S S S S S S S S

P-R (ms) Marker events R-R (ms)

V V V V V V V V V V V T T T T T T T T T T T T T T T T S 4 S 4 S 4 S 4 S 4 S 4 S 4 S 4 S 4 S 4 S 3 S3 S3 S 3 S 3 S 3 S 3 S 3 S 4 S4 S 3 S 3 S 3 S3 S 3 S 3 D 9 9 9 9 9 9 7 8 5 7 8 6 6 6 6 6 4 3 9 9 6 9 6 6 1 0 0 0 0 0 0 0 0 0 0 0 0 0 0 0 0 0 0 0 0 0 0 0 0 0 0 0

300 P-R (ms) 200 100 500 R-R (ms) 400 300

B

1:1 A:V conduction

P-R expected range

Sudden P:R change

R-R expected range VT/VF therapy zone

Gradual R:R change

VT DETECTED

Figure 3-64  A, Discrimination of 1 : 1 tachycardias by P-R pattern using the original sinus tachycardia criterion in Medtronic PR Logic (GEM III, Marquis, Maximo, InSync Marquis, InSync Maximo) ICDs. PR Logic uses patterns of A-V, V-A, V-V, and A-A intervals from consecutive ventricular events to form couple codes describing supraventricular tachycardias (SVTs). These couple codes are based on the number of atrial (A) events within the ventricular (V) interval and their relative timing. Discrimination of sinus tachycardia (ST) from ventricular tachycardia (VT) with 1 : 1 retrograde conduction is critically dependent on the programmable 1 : 1 VT-ST boundary, which is defined nominally as 50% of the current R-R interval (with optional values of 35%, 66%, 75%, and 85%). P-R intervals less than the defined percentage of the current R-R interval and greater than 70 msec are considered antegrade (left panel), and P-R intervals equal to or greater than that percentage of the R-R interval and also greater than 40 msec are considered retrograde (right panel). Rhythms in the detection zone with 1 : 1 pattern and antegrade P-R intervals are classified as ST, and therapies are withheld. Rhythms with 1 : 1 pattern and retrograde P-R intervals are classified as VT. B, New adaptive ST rule in PR Logic (Medtronic Entrust DR ICD) no longer uses 1 : 1 VT-ST boundary to discriminate ST from VT with 1 : 1 retrograde conduction. Rather, it uses a combination of pattern, sudden R-R onset, and sudden P-R onset. Operation of the new adaptive ST criterion is illustrated by this example of a spontaneous episode of ST that converted into VT with 1 : 1 retrograde conduction. During the ST, the R-R and P-R intervals occurred within their expected ranges. After initiation of VT, the R-R and P-R intervals occurred outside their expected ranges. Evidence of ST was lost by the third beat of VT. The sudden change in R-R intervals was small (460 msec), so the R-R intervals during the VT occurred within twice the R-R interval expected range, and adaptation of the RR interval expected range continued. Eventually, the R-R interval expected range adapts to accept the R-R intervals of the VT. Conversely, the P-R intervals occurred outside of twice the P-R expected range, and adaptation was inhibited. VT is appropriately detected after 16 consecutive R-R intervals in the VT detection zone despite gradual R-R onset of a 1 : 1 rhythm, because the P-R intervals had sudden onset and were no longer within the expected P-R interval range. VF, Ventricular fibrillation. (From Stadler RW, Gunderson BD, Gillberg JM: An adaptive interval-based algorithm for withholding ICD therapy during sinus tachycardia. Pacing Clin Electrophysiol 26:1189-1201, 2003.)



3  Sensing and Detection

99

AEGM

VEGM

P-R Intervals A S

V

R-R Intervals

1 5 A0 S

1 5 A0 S V S

7 7 0

1 9 A 0 R V S

4 0 0

V S

3 5 0

1 6 A 0 R T S

3 3 0

1 9 T 0 S T F

3 6 0

1 8 A 0 R T S

3 6 0

1 6 A0 R T S

3 4 0

1 5 T0 S T S

3 2 0

1 6 T 0 S T S

3 4 0

1 6 T0 S T S

3 4 0

1 7 T0 S T S

3 5 0

1 6 T 0 S T S

3 4 0

1 6 T 0 S T S

3 4 0

T S

Interval (ms)

Detection Burst

1 2 T0 S

1 2 T0 S

T 3 F 0 0

3 0 0

1 1 T0 S

T T F 3 F 0 0

1 6 T 0 S

3 6 0

1 1 T0 S T S

4 T0 S

T T 2 F 2 P 9 6 0 0 FVT

Term.

1500 1200 900 600 400 200

ATP

212

210

28

26

24

22 0 2 Time (sec)

4

6

8

10

12

Figure 3-65  Appropriate detection of ventricular tachycardia (VT) with 1 : 1 ventricular-atrial (VA) conduction by abrupt change in P-R pattern and ventricular rate. Upper panel, Dual-chamber stored EGMs and dual-chamber EGM markers. Note that upper numerical values indicate P-R interval, not P-P interval. Lower panel, Interval plot. VT starts in the ventricle with a premature ventricular complex, followed by abrupt change in ventricular rate, ventricular EGM morphology, and the PR-RR relationship (Medtronic PR Logic). Fast VT (FVT) detection at right of EGM is followed by burst antitachycardia pacing (VT Rx 1 Burst), designated by ATP on interval plot. This results in abrupt termination of VT. Note that older versions of this algorithm, which discriminated VT from SVT based only on the PR-RP percentage, may have classified this long-RP tachycardia incorrectly as SVT. AEGM, Atrial electrogram; TF, interval in FVT zone; TS, interval in VT zone; VEGM, ventricular (true bipolar) electrogram; VS, ventricular sensed event.

Figure 3-66  Classification of 1 : 1 tachycardias by chamber of origin. The Sorin Parad algorithm classifies 1 : 1 tachycardias by chamber of onset. Stored ventricular EGM is shown with dual-chamber event markers (top, atrial; bottom, ventricular). Event log and Event summary are shown above each stored EGM. Upper panel, Supraventricular tachycardia (SVT) is diagnosed by atrial onset. Lower panel, Ventricular tachycardia (VT) is diagnosed by ventricular onset.

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SECTION 1  Basic Principles of Device Therapy

Figure 3-67  Ventricular tachycardia (VT) during rapidly conducted atrial flutter discriminated by morphology of ventricular EGM. VT occurs during atrial flutter with 2 : 1 atrioventricular (AV) conduction. Time line marks indicate that the ventricular rate had been in the VT zone for 24 seconds at the beginning of the panel. VT therapy is withheld because of both the 2 : 1 AV relationship and the morphology match of ventricular EGMs to baseline sinus-rhythm EGMs (not shown). Numbers below event markers indicate 98% to 100% morphology match in atrial flutter, compared with 40% to 47% match in VT; “<” (box) marks indicate that ventricular rate is less than atrial rate. VT is detected by AV dissociation and morphology match percentage less than the threshold value of 60%. Morphology classification of each beat as SVT (check marks) or VT (X) is indicated above event markers. (From Swerdlow C, Shivkumar K: Implantable cardioverter defibrillators: clinical aspects. In Zipes DP, Jalife J, editors: Cardiac electrophysiology: from cell to bedside, ed 4, Philadelphia, 2004, Saunders.)

Figure 3-68  Inappropriate detection of rapidly conducted atrial fibrillation (AF) despite rejection by morphology. Right atrial (RA) EGM, event markers, and true bipolar right ventricular (RV) EGM are shown. Values immediately below event markers corresponding to EGMs in ventricular tachycardia (VT) zone (marked T) indicate the percent match between EGM morphology in tachycardia and sinus morphology. Corresponding check marks above event markers indicate that the morphology algorithm classifies the beats as supraventricular tachycardia (SVT) because the match is greater than 60%. Stability and morphology algorithms are combined with “AND” logic in this St. Jude ICD, so that VT is diagnosed if either discriminator classifies the rhythm as VT. The stability algorithm incorrectly classifies the rhythm because the ventricular cycle lengths regularize. The rhythm would have been classified correctly if the morphology discriminator alone had been programmed. F markers indicate ventricular intervals in ventricular fibrillation (“Fib”) zone. S denotes intervals in the “Sinus” zone above the VT detection interval. T denotes intervals in VT zone. DDI, Mode switch to DDI pacing. (From Swerdlow CD, Friedman PA: Advanced ICD troubleshooting. Part I. Pacing Clin Electrophysiol 28:1322-1346, 2005.)



3  Sensing and Detection A-A

Interval (msec)

1800 1500 1200 900 600

101

V-V FVT=240 msec VT=400 msec VF=320 msec

400 200

0

1

2

3

4

5

6

7

8

9

10

11

12

13

14

15

Time (sec) [0=Collection Start] AEGM VEGM

Marker 2 1 1

1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 2 0 9 9 9 9 9 9 9 9 9 9 9 9 9 9 9 9 9 9 9 9 9 9 9 9 9 9 9 9 9 9 9 9 9 9 9 9 9 9 9 0 A 0 A 0 A 0 A 0 A 0 A 0 A 0 A0 A 0 A0 A 0 A0 A0 A0 A0 A0 A0 A0 A 0 A0 A0 A0 A0 A0 A0 A0 A0 A0 A0 A0 A0 A0 A0 A0 A0 A0 A 0 A0 A 0 A0 A 0 A R S R S R S R S R S R S R S R S R S R S R S R S R S R S R S R S R S R S R S R S R S 3 9 0

T S

4 0 0

V S

V 4 S 0 0

3 9 0

T S 3 8 0

T S 3 8 0

T S

3 8 0

T S

3 9 0

T S

3 9 0

T S

3 9 0

T S

3 8 0

T S 3 8 0

T S

3 8 0

T S

3 9 0

T S

3 9 0

T S AF

3 8 0

T S AF

3 8 0

T S AF

3 8 0

T S AF

3 9 0

T S 3 9 0 AF

T S AF

3 8 0

6 3 0

T S AF

6 3 0

A P 7 3 0

V P

6 3 0

A P 6 3 0

V P

Figure 3-69  Appropriate rejection of rapidly conducted atrial flutter by pattern analysis. Upper panel, Interval plot. Lower panel, Dualchamber stored EGM with event markers. Atrial flutter is identified by the consistent 2 : 1 PR association of the regular atrial and ventricular rhythms. AF annotations indicate that the number of intervals in the ventricular tachycardia (VT) rate zone required for detection (16) has been reached, but that therapy is withheld because atrial flutter is diagnosed. Interval plot displays 2 : 1 atrioventricular (AV) association until termination of atrial flutter after 12 seconds. AR, Atrial intervals in pacing refractory period; TS, ventricular intervals in VT zone.

Unknown rhythm detected in VT or VT-1 rate zone

V rate > (A rate + 10 bpm)

ACTIVE DISCRIMINATION

Yes

VT

No VTC rhythm No correlated? Yes SVT

A rate >200 bpm and No V rate unstable (>20 ms)

VT

Yes SVT

Figure 3-70  Boston Scientific Rhythm ID algorithm. The central features of this algorithm are comparison of atrial (A) and ventricular (V) rates (first step) and analysis of ventricular EGM morphology (Vector Timing and Correlation, or VTC). If the V rate exceeds the A rate by at least 10 bpm, the rhythm is classified as ventricular tachycardia (VT). If at least 3 of the last 10 beats are classified as supraventricular by morphology analysis, the rhythm is classified as supraventricular tachycardia (SVT). If morphology analysis does not classify the rhythm as SVT, the A rate and the regularity of the V rate are evaluated. The rhythm is classified as SVT if the A rate exceeds 200 bpm and the ventricular rhythm is irregular (measured ventricular interval stability >20 msec). Otherwise, it is classified as a VT.

Active discrimination represents a paradigm shift in the design of detection algorithms, from “diagnose before intervening” to “treat first; analyze only those tachycardias that persist after treatment.” Although active discrimination presents the risk of proarrhythmia (Fig. 3-75), it can be valuable in discriminating tachycardias with 1 : 1 AV association (Figs. 3-76 and 3-77). One such algorithm was designed and implemented by Biotronik (and incorporated into the SMART-II algorithm) but is not commercialized in Biotronik ICDs.136 Feasibility investigations recently demonstrated a new algorithm using simultaneous atrial and ventricular ATP either to terminate or to discriminate 1 : 1 tachycardias. If the tachycardia persists after the simultaneous AV ATP, the discrimination algorithm considers the rhythm ventricular in origin, if the first sensed event after pacing is on the ventricular channel, and supraventricular in origin otherwise. In 62 ambulatory dual-chamber ICD patients, the algorithm terminated or correctly classified 1379 of 1381 SVT sequences (specificity 99.9%) and 23 of 26 VTs (sensitivity 88.5%).115 FEATURES TO OVERRIDE DISCRIMINATORS Programmable duration-based “safety net” features deliver therapy if an arrhythmia satisfies the ventricular rate criterion for a sufficiently long duration even if discriminators indicate SVT. The premise is that VT will continue to satisfy the rate criterion for the programmed duration, whereas the ventricular rate during transient sinus tachycardia or AF will decrease below the VT rate boundary before the programmed duration is exceeded. The limitation is delivery of inappropriate

102

SECTION 1  Basic Principles of Device Therapy

A-V RATE BRANCH Median atrial rate/ median ventricular rate

V
V=A

Morphology discrimination (MD) Interval stability

Morphology discrimination (MD) Sudden onset

VA

AFib/flutter SVT w/multiblock

VT

ST/SVT

VT

VT/VF

Inhibit therapy

Deliver therapy

Inhibit therapy

Deliver therapy

Deliver therapy

Additional discriminators Diagnosis Figure 3-71  St. Jude A-V Rate Branch + Morphology Discrimination (MD) algorithm. Similar to the Rhythm ID, the central features of this algorithm are a comparison of the atrial (A) and ventricular (V) rates (first step) and analysis of ventricular EGM morphology (MD). Rhythms are classified into three Rate Branches: ventricular rate less than atrial rate (V < A), ventricular rate equal to atrial rate (V = A), and ventricular rate greater than atrial rate (V > A). If V < A, the algorithm applies morphology discrimination (and/or both interval stability and N:1 AV association). If V = A, it applies morphology discrimination (and/or sudden onset). All rhythms in the V > A rate branch are treated as ventricular tachycardia (VT). ST, Sinus tachycardia; SVT, supraventricular tachycardia; VF, ventricular fibrillation.

BIOTRONIK SMART

V rate > A rate

V rate = A rate

V rate < A rate

V unstable

V stable

N:1 multiplicity

A unstable

No N:1 multiplicity

V unstable

V stable

A stable

AV trend

VT

AFlut

VT

AFib

VT

VT

AV irregular

AV regular

No AV trend

Sudden onset

No sudden onset

VT

Sinus T

VT

1:1 SVT

Figure 3-72  Biotronik SMART algorithm. The high-level structure of this algorithm is similar to that of the St. Jude ICD. SMART first compares atrial (A) and ventricular (V) rates. If the V rate is faster, ventricular tachycardia (VT) is diagnosed. If the V rate is slower, ventricular interval stability and A : V ratio analysis results in discrimination of atrial fibrillation (AFib; unstable ventricular rhythm), atrial flutter (AFlut; stable rhythm, N:1 association), and VT (stable rhythm, AV dissociated). If the A rate and V rate are equal, SMART analyzes the interval stability. If the R-R interval is stable and the P-P interval is unstable, VT is diagnosed. If the PP interval is stable, the algorithm sequentially analyzes AV association (“AV Trend”) and sudden onset as discriminators. Sinus T, Sinus tachycardia; SVT, supraventricular tachycardia.



3  Sensing and Detection

A rate < V rate (Many RRs don’t have Ps) A R

Building blocks

Possible classification

T S

T S

A rate = V rate (Most RRs have only 1 P)

A R T S

T S

VT/VF (not SVT)

A R T S

A R T S

A R T S

A R T S

T S

A R T S

T S

A R

A R

A R

T S

A A R R T S

A R T S

A A R R T S

ST/AT

VT/VF + SVT

Other 1:1 SVT

Atrial fib/flutter

VT/VF (not SVT)

ST/AT (w/FFRW OS) VT/VF

• R-R intervals • PR patterns

• RR intervals • PR patterns

Get more data

A rate > V rate (Most RRs have more than 1 P)

A R

103

Process PR logic building blocks (up to 24 R-R intervals and P:R patterns)

A R T S

VT/FVT/VF detection by RR intervals? Yes

No

Yes Median RR < SVTLimit? No

Yes

Double tachycardia identified? No

• RR intervals • PP intervals • PR patterns • RR regularity • PR dissociation • PP regularity

AFib/AFlutter?

Yes

No Sinus tach?

Yes

No Other 1:1 SVT?

Yes

No VT/FVT/VF DETECTION

A

B

Figure 3-73  Medtronic PR Logic algorithm. A, Clinical decision process. The rhythm types are represented using P-R marker diagram examples (top), from least complex (left) to most complex (right). A list of possible PR Logic decisions and the building blocks used are also shown for each of the rhythm types. Note that PR Logic implicitly employs a ventricular (V) rate > atrial (A) rate override (first column), because all rhythms with A rate < V rate and V rate in one of the detection zones are classified as ventricular tachycardia/fibrillation (VT/VF). For tachycardias with A rate = V rate (i.e., 1 : 1 tachycardias, center column), the original PR Logic uses P-R patterns and discriminates 1 : 1 rhythms with critical P-R interval timing zones (see Fig. 3-64, A). The new sinus tachycardia (ST) criterion in PR Logic (Entrust and later) uses the P-R pattern along with P-R and R-R suddenonset criteria (see Fig. 3-64, B). For rhythms with A rate > V rate, PR Logic uses P-R patterns along with R-R regularity, P-R dissociation, and P-P regularity to ensure detection of double tachycardia (VT or VF during atrial fibrillation [AFib]) and to withhold therapy for 2 : 1 atrial flutter, rapid AFib, and ST with far-field R-wave oversensing (FFRW OS). In the newest Medtronic ICDs (Protecta, currently investigational), EGM morphology analysis (Wavelet) is applied for rhythms with A rate ≥ V rate that are identified by SVT by PR Logic pattern analysis alone. In addition, detection of double tachycardia in the Protecta devices requires that abnormal EGM morphology be present. Integration of Wavelet EGM morphology with PR Logic is designed to improve detection specificity compared to PR Logic interval analysis alone. B, PR Logic computational flow diagram. On each ventricular event, PR Logic processes the new P-R, R-P, P-P, and R-R patterns and timing information for the building blocks. If VT/fast VT (FVT) or VF rate detection criteria are satisfied, the ventricular rate override criterion is checked first. If the median R-R interval is less than the supraventricular tachycardia (SVT) limit, detection occurs through the single-chamber detection criteria without considering the PR Logic discrimination algorithm. If the median R-R interval is greater than the SVT limit, and if double tachycardia (VT/FVT/VF + SVT) is not detected, the three PR Logic criteria for identifying SVTs are tested in the order shown. If any one of the PR Logic SVT criteria is satisfied, inappropriate detection is avoided. If an SVT is not positively identified by PR Logic pattern analysis and the A rate ≥ V rate, EGM morphology analysis (Wavelet) is applied and classify the rhythm as SVT if the EGM morphology during tachycardia matches the template. If none of the SVT discrimination criteria is satisfied, VT/FVT/VF is detected when the R-R interval–based criterion is satisfied. If SVT is identified, the entire process repeats itself on each ventricular event until VT/FVT or VF is detected or the rhythm slows out of the ventricular rate detection zones.

therapy when SVT exceeds the programmed duration, which occurs frequently for durations of 1 minute or less and in up to 10% of SVTs at 3 minutes, depending on the VT detection interval and AV conduction.116 Because SVTs are much more common than VTs, programming of discriminator override duration to 3 minutes results primarily in inappropriate therapy of SVT. Programmed durations of 5 to 10 minutes are required to minimize such inappropriate therapy.137 The decision to use a discriminator override should be based on clinical factors, including the probability that discriminators will prevent detection of VT, the likely consequences of failure to detect VT, and the likelihood that SVT in the VT rate zone will persist long enough to trigger inappropriate therapy because of the override. For example, override features may be considered whenever a morphology algorithm is programmed without inducing VT at electrophysiologic study. The Medtronic Protecta ICDs incorporate an additional, separately programmable “VF Zone High Rate Time Out.” It applies only to rhythms that are consistently in the VF rate detection zone. This new algorithm allows SVT discriminators to be applied with

longer-duration (or no) time-out for slower rhythms, with a separate override duration for rhythms with rates in the VF zone. SVT-VT DISCRIMINATORS IN REDETECTION The SVT-VT discrimination during redetection serves two purposes: (1) to prevent inappropriate therapy for SVT after appropriate therapy for VT (Fig. 3-78) and (2) to provide a second chance for the algorithm to classify SVT correctly after inappropriate therapy. Biotronik ICDs (Fig. 3-79) and sorin-ELA ICDs provide essentially equivalent SVT-VT discrimination in initial detection and redetection, except that algorithm building blocks related to tachycardia onset are disabled. Boston Scientific algorithms permit programming discriminators after shocks, but not after ATP, when it is more useful and reliable. In Medtronic ICDs, the single-chamber stability discriminator applies to redetection only if it is “on” for initial detection. The user interface does not indicate that it applies in redetection, and it is rarely programmed for this purpose. St. Jude ICDs do not apply discriminators to redetection.

104

SECTION 1  Basic Principles of Device Therapy

Tachyarrhythmia detected

Unstable

RR Stability

Diagnosis AF No VTLC No Diagnosis VT

Inappropriate Detection, Inappropriate Therapy, and Inappropriate Shocks All failures of SVT-VT discrimination algorithms do not have equivalent clinical outcomes. ATP is delivered immediately after detection so the numbers of detections and therapies are equivalent. In contrast, shock delivery requires capacitor charging, during which therapy may be aborted. However, inappropriate ATP is often unnoticed. It terminates about 40% of SVTs and reduces the ventricular rate below the VT detection interval in another 20%.139 In contrast, inappropriate shocks adversely affect quality of life.140 Active discrimination blurs the distinction between detection and therapy because only tachycardias that persist immediately after diagnostic pacing are classified.

Stable

Associated PR Yes Yes

Diagnosis No AF Diagnosis AFlutter

PR Association 1:1 Gradual

Diagnosis ST

IS VENTRICULAR THERAPY FOR SVT ALWAYS INAPPROPRIATE?

Acceleration Sudden Origin of acceleration Atrial

Diagnosis AT

Ventricular Diagnosis VT

Figure 3-74  Sorin Parad+ algorithm. If most of the detected R-R intervals are in the ventricular tachycardia (VT) zone, ventricular interval stability is analyzed in a first step, using a histogram of R-R intervals. If the rhythm is irregular, atrial fibrillation (AF) is diagnosed and therapy is withheld. If the rhythm is regular, the atrioventricular (AV) association is assessed by comparing peak amplitudes of R-R and P-R interval histograms. If the rhythm is AV dissociated, VT is diagnosed, unless the Long Cycle Search is activated, which inhibits therapy if a long ventricular cycle (VTLC, characteristic of AF) is identified. If N:1 P-R association is identified, the rhythm is classified as supraventricular tachycardia (SVT). In the presence of 1 : 1 P-R association, the Parad+ evaluates the rate of acceleration of the ventricular rate. If acceleration is gradual, sinus tachycardia (ST) is diagnosed. If acceleration is sudden, Parad+ identifies the chamber of origin and withholds therapy if it is the atrium. AFlutter, Atrial flutter; AT, atrial tachycardia.

Neither Medtronic nor St. Jude ICDs provide any discriminators to reject sinus tachycardia in redetection. MEASURING PERFORMANCE OF SVT-VT DISCRIMINATION ALGORITHMS Valid assessments and comparisons of algorithm performance require consideration of multiple ICD-related and clinical factors. A comprehensive assessment requires analysis of all tachycardia episodes, including those not stored in ICD memory and those in the VF zone to which discriminators may not apply.68,69,130 Programmed detection parameters may influence reported algorithm performance.69 In most studies, a few patients contribute a large number of SVT or VT episodes. Therefore, statistical methods such as the generalized estimating equation (GEE)138 should control for such clustered, nonindependent data. Quantitative Considerations Detection algorithms must maintain almost 100% sensitivity for detection of VT; but detection of hemodynamically-stable asymptomatic VT is not necessarily synonymous with therapy. If patient populations and programmed detection boundaries are equivalent, positive predictive accuracy may be the most useful statistical measure of algorithm performance.68 However, it is highly dependent on the ratio of SVT to VT episodes and therefore on the programmed detection rate and patient population (prevalence of SVT and VT). However, it is almost impossible to obtain a clinically meaningful comparison of different detection algorithms based only on their performance on different sets of data.

Persistent, rapidly conducted atrial arrhythmias can cause hemodynamic compromise in patients with LV dysfunction or ischemia in patients with severe coronary artery disease. Thus, algorithmically inappropriate ventricular therapy may fortuitously terminate clinically significant SVT, although inappropriate ventricular therapy for SVT may also be proarrhythmic.141 Further, in patients with ICDs, rapid conduction in AF is often transient, and symptoms are often mild, but ventricular shocks delivered shortly after detection do not permit spontaneous termination of AF or slowing of the ventricular rate. Therefore, shocks will be delivered for AF that may have spontaneously terminated or slowed. Inappropriate shocks for AF may place patients at risk for thromboembolism if they are not anticoagulated. Also, early recurrence is common after transvenous cardioversion of AF.142 Experts differ about whether algorithmically inappropriate ventricular therapy of SVT may be clinically appropriate in specific clinical situations. ICDs designed to deliver both atrial and ventricular therapies may be implanted in patients who are likely to benefit from ICD-based therapy of SVT.

Detection: Programming and Troubleshooting The multiple programmable parameters that affect detection directly or indirectly provide opportunities both for customizing detection and for operator error. A prospective study reported that empirical programming of dual-chamber ICDs provided comparable performance to individualized programming.131 Nevertheless, it is important to understand the many trade-offs and compromises inherent in ICD detection strategies. DETECTION ZONES AND DURATION Zones and Zone Boundaries In patients receiving secondary prevention therapy, typical values for rate zone boundaries are 400 to 360 msec for the sinus-VT boundary, 330 to 240 msec for VT–fast VT boundary, and 300 to 240 msec for the upper limit of the VF zone boundary, with longer cycle lengths for the last boundary usually found in two-zone programming. In primary prevention patients and those whose only arrhythmia is VF, typically one or two zones are programmed, with the sinus boundary set at 330 to 300 msec.143,144 This approach lowers the risk of inappropriate therapy but increases the risk of not treating VT. Programming that permits delivery of ATP is desirable for most patients, even those undergoing implantation for primary prevention and those whose only clinical arrhythmia is VF, because most spontaneous VF begins with rapid VT, and most rapid VT can be terminated by ATP.145-147 Three zones permit different ATP therapies for two distinct rates of VT, as well as ATP for monomorphic VT that overlaps in rate with polymorphic VT. In the largest primary prevention trial,144 ICDs were programmed to a single detection zone at a cycle length of 320 msec, without SVT-VT discriminators, but approximately one



105

3  Sensing and Detection

RA ATP

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6 262 445 281 648 309 289 2 324 285 254 285 285 266 297 313 273 R

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51

Figure 3-75  Atrial proarrhythmia caused by ventricular antitachycardia pacing in 1 : 1 supraventricular tachycardia (SVT). Continuous stored EGM shows right atrial (RA) EGM, event markers, and true bipolar right ventricular (RV) EGM. In the top panel, antitachycardia pacing (ATP) is delivered after rate-only, inappropriate detection of sinus tachycardia as ventricular tachycardia. Ventricular-atrial conduction occurs, initiating atrial fibrillation (AF), which is detected as ventricular fibrillation because redetection does not utilize discriminators. AF requires two shocks (HV) before termination. The first shock is delivered at the end of the middle panel and the second in the middle of the lower panel. Recording is suspended for approximately 1 second after each shock.

106

SECTION 1  Basic Principles of Device Therapy

EGM1: Atip to Aring

EGM2: Can to HVB 3 1 0

A-A interval (ms) Marker annotation

T S

V-V interval (ms)

3 1 0

3 1 0

A A S S

A S

3 2 0

A S

T T T S S S 3 3 3 1 1 1 0 0 0

3 2 0

3 3 3 2 2 2 0 0 0 A A A A S S S S

6 4 0

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T T T T T T T S S D P P P P 3 3 3 2 2 2 2 1 2 2 7 7 7 7 0 0 0 0 0 0 0

3 2 0

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T T T T P P P P 2 2 2 2 7 7 7 7 0 0 0 0

third of shocks were inappropriate, most often for rapidly conducted AF. The sinus-VT rate boundary should be slow enough to ensure detection of all hemodynamically compromising VTs. To prevent underdetection of irregular VT, the VT detection interval should be set with a safety margin at least 40 to 50 msec longer than the slowest predicted VT for consecutive-interval counting and 30 to 40 msec longer for “X out of Y” or “interval + interval-average” counting.134 This safety margin should be greater if rapidly conducted SVT is unlikely or SVT-VT discrimination is reliable at long cycle lengths (Sorin ICDs). The boundary between the two VT zones should be based on the cycle length at which different types or fewer trials of ATP are preferred. The VT-VF rate boundary is based on the cycle length below which ATP or ATP before charging should not be delivered. In Medtronic ICDs, which use consecutive-interval counting above the VF interval and “X out of Y” counting below it, this boundary should be set to prevent underdetection of irregular, polymorphic VT by consecutive-interval counting.

A S

3 6 0

7 1 0

A S

3 1 0

V S

3 1 0

A A S S

3 1 0

T T S F 3 2 6 7 0 0

Figure 3-76  Effect of ventricular antitachycardia pacing on supraventricular tachycardia (SVT) with 1 : 1 atrioventricular (AV) relationship. Atrial EGM (Atip to Aring), high-voltage ventricular EGM (Can to HVB), and dual-chamber event markers are shown. Burst ventricular antitachycardia pacing at cycle length 270 msec is applied to SVTs with cycle lengths of 310 to 320 msec. The atrial interval after the last paced beat is accelerated to the pacing rate, probably indicating entrainment of the atrium. The A-A-V response at termination of pacing is diagnostic of atrial tachycardia.

Duration for Detection of Ventricular Tachycardia Detection duration before ATP should rarely be decreased from nominal values, because therapy is immediate after detection, and undersensing of monomorphic VT is rare. It should be increased in patients who have long episodes of nonsustained VT. Substantial increases in duration for detection of VT probably are safe in St. Jude and Boston Scientific ICDs, which use counting methods that are insensitive to occasional long ventricular intervals or undersensing. In contrast, consecutive-interval counting used by Medtronic ICDs in the VT zone and the Sorin persistence counter (nominally set to 12 consecutive intervals for VT and 6 for VF) may underdetect VT if occasional, long ventricular intervals or undersensing occurs.119 Duration for Detection of Ventricular Fibrillation Duration for detection of VF generally should not be reduced from nominal to prevent inappropriate therapy or aborted shocks for nonsustained VT.146,148 A prospective study of primary prevention patients demonstrated that long-duration detection (30/40 intervals) did not

A-EGM

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3 3 3 4 3 3 3 3 3 3 3 3 3 3 3 3 3 3 3 3 8 8 9 2 1 1 1 1 2 2 3 2 2 2 2 7 6 6 4 6 0 0 0 0 0 0 0 0 0 0 0 0 0 0 0 0 0 0 0 0 A A A A A A A A A A A A A A A A A A A A R R R R S S S S S S S S S S S S R R R R T S

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V S

3 9 0

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3 8 0

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Figure 3-77  Atrial response to ventricular antitachycardia pacing during ventricular tachycardia (VT) with 1 : 1 ventricular-atrial (VA) conduction. Stored atrial (A) and ventricular (V) EGMs are shown with event markers. During antitachycardia pacing, atrial rate accelerates to ventricular rate. The V-A-V response at termination of pacing is characteristic of VT, although atrioventricular (AV) nodal and AV reciprocating tachycardias cannot be excluded.



3  Sensing and Detection

107

RV Tip-RV Ring

RV Coil-Can T F 3 2 0

3 1 0

T F 3 2 0

T F 3 2 0

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VT

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T F 3 4 0

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T F 3 5 0

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ATP

T F 3 6 0

T F 3 6 0

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Sinus Tachycardia T F2 9 0

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ATP T• F

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3 3 6 6 0 0 FVT Rx 1/Sec 2

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14.9 J

A V-V 1,800

VF = 300 ms Burst

V-V interval (ms)

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FVT = 360 ms 20.1 J

VT = 470 ms 35.1 J

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Time (sec) [0 = detection]

Figure 3-78  Redetection. Inappropriate therapy of sinus tachycardia after appropriate therapy for ventricular tachycardia (VT) in a singlechamber ICD. Ventricular sensing and high-voltage leads are shown with event markers. A, Continuous stored EGM strips after detection of VT. The top strip shows successful antitachycardia pacing (ATP) of VT followed by sinus tachycardia in the Fast VT (FVT) zone (TF on event markers). The second strip shows the first inappropriate sequence of ATP. The third strip shows the first of five shocks delivered after ATP. B, Interval plot shows the entire episode. No U.S. ICD manufacturer provides for VT-supraventricular tachycardia (SVT) discrimination after ATP. RV, Right ventricular; VF, ventricular fibrillation. (From Swerdlow CD, Friedman PA: Advanced ICD troubleshooting. Part II. Pacing Clin Electrophysiol 29:70-96, 2006.)

108

AFlut 399

AFlut 398

AFlut 399

VT1 390

VT1 398

VT1 399

AS 719 Ars 195 AS 203 AS 196 Ars 195 AS 203 AS 195 Ars 196 AS 203 Ars 195 AS 196 AS 203 Ars 195 VS 1844

AS Ars 109 VT2 320

20 J 81Ohm

25 AS 985 VS 984

AS 984 VS 984

VS 985

AS 985

AS 975

AS 985 VS 977

AS 984

Shock

VS 984

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AS 984

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AS 726 VS 992

AFlut 391

AFlut 390

AFlut 391

AFlut 391

AFlut 398

AFlut 398

AS 195 AS 204 Ars 195 AS 195 AS 203 Ars 196 AS 203 Ars 195 AS 195 AS 203 Ars AS 125

25 mm/s

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Charge

VT2 328

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AS 1617 VT2 320

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VT2 320

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AS 586 VT2 351

VT2 320

VT1

VT2 320

VT2 320

VT2 328

AS 750

AS 1617

SECTION 1  Basic Principles of Device Therapy

Figure 3-79  Supraventricular tachycardia/ventricular tachycardia (SVT-VT) discrimination in redetection. Dual-chamber EGM markers, atrial EGM, and true bipolar EGM are shown in this continuous strip from a Biotronik Lexos ICD, Model 347000. Rhythm at beginning of upper panel is VT with ventricular-atrial (VA) dissociation. Vertical dotted line indicates detection and start of capacitor charging (Charge). Black horizontal line indicates period of capacitor charging followed by successful shock. Postshock nonsustained atrial flutter begins toward right side of upper panel, continues into lower panel, and shows transient atrial flutter followed by conversion to sinus rhythm. Ventricular rhythm is classified by the lower row of markers as VT2 (faster VT zone) before shock, VT1 (slower VT zone) for the first four ventricular EGMs of atrial flutter, and conducted atrial flutter (Aflut) beginning at the fifth conducted ventricular EGM during atrial flutter (arrowhead). Without SVT-VT discrimination in redetection, VT would have been redetected after 10 intervals.

increase risk of syncope compared to historical controls.143 A significant reduction in heart failure hospitalization (as well as shocks) was reported in a controlled study of primary-prevention, nonischemic heart failure patients comparing detection duration of 12/16 with 30/40 intervals.149

redetection may prevent inappropriate redetection of delayed termination of VT (type II break) or postshock nonsustained VT. ICDs misclassify effective therapy as ineffective if VT/VF recurs before the ICD identifies episode termination and reclassifies the posttherapy rhythm as “sinus” (Fig. 3-80). Misclassification may also occur if SVT begins after successful VT therapy but before ICD-defined episode termination because of frequent premature beats or nonsustained tachycardia. Decreasing the duration for redetection of sinus rhythm (St. Jude) may correct this classification error.

Duration for Redetection Inappropriate therapy for nonsustained VT or SVT may be delivered after ATP or shocks (see Fig. 3-78). Increasing the duration for

5 3 0

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C D 15.0

Figure 3-80  Failure to identify post-therapy sinus rhythm because of rapid reinitiation of ventricular tachycardia (VT) after successful therapy. Atrial EGM, ventricular true bipolar EGM, and EGM markers classify the rhythm as sinus (8 consecutive intervals). Therefore, VT is inappropriately redetected instead of being detected de novo for the second time. This results in delivery of the second programmed Fast VT (FVT) therapy, cardioversion, rather than repeat delivery of the previously successful antitachycardia pacing (ATP) stimulus, which is both painless and more energy efficient. Redetection marker (FV Rx 2 CV) indicates onset of capacitor charging for the second VT therapy. Shock is not shown. Large numbers below right side of upper panel and left side of lower panel show value of sinus rhythm counter, which increments for each interval in the sinus zone (VS) and is reset to zero by premature ventricular complexes (PVCs) with a coupling interval of less than the VT detection interval of 400 msec (TS). VS markers indicate unclassified ventricular intervals during capacitor charging. AS, Atrial sensed event; AR, atrial intervals in pacing refractory period.



3  Sensing and Detection

Recommended Programming of SVT-VT Discriminators in Dual-Chamber ICDs

RANGE OF CYCLE LENGTHS TO WHICH SVT-VT DISCRIMINATORS APPLY

TABLE

The SVT-VT discriminators apply in a range of cycle lengths bounded on the slower end by the VT detection interval and on the faster end by a minimum cycle length that varies among manufacturers (see Fig. 3-51). Usually, SVT-VT discriminators will not withhold inappropriate therapy for SVT if the majority of ventricular intervals (typically 70%80%) are shorter than the SVT limit. Therefore, rapidly conducted AF may be classified as VT even if the mean cycle length is 20 to 40 msec longer than the SVT limit. Programming a sufficiently short, minimum cycle length for SVT-VT discrimination is critical to reliable rejection of SVT. When discriminators are programmed, approximately 25% of inappropriate therapy is caused by SVT with ventricular cycle lengths shorter than that minimum cycle length.68,137,150 In some ICDs, the SVT discriminators can be programmed only in the VT detection zone if two-zone programming is used or in the slowest zone with three-zone programming. In currently approved Medtronic ICDs, the SVT limit is programmable independently, but the performance of SVT-VT discriminators is linked implicitly to the VT/VF zone boundary. AF is not rejected by RR regularity at cycle lengths shorter than the programmed VF detection interval; thus rapidly conducted AF cannot be distinguished from VF in dualchamber and cardiac resynchronization ICDs before the (now investigational) Protecta models. In nominal programming, the “VF Detection Interval” forms the boundary between the VT and Fast VT zones. Therefore, this degradation in discrimination of VT from rapidly conducted AF usually occurs between the VT and Fast VT zones. In these ICDs, conducted AF in the Fast VT zone may be discriminated by programming “Fast VT via VT” rather than the nominal “Fast VT via VF.” The risk is delay in detection of unusual, markedly irregular fast VT with occasional cycle lengths in the sinus zone.

Medtronic PR Logic AFib/AFlutter ON Sinus Tach ON*

Atrial View AFib Rate Threshold 200 bpm

Other 1 : 1 SVTs

Onset 9%

{Wavelet ON Match Threshold = 70%}, Protecta devices only (investigational)

Inhibit if unstable 10%

PROGRAMMING OF SVT-VT DISCRIMINATORS Single-Chamber Discriminators Technical details vary among manufacturers, as do corresponding recommended programmed values, which are summarized in Table 3-4. See earlier Ventricular EGM Morphology for SVT-VT Discrimination for programming and troubleshooting of morphology algorithms. Dual-Chamber Discriminators Dual-chamber algorithms should be programmed “on” in any patient with intact AV conduction and a functioning atrial lead. Specific considerations for each manufacturer are summarized in Table 3-5. In St. Jude ICDs, discriminators may be combined using either the “any” or “all” operators. Using “any,” the algorithm detects VT if any discriminator classifies the tachycardia as VT, resulting in higher sensitivity and lower specificity. Conversely, using “all,” the algorithm detects VT only if all discriminators classify the tachycardia as VT, resulting in lower sensitivity and higher specificity. The “all” operator corresponds to the addition of discriminators in other algorithms.

3-5 

109

Boston Scientific Rhythm ID ON

St. Jude Rate Branch Rate Branch ON A = V Branch: Morphology A > V Branch Morphology; may combine Stability† with “ANY” logic

V rate > A rate ON Sustained rate duration = 3 minutes *In older models before Entrust (Model D153DRG) (*1:1 VT-ST Boundary = 66%). †Stability at 80 msec with AV Association of 60 msec. A, Atrial; AFib, atrial fibrillation; Tach, tachycardia; V, ventricular.

St. Jude Rate Branch.  Morphology should be programmed in both the “V = A” and “V < A” rate branches. Recommended programming adds the stability discriminator in the V < A branch using the “any” operator. This results in a minor increase in sensitivity for detection of VT (98% to 99%) and a similarly minor decrease in specificity (82% to 79%). Addition of the stability discriminator using the “all” operator reduces inappropriate detection of aberrantly conducted AF, but may also reduce sensitivity for detection of VT.114 Boston-Scientific Rhythm ID.  This algorithm has no programmable features. Medtronic PR Logic.  At implantation, rejection rules should be programmed “on” for sinus tachycardia and atrial fibrillation/flutter. The 1 : 1 SVT rejection rule should not be programmed until the atrial lead is stable, because its dislodgement to the ventricle may result in misclassification of VT as a 1 : 1 SVT. This potential problem also applies to the St. Jude Rate Branch algorithm without additional discriminators.

Failure to Deliver Therapy: Undersensing and Underdetection Undersensing and underdetection may be caused by ICD system performance, programmed values (including human error), or a combination of the two. They result in failure to delivery therapy or delay in therapy. UNDERSENSING

TABLE

3-4 

Recommended Programming of SVT-VT Discriminators in Single-Chamber ICDs

Stability* Onset Morphology

Medtronic 40-50 msec, NID = 16 84%-88% 3 of 8 electrograms ≥70% match

Boston Scientific 24-40 msec, duration 2.5 sec 9% Not programmable†

St. Jude 80 msec 150 msec 5 of 8 electrograms ≥60% match

*Less strict values are required for patients taking type I or III antiarrhythmic drugs. †3 of 10 electrograms with Feature Correlation Coefficient greater than threshold. ICDs, Implantable cardioverter-defibrillators; NID, number of intervals to detect VT; SVT-VT, Supraventricular tachycardia–ventricular tachycardia.

Ventricular fibrillation may be undersensed because of combinations of programming (sensitivity, rate, or duration), low-amplitude EGMs, rapidly varying EGM amplitude, drug effects, and postshock tissue changes. At present, the most common causes of VF undersensing are drug or hyperkalemic effects that slow VF into the VT zone, ischemia, and rapidly varying EGM amplitude48,71 (Fig. 3-81). ICDs that adjust dynamic range based on the amplitude of the sensed R wave (Boston Scientific) may be the most vulnerable to rapidly varying EGM amplitude (Fig. 3-82). ICD features that reduce sensitivity to prevent T-wave oversensing may also result in undersensing of VF.60 Prolonged ischemia from sustained VT slower than the VT detection interval may cause deterioration of signal quality, resulting in undersensing of VF. Lead, connector, or generator problems may also manifest as undersensing.

110

SECTION 1  Basic Principles of Device Therapy

Near field V

VF 158

Pre-attempt EGM (10 sec max) Initial Detection VF Zone Pre-attempt AVS Rate 176 bpm

VF 248

Far field V

VF 253 Epsd Grad 1

VF 163

VF 265

VF 163

VF 263 VF 155

VF 280

VF 155

VF 270

VF 270 VF 168

VF 163

VF 265

VF 158

VF VF 283 270 VF VF 163 165

VF 278 VF 143

VF 303 VF 135

VF 305 VF 165

VF 280 VF 168

VF 278 VF 165

VF 290 VF 165

VF 278

VF 175

VF 278

VF 160

VF 300

VF 163

VF 298

VF 168

–– 533

VF 373 –– 98

VF 138 VS 465

VS 335 Dvrt

Chrg

A

End of Episode 00.25

Pre-attempt EGM (10 sec max) Initial Detection Pre-attempt AVS Rate 218 bpm

25 mm/s

VF –– 228 VF 165 230

VS 365

VS 478

VF 140

VS 345

VF 143

VS 345 VF 145

VS 353

VF 153

VS 353

VF 145

VT 310

VF 170

VS 355

VF 145

VS 368

VF 138

VT 315

VF 135 VS 425

VT 310

VF 173

VS 345

VF 145

VF 158 VS 355

VS 368

VS 543

VS 625

VS 578

VF 230

VF 228 Detct Chrg

VF 218

VF 190

VF 188

VF VF VF 135 223 225 VF VF VF VF 188 258 155 195

VS 383

VF 288

VF 243

VF 188

VF 178

VF 280

VF 218

VF 243

–– 170

Chrg

–– 223

VF 170

1

VF 260

VF 180

Shock

Attempt

VF 230

Attempt Type Elapsed Time Therapy Delivered

VF 208

14J Biphasic 00:01 VF Shock 1

Far field V

Pre-attempt EGM (10 sec max) Initial Detection Pre-attempt AVS Rate 60 bpm

Near field V

–– 1363

VS 1573

B Figure 3-81  Underdetection of ventricular fibrillation (VF) caused by hyperkalemia (potassium level, 6.7 mg/dL) in the setting of chronic amiodarone therapy. A, Near-field and far-field ventricular (V) EGM and event markers of a Boston Scientific Prizm VR ICD. The top panel shows a sine-wave ventricular tachycardia (VT) in the far-field channel, which is detected as VF because local EGMs on the near-field channel are doublecounted. At right of upper panel, four intervals greater than the programmed VF detection interval of 316 msec (190 bpm) result in an aborted shock. The lower panel shows persistence of VF after the shock is aborted. Too few intervals are sensed in the VF zone to permit detection of VF. The patient was resuscitated by external shock. B, ICD system testing after correction of hyperkalemia. Induced VF is sensed reliably and is terminated by an ICD shock. The near-field EGM is narrow, indicating that double EGMs present during the clinical arrhythmia were caused by functional conduction block. (Courtesy Dr. Felix Schnoell.)



3  Sensing and Detection

111

ECG

RA

RV (Tip-Coil)

(AS) 1343

PVC 303 VF 265

(AS) 1268

(AS) 1215 PVC 580

VS PVC 305 310

PVP

PVC 283

PVC 483

PVP

(AS) 1185

AS 1158

PVC VP-MT 320 VF 523 170 PVP

VP 555

PVC PVC 288 VF 330 273 PVP

VS 31

PVP

Figure 3-82  Undersensing of ventricular fibrillation (VF) despite normal R wave in sinus rhythm (18.5 mV). Programmer strip recorded at implantation testing of a Boston Scientific Prizm ICD shows electrocardiogram (ECG), right atrial (RA) EGM, and integrated bipolar right ventricular (RV) sensing EGM (Tip-Coil) during implantation testing. The VF EGMs have highly variable amplitudes, resulting in undersensing of low-amplitude EGMs immediately after high-amplitude ones (arrows). Intermittent ventricular pacing (VP) introduces postpacing blanking periods. Slow Automatic Gain Control may contribute to this type of undersensing. AS, atrial sensed event; PVC, premature ventricular complex; PVP→, extension of postventricular atrial refractory period (PVARP); VP, ventricular paced event. (Modified from Dekker LR, Schrama TA, Steinmetz FH, et al: Undersensing of VF in a patient with optimal R wave sensing during sinus rhythm. Pacing Clin Electrophysiol 27:833-834, 2004.)

ICD INACTIVATION If detection is programmed “off ” for surgery using electrocautery, reprogramming must be performed at the end of the procedure, a fact that is easily forgotten, especially with outpatient surgery. One study reported an unexplained 11% annual incidence of transient suspension of detection.151 This problem is addressed in Medtronic ICDs with an audible patient alert that sounds if programmed detection or therapy is “off ” for longer than 6 hours. VENTRICULAR TACHYCARDIA SLOWER THAN PROGRAMMED DETECTION INTERVAL In most ICD patients, VT with cycle lengths greater than 400 to 450 msec are tolerated well, but repeated inappropriate therapies are not. SVT-VT discrimination algorithms, except those in Sorin models,71,134,152 deliver fewer inappropriate therapies if the VT detection interval is programmed to a shorter cycle length, simply because fewer SVTs are evaluated. However, slow VT can be catastrophic in patients with severe LV dysfunction or ischemia,153 and a long VT detection interval is important in some patients with advanced heart failure (Fig. 3-83). The VT detection interval should be increased if antiarrhythmic drug therapy is initiated, particularly with amiodarone or a sodium channel–blocking (type IA or IC) drug.15 It may be prudent to measure the cycle length of induced VT at electrophysiologic testing after initiation of drug therapy.154 However, spontaneous VT often is slower than induced VT.155 SVT-VT DISCRIMINATORS The SVT-VT discriminators may prevent or delay therapy if they misclassify VT or VF as SVT.114,116,119,156 Discriminators that reevaluate the rhythm diagnosis during an ongoing tachycardia (e.g., stability, most dual-chamber algorithms) reduce the risk of underdetection of VT compared with discriminators that withhold therapy if the rhythm is

not classified correctly by the initial evaluation (e.g., onset, chamber of origin algorithms). The minimum cycle length for SVT-VT discrimination should be set to prevent clinically significant delay in detection of hemodynamically unstable VT. PACEMAKER-ICD INTERACTIONS Interactions between ICDs and separate pacemakers have become rare since ICDs incorporated dual-chamber bradycardia pacing. Potential interactions have been reviewed.157-159 The principal interaction that may delay or prevent ICD therapy is oversensing of high-amplitude pacemaker stimulus artifacts. If this occurs during VF, repetitive autoadjustment of sensitivity may prevent detection of VF. INTRADEVICE INTERACTIONS In intradevice interactions, bradycardia pacing features of dualchamber ICDs may interact with and impair detection of VT or VF.54 During high-rate, atrial or dual-chamber pacing, sensing may be restricted to short periods of the cardiac cycle because of the combined effects of ventricular blanking after ventricular pacing and crosschamber ventricular blanking after atrial pacing, which is needed to avoid crosstalk. If a sufficient fraction of the cardiac cycle is blanked, systematic undersensing of VT or VF may occur. When pacing and blanking events occur at intervals that are multiples of a VT/VF cycle length, ventricular complexes are repeatedly undersensed, delaying or preventing detection52-54 (Fig. 3-84). Intradevice interactions have been reported most frequently with the use of the rate-smoothing algorithm.52-54 This algorithm is intended to prevent VT/VF initiated by sudden changes in ventricular rate.160 Rate smoothing prevents sudden changes in ventricular rate by pacing both the atrium and the ventricle at intervals based on the preceding (baseline) R-R interval. As an unintended consequence, it may prevent sensing of VT/VF in some patients, because rate smoothing introduces

112

SECTION 1  Basic Principles of Device Therapy

Atrial EGM RV Coil-Can

Interval (ms) 1800 VF Detected

1720

3 5 0

A P T S

6 8 0

VV SP

4 2 0

V S

F F F F F F S S S S S S 2 1 1 2 1 1 1 8 2 0 2 6 0 0 0 0 0 0

4 6 0

V S

2 1 0

1600

A P

1760 F S

7 2 0

VV SP

4 7 0

V S

1 7 0

F S

4 7 0

V F S S 1 3 0

4 6 0

V S

2 1 0

F S

1 8 0

F S

1 8 0

F S

2 3 0

F S

2 2 0

F S

F F F F V V S S S S S S 2 1 1 1 4 2 5 3 8 8 4 0 0 0 0 0 0 0 VF Rx 1 Defib

5 1 0

V S

4 6 0

V S

1400 1200 1000

3.7 minutes

800 600

TDI

400

FDI 200 0 600

400

0

200

Time Before VF Detection (Seconds)

VF Detected

Figure 3-83  Ventricular tachycardia (VT) slower than the programmed detection interval. The lower panel is a “Flashback Interval” plot of R-R interval cycle lengths before detection of ventricular fibrillation (VF), which occurs at the right side of each panel. The interval number before detection is plotted on the abscissa, and the corresponding interval is plotted on the ordinate. Horizontal lines indicate the VT detection interval (TDI) of 400 msec and the VF detection interval (FDI) of 320 msec. Shortly after the 500th interval preceding detection, regular tachycardia begins abruptly. The constant cycle length indicates reliable ventricular sensing. Atrial flashback intervals (not shown) demonstrated atrioventricular (AV) dissociation. This VT is not detected despite reliable sensing, because the cycle length is greater than the programmed TDI. VT persists for 3.7 minutes until approximately interval 280 before detection, when sensed intervals become highly variable. This indicates degeneration of the rhythm to VF with undersensing that delays detection. During VT and VF, atrial Flashback Intervals (not shown) indicated lower rate limit bradycardia pacing at 40 bpm (1500 msec). The upper panel shows stored atrial and far-field ventricular EGMs immediately before detection with atrial and ventricular channel showing event markers. Specific undersensed EGMs cannot be identified because the rate-sensing EGM was not recorded. However, long sensed R-R intervals ending with ventricular sense (VS) markers indicate undersensing and correspond to long intervals in the upper panel. VF Rx 1 Defib at lower right (arrow) denotes “VF detected.” AP, Atrial paced event; FD, VF detected; FS, intervals in VF zone. (From Swerdlow C, Friedman P: Advanced ICD troubleshooting. Part I. Pacing Clin Electrophysiol 28:1322-1346, 2005.)

repetitive postpacing blanking periods. The algorithm applies rate smoothing to baseline intervals independent of their cycle length, including intervals in the VT or VF zones. Intradevice interactions that result in delayed or absent detection of VT/VF are most common and most dangerous when VT is fast. The parameter interrelationships that result in delayed or absent detection of VT/VF are complex and difficult to predict, but usually elicit a programmer warning. Generally, aggressive rate smoothing (a small, allowable percentage change in RR intervals), a high upper pacing rate, and a long and fixed AV interval favor undersensing and should be avoided. If rate smoothing is required, the AV delay should be dynamic, the upper rates should be 125 bpm or less, and parameter combinations that result in warnings should be avoided. This programming reduces, but does not eliminate, the risk of undersensing.52-54

Detection of SVT/VT as Diagnostic Tool and Basis for Atrial Antitachycardia Pacemakers and Atrial ICDs Detection of SVT and VT provides diagnostics that are useful for pacemaker management.161 Atrial antitachycardia pacing and shocks provide a therapeutic option for some patients with paroxysmal atrial

tachycardia/flutter (AT) and atrial fibrillation (AF). Appropriate delivery of this therapy requires accurate detection and discrimination AT and AF independent of ventricular rate. PACEMAKER DIAGNOSTICS Monitoring for Ventricular Tachycardia Some pacemakers have ventricular high-rate diagnostics that trigger EGM and marker storage of nonpaced rhythms that are faster than a specific rate. Triggers based on ventricular rate alone store episodes that may be rapidly conducted SVT, ventricular oversensing, or VT. Studies of these diagnostics demonstrate the importance of stored EGMs to confirm the diagnosis, because the false-positive detection rate is high.162 Newer dual-chamber pacemakers and antitachycardia pacemakers incorporate tachyarrhythmia detection algorithms that are identical or substantially similar to those in dual-chamber ICDs (Fig. 3-85). The performance of these algorithms may be influenced by differences in pacemaker and ICD atrial sensing characteristics (fixed atrial sensitivity, automatically adjusting sensitivity, or intermediate case) and atrial blanking periods (Fig. 3-86). Even with accurate atrial sensing, Bayes theorem predicts that the fraction of false-positive VT detection in patients with pacemakers is large, because the incidence of true VT is low.



3  Sensing and Detection

Surface ECG Atrial EGM Vent EGM

PABP VBP

AS 1875 VS 935

VP 350

VS 350

VP 350

VP 350

AP 863 VF 283 VF 280

AS 970

VP 350

VP 350

VP 350

Surface ECG Atrial EGM Vent EGM

25-FEB-00 19:30

AP 268 VP 560

VP 570

PVC 448

VP 230

VF 308

VT 415

AP 875

AP 723

VF 305

VP 560

AP 863 VF 283

AP 863 VS 575

VF 283

VF 285

AP 1118

VF 280

VP 570

VF 278

AP 860

AP 260 VP 560

VF 298

VF 285

VP 570

AP 850 VP 560

AP 1135

AP 865

VF 288

VP 580

GUIDANT

VP 570

VF 280

Surface ECG Atrial EGM Vent EGM

25-FEB-00 19:30

VENTAK PRIZM

VS 575

VF 283

VP 270

VP 250

Surface ECG Atrial EGM Vent EGM

25-FEB-00 19:30

AP 1630 VP 350

AP 863

AP 863 VF 293

25 mm/s

VENTAK PRIZM

VP 560

VF 280

113

AP 870 VF 300

AP 855 VF 275

AP 853 VS 573

VF 278

AP 858

AP 260

VF 283

AP 560

VF 788

VP 570

External rescue

AP 878 VP 570

VP 580

VF 298

AP 880 VF 290

AP 8737 VS 583

VF 288

VF 293

AP 260 [VS]

VP 560

[AS] (AS) 733 VF 308

Figure 3-84  Failure to detect ventricular tachycardia (VT) caused by intradevice interaction. The rate-smoothing algorithm introduced atrial and ventricular pacing complexes with associated blanking periods that prevented detection of VT during postimplantation testing. An external rescue shock was required. Top to bottom, Surface electrocardiogram (ECG), atrial EGM, ventricular EGM, and event markers. At top, VT is induced by programmed electrical stimulation with a drive cycle length of 350 msec and premature stimuli at 270, 250, and 230 msec (intervals labeled next to event markers). The first sensed ventricular event occurs 448 msec after the pacing drive (PVC 448). The rate-smoothing algorithm drives pacing to prevent a pause after the premature ventricular complex (PVC), labeled AP↓ 1638. A ventricular paced event does not follow the first AP↓ because a ventricular event is sensed (VT 415). Subsequent rate smoothing generated atrial and ventricular pacing pulses (indicated by AP↓ and VP↓ markers, respectively). The resultant postpacing blanking periods are shown on the figure as horizontal bars. PABP denotes cross-chamber (postatrial pacing) ventricular blanking period. VBP denotes same-chamber (postventricular pacing) blanking period. Together, they prevent approximately four of every six VT complexes from being sensed. Because the VT counter must accumulate 8 of 10 consecutive complexes in the VT zone for detection of VT to occur, VT is not detected. (From Swerdlow C, Friedman P: Advanced ICD troubleshooting. Part II. Pacing Clin Electrophysiol 29:70-96, 2006.)

V-V

A-A

VTM=400 ms

Interval (ms)

Detection

1500 1200 900 600

Term. 9 sec

AEGM VEGM

400 A P

200

230

225

220

215

210

Time (sec)

25

0

5

V S

0

A R

1180 5 6 0

1020

V 3 V 4 V S 7 S 5 S 0 0

6 9 0

9 8 0

A S

A R

9 4 0

A R

9 6 0

A R

9 2 0

A R

8 9 0

A R

8 7 0

A R

EGM for

V 5 V 3 V 3 V 3 V 3 V3 V 3 V3 V 3 V 3 V 3 V 3 V 3V 3 V 3 V3 V 3 V S 0 S 8 S 9 S 7 S 8 S3 S 4 S5 S 4 S 6 S 6 S 7 S 4S 6 S 9 S3 S 6 T 0 0 0 0 0 0 0 0 0 0 0 0 0 0 0 0 0 VTM

A V-V

A-A

VTM=400 ms

Interval (ms)

Detection

1500 1200 900 600

Term. 1.4 min

AEGM

0.5mV

VEGM

400 .2000

200 230

225

220

215

210

Time (sec)

25

0

5

0

V S

5 6 0

V S

5 1 0

A b

.2000

A b

EGM for 1

V V V V V V V V V V V V V V V V V V S 4 S 3 S 3 S 3 S 3 S 3 S 3 S 3 S 3 S 3 S 3 S 3 S 3 S 3 S 3 S 3 S 3 T 1 7 8 7 9 8 7 7 6 8 7 9 5 0 0 8 9 0 0 0 0 0 0 0 0 0 0 0 0 0 0 0 0 0 VTM

B Figure 3-85  Pacemaker diagnostics for monitoring of ventricular tachycardia (VT), showing stored intervals, atrial electrogrm (AEGM), ventricular electrogram (VEGM), and markers from a patient with a Medtronic EnRhythm DR pacemaker. The left panels show the A-A and V-V intervals plotted against time before VT detection, which requires 16 consecutive intervals at less than the VT monitoring interval of 400 msec (horizontal line). The right panels show dual-chamber EGM and markers. A, Onset of spontaneous VT occurs with several premature ventricular complexes (PVCs) and sudden acceleration of ventricular rate with little or no change in A-A intervals (left panel). The VT cycle length is initially 340 msec and progressively lengthens to greater than 400 msec after several seconds. After 8 consecutive intervals of 400 msec or longer, the ICD declares the episode to be “terminated” despite clear atrioventricular (AV) dissociation and ongoing VT (right-hand side of interval plot). The duration of VT is reported as 9 seconds, representing the time that the RR intervals remained less than 400 msec after initial detection. Dual-chamber EGM and markers leading up to detection of VT indicate appropriate sensing in both chambers and AV dissociation (right panel). B, Rapidly conducted atrial fibrillation with severe atrial undersensing reported as VT. The interval plot shows fast and irregular R-R intervals and long A-A intervals (1500 msec), with sporadic A-A intervals shorter than 200 msec. Dual-chamber EGMs indicate appropriate ventricular sensing but severe atrial undersensing of small-amplitude atrial fibrillation, resulting in false-positive detection of VT. AB, Atrial sensed event in postventricular atrial blanking period; AR, atrial refractory sense; AS, atrial sensed event; VS, ventricular sensed event; VT, VT detected.

114

SECTION 1  Basic Principles of Device Therapy

False-positive VT/VF episodes (%)

100%

80%

Patients with AT/AF history and no VT/VF history

60%

40%

20%

Secondary prevention ICD patient population

0% Higher incidence of true VT/VF relative to SVT Figure 3-86  Rate of inappropriate ventricular tachycardia/fibrillation (VT/VF) detection depends on the patient population. Bayes theorem predicts that patients with a higher incidence of true VT/VF relative to supraventricular tachycardia (SVT) will have fewer inappropriate detections than patients with a lower incidence of VT/VF. The receiver operator curve shows the Bayes theorem prediction of detection performance of an algorithm with a fixed sensitivity and specificity. The curve plots percentage of false-positive detections of VT/VF on the ordinate and estimated incidence of true VT/VF on abscissa. Two data points are plotted from clinical studies of two different patient populations with implanted ICDs running the same VT/SVT discrimination algorithm (Medtronic GEM DR68 and AT500 ICDs* with PR Logic). (*Willems R, Morck ML, Exner DV, et al: Ventricular high-rate episode in pacemaker diagnostics identify a high-risk subgroup of patients with tachy-brady syndrome. Heart Rhythm 1:414421, 2004.)

Monitoring for Atrial Tachycardia and Atrial Fibrillation Traditional pacemaker diagnostics for atrial high-rate episodes and automatic mode switches have been validated as a means of monitoring AT/AF.163,164 Diagnostic accuracy of these methods depends on the specific detection algorithm used and the accuracy of atrial sensing. Mode-switching algorithms and their diagnostics have been used as a surrogate marker for AT/AF detection in patients with dual-chamber pacemakers.161,163 A Holter monitoring study of 40 patients with bipolar pacemakers demonstrated that mode-switching diagnostics appropriately identified 53 of 54 true AT/AF episodes (98.1%) with only one short (13-second) false mode-switching episode.163 Other studies have reviewed stored pacemaker EGMs and found high percentages of inappropriate mode switching caused by atrial oversensing.162,165 Most modern pacemakers and ICDs store atrial EGMs for mode-switching events; and important clinical decisions, such as starting anticoagulation, should be based on analysis of atrial EGMs rather than mode-switching events. Further, absence of mode-switching events is not equivalent to absence of AT/AF in pacemakers with long PVAB periods or low atrial sensitivity. The AT/AF detection algorithms in some pacemakers and ICDs provide substantial additional information, including atrial EGMs, classification of AT versus AF, and integrated displays of data from multiple episodes. These include histograms of episode duration and ventricular rate during AT/AF that may be of value in managing atrial antiarrhythmic drugs and AV nodal blocking drugs (Fig. 3-87). The former histogram, combined with patient-activated ICD interrogations, may permit discontinuation of anticoagulation in some patients. A prospective study of 2486 patients with one or more stroke risk factors and an implantable pacemaker or ICD with continuous AT/AF monitoring evaluated the risk of thromboembolic events as a function of AT/AF burden (hours of AT/AF per day). More than 5.5 hours of AT/AF on any day doubled a patient’s risk of stroke.166 Other studies have demonstrated similar increased risk associated with monitored AT/AF episodes.166,167 Computer modeling predicts that home monitoring of patients with automatic ICD-generated alerts for detected AT/AF may reduce the risk of stroke.168 A multicenter randomized trial is designed to test the hypothesis that initiation and withdrawal of oral anticoagulation therapy guided by continuous AT/

AF monitoring reduces the risk of stroke, embolism and major bleeding compared with conventional management.169 DETECTION OF ATRIAL TACHYCARDIA AND ATRIAL FIBRILLATION Principles Bradycardia pacemakers must detect AT/AF rapidly so that mode switching will avoid uncomfortable pacing at the upper rate limit. To determine the atrial rate and rhythm accurately, atrial blanking must be minimized. As long as the atrial rate exceeds the upper tracking limit, accurate determination of atrial rate and rhythm is not important. Ventricular ICDs must discriminate VT from rapidly conducted AF quickly, to permit rapid detection of hemodynamically compromising VT. Discrimination of AT versus AF is less important. In contrast, atrial antiarrhythmic pacemakers and ICDs must detect AF with high specificity to minimize painful and potentially proarrhythmic therapy. Rapid detection is not important, because AF usually is clinically stable and may terminate spontaneously after hours to days. Therefore, atrial antiarrhythmic ICDs should be capable of permitting therapy for long-duration AT/AF while withholding therapy from self-terminating AT/AF. To achieve this goal, they must be capable of detecting AF continuously for extended periods, to discriminate between repetitive self-terminating episodes and persistent episodes. Because atrial EGMs in AF have low and variable amplitudes and slew rates, antiarrhythmic ICDs must have high (automatically adjusting) atrial sensitivity and apply algorithms that are tolerant of some atrial undersensing. They must also discriminate between AT and AF to deliver ATP for AT. Detection of AT/AF for Atrial Therapy High specificity in AT/AF detection has been achieved by multistep methods for detection of AT/AF. Initial detection is based on the presence of an atrial tachyarrhythmia and absence of VT. This is achieved by a measure of atrial rate combined with either comparison of atrial and ventricular rates (Boston Scientific)170 or use of A-V patterns to identify N:1 rhythms (Medtronic) (Fig. 3-88). The trade-off between atrial undersensing and oversensing of FFRWs that applies to dualchamber ICD algorithms is even more important when considering



115

3  Sensing and Detection VENTRICULAR RESPONSE (IN AF)

Long-term trends May 00

Jul 00

Program/interrogate

Sep 00

Nov 00

P

Drug change

Jan 01

Mar 01

P

# beats

May 01 I

Sotalol

CV/ablation/other AT/AF patient check AT/AF total hrs/day

AT/AF episodes/day

24 20 16 12 8 4 0 .25 20 15 10 5 0

ATP change

% ATP success

.25 20 15 10 5 0

% Pacing/day

(0%)

80-100 bpm

3

(0%)

100-120 bpm

124

(4%)

120-140 bpm

175

(6%)

140-160 bpm

344

(12%)

160-180 bpm

912

(33%)

180-200 bpm

880

(32%)

>200

319

(12%)

100

100 75 50 25 0

ATP ON 80

Antici-Pace Change % Pacing/day — A.total — A. ARS/APP

0

DATA - AT/AF EPISODE DURATION HISTOGRAM

Episodes (%)

Treated AT/AF episodes/day

<80 bpm

100 75 50 25 0 100 75 50 25 0

60 40 20 0 < 1 min 10 min 1 hr 4 hr 12 hr 24 hr 72 hr

May 00

Jul 00

Sep 00

Nov 00

Jan 01

Mar 01

May 01

>

Episode duration

Figure 3-87  Diagnostic data stored in atrial antiarrhythmic pacemakers and ICDs assist medical management of atrial arrhythmias. Data are taken from three patients with Medtronic AT500 pacemakers. Left panels, Long-term trend data over 1 year. Panels from top to bottom show total hours per day in atrial tachycardia/flutter (AT) or atrial fibrillation (AF), total AT/AF ICD-detected episodes per day, number of daily episodes treated by antitachycardia pacing (ATP), percentage of ATP therapies classified as “successful” by the ICD, percent atrial pacing, and percent ventricular pacing. ATP ON denotes activation of atrial ATP. Note that, although the many episodes of AT/AF are treated and this treatment is usually classified as successful, there is no detectable change in the number of hours per day of AT/AF until sotalol therapy is initiated. Sotalol decreases the total hourly “burden” of AT/AF more by shortening episodes rather than by preventing them. Percent ventricular pacing is high throughout but increases after sotalol treatment to almost 100%. Top right panel, Ventricular rate during AT/AF as percentage of time spent in AT/ AF; 77% of intervals are shorter than intervals corresponding to 160 bpm, and 12% are shorter than intervals corresponding to 200 bpm, indicating inadequate control of ventricular rate. Lower right panel, Histogram of durations of AT/AF episodes since the last follow-up visit. This patient had 10 episodes of AF since the last visit, but only one episode lasted longer than 10 minutes.

detection of AT/AF for atrial therapy. Minimization of atrial blanking is important to prevent undersensing of AF, and rejection of FFRWs is important to prevent inappropriate detection of AT/AF (see earlier Atrial Sensing in Dual-Chamber ICDs and Atrial ICDs). Once initial detection occurs, the episode timer begins, initiating a sustained-detection mode in which AT/AF remains detected despite a moderate degree of undersensing (Fig. 3-89). When this timer expires, atrial therapy is delivered if AT/AF still persists (Fig. 3-90). Atrial episodes must be interrupted if true VT occurs. The final step in detection is discrimination of AT from AF or, alternatively, determining whether the rhythm is likely to respond to atrial ATP. As shown in Figure 3-88, Medtronic atrial ICDs track dynamic changes in atrial tachyarrhythmia (e.g., transitions between AT and AF) based on the median rate and regularity of the atrial rhythm. Boston Scientific atrial ICDs use a more complex algorithm based on maximum rate, standard deviation, and range of the 12 most recent atrial cycle lengths to plot a point in a three-dimensional space. A decision boundary divides the space into two regions: faster/irregular

atrial cycle lengths (AF) and slower/regular cycle lengths (AT). Classifications are made on a sliding window of 12 consecutive cycles until the end of the episode is reached.170,171 Typically, there may be one timer for ATP and one for shock therapy. In newer Medtronic ICDs, the “reactive antitachycardia pacing” algorithm resets the ATP timer to permit repeat attempts at ATP if changes in atrial rate or atrial rhythm regularity are detected, or after a preprogrammed duration of sustained AT/AF. Changes in rate or regularity of the atrial rhythm are classified as a shift to a new rhythm, which may be more susceptible to ATP termination (Fig. 3-91). Detection Considerations for Atrial Shocks In addition to permitting therapy after hours of continuous AF, atrial ICDs may withhold therapy if the episode duration is sufficiently long that patients may be at risk for thromboembolism if they are not adequately anticoagulated, typically 24 hours. R-wave-synchronized atrial shock therapy is restricted to R-R intervals greater than 400 to 500 msec. This minimizes the risk that a therapeutic atrial shock can

116

SECTION 1  Basic Principles of Device Therapy

Median P-P interval

Auto discrimination zone

AF detection zone

Atrial blanking (100 mS)

AFDI: 280 mS AFDI: 320 mS

ATDImin: 180 mS AT detection zone

AF/AT evidence counter

Atrial fibrillation

P R

A

B

D Count: +1 Count: 28

+1 29

+1 30

A=A range

−1 29

+1 30

+1 31

+1 32

SD

Figure 3-88  Criteria for detection of atrial tachyarrhythmias. A, Medtronic atrial therapy ICDs (Jewel AF, GEM III AT, AT500, and EnRhythm DR) use a combination of atrial cycle length and atrial/ventricular (A : V) patterns to detect atrial fibrillation and atrial flutter. The atrial tachycardia (AT) and atrial fibrillation (AF) detection zones are based on median atrial cycle length (12 P-P intervals) and may overlap. The overlap region is the “autodiscrimination” zone, where regularity of the PP intervals determines whether rhythm classification is AT (regular P-P intervals) or AF (irregular P-P intervals). Detection of AT or AF requires that (1) the median P-P interval is in one of the detection zones and (2) the rhythm is N:1 as determined by the AF/AT evidence counter. The AF/AT evidence counter is an up-down counter (minimum value, 0; maximum value of detection threshold, +15). The counter increments by 1 on each ventricular event if there are two or more atrial events and there is no pattern-based evidence of far-field R-wave oversensing on the atrial channel (see Fig. 3-25, A), and it decrements by 1 if there is strong evidence of lack of N:1 rhythm (e.g., two consecutive 1 : 1 beats). Isolated 1 : 1 beats contribute AF/AT evidence if they are preceded by a confirmed N:1 beat. The AF/AT counter detection threshold is between 24 and 32, depending on the specific ICD. AFDI, Atrial fibrillation detection interval; mS, milliseconds. B, Atrial Rhythm Classification (ARC) algorithm (Boston Scientific) discriminates AF from atrial flutter (AFL) based on the atrial rate and two measurements of variability of atrial rate: the range of the atrial intervals (i.e., the difference between the longest and shortest A-A intervals [A-A range]) and the standard deviation of A-A intervals (SD). The values of these three variables define a point in a three-dimensional space. A curved surface separates the AF region from the AFL region. Points in the AF region have higher atrial rates, a higher range of A-A intervals, and a greater standard deviation of A-A intervals. (From Morris MM, KenKnight BH, Lang DJ: Detection of atrial arrhythmia for cardiac rhythm management by implantable ICDs. J Electrocardiol 33(Suppl):133-139, 2000.)

be delivered into the vulnerable period of the preceding cardiac cycle, but it may prevent shock therapy for rapidly conducted AF. Early or immediate recurrence of AF after shock is an important clinical problem in patients with atrial ICDs. Early recurrence of AF before postshock redetection of sinus rhythm will result in incorrect classification of shock success (Fig. 3-92).

Sensing and Detection Using the Subcutaneous Electrocardiogram IMPLANTABLE SUBCUTANEOUS MONITORING Implantable loop recorders (ILRs) offer patient-activated recordings of subcutaneous ECGs as well as arrhythmia detection algorithms to automatically trigger storage for asystole, bradycardia/tachycardia, and AF. Two sensing electrodes are built into the ILR shell and record a single-lead bipolar ECG that is retrieved with a programmer or remote-monitoring ICD. ECG signals are stored in a circular buffer. Multiple, automatically triggered events may be stored. The memory buffer may also be activated by a hand-held system to record the ECG preceding and following activation. Automated triggers for storage of arrhythmic events in ILRs require robust R-wave sensing and detection algorithms based on RR intervals. The amplitude of R waves recorded by an ILR averages about 0.3 mV, versus 1 to 2 mV recorded in the precordial leads of a surface ECG.

Unlike EGMs, the amplitude of subcutaneous ECGs increases over time to about 0.35 mV at 4 to 6 months, then remains stable.172 R-wave sensing in an ILR uses the same principles as shown in Figure 3-13 with band-pass filtering tuned for lower frequencies and either fixed or auto-adjusting sensitivity. Sensing problems were common with early models that used a fixed sensing threshold, including undersensing caused by abrupt, unexplained decreases in R-wave amplitude and ICD amplifier saturation, as well as oversensing of T waves, myopotentials, and other sources of noise.173-175 Sensing enhancements such as better noise rejection and auto-adjusting threshold reduce both false-positive triggers caused by oversensing and problems related to undersensing.176 Figures 3-93 and 3-94 show examples of rhythms recorded by ILRs. Monitoring of AF burden has been a recent addition to the Medtronic Reveal XT ILRs. The algorithm detects the irregularly irregular ventricular response to conducted AF by analyzing the difference between the present and previous R-R intervals (δRR). Pattern analysis is performed using Lorenz plots, or scatter plots of δRR(i) vs. δRR(i-1) over a duration of 2 minutes.177 The algorithm can distinguish patterns associated with AF and regular AT from sinus rhythm, as shown in Figure 3-95. This algorithm was shown to have sensitivity of 96.1% and specificity of 85.4% for identifying patients with or without AF in a study of 247 patients using Holter ECGs as a “gold standard.”178 AF burden as measured by this algorithm was highly correlated with the Holter standard.



3  Sensing and Detection 13:26:26-1

117

AF - - MS delta

ECG

AEGM

AP

AP

AR FS FS FS FS FS

FS FS FS FS FS FS FS FS FD FD FD

Marker VP PP Median Atrial Evidence

VP 750 0

VS 750 1

VP 750 2

VS 200 30

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VP 210

VP 210 *A

A

ECG

AEGM 1 sec Marker

FD FD FD FD FD FD FD FD FD FD FD FD FD FD FD FD FD FD FD VS VS PP Median Atrial Evidence

VS 170

VS 170

VP 170

VS 180 *A

FD FDFDFD FD FDFD FD FD FD FD VP 170

VS 170

VS 170

VS 170 *A

B Figure 3-89  Holter recording illustrates continuous detection of atrial fibrillation (AF). A, Onset of AF and initial detection of AF after 32 ventricular events (arrow). Double atrial markers (FS) change to triple markers (FD), and the symbol *A appears on the atrial evidence channel to indicate AF episode in progress. B, Rhythms after 5 hours of continuous recording. AP, Atrial paced event; VP, ventricular paced event; VS, sensed R wave. (From Swerdlow CD, Schsls W, Dijkman B, et al: Detection of atrial fibrillation and flutter by a dual-chamber implantable cardioverterdefibrillator. For the Worldwide Jewel AF Investigators. Circulation 101:878-885.)

SUBCUTANEOUS ICDS Investigational subcutaneous ICD (S-ICD) systems, with no transvenous or epicardial leads, are undergoing clinical trials. They rely on subcutaneous ECGs for sensing and detection of VF. The present S-ICD (Cameron Health) uses a 69-cc pulse generator placed in the left midaxillary line over the lower thorax. A single subcutaneous lead is tunneled medially to the left parasternal position and then superiorly along the sternum. It includes two widely spaced sensing electrodes and a defibrillation coil (Fig. 3-96). Sensing The S-ICD automatically selects the best sensing vector based on QRS amplitude and signal-to-noise ratio from among three choices: (1) proximal electrode ring on the lead to the active surface of the pulse generator, (2) distal electrode ring on the lead to the pulse generator, or (3) distal ring to proximal ring. The S-ICD automatically selects one of two amplifier gains (±2 mV unless QRS amplitude is clipped at this setting; otherwise, ±4 mV) and the sensing vector least likely to oversense T waves (Fig. 3-97; see also Fig. 3-96). As in ICDs with intracardiac leads, the S-ICD uses automatic adjustment of sensitivity. Sensitivity increases when rapid rates are detected. In addition to this event detection phase of sensing, the S-ICD has a certification phase that examines sensed events and classifies them as certified QRS complexes or as suspect events. This ensures that accurate ventricular rate is passed to the detection algorithm. An event is

classified as “suspect” if its pattern or timing suggests oversensing, according to specific oversensing rules intended to reject myopotentials, R-wave double-counting, and T-wave oversensing. Sensing and Tachyarrhythmia Detection The detection algorithm (decision phase) examines all certified events and calculates a continuous, running, “four R-to-R interval average” (4 RR average), which is used as a measure of ventricular rate. Rate-based therapy zones include a mandatory rate-only VF (“shock”) zone and an optional VT (“conditional”) zone. In the VT zone the S-ICD uses stable QRS morphology and QRS width to discriminate VT from SVT, as well as varying QRS morphology to identify polymorphic VT. The “START” study compared SVT-VT discrimination using tachyarrhythmias recorded at ICD or S-ICD implant.124 The S-ICD detection algorithm had equivalent sensitivity for VT to the best ICD algorithm (100%), and insignificantly better specificity for rejection of SVT (100% vs. 92%). The S-ICD detection process begins when the “4 RR average” enters either therapy zone. VT or VF detection is a two-step process. The initial detection step uses X of Y counting (18/24 for the first therapy), and the “persistence” step requires that the X of Y criterion remain fulfilled for additional consecutive intervals (nominally two, but automatically increased if shocks are aborted). Capacitor charging begins if persistence is met and if the last two certified sensed events are in a therapy zone. Shock confirmation, redetection, and episode termination function similar to these features in ICDs with intracardiac leads.

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Onset

Detection 12 sec

First Rx

Termination 19 sec

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1 2 1 1 1 1 1 1 1 2 2 1 1 2 1 1 2 1 2 1 1 1 1 1 1 2 2 2 8 1 9 9 9 9 9 8 7 1 0 9 9 0 9 9 0 9 0 8 7 6 5 5 5 6 2 2 T 0 T 0 T 0 T 0 T 0 T 0 T 0 T 0 T 0 T 0 T 0 T 0 T 0 T 0 T 0 T 0 T 0 T 0 T 0 A 0 A 0 A 0 A 0 A0 A 0 A 0 T 0 T 0 T D D D D D D D D D D D D D D D D D D D S P P P P P P S F F V S

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Figure 3-90  Atrial tachycardia (AT) detection and therapy after 1 minute of sustained detection. Spontaneous episode of AT from an atrial therapy ICD. Upper left panel, P-P intervals (open squares) and R-R intervals (closed circles) for this episode. Onset of the AT episode is labeled on the interval plot and on the stored marker channel (top right). The interval plot also labels the initial detection of AT (Detection), delivery of the first therapy (First ATP Rx), and ICD recognition of episode termination (Termination). In this case, therapies were programmed to begin 1 minute after initial detection. Top right panel, Dual-chamber EGM markers preceding detection of AT. Dots indicate discontinuous recording between top and bottom panels. Bottom panel, Composite EGM recorded between the atrial tip electrode (Atip) and the ventricular ring electrode (Vring) and dualchamber EGM markers immediately preceding delivery of successful antitachycardia pacing (First Rx, fast AP events). TS, FS, and TF events refer to atrial intervals in the AT, atrial fibrillation (AF), and overlap zones, respectively. TD (atrial marker), Atrial tachycardia detected; VP, ventricular paced event; VS, ventricular sensed event.

2 2 0

T D

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1373

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Figure 3-91  Example of successfully terminated atrial tachycardia/fibrillation (AT/AF). Top panel, Atrial EGM and marker diagram from AT/ AF episode (median atrial interval, 220 msec) that was initially detected and treated unsuccessfully with antitachycardia therapy (ATP). Bottom panel, After 2.1 hours, the atrial rate slowed, and the reactive ATP algorithm recognized the presence of a slower, regular AT/AF rhythm with a median atrial interval of 370 msec. Atrial ATP was delivered and successfully terminated the AT/AF. AP, atrial paced event; TD, AT/AF detected; TS, AT/AF sensed event; VP, ventricular paced event; VS, ventricular sensed event. (From Mainardi L, Sornmo L, Cerutti S: Understanding atrial fibrillation: the signal processing contribution, San Rafael, Calif, 2008, Morgan & Claypool.)



3  Sensing and Detection

119

Lead II

Shock

Atip-Aring

ERAF

Markers A A A A AAA R R S R S RS

V S

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V C V S E S

A A A AC S S S SD

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V S

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A A A A A A AAAA A AA R R RS R RSRSR S RR

V S

V S

V S

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V P

Figure 3-92  Postshock early recurrence of atrial fibrillation (ERAF) recorded at implantation occurs three RR intervals after shock. Because the ICD requires five consecutive postshock beats of sinus or atrial paced rhythm to detect termination of atrial fibrillation (AF) and store intervals, no data were stored for this clinically unsuccessful shock. Lead II of the surface electrocardiogram, the atrial EGM, and the atrial and ventricular channels with event markers are shown. AR, Atrial refractory event; AS, atrial sensed event; CD, charge delivered; CE, charge end; VP, ventricular paced event, VS, ventricular sensed event. (From Swerdlow CD, Schwartzman D, Hoyt R, et al: Determinants of first-shock success for atrial implantable cardioverter defibrillators. J Cardiovasc Electrophysiol 13:347-354, 2002.)

11:32:07

11:32:20

11:32:33

11:32:46

11:32:59

Figure 3-93  Stored EGM recorded by an implantable loop recorder (Reveal, Medtronic) shows about 45 seconds of asystole, some myopotentials, perhaps one depolarization in the middle of the asystole, very small deflections throughout that are probably P waves, and relative bradycardia on the bottom panel. The amplitude of the subcutaneous electrocardiogram is approximately 0.25 mV.

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12:22:47 P

12:23:00

230 bpm (260 msec) Figure 3-94  Stored EGM recorded by an implantable loop recorder (Reveal, Medtronic) shows nonsustained tachycardia at about 230 bpm with duration of 8 seconds. The underlying rhythm is reset by the high-rate segment. The patient did activate this stored strip, as indicated by the dark triangle and the “P” marker.

RR 670 435 900 520 800 1030 700 675 760 650 515 625 925 δRR

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B Figure 3-95  Algorithm for detection of atrial fibrillation (AF) by implantable loop recorder based on pattern of R-R intervals. A, Illustration showing how R-R intervals derived from a short strip of subcutaneous ECG during AF populate the numerical equivalent of a Lorenz plot of ΔRR intervals. Each point in the Lorenz plot is derived from four R waves that provide three R-R intervals and thus a [δRR(i), δRR(i-1)] value. The markers near the top of each R wave indicate the time at which each R wave is sensed. The magnitude of the vector from the origin of the Lorenz plot to each point encodes the irregularity, and the magnitude coupled with phase encodes the incoherence, of changes in R-R intervals. B, ECG tracings (2-minute duration) of AF (top panel), two examples of atrial flutter (middle and bottom panels), and the corresponding patterns they map to in a Lorenz plot of δRR intervals (right-hand side of each tracing). Note that a collection of δRR intervals on the Lorentz plot during atrial fibrillation/ flutter results in characteristic signatures, including the clustering of points for periods of regularity. These signatures form the basis for the detection/ discrimination algorithm.



3  Sensing and Detection

QRS: T wave: Ratio:

QRS: T wave: Ratio:

1.6 mV 0.4 mV 4

QRS: T wave: Ratio:

0.9 mV 0.1 mV 9

121

1.6 mV 1.0 mV 1.6

Figure 3-96  Electrode system for subcutaneous ICD (S-ICD).

Future Directions: Hemodynamic Sensors for ICDs Current ICDs do not differentiate directly between hemodynamically stable and unstable tachycardias. Morbidity associated with unnecessary shocks for hemodynamically stable VT and repetitive shocks caused by SVT may be reduced using implantable hemodynamic sensors integrated in the therapy decision process. Mixed-venous oxygen saturation, right atrial pressure, RV pressure, subcutaneous photoplethysmography, endocardial accelerometers, and cardiac impedance sensors have been proposed as methods of discriminating hemodynamically stable from unstable tachycardias.57,179-183 Subcutaneous photoplethysmography has been tested in acute and chronic animal models as well as acute human studies. There is a good

correlation between mean arterial pressure and photoplethysmography pulse amplitude; discriminating perfusing (stable) from nonperfusing (unstable) tachycardias is feasible169,184 (Fig. 3-98). An implantable system for ambulatory hemodynamic monitoring using RV pressure has been studied in clinical trials.185 A recent clinical trial studied an ICD with the capability of recording RV pressure waveforms during tachycardias detected using “rate + interval” analysis. Episodes of recorded spontaneous VT-VF demonstrate substantial changes in RV pressure waveforms (Fig. 3-99). Although this clinical trial was terminated because of an unacceptable failure rate of the pressure sensor, it provided valuable data to assist in developing metrics that discriminate hemodynamically stable versus hemodynamically unstable tachyarrhythmias for integration into ICD algorithms.186

Amp mV1

3 0

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0.25

0.3

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3 0.3

0.45 0.5 0.55 0.6 Time Figure 3-97  Automatic selection of best sensing vector for subcutaneous ICD (S-ICD). The potential sensing vectors for S-ICD are shown. The S-ICD uses the ratio of R-wave amplitude to T-wave amplitude to choose the optimal sensing vector.

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O2 index

Pressure (mm Hg)

ECG (mV)

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Figure 3-98  Recordings from investigational sensor during electrophysiology laboratory testing. The sensor is mounted on the housing of an ICD emulator placed in the pocket during implant testing. Inset shows light-emitting diodes (LED) and light-receiving photodiode on housing of ICD emulator. Each panel shows surface ECG (upper tracing), arterial pressure (middle tracing), and photoplethysmographic oxygen (O2) index of tissue perfusion (lower tracing). The first panel shows the onset of atrial pacing at cycle length 400 msec. The arterial pressure remains stable, and the O2 index retains its phasic component with a mean value near the baseline-normalized value of zero. The second panel shows the onset of VF induced by 50-Hz current. With the onset of 50-Hz stimulation, the arterial pressure drops abruptly, and the O2 index loses its phasic component. About 2 seconds later, the O2 index begins a steep, monotonic decline. (Modified from Kroll MW, Swerdlow CD: The future of the implantable defibrillator. In Efimov I, Kroll M, Tchou P, editors: Cardiac bioelectric therapy: mechanisms and practical implications, Berlin, 2008, Springer Verlag.)



3  Sensing and Detection

123

RV pressure (mm Hg)

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VS VS TS TF TF TS TF TF TS TF TF TF TF TF TF TF TF TF TF TF T 438 328 313 297 344 313 313 344 297 313 313 297 297 297 297 297 313 297 297 313

40 20 0

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RV pressure (mm Hg)

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297 266 266 266 266 250 266 250250 266 250 266250250 250 266234 250 250

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VT syncope

RV pressure (mm Hg)

100

C

80 VS VS VS VS VS TF VS VS VS VS VS VSTF TF TF TF TF TF TF TF TF TF TF TF TF TF TF TF TFTPTP TPTP 60 VS 375 375 391 375 375 250 469 391 375 391 375 375 266 250 266 250 266266 266 266 250 266266266 201 266 266 266 266 234 219 234 234

40 20 0

Rapidly-conducted AF 0

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Figure 3-99  Stored EGM and pressure tracing from investigational ICD with hemodynamic sensor (Medtronic Chronicle ICD). In each panel, the upper tracing shows the shock electrogram recorded between right ventricular (RV) coil and left pectoral ICD housing. The marker channel below it shows interval classification and interval duration in milliseconds. The lower tracing shows RV pressure recorded from the implanted hemodynamic sensor in mm Hg. Horizontal axis shows time in seconds. Intervals are labeled VS (sinus zone), TS (VT zone), and TF (fast VT zone). A, Onset of VT at cycle length 300 msec. The patient was asymptomatic prior to antitachycardia pacing. RV systolic pressure falls by less than half. B, Onset of VT at cycle length 250 msec. The patient presented with syncope. RV systolic pressure drops abruptly and remains below a third of the baseline value. C, Rapidly conducted atrial fibrillation at cycle length 260 msec. Inappropriate antitachycardia pacing (TP markers) occurs for the last four beats. There is only a slight reduction in RV systolic pressure at shorter cycle lengths. The patient had no alteration of consciousness. (Tracings courtesy Dr. Michael Sweeney. From Kroll MW, Swerdlow CD: The future of the implantable defibrillator. In Efimov I, Kroll M, Tchou P, editors: Cardiac bioelectric therapy: mechanisms and practical implications, Berlin, 2008, Springer Verlag.)

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Beeman AL, Deutsch G, Rea RF: Paradoxical undersensing due to quiet timer blanking. Heart Rhythm 1:345-347, 2004. 44. Willems R, Holemans P, Ector H, et al: Paradoxical undersensing at a high sensitivity in dual chamber pacemakers. Pacing Clin Electrophysiol 24:308-315, 2001. 45. Castro A, Liebold A, Vincente J, et al: Evaluation of autosensing as an automatic means of maintaining a 2:1 sensing safety margin in an implanted pacemaker. Autosensing Investigation Team. Pacing Clin Electrophysiol 19:1708-1713, 1996. 46. Nowak B, Kampmann C, Schmid FX, et al: Pacemaker therapy in premature children with high degree AV block. Pacing Clin Electrophysiol 21:2695-2698, 1998. 47. Inama G, Santini M, Padeletti L, et al: Far-field R wave oversensing in dual chamber pacemakers designed for atrial arrhythmia

management: effect of pacing site and lead tip to ring distance. Pacing Clin Electrophysiol 27:1221-1230, 2004. 48. Dekker LR, Schrama TA, Steinmetz FH, Tukkie R: Undersensing of VF in a patient with optimal R wave sensing during sinus rhythm. Pacing Clin Electrophysiol 27:833-834, 2004. 49. Callans DJ, Hook BG, Kleiman RB, et al: Unique sensing errors in third-generation implantable cardioverter-defibrillators. J Am Coll Cardiol 22:1135-1140, 1993. 50. Callans DJ, Hook BG, Marchlinski FE: Paced beats following single nonsensed complexes in a “codependent” cardioverterdefibrillator and bradycardia pacing system: potential for ventricular tachycardia induction. Pacing Clin Electrophysiol 14:1281-1287, 1991. 51. Callans DJ, Hook BG, Marchlinski FE: Effect of rate and coupling interval on endocardial R wave amplitude variability in permanent ventricular sensing lead systems. J Am Coll Cardiol 22:746-750, 1993. 52. Glikson M, Beeman AL, Luria DM, et al: Impaired detection of ventricular tachyarrhythmias by a rate-smoothing algorithm in dual-chamber implantable defibrillators: intradevice interactions. J Cardiovasc Electrophysiol 13:312-318, 2002. 53. Cooper JM, Sauer WH, Verdino RJ: Absent ventricular tachycardia detection in a biventricular implantable cardioverterdefibrillator due to intradevice interaction with a rate-smoothing pacing algorithm. Heart Rhythm 1:728-731, 2004. 54. Shivkumar K, Feliciano Z, Boyle NG, Wiener I: Intradevice interaction in a dual chamber implantable cardioverter-defibrillator preventing ventricular tachyarrhythmia detection. J Cardiovasc Electrophysiol 11:1285-1288, 2000. 55. Gunderson BD, Gillberg JM, Wood MA, et al: Development and testing of an algorithm to detect implantable cardioverterdefibrillator lead failure. Heart Rhythm 3:155-162, 2006. 56. Zhang X, Volosin K, Kumar A, et al: Withholding ICD shocks for detected lead fractures. Heart Rhythm 6:S249, 2009. 57. Ellenbogen KA, Wood MA, Kapadia K, et al: Short-term reproducibility over time of right ventricular pulse pressure as a potential hemodynamic sensor for ventricular tachyarrhythmias. Pacing Clin Electrophysiol 15:971-974, 1992. 58. Leitch J, Klein G, Yee R, et al: Feasibility of an implantable arrhythmia monitor. Pacing Clin Electrophysiol 15:2232-2235, 1992. 59. Swerdlow CD: Implantation of cardioverter-defibrillators without induction of ventricular fibrillation. Circulation 103:2159-2164, 2001. 60. Michaud J, Horduna I, Dubuc M, Khairy P: ICD-unresponsive ventricular arrhythmias. Heart Rhythm 6:1827-1829, 2009. 61. Ruetz L, Koehler JL, Jackson TE, Belk P: Sinus rhythm R-wave amplitude does not predict undersensing of ventricular fibrillation by implantable cardioverter-defibrillators. Circ Arrhythm Electrophysiol 120:S650, 2009. 62. Tung L, Tovar O, Neunlist M, et al: Effects of strong electrical shock on cardiac muscle tissue. Ann NY Acad Sci 720:160-175, 1994. 63. Winkle RA, Bach SM, Jr, Echt DS, et al: The automatic implantable defibrillator: local ventricular bipolar sensing to detect ventricular tachycardia and fibrillation. Am J Cardiol 52:265-270, 1983. 64. Goldberger JJ, Horvath G, Donovan D, et al: Detection of ventricular fibrillation by transvenous defibrillating leads: integrated versus dedicated bipolar sensing. J Cardiovasc Electrophysiol 9:677-688, 1998. 65. Yee R, Jones DL, Jarvis E, et al: Changes in pacing threshold and R wave amplitude after transvenous catheter countershock. J Am Coll Cardiol 4:543-549, 1984. 66. Jung W, Manz M, Moosdorf R, et al: Changes in the amplitude of endocardial electrograms following defibrillator discharge: comparison of two lead systems. Pacing Clin Electrophysiol 18:2163-2172, 1995. 67. Kuhlkamp V, Dornberger V, Mewis C, et al: Clinical experience with the new detection algorithms for atrial fibrillation of a defibrillator with dual chamber sensing and pacing. J Cardiovasc Electrophysiol 10:905-915, 1999. 68. Wilkoff BL, Kuhlkamp V, Volosin K, et al: Critical analysis of dual-chamber implantable cardioverter-defibrillator arrhythmia detection: results and technical considerations. Circulation 103:381-386, 2001. 69. Swerdlow CD: Supraventricular tachycardia–ventricular tachycardia discrimination algorithms in implantable cardioverterdefibrillators: state-of-the-art review. J Cardiovasc Electrophysiol 12:606-612, 2001. 70. Gunderson B, Patel A, Bounds C: Automatic identification of implantable cardioverter-defibrillator lead problems using intracardiac electrograms. Comput Cardiol 29:121-124, 2002. 71. Swerdlow C, Shivkumar K: Implantable cardioverterdefibrillators: clinical aspects. In Zipes DP, Jalife J, editors: Cardiac electrophysiology: from cell to bedside, ed 4, Philadephia, 2004, Saunders, pp 980-993. 72. Pinski SL: 2:1 tracking of sinus rhythm in a patient with a dualchamber implantable cardioverter-defibrillator: what is the mechanism? J Cardiovasc Electrophysiol 12:503-504, 2001. 73. Hsu SS, Mohib S, Schroeder A, Deger FT: T-wave oversensing in implantable cardioverter-defibrillators. J Interv Card Electrophysiol 11:67-72, 2004. 74. Cao J, Shrivastav M, Koehler J, et al: Automatic identification of T-wave oversensing by patterns of alternating amplitude and frequency content in implantable cardioverter defibrillator electrograms [abstract]. Heart Rhythm 5:S094, 2008.

75. Garcia-Moran E, Mont L, Brugada J: Inappropriate tachycardia detection by a biventricular implantable cardioverterdefibrillator. Pacing Clin Electrophysiol 25:123-124, 2002. 76. Gunderson B, Swerdlow C, Wilcox J, et al: Causes of ventricular oversensing in implantable cardioverter-defibrillators: implications for diagnosis of lead fracture. Heart Rhythm 7:626-633, 2010. 77. Lloyd M, Hayes D, Friedman P: Troubleshooting. In Cardiac pacing and defibrillation: a clinical approach, Armonk, NY, 2000, Futura, pp 347-452. 78. Swerdlow CD, Gunderson BD, Ousdigian KT, et al: Downloadable algorithm to reduce inappropriate shocks caused by fractures of implantable cardioverter-defibrillator leads. Circulation 118:2122-2129, 2008. 79. Cheung JW, Iwai S, Lerman BB, Mittal S: Shock-induced ventricular oversensing due to seal plug damage: a potential mechanism of inappropriate device therapies in implantable cardioverter-defibrillators. Heart Rhythm 2:1371-1375, 2005. 80. Kallinen LM, Hauser RG, Lee KW, et al: Failure of impedance monitoring to prevent adverse clinical events caused by fracture of a recalled high-voltage implantable cardioverter-defibrillator lead. Heart Rhythm 5:775-779, 2008. 81. Chung EH, Casavant D, John RM: Analysis of pacing/ defibrillator lead failure using device diagnostics and pacing maneuvers. Pacing Clin Electrophysiol 32:547-549, 2009. 82. Hauser RG, Hayes DL: Increasing hazard of Sprint Fidelis implantable cardioverter-defibrillator lead failure. Heart Rhythm 6:605-610, 2009. 83. Pinski S, Eguia L, Sgarbossa E, Trohman R: Incidence and causes of commited shocks in noncommited implantable defibrillators [abstract]. J Am Coll Cardiol 39:93A, 2002. 84. Sweesy MW, Holland JL, Smith KW: Electromagnetic interference in cardiac rhythm management devices. AACN Clin Issues 15:391-403, 2004. 85. Lau CP, Leung SK, Tse HF, Barold SS: Automatic mode switching of implantable pacemakers. II. Clinical performance of current algorithms and their programming. Pacing Clin Electrophysiol 25:1094-1113, 2002. 86. Israel CW: Analysis of mode switching algorithms in dual chamber pacemakers. Pacing Clin Electrophysiol 25:380-393, 2002. 87. Lam CT, Lau CP, Leung SK, et al: Improved efficacy of mode switching during atrial fibrillation using automatic atrial sensitivity adjustment. Pacing Clin Electrophysiol 22:17-25, 1999. 88. Leung SK, Lau CP, Lam C, et al: Programmed atrial sensitivity: a critical determinant in atrial fibrillation detection and optimal automatic mode switching [in process citation]. Pacing Clin Electrophysiol 21:2214-2219, 1998. 89. Sweeney MO, Hellkamp AS, Ellenbogen KA, et al: Adverse effect of ventricular pacing on heart failure and atrial fibrillation among patients with normal baseline QRS duration in a clinical trial of pacemaker therapy for sinus node dysfunction. Circulation 107:2932-2937, 2003. 90. Wilkoff BL, Cook JR, Epstein AE, et al: Dual-chamber pacing or ventricular backup pacing in patients with an implantable defibrillator: the Dual Chamber and VVI Implantable Defibrillator (DAVID) Trial [see comment]. JAMA 288:3115-3123, 2002. 91. Melzer C, Sowelam S, Sheldon TJ, et al: Reduction of right ventricular pacing in patients with sinus node dysfunction using an enhanced search AV algorithm. Pacing Clin Electrophysiol 28:521-527, 2005. 92. Sweeney M, Nsah E, McGrew F, et al: Reduction in ventricular pacing and its long-term clinical outcomes: preliminary results of the Save Pace Trial [abstract]. Heart Rhythm 2:S322, 2005. 93. Olshansky B, Day J, McGuire M, Pratt T: Inhibition of Unnecessary RV Pacing with AV Search Hysteresis in ICDs (INTRINSIC RV): design and clinical protocol. Pacing Clin Electrophysiol 28:62-66, 2005. 94. Deering T, Wilensky M, Tondato F, et al: Auto intrinsic conduction search algorithm: a prospective analysis [abstract]. Pacing Clin Electrophysiol 26:1080, 2003. 95. Savoure A, Frohlig G, Galley D, et al: A new dual-chamber pacing mode to minimize ventricular pacing. Pacing Clin Electrophysiol 28:S43-S46, 2005. 96. Sweeney MO, Shea JB, Fox V, et al: Randomized pilot study of a new atrial-based minimal ventricular pacing mode in dualchamber implantable cardioverter-defibrillators. Heart Rhythm 1:160-167, 2004. 97. Pascale P, Pruvot E, Graf D: Pacemaker syndrome during managed ventricular pacing mode: what is the mechanism? J Cardiovasc Electrophysiol 20:574-576, 2009. 98. Mansour F, Talajic M, Thibault B, Khairy P: Pacemaker troubleshooting: when MVP is not the most valuable parameter. Heart Rhythm 3:612-614, 2006. 99. Gillis AM, Purerfellner H, Israel CW, et al: Reducing unnecessary right ventricular pacing with the managed ventricular pacing mode in patients with sinus node disease and AV block. Pacing Clin Electrophysiol 29:697-705, 2006. 100. De Voogt WG, Vonk BF, Albers BA, Hintringer F: Understanding capture detection. Europace 6:561-569, 2004. 101. Schuchert A, Frese J, Stammwitz E, et al: Low settings of the ventricular pacing output in patients dependent on a pacemaker: are they really safe? Am Heart J 143:1009-1011, 2002. 102. Splett V, Trusty JM, Hayes DL, Friedman PA: Determination of pacing capture in implantable defibrillators: benefit of evoked response detection using RV coil-to-can vector. Pacing Clin Electrophysiol 23:1645-1650, 2000.



103. Crossley GH, Mead H, Kleckner K, et al: Automated left ventricular capture management. Pacing Clin Electrophysiol 30:1190-1200, 2007. 104. Sperzel J, Binner L, Boriani G, et al: Evaluation of the atrial evoked response for capture detection with high-polarization leads. Pacing Clin Electrophysiol 28:S57-S62, 2005. 105. Sperzel J, Milasinovic G, Smith TW, et al: Automatic measurement of atrial pacing thresholds in dual-chamber pacemakers: clinical experience with atrial capture management. Heart Rhythm 2:1203-1210, 2005. 106. Olson W: Safety margins for sensing and detection: programming tradeoffs. In Kroll M, Lehmann M, editors: Implantable cardioverter-defibrillator therapy: the engineering-clinical interface, Norwell, Mass, 1996, Kluwer Academic Publishers, pp 389-420. 107. Anderson MH, Murgatroyd FD, Hnatkova K, et al: Performance of basic ventricular tachycardia detection algorithms in implantable cardioverter-defibrillators: implications for device programming. 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Raedle-Hurst TM, Wiecha J, Schwab JO, et al: Clinical performance of a specific algorithm to reconfirm self-terminating ventricular arrhythmias in current implantable cardioverterdefibrillators. Am J Cardiol 88:744-749, 2001. 113. Gillberg J, Olson W: Dual-chamber sensing and detection for implantable cardioverter-defibrillators. In Singer I, editor: Interventional electrophysiology, ed 2, Philadelphia, 2001, Lippincott– Williams & Wilkins. 114. Glikson M, Swerdlow CD, Gurevitz OT, et al: Optimal combination of discriminators for differentiating ventricular from supraventricular tachycardia by dual-chamber defibrillators. J Cardiovasc Electrophysiol 16:732-739, 2005. 115. Saba S, Volosin K, Yee R, et al: Combined atrial and ventricular antitachycardia pacing as a novel method of rhythm discrimination: the Dynamic Discrimination Download Study. Circulation 121:487-497, 2010. 116. Brugada J, Mont L, Figueiredo M, et al: Enhanced detection criteria in implantable defibrillators. 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Gronefeld GC, Schulte B, Hohnloser SH, et al: Morphology discrimination: a beat-to-beat algorithm for the discrimination of ventricular from supraventricular tachycardia by implantable cardioverter-defibrillators. Pacing Clin Electrophysiol 24:15191524, 2001. 122. Boriani G, Biffi M, Frabetti L, et al: Clinical evaluation of morphology discrimination: an algorithm for rhythm discrimination in cardioverter-defibrillators. Pacing Clin Electrophysiol 24:994-1001, 2001. 123. Gold MR, Shorofsky SR, Thompson JA, et al: Advanced rhythm discrimination for implantable cardioverter-defibrillators using electrogram vector timing and correlation. J Cardiovasc Electrophysiol 13:1092-1097, 2002. 124. Gold MR, Theuns DA, Knight BP, et al: Comparison of arrhythmia discrimination by subcutaneous versus dual chamber transvenous ICD systems: Primary results from START. Circulation 120:S649, 2009. 125. 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3  Sensing and Detection 128. Stadler RW, Gunderson BD, Gillberg JM: An adaptive intervalbased algorithm for withholding ICD therapy during sinus tachycardia. Pacing Clin Electrophysiol 26:1189-1201, 2003. 129. Stein KM, Euler DE, Mehra R, et al: Do atrial tachyarrhythmias beget ventricular tachyarrhythmias in defibrillator recipients? J Am Coll Cardiol 40:335-340, 2002. 130. Aliot E, Nitzsche R, Ripart A: Arrhythmia detection by dualchamber implantable cardioverter-defibrillators: a review of current algorithms. Europace 6:273-286, 2004. 131. Wilkoff B, Sterns L, Morgan JM, et al: Preventing shocks after ICD implantation: can a strategy of standardized ICD programming match physician tailored? (Late breaking clinical trial presentation). Heart Rhythm 2005. 132. Deisenhofer I, Kolb C, Ndrepepa G, et al: Do current dual chamber cardioverter defibrillators have advantages over conventional single chamber cardioverter defibrillators in reducing inappropriate therapies? A randomized, prospective study. J Cardiovasc Electrophysiol 12:134-142, 2001. 133. Theuns DA, Klootwijk AP, Goedhart DM, Jordaens LJ: Prevention of inappropriate therapy in implantable cardioverterdefibrillators: results of a prospective, randomized study of tachyarrhythmia detection algorithms. J Am Coll Cardiol 44:2362-2367, 2004. 134. Bansch D, Steffgen F, Gronefeld G, et al: The 1+1 trial: a prospective trial of a dual- versus a single-chamber implantable defibrillator in patients with slow ventricular tachycardias. Circulation 110:1022-1029, 2004. 135. Friedman P, McClelland R, Bamlet W, et al: Dual chamber versus single chamber detection enhancements for implantable defibrillator rhythm diagnosis (The Detect SVT Study). Circulation 113:2871-2879, 2006. 136. Sinha AM, Stellbrink C, Schuchert A, et al: Clinical experience with a new detection algorithm for differentiation of supraventricular from ventricular tachycardia in a dual-chamber defibrillator. J Cardiovasc Electrophysiol 15:646-652, 2004. 137. Klein GJ, Gillberg JM, Tang A, et al: Improving SVT discrimination in single-chamber ICDs: a new electrogram morphologybased algorithm. J Cardiovasc Electrophysiol 17:1310-1319, 2006. 138. Zeger SL, Liang KY, Albert PS: Models for longitudinal data: a generalized estimating equation approach. Biometrics 44:10491060, 1988. 139. Wathen MS, Volosin KJ, Sweeney MO, et al: Ventricular anti­ tachycardia pacing by implantable cardioverter-defibrillators reduces shocks for inappropriately detected supraventricular tachycardia [abstract]. Heart Rhythm 1, 2004. 140. Irvine J, Dorian P, Baker B, et al: Quality of life in the Canadian Implantable Defibrillator Study (CIDS). Am Heart J 144:282289, 2002. 141. Birgersdotter-Green U, Rosenqvist M, Lindemans FW, et al: Holter documented sudden death in a patient with an implanted defibrillator. Pacing Clin Electrophysiol 15:1008-1014, 1992. 142. Schwartzman D, Musley SK, Swerdlow C, et al: Early recurrence of atrial fibrillation after ambulatory shock conversion. J Am Coll Cardiol 40:93-99, 2002. 143. Wilkoff BL, Williamson BD, Stern RS, et al: Strategic programming of detection and therapy parameters in implantable cardioverter-defibrillators reduces shocks in primary prevention patients: results from the PREPARE (Primary Prevention Parameters Evaluation) study. J Am Coll Cardiol 52:541-550, 2008. 144. Bardy GH, Lee KL, Mark DB, et al: Amiodarone or an implantable cardioverter-defibrillator for congestive heart failure. N Engl J Med 352:225-237, 2005. 145. Schaumann A, von zur Muhlen F, Herse B, et al: Empirical versus tested antitachycardia pacing in implantable cardioverterdefibrillators: a prospective study including 200 patients. Circulation 97:66-74, 1998. 146. Wathen MS, Sweeney MO, DeGroot PJ, et al: Shock reduction using antitachycardia pacing for spontaneous rapid ventricular tachycardia in patients with coronary artery disease. Circulation 104:796-801, 2001. 147. Wathen MS, DeGroot PJ, Sweeney MO, et al: Prospective randomized multicenter trial of empirical antitachycardia pacing versus shocks for spontaneous rapid ventricular tachycardia in patients with implantable cardioverter-defibrillators: Pacing Fast Ventricular Tachycardia Reduces Shock Therapies (PainFREE Rx II) trial results. Circulation 110:2591-2596, 2004. 148. Gunderson BD, Gillberg JM, Olson WH, Swerdlow C: Effect of programmed number of intervals to detect ventricular fibrillation on implantable cardioverter defibrillator longevity and unnecessary shocks [abstract]. Circulation 106:322, 2002. 149. Gasparini M, Menozzi C, Proclemer A, et al: A simplified biventricular defibrillator with fixed long detection intervals reduces implantable cardioverter-defibrillator (ICD) interventions and heart failure hospitalizations in patients with non-ischaemic cardiomyopathy implanted for primary prevention: the RELEVANT study (Role of Long Detection Window Programming in Patients with Left Ventricular Dysfunction, Non-ischemic Etiology in Primary Prevention Treated with a Biventricular ICD). Eur Heart J 30:2758-2767, 2009. 150. Daubert JP, Wojciech Z, Cannom DC, et al: Frequency and mechanisms of inappropriate ICD therapy in MADIT II [abstract]. J Am Coll Cardiol 2004. 151. Kolb C, Deisenhofer I, Weyerbrock S, et al: Incidence of anti­ tachycardia therapy suspension due to magnet reversion in

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implantable cardioverter defibrillators. Pacing Clin Electrophysiol 27:221-223, 2004. 152. Shukla HH, Flaker GC, Jayam V, Roberts D: High defibrillation thresholds in transvenous biphasic implantable defibrillators: clinical predictors and prognostic implications. Pacing Clin Electrophysiol 26:44-48, 2003. 153. Bansch D, Castrucci M, Bocker D, et al: Ventricular tachycardias above the initially programmed tachycardia detection interval in patients with implantable cardioverter-defibrillators: incidence, prediction and significance. J Am Coll Cardiol 36:557-565, 2000. 154. Brode SE, Schwartzman D, Callans DJ, et al: ICD-antiarrhythmic drug and ICD-pacemaker interactions. J Cardiovasc Electrophysiol 8:830-842, 1997. 155. Monahan KM, Hadjis T, Hallett N, et al: Relation of induced to spontaneous ventricular tachycardia from analysis of stored far-field implantable defibrillator electrograms. Am J Cardiol 83:349-353, 1999. 156. Weber M, Bocker D, Bansch D, et al: Efficacy and safety of the initial use of stability and onset criteria in implantable cardioverter defibrillators. J Cardiovasc Electrophysiol 10:145-153, 1999. 157. Cohen AI, Wish MH, Fletcher RD, et al: The use and interaction of permanent pacemakers and the automatic implantable cardioverter-defibrillator. Pacing Clin Electrophysiol 11:704-711, 1988. 158. Geiger MJ, O’Neill P, Sharma A, et al: Interactions between transvenous nonthoracotomy cardioverter-defibrillator systems and permanent transvenous endocardial pacemakers. Pacing Clin Electrophysiol 20:624-630, 1997. 159. Glikson M, Trusty JM, Grice SK, et al: A stepwise testing protocol for modern implantable cardioverter-defibrillator systems to prevent pacemaker-implantable cardioverter-defibrillator interactions. Am J Cardiol 83:360-366, 1999. 160. Wietholt D, Kuehlkamp V, Meisel E, et al: Prevention of sustained ventricular tachyarrhythmias in patients with implantable cardioverter-defibrillators: the PREVENT study. J Interv Card Electrophysiol 9:383-389, 2003. 161. Pollak WM, Simmons JD, Interian A, Jr, et al: Pacemaker diagnostics: a critical appraisal of current technology. Pacing Clin Electrophysiol 26:76-98, 2003. 162. Hammel E, Bertrand B, et al: Appropriate detection of guidant pacemaker stored electrograms assessed by centralized arrhythmia workstation [abstract]. Pacing Clin Electrophsyiol 23:680, 2000. 163. Passman RS, Weinberg KM, Freher M, et al: Accuracy of mode switch algorithms for detection of atrial tachyarrhythmias. J Cardiovasc Electrophysiol 15:773-777, 2004. 164. Seidl K, Rameken M, Breunung S, et al: Diagnostic assessment of recurrent unexplained syncope with a new subcutaneously implantable loop recorder. REVEAL Investigators. Europace 2:256-262, 2000. 165. Israel CW, Gascon D, Nowak B, et al: Diagnostic value of stored electrograms in single-lead VDD systems. Pacing Clin Electrophysiol 23:1801-1803, 2000. 166. Glotzer TV, Daoud EG, Wyse DG, et al: The relationship between daily atrial tachyarrhythmia burden from implantable device diagnostics and stroke risk: the TRENDS study. Circ Arrhythm Electrophysiol 2:474-480, 2009. 167. Capucci A, Santini M, Padeletti L, et al: Monitored atrial fibrillation duration predicts arterial embolic events in patients suffering from bradycardia and atrial fibrillation implanted with antitachycardia pacemakers. J Am Coll Cardiol 46:1913-1920, 2005. 168. Ricci RP, Morichelli L, Gargaro A, et al: Home monitoring in patients with implantable cardiac devices: is there a potential reduction of stroke risk? Results from a computer model tested through Monte Carlo simulations. J Cardiovasc Electrophysiol 20:1244-1251, 2009. 169. Ip J, Waldo AL, Lip GY, et al: Multicenter randomized study of anticoagulation guided by remote rhythm monitoring in patients with implantable cardioverter-defibrillator and CRT-D devices: rationale, design, and clinical characteristics of the initially enrolled cohort. The IMPACT study. Am Heart J 158:364370, 2009. 170. Morris MM, KenKnight BH, Lang DJ: Detection of atrial arrhythmia for cardiac rhythm management by implantable devices. J Electrocardiol 33(Suppl):133-139, 2000. 171. Schuchert A, Boriani G, Wollmann C, et al: Implantable dualchamber defibrillator for the selective treatment of spontaneous atrial and ventricular arrhythmias: arrhythmia incidence and device performance. J Interv Card Electrophysiol 12:149-156, 2005. 172. Krahn AD, Klein GJ, Yee R, Norris C: Maturation of the sensed electrogram amplitude over time in a new subcutaneous implantable loop recorder. Pacing Clin Electrophysiol 20:16861690, 1997. 173. Thaker JP, Patel MB, Jongnarangsin K, et al: Electromagnetic interference in an implantable loop recorder caused by a portable digital media player. Pacing Clin Electrophysiol 31:13451347, 2008. 174. Trigano A, Blandeau O, Dale C, et al: Risk of cellular phone interference with an implantable loop recorder. Int J Cardiol 116:126-130, 2007. 175. Gimbel JR, Zarghami J, Machado C, Wilkoff BL: Safe scanning, but frequent artifacts mimicking bradycardia and tachycardia during magnetic resonance imaging (MRI) in patients with an

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SECTION 1  Basic Principles of Device Therapy

implantable loop recorder (ILR). Ann Noninvas Electrocardiol 10:404-408, 2005. 176. Brignole M, Bellardine Black CL, Thomsen PE, et al: Improved arrhythmia detection in implantable loop recorders. J Cardiovasc Electrophysiol 19:928-934, 2008. 177. Sarkar S, Ritscher D, Mehra R: A detector for a chronic implantable atrial tachyarrhythmia monitor. Biomed Eng IEEE Trans 55:1219-1224, 2008. 178. Hindricks G, Pokushalov E, Urban L, et al: Performance of a new leadless implantable cardiac monitor in detecting and quantifying atrial fibrillation: results of the XPECT Trial. Circ Arrhythm Electrophysiol 3:141-147, 2010. 179. Bordachar P, Garrigue S, Reuter S, et al: Hemodynamic assessment of right, left, and biventricular pacing by peak endocardial acceleration and echocardiography in patients with

end-stage heart failure. Pacing Clin Electrophysiol 23:1726-1730, 2000. 180. Hegbom F, Hoff PI, Oie B, et al: RV function in stable and unstable VT: is there a need for hemodynamic monitoring in future defibrillators? Pacing Clin Electrophysiol 24:172-182, 2001. 181. Kaye G, Astridge P, Perrins J: Tachycardia recognition and diagnosis from changes in right atrial pressure waveform: a feasibility study. Pacing Clin Electrophysiol 14:1384-1392, 1991. 182. Sharma AD, Bennett TD, Erickson M, et al: Right ventricular pressure during ventricular arrhythmias in humans: potential implications for implantable antitachycardia devices. J Am Coll Cardiol 15:648-655, 1990. 183. Ellenbogen KA, Lu B, Kapadia K, et al: Usefulness of right ventricular pulse pressure as a potential sensor for hemo­

dynamically unstable ventricular tachycardia. Am J Cardiol 65:1105-1111, 1990. 184. Turcott R: Detection of hemodynamically unstable arrhythmias using subcutaneous photoplethysmography [abstract]. Heart Rhythm 2:S83, 2005. 185. Bourge RC, Abraham WT, Adamson PB, et al: Randomized controlled trial of an implantable continuous hemodynamic monitor in patients with advanced heart failure: the COMPASS-HF study. J Am Coll Cardiol 51:1073-1079, 2008. 186. Kroll MW, Swerdlow CD: The future of the implantable defibrillator. In Efimov I, Kroll M, Tchou P, editors: Cardiac bioelectric therapy: mechanisms and practical implications, Berlin, 2008, Springer Verlag, pp 551-570.

4 

Engineering and Construction of Pacemaker and ICD Leads HARIS M. HAQQANI  |  LAURENCE M. EPSTEIN  |  JOSHUA M. COOPER*

Tremendous advances have occurred in implanted cardiac device

therapy, from the evolution of pacemakers to the modern era of highvoltage defibrillation and resynchronization devices. Advances in the components, integrated circuitry, size, and software of the implantable pulse generator are readily apparent, but the revolutionary progress in the design and manufacture of device leads is often not appreciated. In any device system, the lead is the critical interface linking the pulse generator and the myocardium; it allows for communication of electrical signals from the heart to the pulse generator through the sensing circuitry, and in the reverse direction, it allows for bradycardia pacing or tachyarrhythmia therapy. Although leadless pacing1 and defibrillation2 device systems are being actively developed, leads clearly will remain an integral component of device therapy for the foreseeable future. Whether used for pacing, resynchronization, or defibrillation, and regardless of their internal structural design, all leads contain several common components, including electrodes, conductors, insulators, fixation mechanisms, and connector pins. Defibrillation leads also contain shock coils for the delivery of high-voltage, high-current electrical discharges to terminate ventricular fibrillation. This chapter examines each of these elements as well as the overall design, construction, and evolution of pacemaker and implantable cardioverterdefibrillator (ICD) leads.

Pacing Leads HISTORICAL MILESTONES IN DEVELOPMENT Permanently implantable pacing leads evolved from the temporary pacing wires that were first used to provide bradycardia support.3 The initial permanent transvenous leads were unipolar and consisted of a basic conductor, an insulator, and a connector pin. The electrodes were large and polished with high-polarization properties, low electrodetissue impedance, and excessive current drain. No lumen was present in these leads to allow for stylet insertion, so lead implantation was a long, difficult task. Also, no fixation mechanism was present, so lead displacement rates were high. The development of bipolar pacing leads minimized far-field oversensing,4 requiring a major reconsideration of lead body design and structure, because two longitudinal conductors needed to coexist inside the lead, separated only by a thin layer of insulation. This and subsequent increases in complexity required advances in materials science so as not to compromise reliability. The development of passive-fixation tines was a huge step forward and dramatically reduced the rate of reoperation for lead dislodgement compared with existing flange-tipped pacing leads.5-7 Active-fixation helices were introduced subsequently and further improved the stability of implanted leads. Modern electrodes were developed with small geometric surface area but large effective area as a result of porous surfaces, and these optimized the trade-off between high current *Disclosures: Dr. Haqqani is the recipient of a Training Fellowship from the National Health and Medical Research Council of Australia (544309) and the Bayer Fellowship from the Royal Australasian College of Physicians. Drs. Cooper and Epstein have received modest honoraria from Boston Scientific, Medtronic, St. Jude Medical, Biotronik, and Spectranetics.

density and low polarization. Chronic threshold rises and late exit block were still major problems in cardiac pacing at that stage, however, and remained so until the revolutionary incorporation of an elutable dexamethasone reservoir into a porous-tip electrode, essentially eliminating this complication.8 Although this likely resulted from the glucocorticoid and not the redesigned porous titanium-tip electrode, this remained unproved until a pioneering randomized double-blind trial compared two otherwise identical leads and electrodes, with and without steroid elution.9 This unequivocally demonstrated the benefit was the result of the dexamethasone. Subsequent advances have included the universal standardization of connector pins (and consequently device headers), improving lead longevity through the use of better component materials, and improving sensing through the use of narrower bipoles. LEAD STRUCTURE AND POLARITY The basic functions of any cardiac device lead—pacing and sensing— require conductors for current flow between the pulse generator and myocardium and insulators to prevent short-circuiting of this current. These elements may be configured in two ways. First, a pace/sense circuit may be composed of a lead containing one conductor that connects to a negatively charged tip electrode, the cathode, from which electrons flow through the myocardium and thoracic cavity back to the active, positively charged pulse generator (IPG), the anode (Fig. 4-1, A). Such a lead, with only one conductor and electrode, is called a unipolar lead because only one electrode is in contact with the heart (although it is part of a bipolar circuit). It has been established that pacing thresholds are lower at most pulse widths when the intracardiac pacing electrode is configured as the cathode rather than as the anode10 (see Chapter 1). Although once the only option, unipolar leads have largely been replaced by bipolar leads. The alternative system consists of a lead body containing two conductors (separated and surrounded by insulation) that connect to a cathode-tip electrode and an anode ring electrode located several millimeters more proximally (Fig. 4-1, B). The IPG is not an active electrode in a bipolar circuit, and such leads are said to have bipolar configuration. Unipolar systems have a large interelectrode distance from the tip of the intracardiac lead to the IPG, and this large field of view makes them vulnerable to the oversensing of myocardial signals (T waves or electrograms from another chamber, known as “crosstalk”) and farfield nonmyocardial signals such as pacing artifacts (another form of crosstalk), skeletal myopotentials, and nonphysiologic electrical “noise” in up to 38% of patients.11 Electrical signal oversensing can have significant consequences on pacing behavior, including ventricular output inhibition and inappropriate mode switching caused by far-field ventricular oversensing on the atrial channel.12,13 Because of the additional consequence of inappropriate shock delivery if oversensing were to occur in an ICD, it is particularly important for ICD leads to have bipolar sensing circuitry. Since the IPG itself serves as one of the electrodes in a unipolar system, skeletal muscle capture and pectoral muscle stimulation may result in the setting of a higher programmed pacing output. Bipolar leads have been shown to have similar but slightly higher, acute and chronic stimulation thresholds to unipolar leads.4 Bipolar leads also carry an element of redundancy in that they

127

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SECTION 1  Basic Principles of Device Therapy

Unipolar

Bipolar

+

RV lead

RV lead +

–

A

–

B

Figure 4-1  A, Unipolar pacing circuit, with an intracardiac cathode located on the lead tip in the right ventricle (RV). The circuit is completed by the active pulse generator, which forms the anode. B, In a bipolar circuit, both cathode and anode are located on the lead, with the anode generally formed by a ring electrode proximal to the tip cathode.

can be used in either the bipolar or the unipolar configuration, as in isolated failure of the proximal, anodal conductor, with no demonstrated difference in performance between the two pacing configurations.14 In addition, the much lower amplitude of the bipolar pacing artifact greatly reduces the likelihood of crosstalk and oversensing. The great advantage of bipolar leads is related to their superior sensing properties. These factors have led to the near-universal adoption of the bipolar lead as the configuration of choice in cardiac device systems.15 The only potential disadvantage of bipolar leads is that reliability is lower than with the less complicated unipolar lead,16 although some studies have not supported this contention.15,17 However, insulation failure usually has negligible effects on unipolar lead performance, a clear reliability advantage. In terms of the structure of the lead body, the simple, single-conductor design of unipolar leads had to be modified to accommodate the greater number of components and increased complexity of the bipolar lead body. There are two main bipolar lead designs, coaxial and coradial (Fig 4-2). Coaxial leads have an inner conductor that extends down the length of the lead to the tip electrode, the cathode, arranged in a coil configuration with a central lumen to allow for passage of a stylet at implantation. This coil is covered by a cylindrical length of inner insulation, which in turn

is wrapped by another coil conductor that also runs down the lead to the ring electrode, the anode. A second, outer layer of insulation and lead covering protect the ring conductor from the outside environment, thereby completing the design. Although this coaxial model has been the industry standard for pacing leads for many years, the resulting bulk and stiffness of this four-layer design are significant compared with the simpler unipolar leads. Coradial bipolar leads addressed some of these concerns with new conductor and insulator technology that was subsequently used in ICD lead design. In coradial leads, a single coil extends down the length of the lead (again with a central lumen to allow for stylet insertion) and consists of two parallel, alternating conductor strands, one of which connects to the cathode and the other to the anode. Each conductor strand is individually coated with a bonded layer of ethylene tetrafluoroethylene (ETFE) fluoropolymer insulation (originally by DuPont, Wilmington, Del), which serves to insulate each strand from the other, despite being intertwined. The single, two-component coil is surrounded by a single, outer insulation covering.18 These leads have comparable bulk (~5F diameter) and flexibility to unipolar leads, often with tines or a fixed, nonretractable helix to minimize size; electrical parameters and reliability have not been demonstrably inferior to coaxial leads.18-21 Defibrillator leads have even greater complexity, with two, three, or four conductors that connect to the pacing/sensing electrodes as well as to the high-voltage shocking coil(s), depending on the specific type of lead. Because of the prohibitive bulk that would result from a coaxial design with more than two conductors, ICD leads generally have a different type of structure, known as a multilumen design. This type of ICD lead consists of a long cylinder of insulating material, with separate internal channels running down its length (Fig. 4-3). Each conductor runs down an individual channel, usually with its own additional covering tube of insulation. The configuration and number of channels differ, depending on the number of conductors, the specific manufacturer, and whether air-filled channels are also

Urethane ETFE

Sense

COAXIAL DESIGN Ring Tip

A

Outer insulation Inner insulation

CORADIAL DESIGN

Compression lumen

Defib

Tip

PTFE

B

Ring

Outer insulation

Figure 4-2  A, Coaxial lead body design consists of two nested coil conductors to the tip and ring electrodes, each wrapped in a separate layer of insulation. The inner coil connects to the tip cathode and contains a central lumen for stylet passage; the outer coil connects to the ring anode. B, In coradial design, a single coil contains the parallel conductor strands to the tip and ring electrodes, with each strand covered by a fluoropolymer insulator. A surrounding layer of outer insulation covers the lead body.

HP silicone

Pace

Figure 4-3  Multilumen ICD lead design consists of a single body of insulation with several internal channels that contain the conductor elements (central pace/sense coil and outer high-voltage cables), each covered by its own layer of fluoropolymer insulation. A surrounding layer of outer polyurethane-based insulation is present. HP, Highperformance. (Courtesy Medtronic, Minneapolis.)



4  Engineering and Construction of Pacemaker and ICD Leads

Deflectable delivery catheter

129

DFT conductor strands are then formed into multifilar (i.e., multiple strands or filaments) coils with a central lumen in coaxial leads, as described earlier, or further coated in a layer of ETFE fluoropolymer before being arranged in bifilar coradial coils,28 again with a central lumen. The alternative way of forming a composite-wire conductor is the drawn brazed strand (DBS) method, generally consisting of six strands of a high-resistance material (e.g., MP-35N) that are tightly molded over a central strand of silver, such that the inner, low-resistance metal is forced between and around the strong outer strands (Fig. 4-5, Drawn filled tube

Figure 4-4  Isodiametric, lumenless, 4.1F pacing lead (Model 3830 SelectSecure). There is no central lumen for stylet passage, so the lead is deployed through an 8.5F steerable catheter delivery system (SelectSite C304) at implantation. (Courtesy Medtronic, Minneapolis.)

incorporated into the design to improve lead handling and protect the conductors from flexion-related damage in the dynamic intracardiac and intravascular environment. One other lead body structure worth mentioning are small (4.1F diameter), isodiametric, lumenless pacing leads (Model 3830 SelectSecure; Fig. 4-4). The lack of a central lumen and the use of a conductor cable (rather than coil) allows for increased insulation redundancy, high tensile strength, and reduced bulk. It also means that the activefixation helix is nonextendable because there is no central coil conductor to transmit torque from the connector pin to extend or retract the helix. Fixation of such leads to the myocardium is achieved by turning the entire lead body. To enable implantation without a stylet, the lead is deployed through an 8.5F steerable catheter delivery system (SelectSite C304). Short-term results from this lead have been reasonable, with no signs of increased lead displacement or fracture risk.22,23 The small external diameter of this lead has been considered to be of some benefit in pediatric patients24 and patients with congenital heart disease,25 but the suggestion of a resulting improvement in chronic venous patency from the smaller size is yet to be substantiated in longterm follow-up. In addition, selective site pacing, in both the atria and the ventricles, is facilitated by the deflectable sheath, although steerable and preformed stylets can also be used with conventional leads to achieve this. No published experience with extraction of chronically placed lumenless leads exists, and the lack of a lumen to accommodate a lead-locking device may be a significant disadvantage of this design.

Silver core

MP−35N

A Drawn brazed strand

CONDUCTORS Both the coaxial and the coradial bipolar pacing lead body designs have been in widespread use and proved to be relatively fracture resistant over time. This durability is at least partially related to the materials and design of the conductor elements used in pacing leads. One study found conductor fracture over a 17-year follow-up in 19 of 561 right ventricular pacing leads (3.4%) and then generally at points of high stress, such as at the lead anchoring sleeve and the costoclavicular ligament (subclavian crush).26 The basic material used for most conductors over the last 3 to 4 decades has been MP-35N (SPS Technologies, Cleveland), an alloy of nickel, cobalt, chromium, and molybdenum. The main advantage of MP-35N is its high strength and resistance to corrosion.27 Its main disadvantage is its high electrical resistance, but this has been overcome with the development of composite-wire conductors that incorporate low-resistance metals such as silver and stainless steel with high-strength materials such as titanium, platinum, and platinum-iridium alloy. In pacing leads, these materials are generally incorporated into a drawn filled tube (DFT) composite-wire conductor strand consisting of a thick, strong body of MP-35N or titanium that has been filled with a central core of softer, low-resistance metal such as silver, often encased in a further outer shell of platinum alloy (Fig. 4-5, A). The

Silver

MP−35N

B Figure 4-5  Composite-wire conductor design. A, Drawn filled tube (DFT) composite wire; thick body of MP-35N is filled with softer, lowresistance silver. B, Drawn brazed strand (DBS) wire; six strands of MP-35N are tightly molded over a central strand of silver. DFT and DBS composite wire can then be further fashioned into multifilar coil and cables. (Courtesy St Jude Medical, Sylmar, Calif.)

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SECTION 1  Basic Principles of Device Therapy

B). This results in a low-resistance, durable composite conductor wire that can once again be formed into a multifilar coil for use in pacing leads.29 However, it is also used to form strong, multifilar, braided cables that are employed as the conductors for the high-voltage circuitry of ICD leads. DBS cables are discussed later.

Environmental stresss cracking

INSULATION Pacing lead insulation is a critical and vulnerable component that is often the cause of lead failure.30,31 One of the problems with the major materials used for lead insulation (polyurethane, silicone rubber, fluoropolymers) is that they were not originally designed specifically for this purpose, and each has disadvantages when used as part of a biologic pacing system. They were incorporated by early lead engineers into their designs for lack of any purpose-built alternatives, and it is only recently that any viable, tailored alternatives have been developed. Poly(ether)urethane is a synthetic, segmented polymer with very high tensile strength and resistance to mechanical abrasion. From a physical perspective, this means that a thinner layer of insulation can be used to cover the lead conductors, thereby reducing the overall lead diameter. Also, polyurethane has excellent lubricity and handling characteristics, with a low frictional coefficient, which can facilitate implantation of two or three lead systems by decreasing the physical interactions between the leads. However, polyurethane leads are stiffer and not fully biostable, being subject to in vivo polymer degradation that can cause insulation and late lead failure.32 The two main types of polyurethane are Pellethane 80A and Pellethane 55D (Upjohn, CPR Division, Torrance, Calif), with 80A particularly prone to biodegradation. Two mechanisms of in vivo, biologically mediated degradation of polyurethane predominate: environmental stress cracking (ESC) and metal ion oxidation (MIO)33 (Fig. 4-6). When polyurethane is subjected to repeated mechanical stresses in the presence of surface-active agents provided by macrophages and α2-macroglobulin, parallel polymer molecules are disordered and interchain bonds weakened. This is the mechanism of ESC,34 manifesting as a deep surface cracking visible with electron microscopy, particularly at the anchoring sleeve and venous entry portions of the lead. By comparison, no mechanical stress is required for MIO to develop because this is a chemical reaction that occurs particularly at the conductor-insulator interface.35 Biologic oxidants, such as peroxide from surrounding inflammatory cells,36 can result in the release of metal cations from the conductor, especially cobalt and nickel ions. These cations cause oxidation of the soft zones in the polymer, starting at the α-carbon of the ether bond and resulting in chain break. The harder Pellethane 55D is much less susceptible to both ESC and MIO than 80A but is significantly stiffer. Nevertheless, the long-term performance of Pellethane 55D has been much better than Pellethane 80A, a material notoriously prone to failure. Silicone rubber (polysiloxane) has none of the disadvantages of polyurethane. It is completely inert and biostable over extended periods and thus not vulnerable to MIO or ESC. Silicone is more flexible, which may decrease the risk of damage to cardiac structures, including lead perforation. It is also more thermally resistant to the effects of electrocautery, which can be a significant advantage during pulse generator replacement or system revision.37 Initially, a softer peroxide-catalyzed silicone rubber, MDX4-4515-50A, was used for pacing leads, but new versions with higher tensile strength and abrasion resistance have been developed. These include high-performance (HP) silicone, extra-tear-resistant (ETR) silicone, and Novus (Med4719, Nusil Technologies, Carpinteria, Calif), produced by hybridizing HP and MDX4 silicone. Unfortunately, no variety of silicone rubber has the positive characteristics of polyurethane, including surface lubriciousness and ease of handling. Although these friction-related problems of silicone can be addressed by an outer layering of polyurethane, the main problem with silicone relates to its lower tensile strength and susceptibility to abrasion and tears (Fig. 4-7, A). This means that silicone insulation is more prone to damage during

A Metal ion oxidation

B Figure 4-6  Mechanisms of polyurethane insulation failure on electron microscopy. A, Environmental stress cracking (ESC) results from mechanical stress, leading to disruption of polymer molecule arrays and interchain bonds. B, Metal ion oxidation (MIO) of the polymer soft segments causes polymer chain break. (Courtesy St Jude Medical, Sylmar, Calif.)

implantation and subsequent interactions in the device pocket, and that a thicker insulation layer must be used to maintain lead reliability, thereby increasing lead bulk. Also, as with glass, silicone is a fluid material and thus prone to a gradual flowing process away from pressure points, known as “cold flow,” which can lead to thinned, denuded, and abraded troughs or depressions adjacent to areas of thickened rubber polymer32 (Fig. 4-7, B). The fluoropolymers polytetrafluoroethylene (PTFE) and ETFE (DuPont) have some of the advantages of both silicone and polyurethane, with good abrasion resistance and biostability.38 Unfortunately, their great stiffness excludes them from functioning as the primary insulating material in any lead, but they can be used, as outlined earlier, as a thin coat on conductor strands, particularly in coradial leads. This inner insulating layer can prevent electrical communication between conductor strands and can also protect them from interacting with an adjacent outer layer of polyurethane, thereby reducing MIO. Recently, the development of a specific insulation material for cardiac leads has suggested that the deficiencies of the standard insulators in pacing leads may finally be addressed. The material is a copolymer hybrid composed of 48% silicone, 40% polyurethane hard segment, and 12% polyhexamethylene oxide soft segment; it is called Elast-Eon (Aortech Biomaterials, Clayton, Victoria, Australia). This



4  Engineering and Construction of Pacemaker and ICD Leads

Abrasion: 4 years

A

Silicone abrasion

Cold flow: 1 year

B

Silicone cold flow

Figure 4-7  Mechanisms of silicone rubber insulation failure. A, Weaker molecular structure of silicone makes it more prone to abrasions and tears at points of higher mechanical stress. B, Glasslike fluid composition of silicone renders it susceptible to cold flow away from areas of higher pressure. (Courtesy St Jude Medical, Sylmar, Calif.)

blending of insulating materials results in the product having the strength, lubricity and abrasion resistance of polyurethane, but the flexibility and biostability of silicone.39 Early bench and clinical testing have been promising to date, with little evidence of cold flow, ESC, or MIO phenomena, the main weaknesses of silicone and polyurethane. The material has been used in the St. Jude Medical Optim family of pacing (Models 1888T and 2088T) and ICD leads (Durata), but no long-term data on its chronic performance are yet available.

131

although vitreous carbon cathodes have particular strength, inertness, and long-term electrical reliability.45 More important than the precise material composition of electrodes seems to be their geometry, particularly the surface area. A large cathodal surface area in contact with the endocardium results in a low current density at the electrode-tissue interface and consequently a higher capture threshold. As tip electrodes were reduced in surface area, apart from better thresholds, higher tissue-electrode impedances were also seen, and this increased pulse generator longevity.46,47 Against this beneficial effect of a smaller electrode is the adverse effect on tissue polarization seen with a smaller electrode surface area. Polarization refers to the electrical afterpotential generated at the electrode-tissue interface by ionic movement induced by the pacing stimulus. At pacing pulse delivery, a layer of positively charged cations is attracted to and surrounds the negatively charged cathode (and a larger layer of negatively charged anions surrounds the cation layer), which opposes current flow from the pacing electrode. This polarization potential is recorded as an afterpotential following the stimulus artifact. To maximize the advantages of small-radius electrodes, yet minimize the disadvantage of the polarization effect more prominent with smaller electrode surface area, current electrodes are small in size but have a complex, porous microscopic surface (Fig. 4-8). Regardless of the method of achieving a complex, highly textured surface, such as coating the electrode with microspheres or a woven mesh of microscopic metallic fibers, the resulting large surface area has proved beneficial with regard to pacing and sensing performance. Titanium nitride–coated or iridium-coated cathodes with fractal surface structure achieve the goal of enhanced surface area with a negligible polarization signal48 (Fig. 4-9). Such electrodes have permitted the development of beat-by-beat automatic capture threshold and pacing output management systems that measure the evoked myocardial response following every pacing stimulus. By verifying myocardial capture for each paced beat, with the immediate delivery of a backup higher-output pulse if capture is lost (Fig. 4-10), the pacemaker can maintain a pacing output only marginally above the capture threshold, thereby prolonging pulse generator battery life.49 Another type of automatic threshold algorithm does not verify beat-to-beat capture, but performs an automatic threshold test daily and adjusts the pacing output accordingly (with larger safety margin above capture

ELECTRODES Electrodes represent the final common electrical and mechanical interface between the lead and myocardium. The design and construction of the electrodes influence the acute and chronic electrical performance of the lead. Because the electrode type also determines the fixation mechanisms employed and the lead tip pressure generated, however, mechanical problems such as lead dislodgement and cardiac perforation can also be affected by the tip electrode. The materials used for lead electrodes have evolved over the past 40 years and have included titanium, anodized platinum, platinumiridium alloys, Elgiloy (an alloy of cobalt, iron, chromium, molybdenum, nickel, and manganese), and vitreous carbon. All these display relative inertness in vivo, which minimizes inflammatory reactions and corrosion over the long term. Minor corrosion is seen, however, with Elgiloy and platinum alloys when in situ for extended periods, although this is of uncertain clinical significance.40 While attempting to control for other differences, several human and animal studies have compared these materials, but the results are conflicting.41-44 In general, applications have been found for these materials in electrode composition, with most having reasonable in vivo performance and longevity,

Figure 4-8  Sintered porous electrode (left) and laser porous electrode (right). (From Hirshorn MS, Holley LK, Skalsky M, et al: Characteristics of advanced porous and textured surface pacemaker electrodes. PACE 6:525-536, 1983.)

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SECTION 1  Basic Principles of Device Therapy

A

B

Figure 4-9  Magnified view (A) and electron micrograph (B) of fractally coated, titanium nitride electrode tip, which optimizes surface area and minimizes polarization resistance. (Courtesy St Jude Medical, Sylmar, Calif.)

threshold). These types of automatic capture-detection algorithms can be unreliable in the setting of a significant polarization afterpotential, because a large polarization signal cannot be distinguished from the local myocardial evoked response.50 Electrodes implanted in myocardium elicit a strong foreign body granulomatous reaction that results in the formation of a local fibrous A

B

capsule at the lead tip, which can have a profound effect on electrical performance. Smooth-surfaced electrodes elicit the formation of a thick, electrically insulating capsule of fibrous and granulation tissue that may impede electrode fixation to the endocardium.51 Complex, porous electrodes, in contrast, are manufactured to have extremely small (20-50 µ in diameter) surface pores that allow for endocardial C

D

II

V5

B 2 1

3

Figure 4-10  Automatic loss of capture algorithm (Autocapture). The atrioventricular (AV) delay is shortened to ensure ventricular pacing (A). The pacing output is automatically decreased until loss of capture is detected by failure to sense a myocardial-evoked response at B2. A backup 5-V pulse is delivered at B3, followed by an automatic gradual increase in pacing output that restores myocardial capture (C), and then resumption of dual-chamber pacing is seen, with the programmed AV delay (D). Modern leads produce negligible polarization afterpotential and thus allow for high-fidelity sensing of the myocardial-evoked response. (Courtesy St Jude Medical, Sylmar, Calif.)



4  Engineering and Construction of Pacemaker and ICD Leads

Steroid plug

133

1699T Optisense, St. Jude Medical) with a standard 10-mm spacing, the close-spaced bipole showed minimal far-field ventricular sensing at a sensitivity of 0.3 mV, compared to 30% 1-year incidence in the standard-bipole group.56 This translated into a large reduction in inappropriate mode switch events, from 23% in the control group to 4% in the narrow-bipole group. FIXATION MECHANISMS

A

Steroid plug

B Figure 4-11  Dexamethasone-eluting reservoirs in passive-fixation (A) and active-fixation (B) leads. (Courtesy St Jude Medical, Sylmar, Calif.)

attachment by way of tissue ingrowth into these regions, which yields a mechanical advantage in addition to the electrical benefits mentioned earlier. A fibrous reaction also occurs with highly textured electrodes, but a lesser insulating effect is observed because the electrically active myocardium is closer to the electrode,51 resulting in improved electrical parameters over time.52 To curtail further the fibrous reaction elicited by the electrode and to decrease the incidence and magnitude of chronic threshold rises, a pharmacologic anti-inflammatory strategy was pursued with glucocorticoid-eluting porous electrodes (Fig. 4-11). These leads showed excellent long-term capture and sensing parameters, with a dramatic reduction in late exit block (i.e., loss of myocardial capture).8 In an elegant and seminal double-blind randomized trial, Mond et al.9 conclusively proved the beneficial effect of steroid elution by comparing two otherwise identical porous titanium electrodes, only one of which incorporated dexamethasone in its tip. There was no difference in capture threshold seen at implant. Although there was an expected late rise in threshold in the standard leads starting at 2 weeks, this was completely abolished in the steroid-eluting lead, both during the trial period and in long-term follow-up.53 Animal studies confirmed that the biologic basis of this electrophysiologic effect was through the glucocorticoid-induced attenuation of the inflammatory reaction, with thinner, less cellular fibrous capsules.54 In bipolar leads, another factor affecting electrical performance, particularly far-field sensing, is the interelectrode distance between the tip cathode and the ring anode. Narrow-spaced bipolar electrodes in atrial leads have been shown to reduce far-field ventricular electrogram sensing without compromising atrial sensing.55 In a randomized trial that compared an atrial lead with a 1.1-mm tip-to-ring spacing (Model

Initial pacemaker leads had no fixation mechanisms, so lead dislodgement rates were high. The first solutions to this problem were wedgeshaped flanges and fins, but these had limitations. The development of scalloped tines was the first robust solution to the problem of lead displacement,5 and comparisons with earlier methods of passive fixation were favorable.57 Tines are usually constructed from the same material as the covering insulation of the lead, typically silicone rubber, and protrude backward from the base of the tip electrode. Their orientation is designed to allow advancement of the lead for initial implantation but to prevent retraction and dislodgement by engaging the myocardial trabeculae of the right atrial appendage and right ventricular apex (Fig. 4-12, A). Although passive-fixation tined leads perform well, with low rates of dislodgement, they are less reliable for pacing at nontraditional sites, and the development of active-fixation leads with a helix at the tip improved the options for pacing site selection (Fig. 4-12, B). These leads generally have an extendable-retractable screw made of platinumiridium or similar alloy that is usually electrically active as part of the cathode. This terminal helix is deployed (i.e., extended out of lead) with clockwise rotation of the connector pin, which transmits torque via the central coil conductor to the helix mechanism at the other end of the lead. A small number of leads use a stylet-driven extension-retraction mechanism. Some smaller-diameter leads have a fixed helix that requires rotation of the entire lead body around the stylet for deployment into the myocardium. Steroid elution from a reservoir at the base of the helix helps to dampen the inflammatory reaction produced by traumatic engagement of the myocardium by the screw. Active- and passive-fixation leads perform similarly overall,58 with some differences related to their construction. Active leads tend to be easier to extract, whereas passive leads tend to have lower chronic thresholds with higher impedances that prolong pulse generator longevity. The clinician must always consider the potential for perforation59 and trauma to surrounding structures60 when using active-fixation leads.61

A

Passive fixation

B

Active fixation

Figure 4-12  Lead fixation mechanisms. A, Passive-fixation tined lead. The tines are constructed of outer insulation material and facilitate advancement of the lead but prevent its retraction by engaging myocardial trabeculae in the right ventricle or right atrial appendage.  B, Extendable-retractable helix is deployed to screw directly into the myocardium in an active-fixation lead. (Courtesy St Jude Medical, Sylmar, Calif.)

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SECTION 1  Basic Principles of Device Therapy

Figure 4-13  Lead anchoring sleeves. Each sleeve contains up to three grooves to allow for suturing the sleeve to the lead body, then to the underlying deep fascia in the pocket. (Courtesy St Jude Medical, Sylmar, Calif.)

IS-1 (pace/sense)

A

LEAD ANCHORING SLEEVES Securing the lead body to the fascia or muscle at the floor of the device pocket at implantation is an important step to prevent lead dislodgement. The sutures used must be fastened firmly but, given the relative fragility of lead insulation as previously outlined, must not be allowed to compromise the lead structure or integrity (Fig. 4-13). Although the anchoring sleeve provides a mechanism to achieve this balance, there has been little development in this area. After being tied down tightly, sleeves made of MDX silicone were found to produce much lower lead deformation than ETR silicone,62 but no long-term in vivo data are available. Suture sleeves that grasp the lead by snapping into place (rather than relying on circumferential deformation of sleeve by encircling suture) have been proposed but are not in current use. CONNECTOR TERMINAL PINS The connector terminal pin forms the interface of the lead with the header of the pulse generator and has gone through significant evolution. With successive generations of pulse generator and lead technology, various standards for connector terminal pins for pacing leads have been used,63,64 as well as two different pin-fixation mechanisms within the device header, the side lock, and the now-universal set screw. Reliability has been high with both pin-header fixation interfaces, although the more forgiving tolerance of the set-screw platform has led to its widespread adoption.65 This design requires a self-sealing grommet to cover the head of the set screw and prevent corrosion of the pin or header connections because pocket fluid enters the header. Sealing fins are required around the lead body or within the header channel to prevent fluid leaking into the header at that site, regardless of the fixation mechanism. Different patterns of connector pins include the 5- and 6-mm-long pins, the 3.2-mm low-profile pin, the VS-1 pin, and the current standard IS-1 connector pin (Fig. 4-14). Compatibility CONNECTOR SIZE OR STANDARD

6 mm, 5/6 mm, 5 mm

3.2 mm

IS−1 Figure 4-14  Evolution of pace/sense connector pins, culminating in current IS-1 standard (bottom).

• 3.2 mm diameter • no sealing rings in header • short receptacle for B lead terminal

IS-1

Figure 4-15  A, IS-1 pace/sense connector pin terminal. B, Cross section of IS-1 header.

between terminal pin type and header channel type is still of major concern when changing pulse generators in patients with older leads. Modern pulse generator headers conform to IS-1 specifications (Fig. 4-15) and are backwards-compatible with VS-1 pins. Occasionally, for example, adapters may be required to downsize a long pin connector to an IS-1 header.

  Resynchronization Leads The symptomatic and prognostic benefits of cardiac resynchronization therapy (CRT) in congestive heart failure patients with electrical dyssynchrony (asynchrony), usually in the form of native left bundle branch block (LBBB) or right ventricular pacing, have been repeatedly demonstrated in those already receiving maximal medical therapy.66,67 Leads used for providing CRT are exposed to different implant conditions than conventional pacing leads and thus have different design parameters. The benefits of CRT were first demonstrated with the placement of left ventricular (LV) epicardial leads. CRT leads can be implanted directly on the epicardium by an open or video-assisted thoracotomy. These leads are, or have similar characteristics to, standard epicardial pacing leads (Fig. 4-16). They are generally sutured or screwed in place on the lateral LV wall. Increasingly, patients who undergo cardiac surgery for coronary disease or valve dysfunction, but who also have heart failure, LV dysfunction, or a chronic need for pacing because of heart block, are having LV epicardial leads implanted at surgery, with the terminal pin end tunneled up to the left infraclavicular fossa for future connection to a CRT device. However, a transvenous approach to LV epicardial lead placement avoids the morbidity and mortality of an open surgical approach and is usually the preferred initial implantation method for patients not otherwise scheduled for cardiac surgery. Therefore, most candidates for CRT therapy have their LV lead implanted transvenously into one of the venous tributaries of the coronary sinus. Compared with standard endocardial pacing, this technique has two challenges:(1) placing the lead into a desirable location despite venous tortuosities, acutely angled branch ostia, obstructive venous valves, patches of myocardial scar with high capture threshold, and regions of phrenic nerve capture, and (2) once placed, achieving stable lead fixation within the thinwalled coronary vein (vs. relative ease of standard tined electrode or active-fixation helix obtaining purchase on endocardial wall of cardiac chamber).



4  Engineering and Construction of Pacemaker and ICD Leads

Figure 4-16  Epicardial active-fixation steroid-eluting leads, implanted surgically. Lead on the left has an active-fixation helix, whereas two electrodes on the right are sutured to the epicardium. (Courtesy St Jude Medical, Sylmar, Calif.)

LEAD PLACEMENT AND FIXATION MECHANISMS For accessing the stimulation site, CRT lead systems come with a range of outer guide introducers to facilitate coronary sinus cannulation and provide support during lead placement. Although the initial transvenous CRT leads were stylet driven, over-the-wire leads were rapidly introduced because of difficulty negotiating venous tortuosities and acute takeoff angles. The development of an over-the-wire system was a major advance,68 and such leads are routinely required to select and place leads in the best possible target venous branches, usually over the posterolateral left ventricle. These leads allow for front or back loading of a 0.014-inch angioplasty wire into the stylet lumen. With or without the assistance of inner guiding catheters, the wire may be used to subselect target venous branches (including those with acute takeoff angles) and the lead may then be tracked over the pre-positioned wire. Obtaining distal vascular access with a wire also allows for venoplasty to facilitate lead placement in the setting of stenotic vascular segments.

135

Once at the target site, and with retraction of the wire or stylet back into the lead body, most lead tips have some degree of memory and assume the form of a preshaped curvature, which may be a smooth curve, an angled cant, or a spiral (Fig. 4-17). These distal lead shapes provide some degree of stability in the target vein by exerting modest lateral pressure on the inner vein surface, thereby increasing friction and a resistance to forward or backward migration. LV lead models that have no distal curvature use small, distal tines to engage the venous wall, but these tines likely have a reduced ability to fix the lead passively in place, particularly in larger vein branches. Thus, stability of tined leads probably relies more on lead flexibility and the cumulative friction along the lead length wherever it is in contact with the venous wall. These straight leads seem to have a higher risk of dislodgement, and subsequent generations of CRT leads have not incorporated tines. In most cases the major contributor to LV lead fixation in the coronary venous system has been the extent to which the outer lead diameter is circumferentially opposed to the inside wall of a venous tributary of given caliber. If the lead has been advanced snugly into a vein branch, significant fixation friction forces may be generated, but lead stability remains variable. The risk of dislodgement also depends on upstream factors, including the number and size of redundant loops and tortuosities in the proximal course of the lead body, cardiac motion, tricuspid valve regurgitation, and interactions with other intracardiac hardware, which can each generate traction forces on the LV lead. Consequently, there is a significant rate of CRT lead displacement of up to 1% to 5% that may require surgical lead revision if there is a significant change in electrode position, ventricular capture threshold, or phrenic capture.66,69,70 The concept of active-fixation leads in the coronary venous system has been brought to clinical practice with the Attain StarFix lead (Model 4195, Medtronic, Minneapolis), which has three sets of deployable lobes that fold out and protrude from the lead body just proximal to the tip (Fig. 4-18). These 12 circumferential lobes are extended by an external sliding-push tubing mechanism, and radiopaque markers are used to indicate whether one, two, or all three sets of lobes are extended or retracted. In a recent multicenter study of 408 implants, Crossley et al.71 found a very low lead dislodgment rate of 0.7% with stable electrical parameters over a mean follow-up of 23 months. Phrenic nerve capture with this unipolar lead required lead revision in 11 patients (2.5%). This reasonable performance must be weighed, however, against potential difficulties of extraction, which fortunately has not been a major issue in the other LV lead models to date. In this study of the active-fixation LV lead, 26 patients (6.4%) underwent an

Figure 4-17  Representative variety of currently available resynchronization leads with differing passive-fixation curvatures. (Courtesy Medtronic, Minneapolis; St Jude Medical, Sylmar, Calif; and Boston Scientific, St Paul, Minn.)

136

SECTION 1  Basic Principles of Device Therapy

output is inadvertently programmed below the true LV capture threshold, because RV anodal capture is mistaken for LV cathodal capture during threshold testing.75 No significant differences have been noted in the short-term and long-term complication rates after CRT implantation among the various leads and CRT systems commercially available.76

ICD Leads

Figure 4-18  Active-fixation resynchronization lead (Attain StarFix, Model 4195) with deployable lobes to grip the lumen of the coronary venous tributary in which it is implanted. (Courtesy Medtronic, Minneapolis.)

attempt at lead revision. In two patients, no attempt to remove the lead was made, and in the remaining 24 revisions, five attempts failed (21%). These patients had maximum implant duration of only 941 days. Of the seven patients who had a lead in situ for more than 1 year, four (57%) were unable to be extracted. This study is similar to the experience of Nagele et al.,72 who encountered extraction difficulty caused by immobile lobes in two patients. The ability to un-deploy the lobes seems to be independent of the duration of implant and may be related to local branch vein thrombosis or an ingrowth of fibrous tissue through the lobe loops. Although both these studies were small and represented early experience with this lead, the difficulties with extraction of the StarFix lead contrast with the relative ease of extraction of chronic LV leads, even after extended periods since implantation.73 DESIGN OF CRT LEADS The original CRT leads were exclusively unipolar because of size considerations, given that LV leads had to be advanced through a long, smaller-caliber sheath that was used to engage the coronary sinus selectively. An early example was the EasyTrak lead family (Models 4510/11/12/13, Guidant, St. Paul, Minn.), which had a 6F lead body diameter and an LV-1 connector terminal pin compatible only with LV-1 headers on Guidant CRT pulse generators. Since then, numerous developments in lead size, polarity, and CRT device programmability have greatly increased the options available to the implanting operator, including lead and device options to manage the problem of phrenic nerve capture, as well as to find suitable pacing configurations in the setting of myocardial scar. Most significantly, bipolar CRT leads were developed that were no larger in diameter than the previous generation of unipolar leads and that could therefore be deployed through the same sheath delivery systems and into venous branches of the same caliber. Newer LV lead versions include 4F to 5F leads that can be advanced through smaller venous subselection sheaths, into more diminutive vein branches, which increases the number of lead positioning options. The interelectrode distance in bipolar CRT leads varies from 8 to 21 mm. The great advantage of bipolar leads is the option of programming between multiple different pacing electrode permutations, including true bipolar and extended bipolar configurations, with the option of selecting whether to use the LV tip or the LV ring electrode as the pacing cathode (Fig. 4-19). These programming options may be invaluable in the patient with posterolateral LV fibrosis or a large field of phrenic nerve capture, where an alternate pacing polarity may obviate the need for lead revision. True bipolar LV pacing also eliminates the problem of simultaneous right ventricular (RV) anodal capture, which can be seen in the setting of unipolar LV lead use and can sometimes partially undermine the benefit of LV pacing.74 Anodal capture from the RV ring or coil electrode can hinder proper CRT delivery by eliminating the ability to stimulate the left before the right ventricle, and occasionally by leading to absent LV pacing if the pacing

The development of the implantable cardioverter-defibrillator was a revolutionary step in the history of medicine.77 Sudden-death prevention became an achievable goal, initially in those who had survived cardiac arrest or malignant ventricular arrhythmias,78 but then in patients at risk of dying suddenly (primary prevention).79,80 Modern ICD systems are extremely sophisticated devices, frequently combined with CRT capability, that feature advanced rhythm-discrimination software, antitachycardia pacing (ATP) during capacitor charging, and high-voltage biphasic shocks, all in a pectorally implanted device of 40 cc or less. Advances in pulse generator technology and reliability are important but of limited relevance if progress cannot be matched in ICD lead design, development, and particularly durability. The reductionist view of ICD leads as simply being “pacemaker leads with additional shock coils” ignores the exponential rise in complexity as more components are added. This complexity puts pressure on long-term reliability in a device expected to cope repeatedly with 800-V highcurrent discharges and move with the heartbeat 37 million times annually. As longer follow-up data are published on both older and newer ICD leads, despite the innovations and developments in recent years, clearly this most vulnerable part of the ICD system continues to be prone to late failure.81-84 HISTORICAL EVOLUTION OF ICD LEADS Epicardial patches placed through thoracotomy were the leads first used in human ICD systems.77 The pulse generators in these systems were large and required placement in an abdominal pocket. Shock energy was delivered through a free-floating 12F nonpacing endovascular titanium spring coil that served as the anode, with the patch as the cathode. Sensing was accomplished initially by epicardial patches (or apical cup electrode in the first implant), but this sensing mechanism proved to be unreliable, and dedicated epicardial and endocardial leads were adopted early. These first-generation ICD systems had a high risk of complications, particularly in the large endovascular coil,

6

4

2

5

1 LV lead

RV lead

3

Figure 4-19  Multiple different pacing polarity configurations are possible with modern bipolar resynchronization leads and generators. Each arrow originates at the cathode and ends at the anode of current left ventricular (LV) lead pacing configurations. RV, Right ventricular.



4  Engineering and Construction of Pacemaker and ICD Leads

including superior vena cava thrombosis, coil fracture, insulation failure, and coil retraction into the subclavian vein caused by lack of a lead anchoring sleeve. Consequently, the second-generation ICD devices in the late 1980s were fully epicardial systems, with shock delivery between two epicardial patches and tachyarrhythmia sensing through an epicardial lead. Epicardial patches also had problems, however, including an ongoing infection and pericarditis risk,85 a possible adverse effect on transthoracic external defibrillation threshold, and a perioperative mortality of up to 4% at implantation in these patients, who were generally at high risk for thoracotomy or median sternotomy.86,87 These patch electrodes also displayed poor in vivo durability over time, with a failure rate of up to 28% at 4 years. More than half of patients with obvious lead failure or fracture were asymptomatic. The epicardial pace/sense leads incorporated in these systems were also suspect, with a failure rate of 14.1% at 8.5 years in one study, with lead fracture accounting for half these failures.88 In contrast to patch failure, patients with epicardial pace/sense lead failure were likely to develop multiple inappropriate ICD shocks from oversensing.89 By the early 1990s, advances in pulse generator design and technology allowed miniaturization for infraclavicular pectoral implantation, although this would require an entirely endocardial defibrillation, sensing, and pacing lead system. The initial transvenous leads were implanted with the pulse generator in an abdominal pocket, and the leads were tunneled subcutaneously from their pectoral vascular insertion site down to the abdomen. Such leads, including the Transvene family (Medtronic), were initially based on coaxial pacing lead platforms. Their development added ATP capability to ICD systems, but also new lead-related problems. Apart from the large bulk of these leads (up to 12F), predisposing to venous occlusion and subclavian crush,90 lead failure was frequent, demonstrating a progressive late incidence related to both conductor fracture and insulation defects.91 Addressing these problems required paradigm shifts in lead body design as well as advancements in materials technology, which ushered in the modern era of multilumen, narrow-caliber transvenous ICD leads. ICD LEAD STRUCTURE The basic elements in pacing leads are also found in transvenous ICD leads: conductors, insulation, electrodes, fixation mechanisms, and connector terminals. In addition, however, ICD leads contain the highvoltage circuitry that links the pulse generator with one or two highvoltage shock coils located on the lead body (Fig. 4-20). One of these coils is located in the right ventricle, a variable distance proximal to the tip pace/sense electrode, whereas the other coil, if present, is positioned on the part of the lead located in the superior vena cava.

A

137

Two basic lead body architectures can be used to incorporate all these elements into an ICD lead. First, the coaxial structure, a design essentially borrowed from the pacing platform of the same name, consists of concentric coiled conductors (to tip electrode, ring electrode, and shock coils) layered inside one another and separated by intervening layers of insulation, with the whole assembly housed in an outer covering layer of insulation. The innermost coil connected down to the tip conductor contained a lumen for stylet insertion that assisted with endocardial lead implantation. The best known examples of this design were the Medtronic Transvene and the Ventritex TVL families of ICD leads, first used in the early 1990s. Although initial reports were positive,92 subsequent longitudinal data showed that these leads were prone to early and late failure, caused by a combination of conductor and insulation problems.83,91 Whether these failures were mainly related to the limited durability of the conductor and insulation (polyurethane 80A) materials, or resulted from an inherent weakness of the bulky coaxial architecture in situations of increased complexity as found in ICD leads, will never be known; this design was no longer in use by the late 1990s. The second main paradigm for ICD lead structure, and still the dominant architecture for modern ICD leads, is the multilumen design, in which parallel cable and coil conductors run to the tip, with ring and shock electrodes in separate channel lumens, all within a single frame of insulation; some models also have additional, empty decompression lumens. The conductor to the tip electrode is conventionally coiled and contains the central lumen for stylet insertion. The coil also permits the transfer of torque when the implanting physician rotates the pace/sense connector terminal pin to deploy an active-fixation helix at the distal lead tip. Each conductor, both cable and coil, is invested with an inner layer of insulation, typically of fluoropolymer type, and the entire lead body is encased in a further outer insulation layer (see Fig. 4-3). ICD LEAD SENSING DESIGN Sensing function is critical in ICD systems to allow for accurate rhythm discrimination and to prevent inappropriate shocks from oversensing as well as failure of tachyarrhythmia treatment caused by undersensing. Therefore, unipolar ICD leads were not developed because of the prohibitive risk of oversensing noncardiac potentials, such as pectoral and diaphragmatic myopotentials. Bipolar ICD lead sensing can be arranged according to one of two paradigms: integrated bipolar and dedicated bipolar (also called true bipolar) designs. Sensing in integrated bipolar leads occurs between the tip electrode and the RV shock coil, which doubles as the sensing anode, the distal extent of which is located 6 to 15 mm proximal to the tip. By comparison, dedicated bipolar leads additionally include a conventional ring electrode (the

B

Figure 4-20  A, Dual-coil ICD lead with second high-voltage shock coil residing in the superior vena cava. B, Single-coil ICD lead with one shock coil in the right ventricle only. The active ICD pulse generator generally serves as the cathode or anode for the first phase of biphasic shocks.

138

SECTION 1  Basic Principles of Device Therapy

Integrated bipole lead

12 mm RV high-voltage electrode

A Dedicated bipole lead RV high-voltage electrode

B

25 mm

Ring for pacing and sensing

Figure 4-21  Design pullback. “Pullback” refers to the distance between the lower active edge of the right ventricular (RV) defibrillation electrode and the RV apex. Design pullback should be minimized for lower defibrillation thresholds. The diagrams show the impact of sensing methods on catheter design and the magnitude of the RV electrode pullback. A, Leads using integrated bipolar sensing/pacing methods (from tip to RV defibrillation electrode) typically have a shorter pullback distance (e.g., 12 mm here). B, Leads using dedicated bipolar sensing/ pacing methods have larger distances from catheter tip to active RV defibrillation electrode because of sensing ring (e.g., 25 mm here).

anode) 10 to 15 mm back from the tip, similar to a bipolar pacing lead (Fig. 4-21). The additional electrode means that an extra conductor is required in dedicated bipolar leads, but the smaller anode in these leads confers superior sensing performance with reduced far-field oversensing.93,94 The larger sensing anode (the distal shock coil) in integrated bipolar leads may be of particular concern in patients with small right ventricles, where the proximal end of the coil may extend across the tricuspid anulus and thereby pose a risk of far-field atrial oversensing.93 Integrated bipolar leads may be more prone to oversensing highfrequency respirophasic noise transients (i.e., diaphragmatic myopotentials), which may cause pacing inhibition and inappropriate shocks.95 The increased current density at the smaller anode of a dedicated bipolar lead does, however, increase the likelihood of anodal myocardial stimulation in CRT systems in which LV pacing is programmed in an “extended bipolar” fashion, with the RV ring doubling as the LV pacing anode,75 although in the current era of bipolar LV leads and pacing vector programmability, this is less of a concern. The main theoretical advantage of integrated bipolar leads, apart from their greater simplicity (which may reduce the chance for component failure), relates to the smaller distance between the RV shock coil in these leads and the RV apex, because there is no intervening ring electrode between the shock coil and the lead tip. This is known as the pullback distance, and data suggest a modestly improved defibrillation threshold with shorter pullback distance,96,97 although this was more important before the current era of high-output, biphasic devices and greater defibrillation safety margins. Theoretical concerns about differences in R-wave amplitudes between the two lead designs have not been borne out by comparative studies,98 but some studies suggest longer sensing latency98 with integrated bipolar leads, as well as the potential for undersensing of postshock ventricular fibrillation, presumably from local myocardial stunning around the shock coil and its proximity to the tip cathode.99,100 This phenomenon was not seen when the shock coil was more than 6 mm from the tip electrode. However, in terms of sensing de novo ventricular fibrillation, no significant differences have been demonstrated between the two lead platforms,101 and both are in clinical use today, although dedicated bipolar systems are implanted more frequently. CONDUCTORS As with pacing leads, similar materials are used in ICD lead conductor construction, with MP-35N being the dominant metal alloy employed. As outlined earlier, its advantages include its corrosion and fracture

resistance, but MP-35N–based conductors can still fracture at points of high stress (e.g., costoclavicular ligament). Recent testing suggests that this risk may be related to the presence of contaminant titanium inclusions, although this is controversial.102 In ICD systems, however, low-resistance conductors are essential to minimize voltage drops across the lead during high-voltage shock delivery. This requirement mandates the use of composite wire technology in which individual conductor strands combine higher-resistance MP-35N with a lowresistance metal such as silver in one of two patterns. The DFT pattern (see Fig. 4-5, A) produces conductor strands that are generally formed into coils, whereas in the DBS pattern (see Fig. 4-5, B), strands are twisted into multifilar braided cables (Fig. 4-22). The latter are particularly important in ICD leads because they have very high tensile strength and are more fracture resistant than coils. Failed ICD leads frequently demonstrate intact high-voltage cables with fractured pace/sense coil conductors. One conductor coil strand is generally required, as outlined earlier, to create a lumen for stylet insertion and to permit active-fixation helix deployment. However, the use of multiple parallel cables in a multilumen ICD lead design produces a stronger lead of smaller diameter than would be possible if only coiled conductors were used (as in the now-obsolete coaxial configurations). Each cable generally is housed in its own lumen of the multilumen lead by coating it with stiff, fluoropolymer insulation, preventing abrasion and damage to the soft, silicone-based insulation of the lead body (see Fig. 4-3). Decompression spaces, either within the lumen that houses a conductor or as separate channels in the housing insulation, may reduce the stress on the conductors, particularly in the setting of repeated lead flexion. ELECTRODES AND HIGH-VOLTAGE COILS Electrode composition and design in ICD leads conform to the same basic principles as outlined earlier for pacing leads, with steroid elution, current density, and polarization considerations again dictating the material composition of the tip cathode and ring anode in dedicated bipolar leads. Fixation mechanisms are also similar, with either electrically active extendable-retractable helices or flexible tines being employed. The high-voltage defibrillation shock coil is the defining feature of the ICD lead. All models include at least one shock coil, wrapped around the distal lead body, 5 to 6 cm in length and positioned in the right ventricle, whereas dual-coil leads have a second shock coil located either 17 cm or 21 cm proximal to the shaft end of the RV coil, with this superior vena cava (SVC) coil being slightly longer, 7 to 8 cm (see Fig. 4-20). The RV coil serves as the defibrillation anode or cathode during phase 1 of a biphasic shock, depending on how the ICD is programmed to deliver high-voltage therapy, with the SVC coil and active pulse generator casing together serving as the other shock cathode or anode. With single-coil ICD leads, the pulse generator solely functions as the second high-voltage electrode. Dual-coil leads may assist in lowering the defibrillation threshold (DFT), although the improvement in DFT with the addition of an SVC coil may be modest

Cable

ICD lead Coil Figure 4-22  Filar composition of unraveled high-voltage conductor cable alongside multifilar central pace/sense conductor coil in dissected ICD lead.



4  Engineering and Construction of Pacemaker and ICD Leads

and of greater benefit when the electrical impedance during single-coil shock delivery is higher.103 Because the high-voltage shock vector plays a role in DFT, some ICDs permit programming changes to eliminate either the SVC coil or the pulse generator itself from the shocking circuit, thereby improving the DFT in certain patients. Shocking coil– only leads are also available for implantation in the coronary sinus or azygous vein, or subcutaneously around the left thorax to improve the shock vector in patients with an elevated DFT and unreliable defibrillation with standard shock configurations. The materials used in shock coil construction are generally platinumiridium alloy or platinum alloy–clad tantalum, because titanium coils produce metallic oxides during high-voltage shocks that prevent their use in ICD leads. Polarization also affects the design of shock electrodes, as it does for pacing electrodes, but unlike the situation described earlier with pacing leads, fractally coating the shock coils does not improve defibrillation efficacy.104 However, some data suggest that iridium oxide coating may improve lead porosity and reduce polarization, even to the point of seeing a clinically meaningful reduction in DFT.105,106 Although an essential part of the ICD lead, the presence of endovascular defibrillation coils poses distinct physical problems, especially related to fibrous tissue ingrowth, which can make lead extraction difficult and potentially hazardous, as discussed later. INSULATION Essentially the same materials used in the manufacture of pacing lead insulation are also used in the context of ICD leads, and their respective advantages and disadvantages are discussed earlier. The multilumen design of ICD leads relies on a single, large-caliber body of insulation that contains the conductor lumens. Other insulation elements may include the jackets investing the conductor cables or an additional outer insulation covering, but the mechanical characteristics of the lead will largely be determined by the material used to construct the lead body. Because ICD leads are larger and more complex than pacemaker leads, silicone rubber is the dominant option for ICD lead body insulation; polyurethane 55D and fluoropolymers are too stiff for this purpose, and polyurethane 80A is too susceptible to biodegradation. The relative structural advantages of polyurethane 55D, lubriciousness and high tear strength, can be used to good advantage instead as an outer insulation layer surrounding the lead body. In this location it is far away from the metal conductors within the thick silicone rubber casing and a fluoropolymer jacket. This barrier prevents exposure to metal ions from the MP-35N conductor that can trigger the oxidative degradation of the soft, polyurethane segment. Likewise, although the major structural weaknesses of fluoropolymers (related especially to rigidity) preclude their use as the primary insulator, inner insulation of the coil and cable conductors is well suited to these highly biostable, abrasion-resistant materials. Generally, ETFE is used for coating of the cable conductors, and PTFE is used to line the central lumen lead body that houses the coil conductor. Although silicone rubber may be biostable and flexible enough to be used as the dominant material for ICD lead body insulation, its softness and high frictional coefficient make lead implantation difficult if the outermost surface of the ICD lead is silicone. Outer coverings of polyurethane may address this disadvantage to some extent, but the longterm mechanical risks related to silicone’s softness remain, as well as its susceptibility to abrasion and cold flow, and require the use of thicker silicone layers and thus increased lead caliber. This weakness is particularly marked at high-stress points such as at the costoclavicular ligament or in the pocket next to the pulse generator, where harsh local mechanical forces can lead to progressive insulation abrasion and local silicone thinning at pressure points. If lead insulation is sufficiently thinned or denuded that a low-impedance dielectric is created, particularly where a high-voltage conductor strand closely opposes the metal casing of the pulse generator in the device pocket, a short circuit with arcing can result during shock delivery, leading to ineffective defibrillation and permanent damage to the lead, the ICD circuitry, or both.107

139

As mentioned in the section on pacing leads, newer insulation materials are being developed for specific application in cardiac device leads, including one already in clinical use, Elast-Eon (Aortech Biomaterials, Clayton, Victoria, Australia). This substance is a hybrid copolymer composed of near-equal amounts of silicone and polyurethane and, in extensive laboratory experience, combines advantages of each.39 A family of pacing and ICD leads uses this material as an outer insulator (Optim, St. Jude Medical), but as yet no leads use this material for the primary lead body insulation. ICD LEAD CONNECTOR TERMINALS As with the initial CRT leads, early ICD leads had high-voltage circuitry that terminated proximally in connector pins compatible only with pulse generators made by the same vendor. This was a major problem, particularly at generator change procedures, when lack of hardware cross-compatibility could limit device selection options and increase procedural complexity. The first solutions involved developing a wide range of adapters and extenders that could be used to switch connector pin design. However, this added further complexity to the system, increased the number of electrical connection points, and introduced another component that was subject to similar modes of failure as the rest of the lead. Eventually, industry-wide standardization occurred in ICD systems, such that the pace/sense coil conductor terminates in a standard IS-1 pacing pin that is placed in the IS-1 port of the ICD header, and each high-voltage shock coil terminates in a standard DF-1 (ISO-11318) connector pin. In a dual-coil ICD lead, therefore, three tails ending in one IS-1 pin and two DF-1 pins emerge from a trifurcation yoke attached to the main lead body (Fig. 4-23). Three separate connector ports are found on the ICD header to accommodate the trifurcated end of a dual-coil ICD lead (one IS-1 port and two DF-1 ports). The yoke and lead tails of the ICD lead are subject to insulation and

A

IS−1

B

DF−1

Trifurcation yoke

C

Dual-coil ICD lead

Figure 4-23  ICD lead connector terminals. A, Standard IS-1 pace/ sense connector terminal. B, High-voltage DF-1 pin connects shock coil(s) to generator header. C, Proximal end of dual-coil ICD lead, with central IS-1 pace/sense pin, two high-voltage DF-1 pins, and trifurcation yoke to interface the three terminal pins with lead body.

140

SECTION 1  Basic Principles of Device Therapy

Pin geometry Low voltage

EXTRACTION CONSIDERATIONS

High voltage DF4−LLHH

Low

A

Low

High

Low = pacing/sensing

B

High High = cardioversion/defribrillation DF−4

Figure 4-24  Schematic representation and photograph of the IS-4/ DF-4 connector pin incorporating high-voltage and low-voltage elements in the one pin and thus obviating the need for an external trifurcation yoke. (Courtesy St Jude Medical, Sylmar, Calif.)

conductor failure, just as in the rest of the lead system. They also increase pocket bulk and may contribute to erosion risk, particularly in the setting of a dual-chamber or biventricular device system, with atrial and LV leads also present. Additional abandoned lead ends might also be present in the device pocket, if nonfunctional or redundant leads remain in place in the setting of lead failure or device upgrade. Besides increasing pocket bulk and interaction between device system components, the greater number of connector pins in a standard ICD lead allows inadvertent connector pin reversal in the device header, and the thick fibrous tissue that frequently encapsulates the lead trifurcation yoke can make a generator change or system revision more challenging. The development of the new IS-4/DF-4 standard connector pin has allowed the integration of two DF-1 high-voltage terminals and one bipolar IS-1 pace/sense terminal, all into one connector pin (Fig. 4-24). This new design addresses the mechanical issues associated with current multiterminal ICD leads, but the long-term reliability of these connector pins is currently unknown.108 In addition, this all-in-one connector pin design precludes the ability to replace a single component of the ICD lead, such as the pace/sense or a shocking coil component, in the setting of pace/sense malfunction or elevated DFT.

A

C

Increasingly, ICD systems are being implanted in younger patients with evolving indications outside the context of acquired cardiomyopathy. Channelopathies, arrhythmogenic RV dysplasia, congenital cardiac abnormalities, and sarcoidosis are indications for implant that affect younger patients, often children, who are expected to live with their ICD leads (since generators are changed at regular intervals) over their lifetime. Because ICD lead failure rates continue to accrue over time with current lead and materials technology,81 and lead infection can occur at any time without warning,109 lead manufacturers and physicians must give due consideration to this aspect of device therapy at implant. In addition, abandoned sterile leads are increasingly being removed to allow for reduced intravascular hardware burden, avoidance of inappropriate oversensing,110 and potentially increased longterm venous patency.111 Lead extraction, even with the modern tools and expertise,112 remains a potentially hazardous endeavor,113,114 particularly in young patients with long-standing dual-coil ICD leads, in which tissue ingrowth into the interstices between the helical turns of the shock coil wire may be causing dense, fibrous adhesions and tissue encapsulation of the lead body (Fig. 4-25). With some data indicating that defibrillation efficacy may not be significantly improved with the addition of the SVC coil,115 one approach is to consider implanting only single-coil ICD leads in patients expected to have longer implant duration.116 However, the real solution lies in the manufacture of shock coils that are easier to extract. All newer ICD lead bodies are isodiametric, with the same or a smaller maximal external diameter as proceeding from the proximal to the distal end of the lead (Fig. 4-26), and are 8-Fr size or smaller, which already improves the ease of extraction over previous transvenous ICD lead systems. Manufacturers also have taken different approaches to the problem of shock coil adhesion to the venous and cardiac walls102 (Fig. 4-27). Silicone backfilling of the interstices between the coils has been used in Medtronic leads to reduce fibrous tissue ingrowth. This approach may be effective117 but is limited by the

B

D

Figure 4-25  A, Chronically implanted ICD lead shows extensive adhesions binding the lead tip and shock coil to the trabeculated ventricular endocardium. B, After lead extraction, extent of shock coil–associated fibrosis is evident. C and D, Cross-sectional photomicrographs of a chronically implanted lead show fibrous tissue invasion into interstices between helical turns of the shock coil. (C and D courtesy St Jude Medical, Sylmar, Calif.)



4  Engineering and Construction of Pacemaker and ICD Leads

141

tissue ingrowth into, and local adherence of, the shocking coils over time. This has made ICD lead extraction easier in animal models, compared with an identical lead without covering sleeves.117 Recent clinical data have corroborated the potential benefit of modified ICD lead shock coils to improve the ease of extraction.119

Step up in diameter

ICD LEAD FAILURE

Figure 4-26  Lead body diameter may not be uniform throughout (top) or may be isodiametric (bottom), which improves the ability to extract chronically implanted leads. (Courtesy Medtronic, Minneapolis.)

surface area of the outer curvature of the round coil strands; the interstices cannot be fully obliterated because a large, exposed coil surface area is required for effective defibrillation. In Biotronik and St. Jude Medical leads, use of flatwire technology allows complete backfilling of the interhelix turns with silicone, allowing for a completely isodiametric shock coil segment while still reducing fibrous tissue ingrowth, as demonstrated in animal testing.102 A third strategy, employed in Boston Scientific leads, is to cover the defibrillation coils with a thin, encapsulating jacket of semiporous expanded polytetrafluoroethylene (ePTFE; Gore-Tex, Gore and Associates, Flagstaff, Ariz). As with all fluoropolymers, this material is abrasion and tear resistant as well as biostable, but it has been molecularly modified to make it semiporous. This allows for free fluid and ion flow so as not to significantly alter high-voltage electrical conductance during shocks from the underlying defibrillation coils. Consequently, no significant difference in clinical defibrillation efficacy has been demonstrated with the use of ePTFE coils.118 In the setting of an integrated bipolar sensing ICD lead and additional abandoned leads in the right ventricle, mechanical interactions between the large RV coil and adjacent lead materials may occasionally result in sensed electrical artifacts. Oversensing from this type of lead-on-lead interaction may be reduced or eliminated if the metal strands of the RV coil are covered by an ePTFE sleeve.110 The real advantage of the semiporous sleeve, however, is that cellular migration cannot occur into the covered underlying coil interstices, thus reducing Figure 4-27  Solutions to the problem of shock coil fibrous tissue ingrowth. Top panel shows the shock coil interstices backfilled with silicone. Middle two panels demonstrate effective backfilling between flatwire helix turns. Bottom two panels show an ePTFE covering sleeve over the shock coil. (Courtesy Medtronic, Minneapolis; St Jude Medical, Sylmar, Calif; and Boston Scientific, St Paul, Minn.)

The ICD lead is one of the most complex and vulnerable parts of the ICD system, and lead failure is a significant problem (Box 4-1). ICD lead failure, perhaps more so than with pacing lead failure, can lead to sudden death. This can occur not only by loss of pacing support in pacemaker-dependent patients, but as most often seen, also by causing recurrent inappropriate shocks due to oversensing of lead noise. These shocks can induce ventricular fibrillation, which the system may then not be able to defibrillate successfully.120 Early data from the transvenous lead experience held great promise,121 but late lead failure was seen at longer-term follow-up across all lead designs and manufacturers.81 Patient-related factors, such as excessive upper body activity, can contribute to ICD lead failure, as can procedural factors at implant, including subclavian (rather than cephalic) access,122,123 rough handling, traumatic transvenous insertion, overtorquing of active-fixation helices, and kinking of lead loops in the pocket. However, none of these factors is positively identified in most cases; the difficulty in extracting ICD leads (possibly with additional lead damage during its extraction) means that the failure mechanism usually remains obscure. A systematic solution to this problem must involve better ICD lead design and manufacture, with advances in materials technology. The Transvene family (Medtronic) of coaxial ICD leads had a structural failure rate of 38% at 8 years,83 which was more often caused by fracture of the outer high-voltage coil conductor than disruption of the polyurethane 80A insulation. Despite changes in design and materials, however, newer generations of ICD leads are also prone to late failure. Kleemann et al.81 found that cumulative survival of 990 ICD leads implanted between 1992 and 2005 was only 60% at 8 years, mainly from insulation defects, including in silicone-based leads. One in five leads implanted for a decade or longer failed. This study was weighted toward the older lead models, however, and 95% of implants were through subclavian access, which may have affected failure rates and mechanisms.

INGROWTH SOLUTIONS

Silicone backfill

Flatwire design and backfill

ePTFE covering sleeve

142

SECTION 1  Basic Principles of Device Therapy

Box 4-1 

SYMPTOMS AND SIGNS OF LEAD PROBLEMS Pacing/Sensing Oversensing (most common) caused by detection of noise Inappropriate shocks Inappropriate aborted shocks Inhibition of pacing Change in pacing impedance Prolonged oversensing immediately after shock delivery Undersensing in sinus rhythm Undersensing or nonsensing of VT/VF, with failure to deliver therapy Elevated pacing thresholds or noncapture Defibrillation Ineffective defibrillation Change in high-voltage impedance Sudden death VT/VF, Ventricular tachycardia/ventricular fibrillation.

In an industry-wide push for increasing miniaturization, smallcaliber ICD leads were introduced in the mid-2000s. Again, although early data gave no cause for alarm,124 more recently125 the small-caliber (6.6F) Sprint Fidelis lead (Models 6930/6931/6949/6948, Medtronic)

has been identified as much more prone than other contemporary lead models to early and late failure, usually presenting with conductor fracture, noise oversensing, and inappropriate shocks. Initial failure rate was 1.29% at 21 months,126 but risk appears to increase with time, up to 5.7% to 12.1% at 3 years,127-129 a failure rate of 3.75% per year, versus 0.58% per year or less for other contemporary ICD lead models.129,130 The mechanism in most cases is a fracture of the pace/ sense hexafilar cathode conductor, either proximally, near the lead anchoring sleeve, or distally, near the anode ring electrode.102,131 Although no longer being implanted, the problem of failure with this lead is ongoing because of the large number of patients with this model in place.

Conclusion The remarkable progress in lead technology over more than four decades has contributed greatly to the modern era of implantable cardiac rhythm devices. Advances such as fixation mechanisms, steroid elution, and transvenous defibrillation lead development have been milestone events, facilitated by ongoing improvements in materials science and lead design. However, the ideal of a small-caliber, easily implanted-extracted lead with reliable long-term performance does not yet exist. Future advances in lead engineering and construction that allow physicians to approach this goal are eagerly awaited.

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4  Engineering and Construction of Pacemaker and ICD Leads 57. Kertes P, Mond H, Sloman G, et al: Comparison of lead complications with polyurethane tined, silicone rubber tined, and wedge tip leads: clinical experience with 822 ventricular endocardial leads. Pacing Clin Electrophysiol 6:957-962, 1983. 58. Hidden-Lucet F, Halimi F, Gallais Y, et al: Low chronic pacing thresholds of steroid-eluting active-fixation ventricular pacemaker leads: a useful alternative to passive-fixation leads. Pacing Clin Electrophysiol 23:1798-1800, 2000. 59. Geyfman V, Storm RH, Lico SC, Oren JW: Cardiac tamponade as complication of active-fixation atrial lead perforations: proposed mechanism and management algorithm. Pacing Clin Electrophysiol 30:498-501, 2007. 60. Van Herendael H, Willems R: Contralateral pneumothorax after endocardial dual-chamber pacemaker implantation resulting from atrial lead perforation. Acta Cardiol 64:271-273, 2009. 61. Vlay SC: Complications of active-fixation electrodes. Pacing Clin Electrophysiol 25:1153-1154, 2002. 62. Li H, Helland J: In vitro evaluation of new design lead anchoring sleeves. Pacing Clin Electrophysiol 15:2005-2010, 1992. 63. Brown S: Connectors for implantable pacemaker leads. Pacing Clin Electrophysiol 3:620-621, 1980. 64. Doring J, Flink R: The impact of pending technologies on a universal connector standard. Pacing Clin Electrophysiol 9:11861190, 1986. 65. Tyers GF, Sanders R, Mills P, Clark J: Analysis of set screw and side-lock connector reliability. Pacing Clin Electrophysiol 15:2000-2004, 1992. 66. Cleland JG, Daubert JC, Erdmann E, et al: The effect of cardiac resynchronization on morbidity and mortality in heart failure. N Engl J Med 352:1539-1549, 2005. 67. Bristow MR, Saxon LA, Boehmer J, et al: Cardiacresynchronization therapy with or without an implantable defibrillator in advanced chronic heart failure. N Engl J Med 350:2140-2150, 2004. 68. Purerfellner H, Nesser HJ, Winter S, et al: Transvenous left ventricular lead implantation with the Easytrak lead system: the European experience. Am J Cardiol 86:157K-164K, 2000. 69. Linde C, Abraham WT, Gold MR, et al: Randomized trial of cardiac resynchronization in mildly symptomatic heart failure patients and in asymptomatic patients with left ventricular dysfunction and previous heart failure symptoms. J Am Coll Cardiol 52:1834-1843, 2008. 70. Beshai JF, Grimm RA, Nagueh SF, et al: Cardiac-resynchronization therapy in heart failure with narrow QRS complexes. N Engl J Med 357:2461-2471, 2007. 71. Crossley GH, Exner D, Mead RH, et al: Chronic performance of an active-fixation coronary sinus lead. Heart Rhythm 7:472-478, 2010. 72. Nagele H, Azizi M, Hashagen S, et al: First experience with a new active fixation coronary sinus lead. Europace 9:437-441, 2007. 73. Hamid S, Arujuna A, Khan S, et al: Extraction of chronic pacemaker and defibrillator leads from the coronary sinus: laser infrequently used but required. Europace 11:213-215, 2009. 74. Freedman RA, Petrakian A, Boyce K, et al: Performance of dedicated versus integrated bipolar defibrillator leads with CRTdefibrillators: results from a prospective multicenter study. Pacing Clin Electrophysiol 32:157-165, 2009. 75. Thibault B, Roy D, Guerra PG, et al: Anodal right ventricular capture during left ventricular stimulation in CRT-implantable cardioverter defibrillators. Pacing Clin Electrophysiol 28:613-619, 2005. 76. Nof E, Gurevitz O, Carraso S, et al: Comparison of results with different left ventricular pacing leads. Europace 10:35-39, 2008. 77. Mirowski M, Reid PR, Mower MM, et al: Termination of malignant ventricular arrhythmias with an implanted automatic defibrillator in human beings. N Engl J Med 303:322-324, 1980. 78. Connolly SJ, Hallstrom AP, Cappato R, et al: Meta-analysis of the implantable cardioverter defibrillator secondary prevention trials. AVID, CASH and CIDS studies. Antiarrhythmics vs Implantable Defibrillator Study. Cardiac Arrest Study Hamburg. Canadian Implantable Defibrillator Study. Eur Heart J 21:20712078, 2000. 79. Moss AJ, Zareba W, Hall WJ, et al: Prophylactic implantation of a defibrillator in patients with myocardial infarction and reduced ejection fraction. N Engl J Med 346:877-883, 2002. 80. Bardy GH, Lee KL, Mark DB, et al: Amiodarone or an implantable cardioverter-defibrillator for congestive heart failure. N Engl J Med 352:225-237, 2005. 81. Kleemann T, Becker T, Doenges K, et al: Annual rate of transvenous defibrillation lead defects in implantable cardioverterdefibrillators over a period of >10 years. Circulation 115:2474-2480, 2007. 82. Ellenbogen KA, Wood MA, Shepard RK, et al: Detection and management of an implantable cardioverter defibrillator lead

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106. Niebauer MJ, Yamanouchi Y, Hills D, et al: Voltage dependence of ICD lead polarization and the effect of iridium oxide coating. Pacing Clin Electrophysiol 23:818-823, 2000. 107. Sweeney MO: Exploding implantable cardioverter defibrillator. J Cardiovasc Electrophysiol 12:1422-1424, 2001. 108. Hauser RG, Almquist AK: Learning from our mistakes? Testing new ICD technology. N Engl J Med 359:2517-2519, 2008. 109. Klug D, Balde M, Pavin D, et al: Risk factors related to infections of implanted pacemakers and cardioverter-defibrillators: results of a large prospective study. Circulation 116:1349-1355, 2007. 110. Cooper JM, Sauer WH, Garcia FC, et al: Covering sleeves can shield the high-voltage coils from lead chatter in an integrated bipolar ICD lead. Europace 9:137-142, 2007. 111. Zartner PA, Wiebe W, Toussaint-Goetz N, Schneider MB: Lead removal in young patients in view of lifelong pacing. Europace 12:714-718, 2010. 112. Wazni O, Epstein LM, Carrillo RG, et al: Lead extraction in the contemporary setting: the Lexicon Study: an observational retrospective study of consecutive laser lead extractions. J Am Coll Cardiol 55:579-586, 2010. 113. Hauser RG, Katsiyiannis WT, Gornick CC, et al: Deaths and cardiovascular injuries due to device-assisted implantable cardioverter-defibrillator and pacemaker lead extraction. Europace 12:395-401, 2010. 114. Saad EB, Saliba WI, Schweikert RA, et al: Nonthoracotomy implantable defibrillator lead extraction: results and comparison with extraction of pacemaker leads. Pacing Clin Electrophysiol 26:1944-1950, 2003. 115. Rinaldi CA, Simon RD, Geelen P, et al: A randomized prospective study of single-coil versus dual-coil defibrillation in patients with ventricular arrhythmias undergoing implantable cardioverter defibrillator therapy. Pacing Clin Electrophysiol 26:16841690, 2003. 116. Cooper JM, Stephenson EA, Berul CI, et al: Implantable cardioverter-defibrillator lead complications and laser extraction in children and young adults with congenital heart disease: implications for implantation and management. J Cardiovasc Electrophysiol 14:344-349, 2003. 117. Wilkoff BL, Belott PH, Love CJ, et al: Improved extraction of ePTFE and medical adhesive modified defibrillation leads from the coronary sinus and great cardiac vein. Pacing Clin Electrophysiol 28:205-211, 2005. 118. Koplan BA, Weiner S, Gilligan D, et al: Clinical and electrical performance of expanded polytetrafluoroethylene–covered defibrillator leads in comparison to traditional leads. Pacing Clin Electrophysiol 31:47-55, 2008. 119. Hackler JW, Sun Z, Lindsay BD, et al: Effectiveness of implantable cardioverter defibrillator lead coil treatments in facilitating ease of extraction. Heart Rhythm 7:890-897, 2010. 120. Undavia M, Fischer A, Mehta D: Fatal outcome in a pacemakerdependent patient. Pacing Clin Electrophysiol 32:550-553, 2009. 121. Fahy GJ, Kleman JM, Wilkoff BL, et al: Low incidence of leadrelated complications associated with nonthoracotomy implantable cardioverter-defibrillator systems. Pacing Clin Electrophysiol 18:172-178, 1995. 122. Roelke M, O’Nunain SS, Osswald S, et al: Subclavian crush syndrome complicating transvenous cardioverter-defibrillator systems. Pacing Clin Electrophysiol 18:973-979, 1995. 123. Kron J, Herre J, Renfroe EG, et al: Lead- and device-related complications in the antiarrhythmics versus implantable defibrillators trial. Am Heart J 141:92-98, 2001. 124. Kupper B, Yee R, O’Hara G, et al: Do small (6.6 Fr) active and passive fixation defibrillation leads perform as well as larger sized leads? A multi-centre analysis. Europace 9:657-661, 2007. 125. Hauser RG, Kallinen LM, Almquist AK, et al: Early failure of a small-diameter high-voltage implantable cardioverterdefibrillator lead. Heart Rhythm 4:892-896, 2007. 126. Krahn AD, Champagne J, Healey JS, et al: Outcome of the Fidelis implantable cardioverter-defibrillator lead advisory: a report from the Canadian Heart Rhythm Society Device Advisory Committee. Heart Rhythm 5:639-642, 2008. 127. Faulknier BA, Traub DM, Aktas MK, et al: Time-dependent risk of Fidelis lead failure. Am J Cardiol 105:95-99, 2010. 128. Beukema RJ, Misier AR, Delnoy PP, et al: Characteristics of Sprint Fidelis lead failure. Neth Heart J 18:12-17, 2010. 129. Hauser RG, Hayes DL: Increasing hazard of Sprint Fidelis implantable cardioverter-defibrillator lead failure. Heart Rhythm 6:605-610, 2009. 130. Porterfield JG, Porterfield LM, Kuck KH, et al: Clinical performance of the St. Jude Medical Riata defibrillation lead in a large patient population. J Cardiovasc Electrophysiol 21:551-556, 2010. 131. Cannom DS, Fisher J: The Fidelis recall: how much pressure can the ICD world bear? Pacing Clin Electrophysiol 31:1233-1235, 2008.

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5 

SECTION 1  Basic Principles of Device Therapy

Implantable Sensors for Rate Adaptation and Hemodynamic Monitoring CHU-PAK LAU  |  CHUNG-WAH SIU  |  HUNG-FAT TSE

Implantable sensors monitor changes in the body’s physiologic con-

ditions. Detection of pathophysiologic changes has been used either to modify the pacing rate of a rate-adaptive pacing system or to allow continuous monitoring of medical conditions such as heart failure or ischemia. This chapter reviews the principles and types of sensors that have been investigated and their clinical applications.

SENSORS FOR RATE-ADAPTIVE PACING Basis of Sensor-Driven Pacing Chronotropic incompetence may occur in up to 30% of patients with sinus node disease. It may also develop over time as a result of degeneration, medications, or comorbidities. In addition, for patients whose atrium is unreliable for pacing, such as during atrial fibrillation, an alternative means is required to simulate the response of the sinus node to exercise or emotion. Implantable sensors for cardiac pacing allow a pacing system to respond to exercise and nonexercise requirements for an increase in heart rate. Sensors are now a programmable option of almost all bradycardia pacemakers. Atrioventricular (AV) synchrony enhances cardiac output by augmenting the stroke volume by 20% to 30% during exercise. However, this increase is relatively small compared with the threefold to fourfold increase achieved by an increase in heart rate. The relative contribution of AV synchrony and rate increase in patients with complete AV block was studied in dual-chamber (DDD) and a rate-matched ventricular pacing (VVI) mode.1 At rest, the cardiac output during DDD pacing was 18% higher than during VVI pacing because of AV synchrony. During exercise, however, the net cardiac output was only 8% higher during DDD pacing compared with VVI pacing at an identical rate. An equivalent exercise capacity was reported in another study,2 and both DDD and rate-matched VVI pacing were superior to fixed-rate VVI pacing during exercise. Cardiac output was similar in the two modes at near-maximal exercise, but at lower workload levels, cardiac output was maintained by an increased arteriovenous oxygen saturation difference and arterial lactate level. In addition, systolic and mean blood pressures were lower when exercise was performed without AV synchrony. These findings suggest that rate response is the primary driver for exercise cardiac hemodynamics, with a much smaller contribution of AV synchrony. EXERCISE RESPONSE IN HEART FAILURE In a patient with heart failure, because the left ventricular (LV) filling pressure is elevated and the heart is working at the flat portion of the Frank-Starling curve, an increase in heart rate is the most important means to increase the cardiac output. In a study of 22 patients with poor LV function and implanted rate-adaptive pacemakers, the benefit of rate adaptation to exercise capacity was greatest in those with the poorest LV function.3 In patients with implanted cardiac resynchronization therapy (CRT) devices,4 chronotropic incompetence (maximum heart rate <70% of age-predicted maximum) was found in 70% of heart failure patients.

144

In this group, sensor-driven CRT pacing improved maximum oxygen consumption and work capacity compared with CRT pacing without rate adaption. This increment is independent of A-V interval shortening during exercise. These data suggest that rate modulation plays a critical role in some patients with impaired LV function who require pacing therapy, even in those with CRT. HEART RATE MODULATION FOR NONEXERCISE NEEDS Exercise is only one of the many physiologic requirements for variation in heart rate. Other changes, such as mental stress, can lead to a rate increase. The sinus rate is also higher when a person assumes the upright posture. Isometric exercise also results in an increase in cardiac output and heart rate in most people. The changes in heart rate that occur during various physiologic maneuvers (e.g., Valsalva) and baroreceptor reflexes may also be important. An appropriate compensatory heart rate response may be especially important in pathophysiologic conditions such as anemia, acute blood loss, or other causes of hypovolemia, and during febrile illness.

Ideal Sensor Characteristics Based on the physiology of the normal sinus node, a sensor system to overcome chronotropic incompetence needs to be sensitive to both exercise and nonexercise needs. It also should be specific, unaffected by internal or external changes that can cause an inappropriate rate change. The sensor should achieve rate modulation at an appropriate speed, with its response proportional to the level of exercise load. The sensor system should be easy to implement (preferably with standard device casing and lead system), should be stable in the body’s internal environment, and should not significantly increase battery consumption. CLASSIFICATION OF SENSORS In a sensor-driven pacing system, a sensor (or combination of sensors) must first detect a physical or physiologic parameter that is related to metabolic demand5 (Fig. 5-1). Second, the rate-modulating circuit in the pulse generator must have an algorithm that relates changes in the sensed parameter to a change in pacing rate. Third, because the magnitude of the physical or physiologic changes monitored by a sensor may differ between patients, clinician input may be necessary to adjust the algorithm, generally by programming one or more rate-responsive variables, to achieve the clinically desired rate response. This requirement for adjustment has decreased with automatic optimization of the rate-responsive settings. Most sensors operate in an open-loop algorithm; that is, the induced rate changes do not induce a negative feedback on the sensed parameter. In a closed-loop sensor system, the induced hemodynamic changes will induce an opposite change in the level of the sensed parameter that is responsible for the initial rate adaptation (negative feedback loop). Theoretically, minimal programming is required in a closed-loop system (see Fig. 5-1). Furthermore, the actual rate response is calculated using rate-response curves after programming a threshold of sensor detection (Fig. 5-2).



5  Implantable Sensors for Rate Adaptation and Hemodynamic Monitoring CLOSED LOOP

OPEN LOOP

Sensor

Sensor

∆ Physiologic parameters

∆ Physiologic/ physical parameters

Algorithm

145

Algorithm

Ph

Ph

ys inp ician ut

ys inp ician ut

∆ Rate

∆ Rate

Figure 5-1  Design of a rate-adaptive system. Closed loop, The physiologic change detected by the sensor is converted to a change (Δ) in rate using an algorithm. The resultant rate change induces a change in the physiologic parameter in the opposite direction, thereby establishing a negative feedback loop. Open loop, The physiologic or physical change detected by the sensor is converted to a change in rate using an algorithm. The resultant rate change does not have a negative feedback effect on the physiologic/physical parameter.

Technical Classification A practical classification is to categorize sensors according to the technical methods that are used to measure the sensed parameter (Table 5-1). Body movements during exercise result in changes in acceleration forces that are transmitted to the pacemaker. Sensors that are capable of measuring the acceleration or vibration forces in the pulse generator Figure 5-2  Types of rate-responsive curves used in rate-adaptive pacemakers employing a single sensed parameter (sensor level). An appropriate filter (F) eliminates unwanted raw signals (e.g., high- or low-pass frequency filters and thresholds). The filtered signals are then appropriately modified (M; e.g., with gains and rectification) before being converted to a rate change. The curves can be linear (A), curvilinear (B), or complex (C and D). The physician can select an appropriate rate-responsive slope from a family of curves (see text). IR, Interim rate.

are broadly referred to as activity sensors. Technically, detection of body movement can be achieved using a piezoelectric crystal, an accelerometer, a tilt switch, or an inductive sensor. Each of these devices transduces motion of the sensor either directly into a voltage or indirectly into measurable changes in the electrical resistance of a piezoresistive crystal. Activity sensing is the most widely used control parameter LINEAR

∆ Rate

F

CURVILINEAR

∆ Rate

M

∆ Sensor

A

F

M

∆ Sensor

B COMPLEX

∆ Rate

∆ Rate

IR

– F

C

M

∆ Sensor

F

D

+

0 M

∆ Sensor

146

TABLE

5-1 

SECTION 1  Basic Principles of Device Therapy

Major Classes of Sensors Used in Implanted Devices, Classified According to Method of Technical Realization Examples

Methods Vibration and acceleration sensing

Physiologic Parameter Body movement

Impedance sensing

Minute ventilation Ventricular inotropic parameter Evoked QT interval* Physical parameters Central venous temperature dP/dt Right ventricular pressure Pulmonary arterial pressure Peak endocardial acceleration Chemical parameters pH Mixed venous oxygen saturation Catecholamine

Ventricular evoked response Special sensors on pacing electrode

Models Sigma, Kappa, EnRhythm Talent Miniswing, Neway Insignia, Pulsar Max, Discovery Actos, Protos, Philos Identity, Integrity, Affinity Kappa 400 Talent Insignia, Pulsar Max Protos, Inos Diamond, Clarity, Selection AF

Manufacturers Medtronic Vitatron ELA-Sorin Boston Scientific Biotronik St. Jude Medical Medtronic ELA-Sorin Boston Scientific Biotronik Vitatron

Chronicle†

Medtronic

Best-Living system

ELA-Sorin

*Currently no longer available in market. †For hemodynamic monitoring only.

in rate-adaptive pacing because of its ease of implementation and its compatibility with standard unipolar and bipolar pacing leads. Additionally, activity sensors can be used for either atrial or ventricular pacing modes. Impedance is a measure of all factors that oppose the flow of electric current and is derived by measuring resistivity to an injected current across a tissue. The impedance principle has been extensively used for measuring respiratory parameters and those associated with ventricular contractility, such as relative stroke volume or the right ventricular preejection interval. Transthoracic impedance can be used to assess respiratory rate and tidal volume by measuring the continuous impedance between the pacemaker casing and an intracardiac electrode. The product of respiratory rate and tidal volume, minute ventilation (MV), is used as the sensed parameter. When using unipolar impedance from a distal electrode of a pacing lead, impedance reflects myocardial contractility, as used in the closed-loop stimulation (CLS) sensor. The intracardiac ventricular electrogram (EGM) resulting from a suprathreshold pacing stimulus has been used to provide several parameters to guide rate modulation. The total duration of depolarization and repolarization can be estimated by the interval from the pacing stimulus to the intracardiac T wave: the QT interval, or stimulus-T interval. This parameter is sensitive to changes in heart rate and circulating catecholamines and can be derived from the paced EGM with conventional pacing electrodes. This was developed as a QT-sensing pacemaker, based on the QT interval shortening during exercise and stress.5 The last group of sensors is incorporated into the pacing lead itself. These specialized leads include thermistors (used to measure blood temperature), piezoelectric crystals (right ventricular pressure), optical

TABLE

5-2 

sensors (mixed-venous oxygen level), and accelerometers at the tip of pacing leads. Some of these sensors measure highly physiologic parameters. For example, oxygen saturation is closely related to oxygen consumption during exercise. Physical activities increase cardiac output and oxygen extraction from the blood, and the tissue arteriovenous oxygen difference widens during increased tissue oxygen consumption. The fall in oxygen saturation will trigger an increase in rate that will improve cardiac output and minimize the decrease in mixed-venous oxygen saturation.6 The mixed-venous oxygen sensor is a true closedloop system, although its technical complexity limits its use for rateadaptive pacing. Sensing of changes in blood pH during exercise has also been suggested as a possible sensor, although the requirement for a specialized lead has impeded its clinical implementation.7 Using an accelerometer incorporated at the tip of a unipolar ventricular electrode, the contractile state of the right ventricle can be assessed indirectly from the endocardial pressure generated during isovolumic contraction of the heart, a parameter known as peak endocardial acceleration (PEA). PEA is significantly increased by exercise and other inotropic stimuli (e.g., mental stress). Over the years, many of these sensors have been implemented in implantable devices. Significant differences in rate response were found among sensors, notably between their sensitivity and specificity (Table 5-2). However, only activity, MV, and CLS sensors are currently used for rate response. The activity sensor has the advantage of reliability and ease of implementation; its lack of proportionality to exercise workload has not been a major clinical disadvantage for most patients. Sensors in special leads are no longer used for rate adaptation because of their uncertain long-term reliability and difficulty during pacemaker replacement. However, these sensors are now used for hemodynamic monitoring in heart failure.

Relative Performance of Different Sensors

Activity Minute ventilation QT interval PEA CLS

Speed High Moderate Low Moderate Moderate

*Refers to technical and clinical realization. CLS, Closed-loop stimulation; PEA, peak endocardial acceleration.

Proportionality Low High Moderate Moderate Moderate

Specificity Low Moderate High Moderate Moderate

Sensitivity Low Low Moderate High High

Reliability* High Moderate Low Moderate Moderate



5  Implantable Sensors for Rate Adaptation and Hemodynamic Monitoring 120

Activity sensing is the most widely used sensor, alone or in combination. Because activity-based pacemakers are operationally simple and do not require a special sensor outside the pulse generator casing, they work with any type of pacing lead, have excellent long-term stability, and are highly reliable. Activity-guided pacemaker implantation is the same as for conventional pacemakers. Although activity sensors may not be excellent proportional sensors, they react promptly to the start and termination of physical exercise. The first activity sensors, piezoelectric crystals, mainly respond to the frequency of vibrations transmitted to the pulse generator, with more variable response depending on the pulse generator and the pocket. The subsequent activity sensors, accelerometers, detect body accelerations and have less interpatient variability and better proportionality (Fig. 5-3). Current Activity-Sensing Devices Medtronic (Kappa 700/900/400, Adapta, Advisa, EnRhythm).  The Medtronic Activitrax and Thera devices were the first generation of activity-sensing pacemakers to use a piezoelectric crystal bonded to the inside of the pacemaker casing to detect activity. Activity levels that exceed a programmable threshold are counted, and the number of counts that exceed a programmable threshold is used to determine the rate response. The clinician adjusts the overall rate response by programming the activity threshold and the rate-response slope. Typically, this instrumentation allows a good response from the programmed lower rate to about 100 beats per minute (bpm) during daily activities. The response to the upper rate is often less accurate and depends on the patient and the programmed settings. Despite their limitations, these piezoelectric devices were clinically highly effective when first introduced. The current Medtronic activity-sensing device uses an accelerometer attached to the circuit board to detect movement in the anteroposterior (AP) plane. Instead of a fixed threshold, a rolling threshold is used to provide a count of activity that depends not only on the frequency of counts that cross a threshold value, but also on the amplitude (strength) of these counts. This rolling threshold minimizes the Electronics Piezoelectric crystal

A

Battery

Body tissue

Piezoelectric accelerometer

B

Electronics

Battery

Piezoresistive accelerometer

C

Electronics

Battery

Figure 5-3  Three types of activity sensors. A, Piezoelectric sensor is bound to the inside surface of the pacemaker case. Vibration is sensed through tissue contact. The mechanical forces are transmitted by the surrounding connective tissue, fatty tissue, and muscles. The extent of contact and the coupling mass of the mechanical forces can vary considerably. B, Piezoelectric accelerometer, and C, piezoresistive accelerometer, mounted on the pacemaker’s hybrid circuitry. The seismic mass of the accelerometer is fixed. Therefore, measured accelerations are independent of surrounding tissue and patient physical properties.

Target rate (bpm)

Current Sensor-Driven Pacemakers ACTIVITY-SENSING AND ACCELEROMETER-BASED PACEMAKERS

147

110 100 90 80

ADL 5 ER 5

70

ADL 3 ER 3

ADL 1 ER 1

60 0

8

16

24

32

40

Sensor counts Figure 5-4  Triphasic rate-responsive curve. Medtronic Kappa 600/700/900, Enrhythm, and Enpulse accelerometer activity-sensing pacemakers; ADL, activity of daily living; ER, exertion response.

detection of low-intensity but high-frequency accelerations, as during motor vehicle travel, but allows detection of high-intensity signals during physical activity. The sensor-driven rate is adjusted by the use of a triphasic rate curve that incorporates two rate ranges, activity of daily living (ADL) and exertion response (ER) programming (Fig. 5-4). The physician determines the lower rate limit (LRL), the ADL rate, and the sensor-driven upper rate limit (SURL). With the nominal ADL response of 3 (I-5), counts between 12 and 20 will give the ADL rate, during which the rate response will be stable. When counts are greater than 20, ER will occur to the maximum programmed SURL. The flat portion of the response curve corresponds to 25% of the total counts. Adjustment of the ADL and ER to a more sensitive setting will decrease the number of counts needed to reach between ADL rate and SURL, and shorten the ADL portion of the response curve. The actual number of counts to effect a change in pacing rate is not fixed and differs among patients. Rather, an ADL response of 3 implies that the patient will spend up to 30 minutes daily at a sensor rate at or above the ADL. This is achieved by adjusting the rate/count relationship to achieve this percentage of rate response, the ADL set point. Similarly, an ER response of 3 means that the patient will be spending about 20 minutes weekly at the higher rates, 105 to 120 bpm, the upper-rate set point. The acceleration signal input can be adjusted by changing the threshold (lower threshold allows detection of more acceleration signals) and slope. These two parameters affect the sensitivity to the acceleration detected. The physician can manually program these parameters to achieve a faster or slower response in both the ADL and ER ranges. With automatic rate response, rate profile optimization will adjust these rate curves automatically, using daily comparison of the rate histogram with the target-rate histogram. The device adjusts more quickly in the first 10 days after a programming change so that the new setting will be achieved more quickly (Fig. 5-5). The Kappa 400 is a combined activity/MV-sensing device. The general hardware and software implementation is similar to the Kappa 700/600 series. However, actual time spent in ADL and ER is longer in the Kappa 400 because a more aggressive rate response is allowed. St. Jude Medical (Identity, Vitality, Zephyr, Accent).  These activity devices use an Omnisense accelerometer bonded to the circuit board and are sensitive to accelerations that occur in the AP axis. Activity counts above a programmable threshold (1 = most sensitive; 7 = least sensitive) are integrated and translated to a rate response using a rateresponse slope (1 = least sensitive; 16 = most sensitive) (Fig. 5-6). The threshold level can be adjusted manually or automatically (Auto). In the Auto setting, the device measures the sensor activity level over the preceding 18 hours to determine the threshold parameter. The Auto threshold can be adjusted using “Auto threshold offset.” Similarly, the slope of response can be calibrated manually or automatically. In the slope Auto adjustment, the device records the activity variance occurring above the threshold over the last 7 days, and the slope is

SECTION 1  Basic Principles of Device Therapy

0.35 0.3 0.25 0.2 0.15 0.1 0.05 0

activity threshold (very low, low, medium-low, medium-high, high, very high); reaction time, and recovery time. Programming can be done manually or automatically using the Automatic Response Factor. In this method, the physician uses the Expert Ease feature in the programmer to decide the LRL, SURL and the sensor rate target, based on the patient’s age, gender, exercise frequency, and target heart rate during exercise, derived from a population average. After inputting these parameters, an initial nominal slope such as 8 (range, 1-16) is programmed. The Automatic Response Factor tracks the maximum sensor rate each day for a week, and the average of this maximum is compared to the sensor rate target. The response factor is then scaled to trigger either a more aggressive response or a conservative response, depending on whether the average rate is below or is above the sensor rate target by 5 bpm, respectively.

Target rate profile Sensor rate profile

LRL

ADL

130–140

120–130

100–110

90–100

80–90

70–80

Exertion response range

60–70

Indicated % of time

ADL response range

110–120

148

Other Activity Sensors.  ELA-Sorin has introduced a gravitational sensor that is used either alone (Swing) or in combination with a PEA sensor (MiniLiving). The gravitational sensor uses the vibrations from a mercury ball to measure body activity. An accelerometer activity sensor (Opus G) uses a half-bridge variable-capacitance accelerometer. The accelerometer senses the AP axis at a frequency range of 0.6 to 6.0 Hz and samples acceleration signals every 1.56 seconds. The device can be programmed manually or automatically using the Autocalibration feature. The level of acceleration at rest is reset daily based on the lowest mean acceleration of 64 consecutive measurements. The maximum mean acceleration of 8 consecutive accelerations is used to match the programmed upper sensor rate.

URL

Rate Figure 5-5  Optimization using target rate histogram. The sensor rate profile is matched against a target rate profile. The activity of daily living (ADL) range includes moderate pacing rates between the lower rate limit (LRL) and the ADL rate. The exertion rates range from the ADL rate to the upper rate limit (URL). By comparing the sensor rate profile with the target rate profile once daily, the pacemaker automatically controls how rapidly the sensor-indicated rate increases and decreases in these ranges. (From Leung SK, Lau CP, Tang MO, et al: An integrated dual sensor system automatically optimized by target rate histogram. Pacing Clin Electrophysiol 21:1559-1566, 1998.)

Clinical Results of Activity Sensors

adjusted to achieve a rate response such that almost 1% heart rate occurs beyond 23% of heart rate reserve. Typically, the slope adjusts by a factor of 2 each week. If the patient has been immobile for the previous week, as indicated by the small activity variance, the slope setting will be held constant, to avoid undue rate response when the patient resumes activity. In addition, rapid manual adjustment of the slope can be achieved by instructing the patient to perform an exercise at a prescribed slope. The Prediction Model will attempt to suggest an alternative slope based on the detected rate response. In addition, the onset and recovery kinetics are separately programmable. These devices also feature a “rest rate” that may be programmed below the LRL. The rest rate is attained when the acceleration sensor does not register significant counts (the “activity variance” level) over an extended period, implying the patient is resting or asleep.

Clinical studies have demonstrated that activity-based pacing systems improve exercise capacity and reduce symptoms compared with VVI pacemakers. In an early study of a piezoelectric vibration-based sensor pacemaker, Benditt et al.8 compared exercise tolerance during VVI and VVIR pacing using treadmill tests. Sensor-driven pacing prolonged exercise duration by 35% and led to similar increments in peak oxygen consumption and oxygen consumption at anaerobic threshold. Activity sensor–driven pacing also reduced the patient’s perception of exertion at comparable exercise levels, and the benefit was sustained when exercise testing was repeated after an average 5 months of follow-up. Further, at follow-up exercise testing, reversion of the pacing system to a VVI mode resulted in prompt deterioration of both observed oxygen consumption and exercise duration. Crossover studies comparing VVIR mode to DDD pacing found no significant difference in symptom scores, maximal exercise performance (treadmill), or plasma concentrations of epinephrine, norepinephrine, and atrial natriuretic peptide, suggesting rate adaptation was the most important factor for exercise in the patient population studied.9,10

Boston Scientific (Insignia, Pulsar Max, Altrua).  An accelerometer mounted in the circuit board is used to detect AP acceleration. In addition to LRL and URL, four other parameters are used to determine the rate response: response factor (1 = least sensitive; 16 = most sensitive); RATE RESPONSE SLOPES SURL

AUTO SLOPES

16

30% % of time

Rate response

8 20% 10%

1 LRL Low activity

A

Lower rate

High activity Sensor levels

B

Heart rate

Upper sensor rate

23% HRR Figure 5-6  Rate-response and “Auto” slopes. A, Rate-response curves from St. Jude Medical activity sensors. B, Automatic threshold (Auto) slope ensures that about 1% of the pacing rate will reach more than 23% of the heart rate reserve (HRR) over time. Note that the sensor-driven upper rate limit (SURL) may not be reached at a low slope setting. LRL, Lower rate limit.



5  Implantable Sensors for Rate Adaptation and Hemodynamic Monitoring

In a study of the efficacy of automatic “rate profile optimization” over time, 11 patients with Kappa 700 pacemakers performed treadmill testing at 1 month, 1 year, and 2 years after implantation.11 Based on the sinus profile at follow-up, a more aggressive slope was needed to match the sinus rate profile. This required a change in ADL response from 3 to 4 and in ER from 3 to 4 at 1 year, and a change in activity threshold from medium-low to low at 2 years. These adjustments enabled better approximation of pacing rate to sinus rate during treadmill exercise. Exercise capacity was maintained during the 2 years of follow-up. Variation in activity level (Activity Variance) as detected by an accelerometer is used to determine the rest rate in the St. Jude Medical devices,12 with an overall appropriate rate response. Limitations of Activity-Sensing Devices Activity sensing may give inappropriate heart rate responses when subjected to environmental vibrations, such as from motor vehicle movement over rough terrain, air travel, or use of vibrating appliances or machinery. The piezoelectric vibrational sensor also responds to the application of static pressure on the pulse generator, which may cause inappropriate rate increase when the patient lies prone. This falsepositive response to pressure is no longer a problem with pacemakers that incorporate an accelerometer. The effects of various means of locomotion on pacing rate were assessed for different activity-based pacemakers.13 Three different activity-based pacing systems (peak counting algorithm, integration type, and accelerometers) were strapped to the chests of volunteers. Bicycling on the street resulted in higher pacing rates than did stationary bicycling for each type of pacemaker, although none of the pacemakers reached the heart rate achieved by the normal sinus node. During driving, the pacemakers increased the pacing rate, although the intrinsic sinus rate continued to be higher. In passively riding passengers, the pacemakers tended to produce a higher pacing rate than that of the normal sinus node. Independent of the sensor, activity-initiated rate response depends on the manner in which activity is performed rather than the exercise workload, and proportionality is generally limited.14 Activity sensors typically produce a higher heart rate walking downhill than uphill because of the greater force of heel-strike in the descending direction. Nonexercise stresses such as emotional changes are not detected. MINUTE VENTILATION SENSING Although a respiratory rate–sensing pacemaker was in use as early as 1983, it was limited by the need for an auxiliary subcutaneous electrode, as well as easy interference by arm movement because of unipolar impedance sensing.15 All subsequent generations of respiratory sensor detect impedance as a surrogate measure of minute ventilation for rate adaptation. Again, MV is the product of tidal volume and respiratory rate. Although MV is not actually measured, both the amplitude of changes in transthoracic and intravenous impedance (correlating with tidal volume) and the frequency of changes in these parameters (correlating with respiratory rate) can be measured by injecting a very-low-amplitude and short-duration electrical stimulus between an electrode on a pacemaker lead and the pulse generator casing. Thus, changes in MV can be estimated by changes in minute impedance. The advantage of this sensor is that it only requires a standard bipolar pacing lead in either the right atrium or right ventricle, without the need for additional hardware. The response of MV during exercise is proportional to oxygen consumption (below anaerobic threshold) and responds at a reasonable speed; battery drain is not excessive. Current Minute Ventilation–Sensing Devices Medtronic (Kappa 400).  The Medtronic MV pacer was marketed as a dual-sensor device. The MV component requires a bipolar lead. Subthreshold, 30-µV biphasic pulses at 1 millisecond (msec) are injected between the electrode (atrial or ventricular) and pacemaker case at 16 Hz, and the resulting impedance is measured between the

149

lead tip and the case. Changes in MV are measured as MV counts over a 32-second, short-term moving average to avoid transient changes in impedance, such as those caused by coughing and speech. The shortterm average is compared over an 18-hour, long-term average to effect a change. Medtronic Kappa 400 has been evaluated using symptom-limited treadmill testing.16 All patients had cardiopulmonary gaseous analysis in conjunction with simultaneous recording of MV-induced impedance changes. Calibrations were made during resting, supine, sitting, and different types of exercise (bicycle or treadmill). The impedancederived MV was smaller for sitting versus supine position, for shallow versus slow breathing, and for bicycle versus treadmill exercise, possibly because of change in chest cage geometry. Correlation coefficients of the best-fit line for measured and device-based MV were high (>0.9 in first-, second-, and third-order polynomial equations).17 Furthermore, the calibration between measured MV and impedance MV changes over time (1 week to 1 month). However, the changes were not correlative, with implications for the need of continual automatic adaptation. These findings also suggest the potential use of the sensor for MV monitoring. Boston Scientific (Pulsar Max, Insignia, Altrua).  In the Boston Scientific devices, MV is available in conjunction with an accelerometer sensor, although either sensor can be used alone or blended with the other. MV collected by the atrial or ventricular lead (programmable) over 24 hours is used as an average against which future changes in MV are compared for a rate response. The minimum time for achieving a baseline is 4 minutes (“4 → On”). A linear response curve is used, and a total of 16 response factors can be chosen below the pacing rate estimated for anaerobic threshold. This rate is also programmable. In between the pacing rate for anaerobic threshold and the maximum predicted heart rate, a gentler “high rate response” can be programmed. The adaptation can be made faster by activating the 4-minute walk within 30-minute option, in which the subject is instructed to exercise to achieve the maximal MV change. The telemetered MV impedance signal from a Boston Scientific device was compared with measured MV in 20 patients.18 Respiratory rate was accurately measured by the device during hyperventilation, with difference less than 0.2 breath/min. During 10-minute cycle ergometry at 50 W, the correlation between MV measured directly and by the device was 0.99. Large, individual variations exist between the measured MV and the impedance MV slope, requiring individualized rate-response curves. ELA-Sorin (Chorus, Talent, Opus, Symphony, Rhapsody, Reply).  Minute ventilation is determined from a default bipolar atrial lead, with current injection from distal electrode to casing, and current collection between the proximal pole and casing. In the event of a unipolar atrial lead, the distal electrode of the ventricular lead is used as the current-injecting electrode. Sourcing and injection can also be performed with a bipolar ventricular lead. The automatic slope algorithm in the Chorus determines the resting and maximum MV values daily. If MV over 128 cycles remains below the 24-hour MV average, the lower rate will decrease to the LRL. The device calculates the exercise MV signal by looking for the maximal MV signal and recalculates this value every eighth cycle. The exercise MV value is increased or decreased in 6% intervals. The mean resting and exercise MV values are used to adjust the rate-response slope automatically over the range of values from 1 to 15 in steps of 0.1. The response of the Talent DR during exercise was compared in 81 patients. The correlation coefficient between the sensor rate and programmerderived sensor rate was 0.983 ± 0.005, and a linear relationship was observed between the heart rate reserve and MV reserve.19 Clinical Experience of MV Sensor Pacemakers Minute ventilation is an indirect but reliable marker of metabolic demand, and in the first generation of MV pacemakers (Meta MV, Telectronics), the rate response was proportional to the level of

SECTION 1  Basic Principles of Device Therapy

Limitations of Minute Ventilation Sensing As with all impedance systems, the MV sensor will likely be affected by electromagnetic interference, arm swinging, coughing, and hyperventilation.16 Artificial ventilation will induce an unphysiologic rate, so MV sensing needs to be disabled in such patients (e.g., general anesthesia, mechanical ventilation). The use of MV devices in patients with lung disease and heart failure remain controversial. Early generations of MV-sensing devices limited the maximum detectable respiratory rates to less than 48 to 60 breaths/min to minimize the cardiac effects of changes in impedance. Because of this filtering of the cardiac component of changes in impedance, MV sensors may not accurately reflect rate adaptation at very high respiratory rates, as may occur in children. A bipolar ventricular lead is needed for MV sensing in several models. The battery current for MV sensing may consume approximately 2% of the total current of a dual-chamber pacemaker. Some amplifiers may sense the small impedance pulses unless preventive steps are taken. In some products, a blanking period of about 1 msec is applied to the amplifiers to “blind” them to these pulses. In other cases, special balancing of the pulse achieves the same objective. Sensing of the impedance pulses by a surface electrocardiogram (ECG) monitor may occur, depending on the ECG machine’s frequency response as well as pulse width and balancing of the impedance pulse. Also, frequencies above a few kilohertz (kHz), such as generated by electrocautery and electrosurgery, are detected by the rate-response circuitry and could drive an MV-sensing pacemaker to its maximum rate. It is recommended that the rate-response function be inactivated whenever electrosurgery will be used. Respiration may also be influenced by phonation and coughing that have no direct relevance to change in cardiac output. UNIPOLAR VENTRICULAR IMPEDANCE: CLOSED-LOOP STIMULATION SENSOR Contractility of the right ventricle increases during catecholamine stimulation, as occurs in exercise and emotional stress. In the absence of an adequate rate response, exercise will induce a higher contractility. When rate response is adequate, the change in contractility is less, thus establishing a negative feedback loop and a new, steady contractility state. On the other hand, elevated pacing rates can increase contractility, the so-called Treppe effect, although this has not been important in clinical practice. Right ventricular (RV) impedance can be derived from a standard pacing lead and can be used to monitor heart function and catecholamine state. Sensor and Algorithm The CLS sensor is based on unipolar impedance at the tip of a pacing lead.22 Subthreshold pulses of automatically selected outputs, ranging from 100 to 400 microamperes (µA), with biphasic duration of 46 msec are emitted 50 to 300 msec after a sensed or paced ventricular event. During late diastole immediately after a ventricular paced (Vp)

or sensed (Vs) event, the blood volume of the right ventricle is highest and RV impedance is lowest. On the other hand, as contraction occurs, the walls surrounding the electrode tip draw closer, and impedance rises (Fig. 5-7, A). A baseline waveform will depend on the conduction state of the heart: AsVs, AsVp, ApVs, ApVp (As, atrial sensed; Ap, atrial paced event). Baseline CLS waveforms will be acquired only when the associated accelerometer indicates “no activity,” and a waveform will be discarded within 48 hours if not referenced. An average template of the baseline CLS waveform will take 2 to 3 days to optimize. As contractility increases during exercise, unipolar impedance will change. The time-integrated difference between the exercise and baseline impedance waveforms is converted to a pacing rate using an auto response factor, which is continually adapted and patient specific. The magnitude of rate response is then determined by a programmable exertion threshold rate (ETR; very low, low, medium, high, and very high) (Fig. 5-7, B). The ETR, acting through the auto response factor, will determine that 80% of the heart rate will occur below the ETR and 20% above the ETR. An active young individual will probably require a higher ETR than an inactive elderly patient. Again, the type of rate profile attained will be determined by the patient’s cardiac condition, physical state, and specific automatic response factor. Although not strictly a dual-sensor pacemaker, the incorporated accelerometer performs an “on-off ” function to decide on the acquisition of the baseline CLS waveform and the CLS-driven response. Full CLS-driven rate according to the programmed ETR will be allowed only if the accelerometer registers “exercise”; otherwise, only 20 bpm above the LRL is allowed to enable a nonexercise rate increase. In neurocardiogenic syncope, this rate-limiting algorithm can be inactivated for full overdrive response, as during a syncopal episode. TIME (ms) 240 Resting

200 Units

exertion and correlated with the normal sinus response.20,21 Compared with VVI pacing, MV-based VVIR pacing increased exercise capacity by 33%,21 and maximal oxygen consumption and cardiac output were significantly improved. In one study of 10 patients with the first version of MV-driven VVIR pacing, pacing rate was highly correlated with measured MV (r = 0.89), respiratory quotient (r = 0.89), Vco2 (r = 0.87), tidal volume (r = 0.87), Vo2 (r = 0.84), and respiratory rate (r = 0.84). Maximum oxygen consumption also increased from 13.4 ± 3.4 to 16.3 ± 4.1 mL/kg/min (P = .0004) using MV-driven over VVI pacing. Improvement in symptoms were also documented in the VVIR mode.21 The MV sensor has good long-term stability; programming of the sensor is relatively simple; and the rate response is appropriate during daily activities. Compared with activity pacing, MV is significantly better in achieving a near-normal pacing rate–workload relationship, whereas activity sensing tends to overpace at low-level exercise and underpace at peak exercise and in the recovery period.

160 120

Exercise

80 40 50

A

100

150

200

250

300

Milliseconds after a sensed or paced ventricular event 8

ETR = 73

7 6 Rate (%)

150

ETR = 80

ETR = 92

Very low

5 4

Medium

3

Very low

2 1 0

B

60

70

80

90

100

110

120

130

Figure 5-7  Closed-loop stimulation (CLS) sensor and algorithm. A, Changes in CLS parameter during exercise. Hatched area represents the difference between baseline and exercise CLS waveforms and is converted to a rate using the exertion threshold rate (ETR). B, Impact of ETR on rate response of CLS. “Medium” ETR corresponds to a rate of 80 beats per minute (bpm; 20 bpm above lower rate limit). Programming this rate will ensure a rate response of 80% less than this rate and 20% above this rate. A higher or lower ETR will result in different rate responses.



5  Implantable Sensors for Rate Adaptation and Hemodynamic Monitoring

151

Micro mass PEA Pace/sense Electronics Figure 5-8  Biomechanical Endocardial Sorin Transducer. The BEST sensor (Sorin Biomedica Cardio) consists of a microaccelerometer located inside a rigid capsule at the tip of a standard ventricular pacing lead, which is connected to a triple header. PEA, Peak endocardial acceleration.

Current Closed-Loop Stimulation Devices: Biotronik (Inos, Protos, Cyclos, Entovis, Evia) Acute measurements of the CLS parameter (previously known as “ventricular inotropic index”) were taken in 82 patients with chronically implanted unipolar ventricular leads at pulse generator replacement.22 A wide fluctuation of baseline impedance was observed (500-1500 Ω), whereas CLS fluctuated by about 4 to 25 ohms (Ω), with a good correlation between CLS and the baseline impedance. A clinical study of 205 patients evaluating the CLS pacemaker included a significant number of young subjects with complete AV block caused by Chagas’ disease.23 Satisfactory rate modulation was reported in 93% of patients. In the remaining 7% of patients, rate adaptation could not be achieved because of poor exercise tolerance, severe myocardial dysfunction, and intermittent intrinsic AV conduction. In a multicenter study that included 178 VVIR (Biotronik Neos-PEP) and 84 DDDR (Biotronik Diplos-PEP, InosDR) devices, physiologic rate adaptation was possible in 93% and 96% of patients with these devices, respectively.24 Apart from exercise rate response, this study also involved mental stress testing using color-word matching and the infusion of inotropic agents. A moderate level of rate response was documented in some patients with CLS pacemakers during these nonexercise conditions that was greater than during accelerometer-guided pacing. The benefit of CLS sensor pacing during mental stress and cognitive function has also been studied.25 Considerable interest surrounds the use of the CLS sensor to detect changes in posture. Passive head-up tilt, which depletes intravascular volume, increases the inotropic state of the heart. In a multicenter study, Inotropy Controlled Pacing in Vasovagal Syncope (INVASY), 50 patients with severe vasovagal syncope and positive head-up tilt test were randomized between DDD-CLS and DDI mode at 40 bpm.26 Whereas 7 of 9 patients in the DDI arm experienced syncope within 1 year, only 4 of 41 patients in the DDD-CLS arm had presyncope. The authors suggested the efficacy of this approach, although a placebo effect of pacing was suspected in 22% of patients. In 131 patients with chronotropic incompetence, CLS pacing resulted in a higher pacing rate compared with accelerometer sensor, but no difference in 6-minute walking distance (6MWD) test.27 However, twice as many patients preferred the CLS mode. The use of CLS in the left ventricle, as in a CRT device for monitoring, may be a useful application of this technology (see later discussion).

PEAK ENDOCARDIAL ACCELERATION Sensor and Algorithm The contractile state of the heart can be identified by the maximal velocity of shortening of unloaded myocardial contractile elements, which can be measured with a catheter tip accelerometer attached to the ventricular wall. The “peak endocardial acceleration” represents the endocardial vibration measured by the accelerometer in the right ventricle during isovolumetric contraction phase of the ventricles. This signal is closely associated with the intensity of the first heart sound. The ELA-Sorin BEST (Biomechanical Endocardial Sorin Transducer) sensor is a microaccelerometer consisting of an acceleration sensor built into a nondeformable capsule on the tip of a standard unipolar ventricular pacing lead. The lead is placed against the RV wall so as to be sensitive to its acceleration and insensitive to the pressure of blood and myocardium (Fig. 5-8). This system has a frequency response up to 1 kHz and a sensitivity of 5 mV/G (1 G = 9.8 m/sec/sec). In early animal experience using an external system and an implantable radiotelemetry system, the PEA was not affected by heart rate but significantly increased by emotional stress, exercise stress testing, and inotropic stimulation.28 The PEA signal changes in parallel to the maximal LV dP/dt and appears to measure the global LV contractile performance rather than the regional mechanical function. The PEA signal that occurs 150 msec after the R wave corresponds to the isovolumetric contraction phase of the left ventricle (PEA-I)29 (Fig. 5-9). A smaller signal also occurs in the 100-msec period after the T wave,

ECG

LV dP/dt

Ao Press

LV Press

Advantages and Limitations Closed-loop stimulation appears to be a good sensor to measure cardiac contractility using a conventional ventricular pacing electrode. The demand on pacing energy is acceptable. As a contractility sensor, CLS is sensitive not only to exercise but also to nonexercise requirements, and it may therefore be used for monitoring cardiac contractility for non–rate augmentation purposes. CLS can be used only in a pacing mode that incorporates a ventricular lead. Ventricular ischemia and cardioactive medications likely influence CLS rate adaptation, although automatic adjustment may permit long-term function. A small proportion of patients are not suitable for CLS sensors because of severely impaired RV function. Such patients may be identified preoperatively, and a backup activity sensor is always available.

EA PEA-I

PEA-II

Figure 5-9  BEST tracings. Electrocardiogram (ECG), left ventricular (LV) dP/dt, aortic pressure (Ao Press), LV pressure, and endocardial accelerometer tracing (EA) recordings in a sheep with the BEST sensor (see Fig. 5-8). The peak of the accelerometer tracing during the isovolumic diastole (PEA-II) occurs in a 100-msec period after the T wave on the ECG. (From Plicchi G, Mercelli E, Parlapiano M, et al: PEA-I and PEA-II based implantable haemodynamic monitor: preclinical studies in sheep. Europace 4:49, 2002.)

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SECTION 1  Basic Principles of Device Therapy

the so-called PEA-II, which corresponds to the isovolumetric LV relaxation. PEA-II is related to peak negative dP/dt (r = 0.92) and aortic diastolic pressure (r = 0.91).30 To enable the PEA sensor to be used in patients requiring an ICD lead, the sensor has been incorporated into a right atrial (RA) lead (SonR sensor). Preliminary results suggest that the SonR sensor is similar to the RV PEA signal and that it might be useful in patients receiving cardiac resynchronization therapy with defibrillation (CRT-D).29 This offers the potential to measure cardiac hemodynamics in heart failure patients with an ICD. Current PEA Devices: BEST Sensor Rate–Adaptive Pacemakers (Sorin Best-Living Systems: Miniliving D and S) These PEA devices use a dedicated unipolar ventricular pace/sense lead, which is IS-1 compatible. Together with the two electrodes for PEA signals, a tripolar connector is necessary. These devices also incorporate an activity sensor detecting the vibration (gravitational sensor). In the PEA mode, continuous PEA signal collected over time is used as a reference, against which the change in PEA is compared. A linear pacing rate response curve is used to relate the change in PEA signal and the rate response. In the dual-sensor mode, the combined output from both sensors is used to drive the rate to an intermediate level, and further rate increases are guided by the PEA sensor alone. In addition, prolonged absence (about 45 minutes) of signals from both the gravitational and the PEA sensor will allow the basic rate to decrease to the programmable rest rate. Clinical Results Clinical studies show good correlation between the sinus rate and PEA sensor–indicated rate during daily life activities and submaximal stress testing.31,32 Similar results were obtained in patients tested during electrophysiologic studies using an external system; changes in PEA were linearly related to the RV dp/dt during dobutamine infusion. PEA signals have been used to monitor hemodynamic function and program the A-V interval. In 13 patients with end-stage heart failure implanted with a DDD-PEA device with a custom lead arrangement, PEA level during RV, LV, and biventricular (BiV) pacing were compared.32 Both LV and BiV pacing increased stroke volume (21% and 37%, respectively) compared with RV pacing, and mean PEA changes over 15 minutes were also higher (43% and 38%). In addition, an apparent minimum PEA level at the optimal A-V interval has shown promise in automatically detecting the optimal A-V interval in a dual-chamber device.33 The algorithm is now automatic. An increase in PEA during head-up tilt-table testing has been observed, and the use of PEA-driven overdrive pacing in patients with vasovagal syncope has been reported.34 Patients randomized to DDDR have a reduced frequency of syncope compared with DDI pacing. In 15 patients with a CRT pacemaker, the PEA correctly indentified the optimal AV interval in 75% of patients.35 These data suggest the potential use of PEA sensor for hemodynamic monitoring (see later). Advantages and Limitations PEA is a proportional sensor that shows good correlation to cardiac workload, especially at higher ranges. The PEA sensor is limited by the need for a specialized lead, leading to concerns about its long-term stability and its use at pacemaker replacement (when a standard header pacemaker is used). Development of the PEA sensor in an atrial lead allows its use in patients requiring an ICD lead.

Current Combined Sensor Devices Experience with sensors has suggested that rapidly responding sensors (e.g., activity) are not proportional at higher levels of cardiac workload, whereas proportional sensors are usually slow in response. As a result, an activity sensor overpaces at low activity levels but underpaces at higher exertional levels (Fig. 5-10). Furthermore, single sensors may be limited by insensitivity to nonexercise stress and are prone to

interference by nonphysiologic causes. Thus, it is logical to enhance their rate-response profile by combining two or more sensors. Two principles guide the combining of sensors: sensor blending and sensor cross-checking. Sensor blending involves combining sensordriven rates from individual sensors in a certain ratio. This can be the “faster win” method, in which the higher rate indicated by either sensor is chosen as the pacing rate, or ratios of the individual rates are added together to compute the actual rate response. Sensor cross-checking enhances the specificity of each sensor. If a more specific sensor registers “no exercise” or physiologic stress, changes in the other, less specific sensor can be ignored or its response attenuated (Fig. 5-11). Table 5-3 summarizes instrumentation of current dual-sensor devices. Details of the combined CLS/activity and PEA/activity devices are discussed earlier. CURRENT DUAL-SENSOR DEVICES Medtronic Pacemakers Medtronic marketed two dual-sensor pacemakers: the combined activity and QT sensors (Vitatron) and the combined activity and MV sensors (Kappa 400). Both are no longer manufactured. In the activity/ QT device, a “faster win” algorithm was used, with highest rate of two sensors used to determine actual pacing rate. Since the QT sensor is more specific for physiologic changes, an increase in activity sensor level in the absence of QT documenting exercise was used as a crosscheck, resulting only in a brief and limited rate response. Using an adaptive sensor level to the SURL, the QT sensor level was continuously scaled down if, at the SURL, QT continued to shorten, suggesting the sensor level had reached the upper rate too soon. The converse was also operative. The combined sensor rate was found to be closer to the sinus rate during daily activities.36 In the Kappa 400, a piezoelectric sensor is used for activity sensing, and MV is sensed from a bipolar ventricular lead. Differential sensor blending is used. Up to the ADL rate, activity input predominates, whereas MV-driven pacing will predominate at the SURL. Activity and MV sensors were checked against one another. In the absence of piezoelectric sensor indicating exercise, MV pacing only reached the ADL rate, and vice versa. Only when both MV and activity signified exercise would pacing occur above the ADL rate. The dual-sensor rate response has been reported to be reliable for both maximal and submaximal activities, as well as resistant to nonphysiologic interference.37 Compared with MV sensor alone, dualsensor mode reduces oxygen deficit acquired during exercise by enhancing the initial rate response.38 “Quantitative rate adaptation” is also superior in the dual-sensor mode.39,40 “Rate profile optimization” was found to be a useful method for rate-adaptive programming, comparable to manual programming,41 and may be superior to accelerometer alone for rate optimization.11 Boston Scientific (Insignia, Pulsar Max, Altrua) An accelerometer activity sensor is integrated with the MV sensor, using differential sensor blending. At low heart rates, the blended sensor rate is approximately 80% accelerometer and 20% MV sensor. This ratio changes to 40%/60% near the SURL. In addition, if the MV rate is higher than the accelerometer rate, the dual-sensor rate follows the MV level. In a study involving 120 patients with Insignia, the programmed sensor modes were randomized to the accelerometer sensor alone, the MV sensor alone, or the dual-sensor mode, each for a 3-month period.42 Using the implanted Activity Log to determine the mean percentage and intensity of activity, quality of life (QOL), and New York Heart Association (NYHA) classes were assessed at the end of each period. Overall, either single-sensor DDDR mode improved Activity Log, QOL, and NYHA scores compared with DDD pacing, but with no difference between the two sensors. The dual-sensor mode did not improve these measurements. This study may have been limited by the prolonged, triple-crossover design but does suggest that clinical differences are likely minimal between sensors and their combinations.



5  Implantable Sensors for Rate Adaptation and Hemodynamic Monitoring 1.00

1.0

HR (normalized)

Observed

Expected Quartile 1 137%

0.25 Quartile 2 81%

Quartile 3 66%

A

0.25

0.50

1.00 Max VO2

0.75

Observed

1.0 Total area 122%

0.8

0.6

HR (normalized)

0.4 Quartile 3 100%

0.4

0.2

0.0

Metabolic equivalents (normalized)

Total area 30%

0.8

0.6

Quartile 2 166%

0.8

1.0

Expected

Observed

Quartile 1 352%

1.0

B

Metabolic equivalents (normalized)

HR (normalized)

Expected 0.4

0.0

0.00 Min VO2

0.6 Expected 0.4

Observed

0.2

Quartile 4 75%

0.0

0.0

0.00 Min VO2

C

0.6

0.2

Quartile 4 102%

0.00

0.2

Total area 85%

0.8 HR (normalized)

Total area 89%

0.75

0.50

153

0.25

0.50

0.75

1.00 Max VO2

Metabolic equivalents (normalized)

1.0

D

0.8

0.6

0.4

0.2

0.0

Metabolic equivalents (normalized)

Figure 5-10  Quantification of pacing percentage of minute volume (MV) and activity pacer during graded treadmill exercise. A, Results of the normalized heart rate and workload in one patient. There is overpacing in quartile 1 (137%), underpacing in quartiles 2 (81%) and 3 (66%), and near-ideal pacing in quartile 4. B, Near-ideal rate recovery of the MV sensor. C and D, Exercise response of an activity sensor, with overpacing most of the time, but inadequate rate recovery. (From Kay GN: Quantitation of chronotropic response: comparison of methods for rate-modulated permanent pacemakers. J Am Coll Cardiol 20:1533, 1992.)

TABLE

5-3 

Types of Dual-Sensor Pacemakers in Current Use

Manufacturer Medtronic

Models Kappa 400

Sensors ACT: piezoelectric MV: impedance

Algorithms Blending ≤ADL range: ACT + MV ADL-ER range: mainly ACT

Boston Scientific

Pulsar Max Insignia Altrua

ACT: accelerometer MV: impedance

ELA-Sorin

Chorus Talent Symphony Rhapsody Talos Cylos Entoris

ACT: accelerometer MV: impedance

Blending Low heart rate: ACT 80%, MV 20% High heart rate: ACT 40%, MV 60% No blending MV-determined rate response if ACT indicates exercise No blending No ACT rate contribution Rate response determined only by CLS

Biotronik

CLS: unipolar ventricular impedance ACT: accelerometer

Cross-checking ACT (0) and MV (+): up to ADL rate ACT (+) and MV (0): up to ADL rate ACT (0) and MV (+); and ACT (+) and MV (0): limited rate ACT (+) and MV (0): initial limited rate response ACT (0) and MV (+): rate recovery ACT (0) and CLS (+): limited rate response ACT (+) and CLS (0): no rate response

Automaticity Rate Profile Optimization

AutoLifestyle

Automatic matching MV sensor to LRL and SURL Auto Response Factor adjusts CLS data to reach rate distribution determined by the programmed Exertion Threshold Rate

ACT, Activity level; ADL, activity of daily living; CLS, closed-loop stimulation; ER, exertion response; LRL, lower rate limit; MV, minute ventilation; SURL, sensor-driven upper rate limit.

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SECTION 1  Basic Principles of Device Therapy

TABLE

SR

5-4 

Select Patients for Dual-Sensor Pacemaker

Clinical Situation High exercise heart rate preferred

S1

Avoidance of undue rate acceleration Need for nonexercise rate response

S2

Examples Young, athletic individuals may require a more physiologic sensor. Working in a vibrational environment Vasovagal syncope Stress response

Hemodynamic or other monitoring

Exercise

Nonexercise stress

Interference

Proportionality and speed

Sensitivity

Specificity

Figure 5-11  Algorithms for sensor combinations needed to achieve better proportionality and speed of response, sensitivity, and specificity. Top to bottom, Graphs depict the responses of the sinus node (SR), sensor 1 (S1), sensor 2 (S2), and combined rate profile (S1 + S2). SR shows ideal proportionality, speed of rate response, and freedom from interference. S1 is a rapidly responding sensor, although it is neither proportional nor sensitive and is susceptible to interference. S2 is a proportional and sensitive sensor, although it has a slow response. It is also specific to exercise. Note the improved ability of the combined sensor approach in simulating the sinus rate under different conditions.

The Limiting Chronotropic Incompetence for Pacemaker Recipients (LIFE) study of 1256 patients with chronotropic incompetence found that the blended sensor improved metabolic-chronotropic slope compared with the activity sensor alone43 (Fig. 5-12). However, besides improved physical activity time/energy expenditure, there was no measurable difference in QOL between the blended sensor and the activity sensor alone. A prospective randomized parallel study on activity and MV with the primary outcome of Vo2 is planned (APPROPRIATE study).44 This study will randomize 1000 patients who will be screened for CI using a 6MWD and treadmill heart rate at 1 month. Patients with confirmed CI will have their sensors optimized with a brief walking test. Apart from Vo2 max, rate changes during daily activities will be assessed. This study will shed light in the role of different sensors on exercise oxygen uptake. (Unfortunately, the study was prematurely terminated because of suboptimal recruitment.) ELA-Sorin (Chorus, Talent, Symphony, Rhapsody, Reply) An accelerometer and a MV sensor are combined in these devices. Sensor blending and cross-checking are both operative to effect rate adaptation during exercise (Fig. 5-13). When the accelerometer is active but MV has not increased, as may occur at the beginning of exercise, rate response occurs according to a fixed activity response curve to a limited rate. When MV increases, rate response follows the MV sensor–driven rate. Persistent absence of accelerometer signal is considered the cessation of exercise, and this allows a decrease in the pacing rate using a recovery curve to the LRL, even though MV remains higher than baseline. A multicenter study involved 81 patients with the ELA dual-sensor pacemakers.19 In patients who underwent exercise stress testing at 1-month follow up, sensor-driven rate had good correlation with the sinus rate (r = 0.92 ± 0.07; P < .001), with the slope of linearity at 1.0 ± 0.2. Using metabolic reserve to relate to heart rate reserve, a slope of 1.1 ± 0.2 was obtained, suggesting a close relation between the dualsensor rate and metabolic workload. ARE DUAL SENSORS JUSTIFIED? Dual-sensor implementation is technically and clinically feasible and provides a rate profile closer to the normal sinus rate. In practice, a more specific sensor is added to the activity sensor to enhance its

specificity. However, apart from physiologic parameters such as oxygen uptake and transport kinetics, no clinical benefit over single sensor has been seen in large groups of patients, likely reflecting the limited exercise activities of most patients meeting standard indications for pacing. For select patients who are highly active, however, these dual-sensor pacemakers may provide important clinical benefits (Table 5-4).

Clinical Benefits of Sensor-Driven Pacing Controversy remains about the clinical benefits of sensor-driven pacing, especially the lack of symptomatic benefit observed in some comparative studies. Factors that may affect such comparisons include (1) definition of CI, (2) programming of sensor (extent of rate adaptation), (3) patient populations including their comorbidities, (4) pacing modes, and (5) site of ventricular pacing (Table 5-5). CHRONOTROPIC INCOMPETENCE Although factors contributing to CI can be obvious—age, coronary artery disease, sinus node dysfunction, and cardioactive medications— what constitutes CI remains uncertain. CI can vary from 9% to 84%, MCR SLOPE

1.0

Heart rate reserve (%)

S1 + S2

0.8 p <0.001

0.6

0.4

0.2

0.0 0.0

0.2

0.4

0.6

0.8

1.0

Metabolic rate reserve (%) Figure 5-12  Sensors in patients with chronotropic incompetence (CI). Overall sensor response (blended sensor [BS] and activity sensor [XL]) from 1 to 6 months in CI patients with adequate exercise tests. Blue lines indicate paired BS responses from 1 month (dashed line) to 6 months (solid line). Orange lines indicate paired XL responses from 1 month (dashed line) to 6 months (solid line). The black solid line indicates a normal metabolic-chronotropic relationship (MCR) slope; black dashed lines, 96% confidence interval CI (0.8-1.3). (From Coman J, Freedman R, Koplan BA, et al: A blended sensor restores chronotropic response more favorably than an accelerometer alone in pacemaker patients: the LIFE study results. Pacing Clin Electrophysiol 31:1433-1442, 2008.)



5  Implantable Sensors for Rate Adaptation and Hemodynamic Monitoring SENSOR BLENDING

155

AUTOCALIBRATION

Rate

Rate

Y SURL

SURL

LRL

X LRL

Rest

Exercise

Rest

Rest MV/ACT

B

Exercise MV/ACT

ACT

A

MV

Figure 5-13  Algorithm of dual-sensor device (ELA-Sorin). A, Combination of activity (ACT) and minute ventilation (MV) inputs during exercise. B, Autocalibration of the sensor. If the maximum sensor level is reached before the sensor-driven upper rate limit (SURL; indicated by X), then the slope is adjusted in steps to Y to match the SURL to the maximum sensor level. LRL, Lower rate limit.

TABLE

5-5 

Factors Affecting Clinical Benefit of Sensor-Driven Pacing

Issues Patient related

Device related Programming Methods of measurements

Factors Degree of chronotropic incompetence Left ventricular function Sinus vs. atrial fibrillation Age and comorbidities Sensor type Site of ventricular lead(s) Rate adaptive pacing mode Sensor programming Lower and upper rate Percentage of ventricular pacing Protocol: crossover or parallel arm Exercise Symptoms Quality of life

This study did not address whether rate adaptation improves exercise or QOL. However, patients with CI had a lower QOL (physical and mental component) than those who were chronotropic competent. With correction of CI in the DDDR mode, there was an increase in physical activity time and kilocalories expanded per week compared to baseline DDD at 1 month, with a greater tendency toward higher levels for the dual sensors than the single, activity sensor (mean change for activity score from baseline to 6 months: activity sensor = 1.45 ± 17.12, dual sensor = 2.69 ± 15.3; and energy expenditure change: activity sensor = 359 ± 367, dual sensor = 716 ± 3810

100 90

86%

87%

%

80

70%

70 Percent

depending on the definition used.43,45-57 Based on cardiac mortality, CI defined as “inability to reach 85% maximum age-predicted heart rate” will identify 2.2 times as many individuals with increased cardiac mortality.45 From the patient’s perspective, CI should be defined as a level at which its reversal by a sensor will improve exercise capacity and QOL. There are few long-term studies in pacemaker patients, although it is generally accepted that inability to reach a rate greater than 100 bpm represents severe CI.48 The metabolic-chronotropic relationship may be a scientific way to classify CI at peak and each level of exercise, with a slope value of the plot of heart rate reserve to metabolic reserve of less than 0.8 identifying CI.49 In 547 pacemaker patients, exercise testing was used to identify CI at 1 month after pacemaker implantation.43 Patients were recruited in this study if they could complete three or more stages of CAEP protocol and attained a perceived-exertion Borge scale of 16 or higher. CI is defined as a heart rate reserve less than 80% of metabolic reserve at each stage of the CAEP exercise test, and if more than 0% of these stages were incompetent, CI is identified. Based on this study, 50% of these patients had CI. The study further assessed the prevalence of CI using different definitions in the same cohort according to the previous definition, and found that the incidence is similar to using maximum predicted heart rate of less than 70% as CI. The other definitions tend to be more sensitive and identify 70% to 87% of recipients with CI (Fig. 5-14).

60 50 40

34%

35%

LIFE

70% MPHR

30 20 10 0 85% Metabolic MPHR chronotropic slope <0.8

80% HRR

Figure 5-14  Prevalence of chronotropic incompetence (CI). Results in 547 pacemaker recipients during maximum exercise according to published criteria. See text for LIFE study CI definition (Fig. 5-12). HRR, Heart rate reserve; MPHR, maximum [age] predicted heart rate. (Data from Coman J, Freedman R, Koplan BA, et al: A blended sensor restores chronotropic response more favorably than an accelerometer alone in pacemaker patients: the LIFE study results. Pacing Clin Electrophysiol 31:1433-1442, 2008.)

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SECTION 1  Basic Principles of Device Therapy

Kcal).43 These data suggest that a more stringent criterion for CI may identify a group of pacemaker recipients who are more symptomatic and benefit from dual sensors, although the benefit of accurate rate adaptation remains controversial in large-scale studies. UPPER RATE LIMIT The peak heart rate (HR) is inversely related to age. In large studies, mean peak HR has been traditionally estimated by the formula: HRmax = 220 − age. The variance is approximately 15 bpm. This is a useful guide to prescribe physical training, but its application to pacemaker recipients remains controversial. Kindermann et al.50 exercised 49 patients with conventional DDD pacemaker using cardiopulmonary exercise results and Doppler to define the optimal upper rate. In patients with normal heart function (left ventricular ejection fraction [LVEF] ≥55%), the URL was 86% of the maximum predicted HR. In those with impaired heart function (LVEF ≤45%), a 75% maximum predicted HR was optimal, beyond which cardiac output tended to decrease. This may be related to decreased ventricular compliance and other factors, such as ischemia that could occur at higher HR. Of interest, contractility augmentation as a function of increased HR appeared to be attenuated with RV pacing. In 22 patients with heart failure and a wide QRS complex, Vollmann et al.51 found that BiV pacing restored the rate-related contractility when pacing at 80 to 140 bpm, whereas the change in RV pacing was minimal, and the changes during LV-only pacing was intermediate. Thus, the site of ventricular pacing is an important determinant of the URL. RATE-ADAPTIVE PACING MODES VVIR vs. VVI The first rate-adaptive pacing mode introduced was VVIR pacing. This mode was most extensively studied, and often applies to patients who had complete CI, as in those with complete AV block. Independent of the type of sensors, a 69% and 32% increase in heart rate and exercise capacity has been documented over VVI pacing.52 In addition, VVIR pacing has been shown to improve most aspects of QOL over fixed-rate pacing.53 AAIR vs. AAI/VVI An infrequently used pacing mode, AAIR has the theoretical advantage of preserving native conduction and avoiding ventricular pacing altogether. The disadvantages include progression of AV conduction disease, even in patients with normal AV conduction at baseline, and potential induction of Wenckebach AV block by inappropriately high atrial pacing rates. How quickly AV conduction can adapt with exercise may be an issue, because lack of AV adaptation during an appropriate sensor-driven atrial rate may prolong the P-R interval, resulting in AV dyssynchrony. If the P-R interval is greatly prolonged, the pseudo– pacemaker syndrome may result.54 An early study on AAIR pacing compared AAIR, DDDR, and VVIR mode in patients with intact intrinsic AV conduction (1 : 1 A-V ≥100 bpm; atrial-R ≤220 msec).55 Patients with VVIR pacing had worse QOL than AAIR patients, as well as more palpitations and depression despite a lower frequency of ventricular pacing than DDDR patients (39% ± 7% vs. 64% ± 11%). All patients initially had normal LV function, which may have obscured minor differences between modes. In a further study, AAIR pacing that increased the maximum heart rate to 142 bpm did not increase oxygen uptake compared with VVI backup at 40 bpm in patients with impaired LV function.55 The study was limited by an intrinsic maximum rate of 130 bpm even without rate adaptation. DDDR vs. DDD Much controversy surrounds the potential benefit of DDDR versus DDD pacing. In a small cohort of 10 patients, DDDR improved exercise capacity over DDD pacing in patients with severe CI.46 However, in a more recent trial including 873 patients, despite an increase in HR

from 101.1 ± 21.1 bpm to 113.3 ± 19.6 bpm at 6 months, there was no difference in exercise capacity between the two pacing modes.56 More than 90% of beats in the two modes were ventricularly paced, and a potential reason for lack of benefit may be related to the potentially negative effects of RV apical pacing. In a small study, van Hamel et al.57 compared DDDR versus DDD mode in 99 patients with either AV block or sinus node disease. They found no difference between DDDR and DDD mode in terms of QOL, despite significantly higher atrial pacing rate in the sensor-driven arm. Again, this study documented a high percentage of RV pacing in both sinus node patients (87.9%) and AV block patients (96.8%). Pacemaker Operations That Minimize Ventricular Pacing With the recognition of disadvantages of excessive RV pacing, most manufacturers provide some form of A-V interval extension to reduce RV pacing and allows intrinsic conduction. These features are usually inactivated during the DDDR mode. In the minimal ventricular pacing algorithm of the Medtronic, the device functions in AAI ↔ DDD, and although rate response can be programmed, it will only be operative in the DDIR mode that occurs after a mode switch for atrial fibrillation. It will be of interest to see if this mode can be sensor driven to allow rate adaptation with minimal ventricular pacing. VENTRICULAR PACING SITE The role of sensor-driven pacing in CRT/CRT-D is discussed earlier for patients with heart failure. Recent observations indicate that RV apical pacing may be associated with impaired ventricular function in the long term, especially in patients with a depressed LVEF who do not require pacing. Thus, the benefit of rate increase might be mitigated by the harmful effect of RV apical pacing. Inducing a narrower QRS complex in RV septal pacing better preserves LV function.58 A recent study in 12 patients with high-level AV block (8 of 12 with permanent atrial fibrillation) in the VVIR mode received an upgrade from chronic RV apical to RV septal VVIR pacing.59 Over 18 months, these patients showed a significant increase in LVEF (55.2 ± 2.6 to 60.4 ± 2.9%) and improvement in 6 MWD. A control group with simple pacemaker replacement had no change in LV function. This study suggests that the potentially harmful effect on LV function by VVIR pacing can be ameliorated with septal VVIR pacing.

Conclusion Sensor-driven pacing is now an integral part of virtually all pacemakers and ICDs. Activity sensing, despite several theoretical disadvantages, is widely used, followed by the closed-loop stimulation and minute ventilation sensors. The clinical benefit of dual sensor over single sensor is likely to be small, but may be useful for certain patient cohorts. The lack of observed clinical benefit of sensor-driven pacing in DDDR mode may be related to patient co-morbidities, the definition of CI, excessive RV apical pacing, and the site(s) of pacing used.

SENSORS FOR HEART FAILURE MONITORING Acute Decompensated Heart Failure About 5 million Americans suffer from heart failure, with an estimated 550,000 new cases each year and resulting in significant mortality and morbidity. Furthermore, estimated annual treatment costs $33.2 billion, most incurred during hospitalization for acute decompensated heart failure (ADHF).60 ADHF refers to a clinical condition of worsening heart failure with dyspnea and often with evidence of fluid overload.61 ADHF is often triggered by one of four main factors: atrial fibrillation, anemia, hypertension, or medication/dietary indiscretion. In the 1991–1994 Connecticut Medicare beneficiaries review, ADHF resulting in hospitalization carried an 8% in-hospital mortality.62 Importantly, of the 17,448 survivors, 44% were readmitted once, 18%



5  Implantable Sensors for Rate Adaptation and Hemodynamic Monitoring

of these for recurrent heart failure. Overall, 24% died within 6 months of the first ADHF event, and 53% either died or were readmitted over the same period. The prognosis for heart failure admission has improved over time,60 partly because of adherence to optimal doses of evidenced-based therapy. Despite this, prevention of ADHF can have significant prognostic impact for the patient, in addition to reducing the cost of heart failure management. LIMITATIONS OF SYMPTOMS, SIGNS, AND INVESTIGATIONS Although the most common presenting complaint for hospitalization in ADHF patients, dyspnea occurs relatively late in relation to changes in hemodynamics and intravascular volume. Until the era of implantable devices, it was not possible to evaluate the hemodynamic changes leading to ADHF except by repeated catheterization. Adamson et al.63 implanted an RV pressure sensor to measure RV systolic and diastolic pressure during heart failure exacerbation in 32 patients. They found that at a mean of 4 ± 2 days before admission, RV systolic pressure started to increase in 9 of 12 heart failure events (Fig. 5-15). As a group, RV systolic pressure increased by 25% ± 4% and heart rate by 11% ± 2% during ADHF, with substantial changes also during minor exacerbation. This study suggests that intracardiac pressure changes precede symptoms of heart failure exacerbation. Likewise, an implantable intrathoracic impedance sensor to assess pulmonary fluid showed that fluid overload occurred 18.3 ± 10.1 days before dyspnea.64 Thus, 40

Percent change

30

Hospitalization

Pressure change (9 of 12)

20 10 0 –10 Baseline –7 –6 –5 –4 –3 –2

A

Recovery

Days relative to event 40 30

Percent change

–1

Pressure change

Event

20 10 0 –10 Baseline –7 –6 –5 –4 –3 –2

B

–1

Recovery

Days relative to event

Figure 5-15  Right ventricular (RV) sensor to measure RV pressure during heart failure exacerbation in 32 patients. A, Pressure changes occurred in 9 of 12 events 4 ± 2 days before hospitalization (major events). B, Pressure changes before minor exacerbations (n = 24) that did not require hospitalization. Percent changes in RV systolic pressure (black circles), estimated pulmonary artery diastolic pressure (green diamonds), and heart rate (red triangles). (From Adamson PB, Magalski A, Braunschweig F, et al. Ongoing right ventricular hemodynamics in heart failure: clinical value of measurements derived from an implantable monitoring system. J Am Coll Cardiol 41:565-571, 2003.)

157

dyspnea is a late event and does not allow either the clinician or the patient enough time for intervention to avert hospitalization. The cardinal physical signs of congestive heart failure are a third heart sound, pulmonary crackles, raised jugular venous pressure, and pedal edema. However, these signs have poor sensitivity to detect heart failure. In a study of 50 patients with increased pulmonary artery wedge pressure (PAWP ≥22 mm Hg), lung crackles were identified in 19% of patients, and higher jugular venous pressure and peripheral edema were present in only 50% and 20% of patients, respectively.65 A third heart sound was heard in most patients, but it was also detected in those with a low PAWP. This combination of signs carries a sensitivity of 58% and specificity of 100% for congestive heart failure. Physical examination of jugular pressure is difficult and often inaccurate.66 Body weight change is both a symptom and a sign that predicts ADHF. Some suggest this may be an unreliable sign because body weight not only reflects body fluid, but also may depend on amount of food/fluid intake and other causes of weight loss or weight gain. In a recent study, 134 patients with heart failure hospitalization were compared with a case-matched group without hospitalization.67 Body weight gain 1 week before hospitalization was associated with increasing risk of hospitalization (relative risk: Δ2-5 lb = 2.77; Δ5-10 lb = 4.46; and >10 lb = 7.65). Therefore, monitoring of body weight remains useful to identify a high-risk group for intervention and provides supportive evidence for ADHF. Radiologic evidence of ADHF tends to occur late. Brain natriuretic peptide (BNP) has been proposed as a means to monitor and improve heart failure management. In a randomized study on the N-terminal BNP to guide heart failure therapy (TIMECHF) study, 499 patients 60 years or older with systolic heart failure were randomized either to monitoring of BNP to maintain a level twice or less the upper limit or to conventional management without BNP guidance.68 There was no significant difference in the probability of remaining free from all-cause hospitalization between BNP-guided and conventional therapy (41% vs. 40%). Both groups had similar quality of life (QOL). However, the secondary endpoint of heart failure hospitalization was significantly reduced (72% vs. 62%), and outcomes were better in patients age 60 to 74 but not in those 75 or older. Although vigilant monitoring of symptoms and signs (and BNP levels) is useful, this strategy is not a guarantee for accurately predicting ADHF. On the other hand, frequent monitoring of some signs and symptoms and the use of external physiologic (and implantable) data may be superior to conventional care. A meta-analysis of both cohort (2354 patients) and randomized trials (6258 patients), with 6 to 12 months of follow-up, found a significantly lower rate of mortality and hospitalization compared with conventional management.69 Relative reduction in heart failure hospitalization of 0.93 and 0.52 was seen in randomized and cohort studies, respectively, and mortality was reduced by 0.83 and 0.53. The Cochrane Database Systematic Review69a reviewed 25 studies and 5 abstracts and suggested that telephone support and telemonitoring are effective in reducing all-cause mortality and heart failure hospitalization, and thus improve QOL and reduce cost and evidence-based prescribing. Several prospective randomized studies have been published since these meta-analyses. The Randomized Trial of Phone Intervention in Chronic Heart Failure (DIAL) trial69b randomized 1518 outpatients to either nurse-led telephone-based intervention or conventional treatment. The three main supervision areas were weight, diet, and medication compliance, and nurse specialists were allowed to change diuretic doses according to specified criteria. Patients were organized to call once every 14 days for three times, and thereafter at a frequency depending on the severity of heart failure. This protracted study followed patients up to 3 years after trial completion. The primary end point (death or heart failure hospitalization) was reduced in the intervention versus conventional arm, with the main benefit being a reduction in hospitalization (28.5% vs 35.1% at 3 years). An educational effect and adherence to the above three supervision areas are considered the main reasons for improvement. The Telemedical Interventional Monitoring in Heart Failure (TIMHF) trial69c also randomized stable Class II or III heart failure patients

158

TABLE

5-6 

SECTION 1  Basic Principles of Device Therapy

Monitoring for Heart Failure Pathophysiology

Pathophysiologic Area Electrical Remodeling Atrium Ventricle Conduction Mechanical Remodeling Atrium Ventricle Pressure changes

Monitoring Examples Rate, atrial ERP, atrial fibrillation Rate, ventricular ERP, ventricular tachyarrhythmias Atrioventricular node conduction time and capacity Intra-atrial and interatrial conduction Intraventricular and interventricular conduction Size, function, structure Size, function, structure End-diastolic pressure, pulmonary artery pressure (and wedge pressure), venous pressure

Neurohormonal Changes Sympathovagal Heart rate variability imbalance ERP, Effective refractory period.

who had a history of heart failure hospitalization within 2 years to either remote data monitoring (ECG, blood pressure, body weight) and sent these data via a personal digital assistant to the monitoring center daily, with physician-led response. Over a median follow up of 26 months, there was no difference in heart failure hospitalization or death, nor in overall QOL score. However, physical functioning was improved. The study documented an 81% compliance in ≥70% daily transmission of data. The Telemonitoring in Patients with Heart Failure (TELE-HF) trial69d randomized 1653 patients who had recently been admitted with ADHF. A daily telephone-based interactive voice response system on symptoms and weight was reviewed by the patients’ physicians. At the end of 180 days there was no significant difference in the primary end point of death and heart failure hospitalization between the telephonebased intervention versus the usual care group (52.3% vs 51.5%, respectively). Again, this study shows only a 55.1% adherence to interventional therapy at the end of 26 weeks in the 85.6% of patients who made at least one call. The different efficacies of telemonitoring reported between metaanalyses and prospective studies may be attributable to differences in severity of patients involved in the studies, the types of clinical variables measured, and the intervention offered. Among these, suboptimal compliance and lack of rapid therapeutic intervention (which often varies between studies) are important drawbacks. Furthermore, these studies have mainly been limited to easily measurable clinical variables that occur relatively late in the course of ADHF. Because many patients with LV dysfunction require cardiac implantable electronic devices (CIEDs) for arrhythmia prevention or therapy (implantable cardioverter-defibrillator, ICD) or for heart failure treatment (cardiac resynchronization therapy [CRT] or CRT-defibrillation [CRTD]), implantable sensors for heart failure monitoring may be added in these devices. The use of implantable sensors allows continuous monitoring of physiologic parameters with less labor-intensive efforts. Furthermore, the ability of these sensors to detect earlier physiologic changes before ADHF may help prevent heart failure hospitalization. Such sensors can also measure hemodynamic values in different body positions and on an ambulatory basis, and may obviate the need for invasive monitoring on admission for ADHF therapy. MONITORING OF PATHOPHYSIOLOGIC CHANGES The three pathophysiologic areas for heart failure monitoring are electrical remodeling, mechanical remodeling, and neurohormonal changes that occur with heart failure (Table 5-6). Electrical remodeling in either the atrium or the ventricle will result in changes in normal automaticity, conduction properties, and refractory period and will predispose to atrial fibrillation (AF) and ventricular tachyarrhythmias. Arrhythmias are routinely monitored and treated by CIED, either by pacing or defibrillation. The effective refractory

period (ERP) can be monitored by physician-activated electrical stimulation through a programmer. Monophasic action potential duration could be assessed by telemetry in some early devices.70 Using fractalcoated iridium electrodes, the paced unipolar ventricular signal was found to correlate with acutely measured monophasic potential. However, the implications of ERP monitoring remain uncertain. Intrachamber and interchamber conduction times in heart failure are important because they are either the result or the cause of electrical remodeling in heart failure. A wide left bundle branch block (LBBB) QRS complex is associated with ventricular dyssynchrony and impaired LV function. In a postmortem series of 34 patients with premorbid serial ECGs, progressive PR and QRS prolongation occurred with worsening of heart failure. Cardiac mass index was also significantly correlated with V1-V6 R-wave amplitudes.71 Optimal left-sided atrioventricular (AV) and cross-chamber sequencing are important for the proper systolic function of the heart (see Chapter 30). Although relatively simple, these conduction parameters have not been examined as long-term predictors of cardiac decompensation. The major development of sensors in heart failure monitoring is to track mechanical remodeling, including chamber size, pressure changes, and pulmonary congestion consequent on heart failure. Such sensors use either impedance or specialized pressure-sensing leads or devices. A variety of neurohormonal changes result from the compensating mechanisms of heart failure. These changes are not easily measurable directly but may be reflected through heart rate variability and electrical repolarization such as the QT interval. This will be an interesting new area for monitoring and as a guide to therapy.

Sensors for Heart Failure Sensors for rate adaptation can be classified according to the technical instrumentation. Table 5-7 summarizes the sensors that have been used for or proposed to monitor the events of heart failure. The paced QRS complex enables the QT duration to be evaluated. In addition, the RV-evoked response, measured as the ventricular depolarization gradient, has been proposed as a marker that is sensitive to catecholamines.72 These sensors have long been used as surrogate markers of sympathetic level that increases with exercise, with a view toward rate-adaptive pacing. It is uncertain if they are also good markers of increase sympathetic tone that accompanies heart failure. Heart rate variability (HRV) can be measured in a non-pacing-dependent patient and is a well-established prognostic marker of heart failure because it reflects sympathovagal imbalance.73 Reduced HRV precedes heart failure events. Although percentage of biventricular pacing and HRV are rhythm diagnostics rather than implantable sensors, they provide important indications of the prevailing heart failure status. Piezoelectric crystals or accelerometers are used to monitor body motion. Decreased body activity is an intuitive consequence of heart failure exacerbation. Impedance can be used to measure pulmonary TABLE

5-7 

Sensors for Monitoring Heart Failure

Sensors Electrogram

Intraventricular impedance Specialized lead

Transthoracic impedance Accelerometers

Parameters Percentage of pacing ST segment AF Ventricular arrhythmias Paced QRS Paced QT interval Heart rate variability LV size and function PEA to monitor LV function PA pressures PA wedge pressures LA pressures Pulmonary edema Minute ventilation Activity level

AF, Atrial fibrillation; LA, left atrial; LV, left ventricular; PA, pulmonary artery; PEA, peak endocardial acceleration.



5  Implantable Sensors for Rate Adaptation and Hemodynamic Monitoring

edema and MV in a transthoracic manner. Intraventricular impedance is used to monitor LV size and contractility. Direct measurements of intracardiac pressure are now possible with pressure sensors instrumented in the right ventricle, pulmonary artery, or left atrium. A decrease in the mixed-venous oxygen saturation is highly associated with reduced cardiac output (or exercise) and has been measured by implantable devices. ACTIVITY MONITORING Kadhiresan et al.74 reported the use of externally attached accelerometers at the chest wall to monitor walking distance in heart failure patients. Using an acceleration threshold of 50 mG, a walking speed of 2 mph (≅2.8 METS) could be detected in a group of 30 patients. The activity log index so defined was closely related to the walking distance and increased when patients were tested during CRT-on versus CRT-off phase. Similarly, time per day with physical activity greater than a threshold of 70 steps/min was detected in an implanted accelerometer sensor,75 and the activity level trend was similar to HRV trending during ADHF and correlated with NYHA functional class at baseline (Fig. 5-16). 350

MDPA (min/day)

300 ***

250 200

***

150 100 50 0 0

1

2

3

4

A

5

6

7

8

9

10 11 12

Week after implant NYHA at baseline II

III

IV

100

SDANN (ms)

90 80

***

70

*

60 50 40 0

B

1

2

3

4

5

6

7

8

9

10 11 12

Week after implant

Figure 5-16  Activity and heart rate variability (HRV) after CRT implantation. Trend data analysis performed weekly: A, mean daily physical activity (MDPA; mean ± SEM; ***P < .001); B, standard deviation of intrinsic intervals in 288 5-minute segments of day averaged over a week (SDANN; mean ± SEM; *P < .005). Results are grouped by preimplantation assessment of disease severity; New York Heart Association (NYHA) classification. Both activity and HRV improve with CRT; the level of improvement is related to NYHA class. (P = repeated-measures analysis of variance for differences between trends.) (From Braunschweig F, Mortensen PT, Gras D, et al: Monitoring of physical activity and heart rate variability in patients with chronic heart failure using cardiac resynchronization devices. InSync III Study Investigators. Am J Cardiol 95:1104-1107, 2005.)

159

Although a crude index, activity sensing is a good reflection of general well-being of the patient with heart failure. It is a readily available sensor in most CIEDs and uses minimal battery energy. In addition, the absence of activity signifies the patient is at rest, allowing determination of other measurements (e.g., respiration). The use of accelerometer rather than piezoelectric crystal avoids some of the external vibration interference on measuring body activity. An activity sensor is used with other sensors for heart failure monitoring. HEART RATE VARIABILITY In patients with heart failure, neurohormonal activity predicts cardiovascular outcome.74 HRV is an indirect measure of autonomic tone and predicts both sudden and nonsudden cardiac death.76,77 HRV has been proposed not only to predict heart failure severity, but also to guide treatment and perhaps predict ADHF. In a randomized study of 50 patients with an implanted CRT device, HRV was measured in the atrial sensing mode (VDD at 30 bpm) with either CRT-on or CRT-off.63 HRV, measured as the standard deviation (SD) of atrial cycle length of all atrial sensed beats, was 25% higher in the CRT-on versus the CRT-off arm and 27% higher when the patients were receiving a beta-adrenergic blocker. Thus, time-domain HRV improves with CRT, likely reflecting changes in both sympathetic and parasympathetic activities with heart failure improvement. In Boston Scientific devices, HRV is measured by the so-called SDANN, the SD of the intrinsic intervals in the 288 5-minute segments of a day, averaged over a week. If the percentage of intrinsic beats is less than 67% for 24 hours, the data for that day are discarded. Using the SDANN in 113 heart failure patients with CRT, CRT reduced the mean heart rates and increased SDANN (from 69 ± 23 msec to 93 ± 27 msec) after 3 months (see Fig. 5-16). Furthermore, lack of HRV improvement predicts nonresponders to CRT.78 The HRV can be plotted at each heart rate over a 24-hour period, resulting in a so-called footprint79 (Fig. 5-17). The normalized size of the footprint is the “footprint number,” and the graph and number give an easy understanding of the HRV level: “the larger, the better.” In another cohort study, HRV using either SDANN or footprint predicted mortality in 842 patients implanted with a CRT device during 11.8 months of follow-up.80 A clinical score was derived based on 2-week postimplant diagnostic data (SDANN <43 msec, mean heart rate >74 bpm, footprint number <29, activity percent <5%) from a 436-patient cohort in the CRT RENEWAL study.81 Using this scoring system, patients could be triaged to low, moderate, and high risk. When applied to a separate group of CRT recipients in the HF-HRV cohorts, this scoring system accurately predicted the level of mortality risk (low 2.8%, moderate 10.1%, and high 13.4%) depending on tertiles of their score. In Medtronic devices, a long-term measure of HRV, the SDAAM (SD of 5-minute median atrial sensed intervals) is used. The algorithm averages the 24-hour SD of intrinsic atrial cycle length and will exclude the day’s data if the percentage of atrial pacing exceeds 80% or an atrial high rate episode (AHRE) is detected. The change in SDAAM is compared to a rolling average of the preceding 6 months. The SDAAM, nighttime heart rate, and activity level were used to predict outcome and to detect heart failure hospitalization in 397 patients.82 Figure 5-18 shows that SDAAM less than 50 msec predicts overall mortality. Furthermore, the absolute value of SDAAM remained low in patients who were hospitalized or who died. SDAAM declined from 76 ± 27 msec to 64 ± 26 msec at hospitalization, and the change was apparent up to 3 weeks before the event. An algorithm was developed to use SDAAM to predict hospitalization (Fig. 5-19). At a threshold of 200-msec days, a sensitivity rate of 70% was associated with 2.4 false-positive events per patient-year of follow-up, at a median of 16 days before the event. The sensitivity is not affected by the use of beta blockade. The SDAAM is significantly better than the nighttime heart rate and activity level change for predicting heart failure. These data suggest that autonomic surveillance such as the use of HRV is a good way to monitor both heart failure prognosis and to predict ADHF. However, HRV is limited

160

SECTION 1  Basic Principles of Device Therapy

Counters - heart failure/Brady

Counters - heart failure/Brady

04-SEP-2008 to 15-JAN-2009

14-FEB-2009 to 28-APR-2009

Percent

Since last reset

Atrial

0% Paced 100% Sensed ATR switches Minimum duration Maximum duration Mode switch time PMT

2.9 K 14.2 M 1 00:31 m:s 00:31 m:s 0% 2

Paced Sensed ATR switches Minimum duration Maximum duration Mode switch time PMT

Atrial

Right ventricular Right paced 88% Tracked 100% BiV triggered 0% Device determined 0% Right sensed 12%

13.2 M 13.2 M 1 19.2 K 1.8 M

Amio

Left ventricular

Since last reset

Percent 67% 33%

4.9 M 2.5 M 0 00:00 m:s 00:00 m:s 0% 0

Right ventricular Right paced Tracked BiV triggered Device determined Right sensed

7.4 M 2.3 M 0 5.1 M 191.3 K

97% 31% 0% 69% 3%

Left ventricular

Left paced 88% Tracked 100% BiV triggered 0% Device determined 0% Left sensed 12%

13.2 M 13.2 M 0 19.2 K 1.7 M

Single or double PVCs Three or more PVCs

Left paced Tracked BiV triggered Device determined Left sensed

1.3 M 759

97% 31% 0% 69% 3%

7.4 M 2.3 M 0 5.1 M 199.3 K

Single or double PVCs Three or more PVCs

176.0 K 0

A HEART RATE VARIABILITY – PLOT Last measured

HEART RATE VARIABILITY – PLOT

Mean

Last measured

Variability (ms)

Variability (ms)

Amio

0

0 40

60

80 100

200

(min–1)

B

Mean

120

120

Less More frequent frequent

40

60

80 100

200

(min–1) Less More frequent frequent

Figure 5-17  Heart rate variability (HRV) and “footprint number.” A, Patient with dilated cardiomyopathy receiving chronic resynchronization therapy with defibrillation (CRT-D; Boston Renewal). Deterioration of heart failure results from loss of biventricular (BiV) pacing secondary to premature ventricular contractions (PVCs) (%Vp = 88%). After amiodarone (Amio), with reduction of PVCs, BiV pacing increased to 97% with clinical heart failure improvement. B, HRV presented at each heart rate, resulting in two-dimensional plot known as “footprint.” Overall size of the footprint reflects HRV. Footprint improved after a period of CRT, with a greater range of heart rate experienced. Atrial tracking responses (ATR) and pacemaker-mediated tachycardia (PMT) were both suppressed after amiodarone.

when there is a high percentage of atrial pacing and during atrial tachyarrhythmias, and medications may affect HRV measurements. PERCENTAGE BIVENTRICULAR PACING In a cohort retrospective analysis of two heart failure trials involving CRT-D (1812 patients), Koplan et al.82a analyzed the relationship between percentage of biventricular (BiV) pacing and the outcome of death and heart failure hospitalization. When subjects were divided

into quartiles, there was a lower event rate with an increasing percentage of BiV pacing: 44% reduction in event rates in those paced 100% of the time versus those paced less than 92%. Even subjects paced 93% to 97% had a higher event rate than those paced between 98% and 99% (22% higher). The main reason for an inadequate percentage of pacing was atrial arrhythmias. Thus, a high percentage of BiV pacing (≥98%) is an important goal to achieve in patients with CRT. A low percentage of BiV pacing is particularly the problem in those with AF, especially when AF is permanent. Because of frequent



5  Implantable Sensors for Rate Adaptation and Hemodynamic Monitoring 1.00

Survival

using a template matching system.83 Only 9 to 19 (47%) had adequate complete BiV pacing, the remaining had either fusion or pseudofusion beats. Patients with 86.4% ± 17.1% full capture had one or more NYHA class improvement than those fully captured at only 66.8% ± 19.1% of the time. This study suggests careful evaluation of the percentage of BiV pacing using multiple surface ECG leads will be needed in patients with AF and conducted response, and sensors to detect full BiV capture will be an interesting development.

SDAAM >100 ms SDAAM 50–100 ms

0.95

SDAAM <50 ms

0.90

RIGHT VENTRICULAR PRESSURE

0.85 SDAAM <50 ms versus SDAAM >100 ms: Hazard ratio = 3.2; p = 0.02

0.80 0

2

4

6

8

10

12

Months Figure 5-18  Kaplan-Meier survival curve for all-cause mortality in heart failure patients. Survival plot based on average value of standard deviation of 5-minute median atrial sensed intervals (SDAAM) for weeks 5 to 8 after implantation. There were 36 patients in SDAAM <50-msec group, 227 in 50- to 100-msec group, and 167 in >100-msec group. SDAAM <50 msec when averaged over 4 weeks was associated with increased mortality risk. (From Adamson PB, Smith AL, Abraham WT, et al: Continuous autonomic assessment in patients with symptomatic heart failure: prognostic value of heart rate variability measured by an implanted cardiac resynchronization device. InSync III Model 8042 and Attain OTW Lead Model 4193 Clinical Trial Investigators. Circulation 110:2389-2394, 2004.)

intrinsic conduction and fusion beats in AF, a high percentage of BiV pacing may still not reflect the percentage of true mechanically synchronized beats. In one study of 19 patients with permanent AF with diagnostic counters showing greater than 90% BiV pacing, 12-lead Holter was used to monitor “true” complete BiV capture 1

0.8

Sensitivity

161

Feasibility Study

SDAAM

0.6 Activity 0.4

Night heart rate

0.2

0

1

2

3

Pulmonary artery (PA) pressure and PAWP monitoring have been extensively studied in patients admitted with advanced heart failure.8486 An implantable sensor that measures pressure has been incorporated into a pacing electrode with an initial application for rate adaptive pacing.87 This sensor is a hermetically sealed piezoelectric crystal with a diaphragm facing the bloodstream. Early experiences showed the ability of this sensor to record RV pressure continuously when connected to an implanted hemodynamic monitor.88,89 The Medtronic Chronicle IHM (Model 9520) is a nonpacing implantable pulse generator capable of external radiofrequency connection and integration in a Web-based data system, with 128 kilobytes of RAM for continuous storage of sensor data. Piezoelectric activity from a passive-fixation lead in the RV outflow tract is sampled up to once every 2 seconds, timed to the sensed unipolar RV electrogram (EGM). The sensor frequency is linear up to 100 Hz, and a 60-Hz high-pass filter is used. The frequency and timing of RV systolic pressure sampling are programmable, often at early morning when the subject is likely to be supine and rested. In addition, instantaneous high-resolution recordings can be made using an external triggering device. RV pressures (systolic, diastolic, dP/dt) are recorded, and RV systolic reflects PA systolic pressure. Prior studies have assessed the ability to estimate PA diastolic pressure from the RV pressure waveform.90 The minimum PA pressure occurs at pulmonic valve opening. This is estimated at the time of the maximum positive RV dp/dt, which is simultaneous with the RV preejection period (Fig. 5-20). Preimplant calibration pressure is required to allow absolute pressure measurement. An external barometric pressure measurement device provides external pressure reference against which sensor data can be calibrated.

4

5

False positives per patient-year of monitoring (85.7 patient-years total) Figure 5-19  Comparison of detector performance for SDAAM, night heart rate, and patient activity. Plots show sensitivity to detect impending cardiovascular hospitalization versus number of falsepositive detections per patient-year of monitoring, as a function of detection threshold. (From Adamson PB, Smith AL, Abraham WT, et al: Continuous autonomic assessment in patients with symptomatic heart failure: prognostic value of heart rate variability measured by an implanted cardiac resynchronization device. InSync III Model 8042 and Attain OTW Lead Model 4193 Clinical Trial Investigators. Circulation 110:2389-2394, 2004.)

The long-term stability of the pressure sensor has been previously reported.89,90 In one study, serial Swan-Ganz catheterizations at 3, 6, and 12 months after implant showed a small baseline error at 12 months of less than 1 mm Hg from the time of implantation.91 Furthermore, the accuracy of pressure measured is not affected by body posture. As mentioned earlier, Adamson et al.63 studied 32 patients with heart failure who received a Chronicle IHM and found that long-term RV pressure was stable in most patients. In some, however, it showed significant variability despite measurements being taken at 4 am with no activity level registered. During a total of 36 volume-overloaded events, RV systolic pressure increased by 25% ± 4%, heart rate by 11% ± 2%, and estimated PA diastolic pressure by 26% ± 4%. In all events, pressures were increased 1 day before the required clinical intervention. In patients with heart failure hospitalization, increases in one of the pressures occurred in 9 of 12 events, whereas this occurred in 9 of 24 patients during a nonhospitalized episode (see Fig. 5-15). During a volume-depleted state in seven patients, RV pressure parameters were reduced. All patients returned to baseline levels after therapeutic intervention. The authors further reported that sustained increase in one of the pressures (>20% from baseline) occurred in patients subsequently admitted, at a mean of 4.2 days before admission. When the device data were available to the monitoring physician, a reduction in heart failure hospitalization was subsequently observed. This important study also documents the time sequence of RV pressure changes that occurred before heart failure exacerbation, suggesting that

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SECTION 1  Basic Principles of Device Therapy

ECG, ECG: (105–2 mV/CM)

EGM, ECG: (85–10 mV/CM)

RVP, pressure: (30–10 mm Hg/CM)

RVP, derivative: 15–500 mm Hg/CM

1 EGM 3 1 = R-R interval 2 = RVDP, at QRS detection

4

3 = RVSP, at peak waveform

2

4 = ePAD, at maximum positive dP/dt RVP

dP/dt Figure 5-20  Real-time right ventricular (RV) pressure waveforms. Pressure waveforms assessed from an implantable hemodynamic monitor (IHM) system using a programmer, with illustration of how IHM determines the pressure values. ECG, Electrocardiogram; EGM, electrogram; ePAD, estimated pulmonary artery diastolic pressure; RVDP, RV diastolic pressure; RVSP, RV systolic pressure; dP/dt, contractility as measured by change in pressure over change in time. (From Adamson PB, Magalski A, Braunschweig F, et al: Ongoing right ventricular hemodynamics in heart failure: clinical value of measurements derived from an implantable monitoring system. J Am Coll Cardiol 41:565-571, 2003.)

pressures increase for several days and reduce the patient’s “tolerable reserve” before the final increase in pressures leading to clinical heart failure exacerbation. Zile et al.92 compared the ongoing RV hemodynamics between patients with systolic and diastolic heart failure during heart failure events. RV diastolic pressure was elevated in both conditions, although there was a trend for more rapid RV diastolic pressure elevation to occur in diastolic heart failure with less compliant ventricles than during systolic heart failure. Clinical Outcome Studies The COMPASS-HF (Chronicle Offers Management to Patients with Advanced Signs and Symptoms of Heart Failure) study prospectively randomized 274 NYHA Class III or IV heart failure patients into a conventional care group or a group with therapy also guided by Chronicle-derived RV pressure parameters.93 The two safety endpoints documented no pressure sensor failure and an 8% implanted system complication. During 6-month follow-up, there was a statistically insignificant trend for reduction of either heart failure hospitalization or the need of intravenous diuretics (primary endpoint) by 21%. Later

analysis showed a 36% prolongation in time to the final heart failurerelated hospitalization in the Chronicle-guided treatment group (Fig. 5-21). Bourge et al.93 suggest that the lack of significance in the primary endpoint may be caused by the lower-than-expected number of heart failure events (predicted 1.2, actual 0.85 per 6 patientmonths), thus decreasing the power of the study. The lower event rate may result from enrollment in specialized centers with high compliance for heart failure treatment. Importantly, Chronicle-guided therapy resulted in 54% more adjustment in diuretic doses, without an increase risk in hypovolemic heart failure events. The study also included 70 of 274 patients with diastolic heart failure. A subgroup analysis found a 20% nonsignificant reduction in heart failure events when diastolic heart failure patients were managed with Chronicleguided information.94 A large outcome study involving 850 patients with Class I to III patients with a single-lead ICD integrated with IHM using weekly transmission will examine the effect on heart failure events.94a The Chronicle ICD (Model 7286, Medtronic, Inc) uses a single RV apical shock coil and a pressure sensor electrode in the RV outflow track for pressure measurement from which PA diastolic pressure is estimated.

Freedom from heart failure hospitalization



5  Implantable Sensors for Rate Adaptation and Hemodynamic Monitoring 1.0 0.8 0.6 HR = 0.64 (0.42–0.96), p = 0.03

0.4 0.2

Chronicle (37 patients with event) Control (57 patients with event)

0.0 0

50

100

150

200

Days from randomization Number at risk Chronicle Control

124 132

120 119

108 110

101 91

93 87

89 80

84 77

4 3

Figure 5-21  Time to first heart failure hospitalization in COMPASSHF Study. Results showed a 33% improvement in the time to the first heart failure events. Hazard ratio (HR) from Cox proportional hazards regression model. (From Bourge RC, Abraham WT, Adamson PB, et al: Randomized controlled trial of an implantable continuous hemodynamic monitor in patients with advanced heart failure: the COMPASS-HF study. J Am Coll Cardiol 51:1073-1079, 2008.)

OTHER PRESSURE SENSORS Left Atrial Sensor The PA diastolic pressure is an indirect assessment of the LV filling pressure, and measurement of left atrial (LA) pressure may reflect an earlier pressure change that triggers pulmonary congestion. This may allow a longer time window for physician intervention to avert ADHF. The HeartPOD LA pressure monitoring device (St. Jude Medical) comprises an implantable sensor lead that is attached to a coil antenna for telemetry of sensor signals (Fig. 5-22). The sensor has a titanium pressure-sensing membrane of 3 × 7 mm and is capable of highfidelity pressure, temperature, and EGM measurement. An external patient advisory module (PAM) sends a 125-kHz radiofrequency transmission to the antenna, which then captures 20 seconds of sensor data. The PAM has 13 megabytes of memory and can store about 3 months’ data of six recordings per day. In a single-center feasibility study, eight patients received the LA pressure device using a transseptal approach from the femoral vein.95

Venous entry at 1 cm above the inguinal ligament was achieved by conventional guidewire introduction in the femoral vein below the ligament as a landmark. After standard transseptal puncture, an 11-French sheath was placed in the LA, and the sensor was deployed until the distal set of anchors was on the left side of the interatrial septum. The sheath was then removed, and the sensor membrane was positioned 1 to 2 mm on the LA septum. All patients received dualantiplatelet therapy for 6 months (aspirin, 160 mg, and clopidogrel, 75 mg, daily). The device was calibrated over time using Valsalva maneuver, with the PAM-measured expiratory pressure to equal LA pressure. Over 6 months, the net drift was 0.2 ±1.9 mm Hg per month, although much more in one patient. The authors emphasized the need to maintain the sensor orthogonal to the septum to avoid distortion of LA pressure waveform. The Hemodynamically Guided Home Self-Therapy in Severe Heart Failure Patient (HOMEOSTASIS) trial enrolled 40 patients with NYHA Class III to IV heart failure.96 All patients had the LA pressure device implanted successfully, although sensor failure occurred in four, which was subsequently replaced in three patients. The study started after an initial 3-month period in which heart failure hospitalization, related events, and medications were documented. LA pressure data were then used to adjust the diuretic dosages. Survival without heart failure events occurred in 61% of patients at 3 years. Mean daily pressure therapy reduced LA pressure (from 17.6 to 14.8 mm Hg), frequency of high-pressure events (>25 mm Hg by 67%), and diuretic dosage, with better NYHA class, LVEF, and more frequent uptitration of angiotensinconverting enzyme (ACE) inhibitors. The role of LA pressure monitoring needs to be confirmed in prospective randomized trials. At present, issues remain about sensor stability, ease of implant (subclavian implant tools are now available), and risk of thromboembolism (which is probably low and comparable to occluder for closing atrial septal defect). Pulmonary Arterial Pressure Sensor Diastolic PA pressure is often used as a surrogate measure of PAWP. An implantable wireless PA pressure transducer (W-IHM) has been developed (Cardio-MENS, Atlanta, GA, USA), which is capable of transmitting continuous PA pressure readings when powered by radiofrequency signal from outside (Fig. 5-23, A). In one study, the device has been shown to correlate with direct PA pressure measured invasively and with echocardiographic Doppler data.97 The cardioMENS Heart Sensor Allows Monitoring of Pressure to Improve Outcomes in NYHA Class III Heart Failure Patients (CHAMPION)97a,97b trial investigated the use of such a sensor in 550 patients in the U.S. Compared to 280 patients randomized to have the sensor information blinded and treated in the conventional manner, daily measurement of PA pressure with W-IHM in the 270-patient active group resulted in significant 39% reduction of heart failure hospitalizations (Fig. 5-23, B). There was no device failure,

Figure 5-22  Left atrial (LA) pressure sensing device. A, Patient advisory module. B, LA pressure sensing device. C, Close-up image of the sensor tip. (From Ritzema J, Melton IC, Richards AM, et al: Direct left atrial pressure monitoring in ambulatory heart failure patients: initial experience with a new permanent implantable device. Circulation 116:2952-2959, 2007.)

A

163

B

C

164

SECTION 1  Basic Principles of Device Therapy

Control group (253 hospital admissions for heart failure) Treatment group (153 hospital admissions for heart failure)

260 240 220 200 180 160 140 120 100 80 60 40 20 0

Hazard ratio 0.64 (95% CI 0.55–0.75) p <0.0001

0

B

90

90 80 70 60 50 40

Number at risk Control group 280 267 252 215 179 138 105 67 Treatment group 270 262 244 210 169 131 108 82

Hazard ratio 0.71 (95% CI 0.55–0.92) p = 0.0086

30 20 10 0

180 270 360 450 540 630

Time from implant (days)

Control group (138 patients with event) Treatment group (107 patients with event)

100 Freedom from hospital admissions or mortality

Cumulative hospital admissions

A

0

C

90

180 270 360 450 540 630

Time from implant (days) 280 223 186 146 113 80 270 226 202 169 130 104

and system-related complications occurred in only eight patients, with a 1% implant related complication. All patients were either continued on warfarin or aspirin after 1 month of dual antiplatelet therapy. No episodes of pulmonary infarction or embolism were reported related to the sensor. The CHAMPION study is the first large study to confirm the role of implantable pressure sensor data to reduce heart failure morbidity. Unlike telemonitoring that involves only clinical and non-invasive heart failure parameters, the invasively derived PA pressure allows the physician to adjust neurohormonal, diuretic, or vasodilator drugs, with an aim to reduce PA pressure. In addition to the primary end point of hospitalization rate, there are also encouraging secondary outcomes, such as reduction in mortality (Fig. 5-23, C), improvement in QOL, and shorter duration of heart failure–related hospitalizations. There were more changes in medications, and the benefit occurred in both the systolic and non-systolic heart failure groups. Other implanted PA pressure devices are also under development. For example, an external ultrasound-activated PA pressure sensor in the right pulmonary artery that is fixed with a self-expanding wire mesh has been tested in 31 patients with chronic heart failure (PAPIRUS study) (InPressure, Remon Medical Technologies, Boston Scientific). Pressure values are reported to be similar to Millar catheter–measured data with no short term acute issues.97c

57 84

39 62

Figure 5-23  Wireless implantable pulmonary arter (PA) pressure sensor. A, CardioMENS sensor for implanting in a distal branch of PA artery. B, Cumulative heart failure– related hospitalizations during entire period of randomized single-blind follow-up. C, Freedom from first heart failure–related hospitalization or mortality during the entire period of randomized follow-up. (From Abraham WT, Adamson PB, Bourge RC, et al: Wireless pulmonary artery haemodynamic monitoring in chronic heart failure: a randomized controlled trial. Lancet 377:658-666, 2011.)

every 20 minutes is averaged to establish a baseline impedance level. The algorithm is inactive for the first 34 days after device implantation to allow time for postimplant pocket healing and electrode stabilization. When pulmonary fluid has accumulated beyond a programmable threshold, an alarm alerts the patient (or the physician via remote monitoring) (Fig. 5-25).

INTRATHORACIC IMPEDANCE FOR PULMONARY FLUID STATUS Device Description The concept of impedance to monitor pulmonary congestion is based on a canine experiment.98 The Medtronic OptiVol fluid management system uses transthoracic impedance to measure pulmonary fluid (Fig. 5-24). Low-voltage, nonstimulating currents are injected between RV ICD electrode (e.g., InSyn Sentry) or RV bipolar pacing electrode (e.g., Advisa) to the ICD or pacemaker casing between noon and 5 pm, at a time when the subject is presumably ambulant and upright. Sampling

Figure 5-24  Intrathoracic impedance measurement by implantable system. (From Yu CM, Wang L, Chau E, et al: Intrathoracic impedance monitoring in patients with heart failure: correlation with fluid status and feasibility of early warning preceding hospitalization. Circulation 112:841-848, 2005.)



5  Implantable Sensors for Rate Adaptation and Hemodynamic Monitoring

165

INITIAL INTERROGATION: CARDIAC COMPASS TRENDS >200

P P

P

P

I

160 OptiVol fluid index OptiVol threshold

120 80 40 Fluid

0 Sep 2008

Nov 2008

Jan 2009

BNP Normal

Mar 2009

May 2009

Jul 2009

Sep 2009

Raised BNP

Figure 5-25.  OptiVol level changes in patient with recurrent heart failure. At the nominal level of 60 ohms per day (Ω/day, Fluid), the algorithm detected an episode of impedance fall which was not corroborated with brain natriuretic peptide (BNP) level. A higher OptiVol threshold accurately detected all heart failure episodes.

Feasibility Studies Fluid index (W • days)

120

60

Impedance (W)

0 90 80 70 Reference impedance 60 0

40

A

80

120

160

200

Days after implant

Impedance (W)

90 Reference baseline

80 70

Impedance reduction

60 Duration of impedance reduction

–28

B

–21

–14

–7

0

Days before hospitalization

Figure 5-26  Intrathoracic impedance to monitor fluid status in patients with NYHA Class III or IV heart failure (MID-Heft Study). A, Operation of algorithm for detecting decreases in impedance over time. Differences between measured impedance (bottom) and reference impedance (solid line) are accumulated over time to produce fluid index (top). Threshold is then applied to fluid index to detect sustained decreases in impedance. B, Example of impedance reduction before heart failure hospitalization (arrow) for fluid overload and impedance increase during intensive diuresis during hospitalization. Label indicates reference baseline (initial reference impedance value when daily impedance value consistently falls below reference impedance line before admission). Magnitude and duration of impedance reduction are also shown. Days in hospital are shaded. (From Yu CM, Wang L, Chau E, et al: Intrathoracic impedance monitoring in patients with heart failure: correlation with fluid status and feasibility of early warning preceding hospitalization. Circulation 112:841-848, 2005.)

The feasibility of using intrathoracic impedance to monitor fluid status has been tested in 34 patients with NYHA Class III or IV heart failure in the MID-Heft Study.64 A defibrillation lead in the RV apex was connected to a special pacemaker capable of injecting and sourcing impedance. Three impedance current vectors were tested: RV lead ring to device case and sampled from coil to device case; RV coil to device case and sampled from RV coil to device case; and RV ring electrode to device case and sampled from RV tip electrode to device case. The electrode pairs that use RV coil to device case are most suitable. An initial decrease in impedance, followed by an increase to steady state, is observed during pocket healing, which peaks by about 4 weeks. During a follow-up period of 20 ± 8.4 months, 25 adjudicated heart failure hospitalizations occurred in 10 patients (Fig. 5-26). Intrathoracic impedance started to decrease before worsening symptoms at 15.3 ± 10.6 days, whereas dyspnea was reported, at the earliest, only 3 days before hospitalization. During 17 heart failure hospitalizations, PAWP was found to correlate significantly with impedance (r = −0.61, P < .001), and impedance increased as fluid removal occurred with diuresis (Fig. 5-27). An algorithm for detection of heart failure was derived from 6.2 patient-years of monitoring from 10 patients and tested in the remaining cohort (see Fig. 5-27). Depending on the set threshold for impedance crossing, a receiver-operator curve was constructed (Fig. 5-28). Lower detection thresholds lead to higher sensitivity to detect ADHF, but also to a higher false-positive rate of detection. Conversely, a higher threshold will be more specific but may underdetect some ADHF events. A level of 60 Ω/day was suggested, to give a sensitivity of 76.9% at the cost of 1.5 false-positives per patient-year of monitoring. Early warning occurred at 13.4 ± 6.2 days of heart failure hospitalization event. This landmark study shows that intrathoracic impedance is correlated with pulmonary congestion and changes days before ADHF. Clinical Outcome Studies The impact of automatic OptiVol alert to guide therapy was observed in 532 patients with heart failure in the Italian OptiVol Study.99 With the OptiVol alert “on,” 67% of heart failure patients were detected who required hospitalization or therapy adjustment. Interestingly, in the patients with OptiVol alert “off,” heart failure hospitalization was significantly more frequent (20% vs. 7%). The study also showed a falsepositive detection rate of 0.5 per patient-year of follow up. Although impedance monitoring from RV coil to device case detects lung edema, measurement of RV and LV lead vectors may be able to assess LV volumes better. In 170 patients, 43 with replacement devices, Maines et al.100 showed that a 12-week maturation period was required

PCWP (mm Hg) Impedance (W)

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SECTION 1  Basic Principles of Device Therapy

Figure 5-27  Data form heart failure patients (MID-Heft Study). A, Example from one patient. Relationship between intrathoracic impedance, pulmonary capillary wedge pressure (PCWP), and net fluid loss (I/O) during 4 days of intensive diuresis in cardiac care unit (CCU). B, Pooled data from all patients. Relationship between intrathoracic impedance and PCWP during intense diuresis in CCU pooled from 14 heart failure hospitalizations in five patients. C, Relationship between intrathoracic impedance measurement at admission to CCU and overall amount of fluid loss that occurred during CCU stay from 17 heart failure hospitalizations in six patients. (From Yu CM, Wang L, Chau E, et al: Intrathoracic impedance monitoring in patients with heart failure: correlation with fluid status and feasibility of early warning preceding hospitalization. Circulation 112:841-848, 2005.)

70 65 60 55 30 20 10

Fluid I/O (L)

0 –2 –4 –6 0

1

A

2

3

4

Days in cardiac care unit 100 r = –61, p <.001

Impedance (W)

90 80 70 60

for the pocket and the leads in first-time implants. However, a similar adaptation, but stabilizing earlier, was still required for the replacement group, suggesting the adaptation partly results from the leads or algorithm itself. Using the RV-LV impedance vectors, this study suggests that significant differences exist in impedance in the volume responder versus nonresponder groups, with a higher BiV-measured impedance in the responders, indicating a role for impedance in tracking responders to CRT. The ability of OptiVol to reflect heart failure outcome has been examined in several cohort studies. In a multicenter studies of 558 heart failures with InSync Sentry from 34 Italian centers, devicerecorded OptiVol fluid index above 60 Ω/day was associated with a 36% increased risk of heart failure hospitalization over 326 ± 216 days of follow-up.101 Multivariate analysis showed that in addition to OptiVol level, a higher percentage of days with low activity, low HRV, and increased nighttime heart rate were independent predictions of hospitalization. In 62 patients with OptiVol algorithm, prospective observational measurements of NT-pro-BNP and clinical heart failure status were collected over 27 ± 2 weeks.102 There was a significant but weak relationship in all pooled change in impedance and change in BNP level over all the visits (r = 0.30; P < .001). However, BNP level increased significantly when an OptiVol alert was associated with

50 100 0

20

30

40

50

60 80

PCWP in cardiac care unit (mm Hg) 100 r = 86, p <.001

90

60

40

80 20 70 0 0

60

1

2

3

4

5

False positives per patient-year 50 –10

C

10

Sensitivity (%)

Impedance upon entry to cardiac care unit (W)

B

–8

–6

–4

–2

Total fluid I/O in cardiac care unit (L)

0

Figure 5-28  Algorithm performance on validation data set. Detector performance curve shows trade-off between sensitivity and falsepositive rate as threshold for detection changes. For validation data set, nominal threshold of 60 Ω/day resulted in sensitivity of 76.9% and falsepositive rate of 1.5 false-positive per patient-year of monitoring, as highlighted by circle. (From Yu CM, Wang L, Chau E, et al: Intrathoracic impedance monitoring in patients with heart failure: correlation with fluid status and feasibility of early warning preceding hospitalization. Circulation 112:841-848, 2005.)

5  Implantable Sensors for Rate Adaptation and Hemodynamic Monitoring 70 60 50 40 30 20 10 0

n = 32

n = 12

n=8

< – 10%

B

>10%

NT-proBNP change (%) p <0.05

15 10 5 0 –5 –10 –15

< – 10%

0 + 10%

>10%

NT-proBNP change (%)

Figure 5-29  Increased brain natriuretic peptide (NT-proBNP) concentration and decreasing impedance suggest worsening heart failure status. A, Algorithm alerts and NT-proBNP changes in subcategories. Percentage of alerts associated with a decrease in NT-proBNP concentration (<−10%; n = 8), number of alerts with almost stable NT-proBNP concentration (0 ± 10%; n = 12), and number of alerts with an increase in NT-proBNP concentration (>10%; n = 32). B, Intrathoracic impedance change from the alert onset to patient evaluation in the three subgroups in A with a decreasing, stable, or increasing NT-proBNP concentration compared with the preceding visit. Upper and lower margins of the box represent the upper and lower quartiles; the line within, the median; and the whiskers, the data spread. (From Lüthje L, Vollmann D, Drescher T, et al: Intrathoracic impedance monitoring to detect chronic heart failure deterioration: relationship to changes in NT-proBNP. Eur J Heart Fail 9:716-722, 2007.)

clinical signs of ADHF (89 ±25% increase) versus an insignificant increase in those without signs (25% increase). In patients with an alert, BNP increased by more than 10% in most events. By the time of patient medical visit, impedance continued to fall, suggesting worsening heart failure status (Fig. 5-29). In 42% of cases, a device alert was not related to clinical signs of heart failure, although BNP overall increased by 25% in these patients. Furthermore, the number of days and the duration of time that OptiVol fluid level exceeds threshold are reported to have positive prediction of the subsequent frequency of ADHF102a (Fig. 5-30). The usefulness of OptiVol to prevent heart failure admissions was reported in a single-center case-control study of 27 patients with InSync Sentry compared with a clinically similar group of 27 patients with CRT-D without OptiVol.103 In patients with OptiVol monitoring, 12 of 27 had alerts resulting in intervention. Hospitalization occurred in only 1 of 27 over 1-year follow-up. In contrast, in the control patients, 7 of 27 had heart failure hospitalizations. When coupled with remote patient monitoring (Medtronic Carelink), 20 of 28 OptiVol alerts (71%) in 67 patients could be remotely managed.104 In one study, crossing the OptiVol threshold (>60 Ω/day) was related to a higher risk of atrial fibrillation.105 Atrial tachycardia occurred in 43% of patients before and in 29% after OptiVol level crossing. This relationship was not observed in another study, which did show a higher prevalence of ventricular tachycardia and ventricular fibrillation (VT/VF) at a lower level of OptiVol (15-45 Ω/day), suggesting VT/VF occurred at a time of volume overload.106

ADHF hospitalization rate (yr–1)

Impedance change (Ω) (alert onset → alert evaluation)

A

0 + 10%

167

The threshold for OptiVol alert was also tested in 115 patients implanted with the Medtronic InSync Sentry CRT-D.107 During follow-up of 9 ± 5 months, 45 OptiVol alerts occurred in 30 patients; 15 alerts (33%) were correlated with clinical signs and symptoms of heart failure, and the authors suggested an increase of threshold to 90 Ω/day may increase the specificity to 73%. They found no cause for falsepositive alerts, but heart failure patients had significantly higher OptiVol level versus those without. On the other hand, in the European InSync Sentry Observational Study on 373 subjects,108 the level of 60 Ω/day was associated with a 60% sensitivity and 60% positive prediction of ADHF. Nine of 53 events (17%) were not associated with an OptiVol alert, and in an additional 11 events, an increase in OptiVol level occurred but did not exceed the programmed detection threshold. When compared to daily body weight measurement, OptiVol crossing has a higher sensitivity (76% vs 23%), a lower rate of unexplained detection (1.9 vs 4.3/ patient-year), suggesting that impedance is a useful and better marker to weight change for heart failure monitoring.108a These studies confirm the usefulness of intrathoracic impedance to monitor ADHF but emphasize the need for fine-tuning the detection algorithm and individual programming of OptiVol detection level.

p = 0.001

1.4

0.7

0.0 0

>3

p = 0.005

1.0

0.5

0.0 0

B

1–3

Intrathoracic impedance fluid index threshold crossings (yr–1)

A ADHF hospitalization rate (yr–1)

Alerts (%)



1–3

>3

Days with intrathoracic impedance fluid index above threshold (yr–1)

Figure 5-30  Comparison (Poisson regression) of acute decompensated heart failure (ADHF) hospitalization rates between groups of subjects. A, Subjects with no intrathoracic impedance fluid index threshold crossing events (TCEs; n = 197), subjects with one to three TCEs (n = 79), and subjects with more than three TCEs (n = 28). B, Subjects with 0 (n = 1910), subjects with 1 to 30 (n = 31), and subjects with more than 30 (n = 82) intrathoracic impedance fluid index threshold crossing days per year. Error bars indicate 95% confidence intervals. (From Small RS, Wickemeyer W, Germany R, et al: Changes in intrathoracic impedance are associated with subsequent risk of hospitalizations for acute decompensated heart failure: clinical utility of implanted device monitoring without a patient alert. J Card Fail 15:475-481, 2009.)

168

SECTION 1  Basic Principles of Device Therapy

Advantages and Limitations

TABLE

Intrathoracic impedance is one of the most extensively used sensors for monitoring heart failure. Its advantages include the relative ease of instrumentation, requiring no additional leads or complex implantation. The battery energy expenditure is low. Intrathoracic impedance has been relatively well characterized in acute setting and for long-term monitoring. Although sensitive, its specificity may be limited because impedance in the vector used may be affected by clinical events that do not indicate pulmonary congestion, such as pleural effusion109 (Table 5-8). INTRACARDIAC IMPEDANCE Impedance signals derived from fully intracardiac electrodes better reflect cardiac volume changes than transthoracic impedance. Indeed, Salo et al.110 reported the use of a tripolar RV lead to measure RV volume changes during the cardiac cycle, from which RV volume and contractility can be derived for rate-adaptive pacing. With the addition of an LV lead in CRT, more accurate measurement of LV volume is now possible for heart failure monitoring. Unipolar Impedance from Right Ventricle Unipolar impedance from the RV apex to the CIED casing samples a small region in the cardiac apex. This results in a signal recorded by the closed-loop stimulation (CLS) sensor (see earlier). Because most current is lost over a distance of about 1 cm from the apex, the signal reflects regional contractility of the ventricle rather than a change in stroke volume. While during acute induction of ventricular fibrillation, the fall in unipolar RV impedance reflects the fall in arterial pressure, the sensitivity is insufficient to discriminate between hemodynamically stable ventricular tachycardia and supraventricular tachycardia.111 These results suggest that unipolar impedance reflects LV contractility only when the changes are significant, and it may not be able to detect small changes in cardiac contractility, as in heart failure monitoring. Multipolar Impedance Several groups and manufacturers have investigated the optimal electrode arrangement for detecting ventricular volumes. With currents flowing between intracardiac electrodes (RV, LV, and RA) and to the CIED casing, enlargement in ventricular volumes will decrease impedance as more of the heart is encompassed. Examining four intrathoracic and two intracardiac vectors in 16 dogs and five sheep, Panescu et al.112 monitored cardiac function with biweekly cardiac catheterization and echocardiography and LA pressure with an implantable

5-8 

Mechanism Blood viscosity Extrapulmonary changes Right-sided heart failure Pulmonary changes

35

–30

LV-Can LAP LV-RV

–25

LV-RA

–20

RV-Can RVcolCan

30 25

–15 –10

20 15 10

–5

RA-Can

0 –7 –6

3 4 5 6 7 8 9 10 11 12 13 14 15 16 Days

5

LAP (mm Hg)

Start of RV pacing

–35

Examples Anemia Pneumothoracic Pleural effusion Peripheral edema not detected Pneumonia

pressure sensor. After several weeks of high-rate pacing to induce heart failure, LVEF decreased significantly (52% to 34%), with increases in LV end-diastolic volume (65 to 97 mL), LV end-diastolic pressure (7 to 16 mm Hg), and LA pressure (7 to 26 mm Hg). Impedance value measured by all vectors decreased with the onset of heart failure (Fig. 5-31). The maximum decrease occurred with LV-can and LV-RV vectors. Importantly, LV-can impedance changes more with heart failure than vectors involving the right-sided heart electrodes (RA-can, RV-can, and RV coil-can), and RV-LV and LV-RA changes were intermediate. LA pressure correlated best with LV-can impedance (r2 = 0.73) than RV-can (r2 = 0.43) and RV coil-can (r2 = 0.52). Circadian variation in impedance also decreased in heart failure (5% ± 2% to 2% ± 1%). Thus, in these animal models, incorporation of an LV vector significantly improved the detection of LV volume increase in heart failure. Biventricular impedance has been measured using a quadripolar electrode arrangement.113,114 In nine mini-pigs with pacing-induced heart failure, biphasic pulses (15 µsec pulse width, 600 µA constant current amplitude) were injected between the RV ring and tip electrode, and impedance was sourced using the LV ring and tip electrode.114 The impedance signal (measured as voltage divided by 600 µA current) was recorded using 8-bite resolution, and the mean impedance was calculated over the entire cardiac cycle. “Stroke impedance” was calculated by the difference between impedance values during systole and diastole. “Systolic impedance” was defined as the highest impedance 50 to 500 msec after the R wave, whereas diastolic impedance was measured by a 20-msec window within the R wave. After 20 days of heart failure induction by rapid pacing in these animals, the increase in LV end-diastolic pressure was found to be significantly correlated with the end-diastolic impedance, which decreased by 30% (r = −0.81; P < .001). End-diastolic volume also trended in the same direction as the impedance value, which decreased by 20%. The corresponding measured intrathoracic impedance decreased by 8%, which had a poorer correlation with the end-diastolic pressure. The less striking change of intrathoracic impedance versus BiV-measured impedance might be attributed to the countering effect of lead/device

–40

Change in impedance (%)

Possible Causes of Increased OptiVol Fluid Index in Absence of Pulmonary Congestion

Figure 5-31  Impedance value measured by all vectors. Animal model shows left ventricular (LV) vector significantly improves detection of LV volume increase in heart failure. The changes in intracardiac impedance trend inversely with left atrial pressure (LAP). (From Panescu D, Naware M, Siou J, et al: Usefulness of monitoring congestive heart failure by multiple impedance vectors. Proc IEEE Eng Med Biol Soc 2008:5668-5670, 2008.)



5  Implantable Sensors for Rate Adaptation and Hemodynamic Monitoring

TABLE

5-9 

Intrathoracic vs. Intracardiac (Biventricular) Impedance

Norm (ITZ)

1.05 Heart failure parameters Electrode arrangement

1

Lead/casing maturation Influence of lung disease Influence of lead location Circadian and postural effect Sensitivity and specificity

0.95

00:00 AM 04:00 AM 08:00 AM 12:00 PM 04:00 PM 08:00 PM Time of day Figure 5-32  Variation in intrathoracic impedance depends on time of day. In this animal model (pig #12), circles depict the 4-hour-averaged impedances as recorded by the pacemaker, normalized to the daily mean. Squares depict the means of summarized 4-hour averages. The line estimates diurnal variation by spline interpolation of the means. (From Stahl C, Beierlein W, Walker T, et al: Intracardiac impedance monitors hemodynamic deterioration in a chronic heart failure pig model. J Cardiovasc Electrophysiol 18:985-990, 2007.)

casing maturation, which did not occur with BiV impedance, and pulmonary fluid collection being relatively less in the porcine model of mild heart failure studied (Fig. 5-32). These animal experiments suggest that biventricular impedance can be used to monitor the LV size and pressure changes that occur with heart failure. The theoretical advantages over transthoracic impedance are no significant time lag for lead/pocket maturation and no influence of changes in pulmonary condition on impedance values (e.g., pleural effusion, pneumonia). Because LV end-diastolic pressure increases before pulmonary edema occurs, this sensor can be used to detect deterioration of early heart failure that has not resulted in significant pulmonary fluid accumulation, and to monitor LV function. The limitations include the need for an LV lead (which restricts its use to a CRT device), dependence on the relative position of RV-LV leads (only relative changes rather than absolute value can be detected), and significant diurnal (and possibly postural) changes that must be considered in an implantable system (see Fig. 5-32). Clinical Studies

AortaP (mm Hg) Z (ohm)

Figure 5-33  Biventricular (BiV) impedance in study of 14 heart failure patients. Example of raw data recording: ECG, surface electrocardiogram; AortaP, aortic blood pressure; Z, intracardiac impedance; vertical lines 100 msec in front of QRS complex. The respiratory influence is clearly visible in the pressure and impedance tracings. (From Bocchiardo M, Meyer zu Vilsendorf D, Militello C, et al: Intracardiac impedance monitors stroke volume in resynchronization therapy patients. Europace 12:702-707, 2010.)

ECG (mV)

An acute study of BiV impedance in 14 heart failure patients during implantation of a CRT device also tested the effect of different LV lead locations on BiV impedance measurements; changes in stroke

169

Applicability Clinical evaluations

Intrathoracic Impedance Pulmonary edema RV lead or coil to casing (tripolar) Takes about 1 month Yes Less Yes About 70% (depends on threshold) Pacemakers and ICD Relatively extensive

Biventricular Impedance LV volume RV-LV bipoles (quadripolar) Less Less Likely significant Yes N/A CRT-P or CRT-D Limited

CRT, cardiac resynchronization therapy; D, defibrillation ; ICD, implantable cardioverterdefibrillator; LV, Left ventricular; N/A, not available; P, pacing; RV, right ventricular.

volume were induced with overdrive pacing115 (Fig. 5-33). Data from 20 overdrive pacing episodes and six lead locations showed good correlation between measured stroke impedance and stroke volume (r = 0.82 ± 0.10) and pulse pressure (r = 0.81 ± 0.16). The authors reported no significant effect of LV lead positions on the efficacy of impedance measurement, although accuracy and signal size tended to be better in the midventricular region than in either the basal or the apical region. Lack of lead fixation was suggested for the one outlier in the study. At present, limited data are available on long-term outcome of these devices. Table 5-9 summarizes the relative merits and limitations of intrathoracic and biventricular (intracardiac) impedance. In patients with a suitable device (i.e., LV lead), combined transthoracic and BiV impedance likely can be used. MINUTE VENTILATION Because heart failure leads to compensatory hyperventilation (especially in the resting state), an expert system has been tested that combines activity and MV to predict heart failure.116 The algorithm includes (1) mean daily resting MV and MV during activity and (2) mean daily activity level. A stable MV and activity level will suggest stable clinical heart failure, whereas an increase in MV, especially at rest and combined with decrease in activity, suggests deteriorating heart failure. Conversely, a stable MV level with increase in activity indicates recovery from heart failure. A total of 19 patients with no heart failure history who had received a Talent DDDR (Sorin-ELA) were compared with 48 heart failure patients with a Talent CRT pacemaker. While mean activity was similar, the resting and activity MV levels were

1 0.5 0 –0.5 120 100 80 60 0.8 0.6 0.4 0.2 0 580

585

590

595 Time (s)

600

605

170

SECTION 1  Basic Principles of Device Therapy

higher in the CRT group. Overall, the expert system resulted in a sensitivity of 88%, specificity of 94.7%, positive predictive value of 71%, and negative predictive value of 98.2% for heart failure detection. ST-SEGMENT SHIFT An ST-segment deviation either heralds ischemia or myocardial injury. Myocardial ischemia requires medical therapy or revascularization, especially in symptomatic individuals. Myocardial injury, on the other hand, calls for medical emergency reperfusion. Prompt treatment of a myocardial infarction (MI) will significantly reduce mortality. Delay in patient recognition of MI-induced chest pain (or “silent” MI) contributes significantly to delayed MI presentation.117 Because long-term external ambulatory electrocardiographic recording is unlikely to be practical, electrograms have been proposed to reflect infarct and ischemia in an animal model.118 Incorporation into a CIED with patient alert and remote monitoring allows ST-segment monitoring. The Angel Med Guardian (now under St. Jude Medical) is a singlechamber device with an RV apical lead. An intracardiac electrogram (EGM) was derived from RV apex to the device casing. The device records 10 seconds of EGM data once every 90 seconds for normal sinus beat within 50 to 90 bpm. Data are amplified with a gain of 62.5 to 625 times and band-passed between 0.25 to 45 Hz, followed by A/D conversion at 200 Hz. ST-segment level is compared to the corresponding PQ-segment level, and a baseline ST level is calculated as a rolling 24-hour average, acquired hourly. The extent of the deviation is normalized by the average R-wave voltage. The “normalcy” of ST-segment deviation at different heart rates is further tested during an exercise test, because some ST shifts tend to occur at different heart rates in the absence of ischemia. A rate-adjusted spontaneous ST deviation then will be considered an ischemic event and an alert triggered. A limited number of Guardian devices have been implanted in humans.119 During angioplasty with temporary coronary artery occlusion, ST-segment deviations occurred, with a negative shift during left anterior descending artery occlusion and a positive shift in other arteries (Fig. 5-34). Ten abnormal alerts occurred in six patients, leading to coronary artery interventions. During stress testing, the EGM showed a much cleaner signal and greater shift than the corresponding surface ECG. Typically, a 40% ST depression in the EGM corresponds to −0.8

mV on the surface ECG. Because the EGM is recorded at the RV apex close to left anterior descending artery territory, ischemia in this artery leads to ST depression, whereas ischemia in other territories results in a reciprocal ST elevation. Further work is required to explore the ST-segment sensor’s ability to detect the problem artery. PEAK ENDOCARDIAL ACCELERATION The PEA sensor has been used for rate adaptation. The PEA signal measures the closure sound of the mitral valve and reflects cardiac contractility. A minimal PEA signal occurs during optimal A-V interval in DDD devices120 and reflects the optimal A-V interval in most patients. A new CRT-P (New Living CHF, Sorin) is now available to monitor heart function and to program A-V interval in CRT device. Both contractility and LV filling contribute to PEA, and an index known as “PEA area” is derived from the PEA values at different A-V and V-V intervals. The maximum PEA area will define the optimal V-V and A-V interval for the patient. In 15 patients implanted with CRT with PEA sensor, cardiac catheterization with LV dP/dt was measured with PEA area determined.35 A-V interval was scanned between 60 and 220 msec. The authors found a responder rate to CRT (defined as 10% increase in dP/dt) in 75% of patients. Concordance of PEA area versus dP/dt methods occurred in 8 of 12 patients. These data are interesting, although the role of PEA for A-V interval programming in the long term is uncertain, and the ability of the sensor to monitor LV function remains to be tested.

Combined Heart Failure Diagnostics The Program to Access and Review Trending Information and Evaluate Correlation to Symptoms in Patients with Heart Failure (PARTNERS HF) is an observational study on the use of diagnostics to predict heart failure.121 A total of 100 U.S. sites prospectively recruited 694 CRT-D patients and followed them for 11.7 ± 2 months. Table 5-10 shows the diagnostic data considered important in an algorithm to predict ADHF. A positive algorithm was defined as the occurrence of 22 events in eight variables during a 1-month period, including long duration of atrial fibrillation (AF), rapid AF rate, increased OptiVol fluid index,

Baseline LAD

*

Inflation

Baseline LCX Inflation

Baseline RCA Inflation

0

5 Time in seconds

10

Figure 5-34  Guardian device recordings. Before (baseline) and 2 to 3 minutes after (inflation) balloon inflation, which occurred during angioplasty. Data are shown from three patients who underwent balloon occlusions of the left anterior descending (LAD), left circumflex (LCX), and right coronary artery (RCA), resulting in ST-interval shifts of −42%, 34%, and 18%, respectively. In the LAD occlusion recording, the arrow points to the time at which the balloon was deflated. The negative ST shift begins to shift positive, toward the shape it demonstrated during the preocclusion baseline recording, within 2 beats. (From Hopenfeld B, John MS, Fischell DR, et al: The Guardian: an implantable system for chronic ambulatory monitoring of acute myocardial infarction. J Electrocardiol 42:481-486, 2009.)



171

5  Implantable Sensors for Rate Adaptation and Hemodynamic Monitoring

TABLE

Eight Cardiac Compass Heart Failure Device Diagnostic Parameters and Algorithms Used in PARTNERS Heart Failure Study

5-10 

Diagnostic Parameter Atrial fibrillation (AF) duration Ventricular rate during AF Fluid index (OptiVol) Patient activity Night heart rate Heart rate variability (HRV) Percentage of pacing CRT ICD shock for potentially lethal VT/VF

Algorithm AF ≥6 hours on at least 1 day in patients without persistent AF (7 consecutive days with ≥23 hours of AF) AF = 24 hours and average ventricular rate during AF ≥90 bpm on at least 1 day High fluid index on at least 1 day; thresholds included ≥60, ≥80, and ≥100 Ω/day Average patient activity <1 hour over 1 week (nonoverlapping weekly windows) Average night heart rate >85 bpm for 7 consecutive days (nonoverlapping weekly windows) HRV <60 msec every day for 1 week (minimum 5 measured days) (nonoverlapping weekly windows) Ventricular pacing ≤90% for 5 of 7 days (nonoverlapping weekly windows) ≥1 shocks during the evaluation period

Data from Whellan DJ, Ousdigian KT, Al-Khatib SM, et al; PARTNERS Study Investigators: Combined heart failure device diagnostics identify patients at higher risk of subsequent heart failure hospitalizations: results from PARTNERS HF (Program to Access and Review Trending Information and Evaluate Correlation to Symptoms in Patients with Heart Failure) study. J Am Coll Cardiol 55:1803-1810, 2010.) CRT, cardiac resynchronization therapy; HF, Heart failure; bpm, beats per minute; ICD, implantable cardioverter-defibrillator; VT/VF, ventricular tachycardia/ventricular fibrillation.

Conclusion Sensors have been introduced to optimize pacing rate in patients with chronotropic incompetence. Technology has now evolved such that sensor-driven rate-adaptive pacing is a programmable parameter of almost all CIEDs. With the increasing use of CIEDs to treat and monitor heart failure, sensors have now been added to optimize programming of CRT devices and to monitor heart failure progression.

>2 diagnostic criteria met % triggered evaluations (N = 1324)

6% P <0.0001 Hazard ratio = 5.5 (95% CI: 3.4–8.8)

5%

+ Diagnostic

4% 3% 2% 1%

– Diagnostic

% of evaluations when >2 diagnostic criteria met (N = 980)

Monthly evaluations with subsequent heart failure hospitalization (pulmonary)

low patient activity, abnormal autonomic tone, and device therapy. A very high OptiVol fluid level (>100 Ω/day) alone is considered positive diagnosis of ADHF. In PARTNERS heart failure, 90 patients had 141 adjudicated heart failure events, occurring 60 days after implantation. A positive combined diagnostic set predicts a 5.5-fold increased risk of hospitalization in the next month, even after adjusting for the clinical variables (Fig. 5-35). Figure 5-36 shows the main diagnostic parameters are OptiVol level of 60 Ω/day or greater, low activity, and heart rate variability (HRV). Additional OptiVol (≥100 Ω/day; 28% of patients) is also predictive of ADHF. Further subgroup analysis suggests that the specificity of ADHF improves with a higher fluid index and using more nonfluid-related indices at the expense of lower specificity. Interestingly, in patients with a prior heart failure history, diagnostic parameters are no longer predictive. The reason for this post-hoc analysis is uncertain. Diagnostic accuracy improves if sampling is performed every 15 days versus less often (Fig. 5-37).

10

20

30

Days after diagnostic evaluation Figure 5-35  Kaplan-Meier estimates of heart failure risk. Percentage of monthly evaluations with a subsequent heart failure hospitalization caused by signs or symptoms of pulmonary congestion. Patients with a positive combined heart failure device diagnostic algorithm had a 5.5-fold increased risk of a subsequent heart failure event within 30 days. CI, Confidence interval. (From Whellan DJ, Ousdigian KT, Al-Khatib SM, et  al; PARTNERS Study Investigators: Combined heart failure device diagnostics identify patients at higher risk of subsequent heart failure hospitalizations: results from PARTNERS Heart Failure (Program to Access and Review Trending Information and Evaluate Correlation to Symptoms in Patients with Heart Failure) study. J Am Coll Cardiol 55:1803-1810, 2010.)

29%

28%

N = 980 evaluations 80

67%

60

62% 43%

40 20

21% 7%

14%

18% 5%

0 AF

0 0

43%

OptiVol >100 met

AF

AF OptiVol Low High Low Low ICD + >60 activity night HRV pacing shock(s) heart % RVR rate Fluid Activity Autonomics Device Therapy

Figure 5-36  Combined heart failure device diagnostics triggered. Venn diagram shows that 72% of evaluations had two or more (≥2) heart failure device diagnostics triggered, with the remaining 28% triggered by OptiVol fluid index of ≥100 Ω/day. OptiVol fluid index, low activity, and low heart rate variability (HRV) were the most common reasons for triggers. AF, Atrial fibrillation; ICD, implantable cardioverter-defibrillator; RVR, rapid ventricular response. (From Whellan DJ, Ousdigian KT, Al-Khatib SM, et al; PARTNERS Study Investigators: Combined heart failure device diagnostics identify patients at higher risk of subsequent heart failure hospitalizations: results from PARTNERS Heart Failure (Program to Access and Review Trending Information and Evaluate Correlation to Symptoms in Patients with Heart Failure) study. J Am Coll Cardiol 55:1803-1810, 2010.)

172

SECTION 1  Basic Principles of Device Therapy

Algorithms

P-value 2.7

Fluid index >60

<0.0001

3.3

Fluid index >80

<0.0001

3.9

Fluid index >100

<0.0001

3.2

1 criterion (excluding fluid)

<0.0001 6.4

>2 criterion (excluding fluid)

<0.0001

5.5

Combined algorithm 0.25

0.5

1.0

A

2.0

4.0

<0.0001 8.0

16.0

Hazard ratio

Prior HF event?

P-value 0.9

Yes

0.848

5.4

No 0.25

0.5

1.0

B

2.0

<0.0001

4.0

8.0

16.0

Hazard ratio

Evaluation frequency

P-value

6.9

15 days (semi-monthly)

<0.0001

5.5

30 days (monthly)

<0.0001

3.1

90 days (quarterly) 0.25

0.5

1.0

C

2.0

4.0

<0.0001 8.0

16.0

Hazard ratio

Figure 5-37  Risk stratification for different algorithms, heart failure event history, and evaluation frequencies. A, Various algorithms. An increasingly higher fluid index threshold identifies patients at higher risk, as does combined heart failure device diagnostic algorithms.  B, History of heart failure event. In patients without a previous heart failure event, the combined heart failure device diagnostic did not identify patients at high risk of a subsequent event; however, in patients with a previous event, the combined heart failure device diagnostic algorithm identified patients at a 5.4-fold increased risk of a subsequent heart failure event within 30 days. C, Evaluation frequency. Increasing the frequency of reviewing the heart failure device diagnostics from quarterly to monthly  will substantially increase the ability to identify patients at higher risk, whereas changing from monthly to semimonthly provides a less notable increase. Reproduced with permission. (From Whellan DJ, Ousdigian KT, Al-Khatib SM, et al; PARTNERS Study Investigators: Combined heart failure device diagnostics identify patients at higher risk of subsequent heart failure hospitalizations: results from PARTNERS HF (Program to Access and Review Trending Information and Evaluate Correlation to Symptoms in Patients with Heart Failure) study. J Am Coll Cardiol 55:1803-1810, 2010.)

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Yu CM, Wang L, Chau E, et al: Intrathoracic impedance monitoring in patients with heart failure: correlation with fluid status and feasibility of early warning preceding hospitalization. Circulation 112:841-848, 2005. 65. Stevenson LW, Perloff JK: The limited reliability of physical signs for estimating hemodynamics in chronic heart failure. JAMA 261:884-888, 1989. 66. McGee SR: Physical examination of venous pressure: a critical review. Am Heart J 136:10-18, 1998. 67. Chandhry SI, Wang Y, Concato J, et al: Patterns of weight change preceding hospitalization for heart failure. Circulation 116:15491554, 2007. 68. Pfisterer M, Buser P, Rickli H, et al: BNP-guided vs. symptomguided heart failure therapy. The Trial of Intensified vs. Standard Medical Therapy in Elderly Patients with Congestive Heart Failure (TIME-CHF) randomized trial. JAMA 301:383-392, 2009. 69. Klersy C, De Silvestri A, Gabutti G, et al: A meta-analysis of remote monitoring of heart failure patients. J Am Coll Cardiol 54:1683-1694, 2009. 69a.  Inglis SC, Clark RA, McAlister FA, et al: Structured telephone support or telemonitoring programmes for patients with chronic heart failure. Cochrane Database Syst Rev CD007228, 2010. 69b.  Ferrante D, Varini S, Macchia A, et al: Long-term results after a telephone intervention in chronic heart failure: DIAL (Randomized Trial of Phone Intervention in Chronic Heart Failure) follow-up. J Am Coll Cardiol 56:372-378, 2010. 69c.  Koehler F, Winkler S, Schieber M, et al: Impact of remote telemedical management on mortality and hospitalizations in ambulatory patients with chronic heart failure: the telemedical interventional monitoring in heart failure study. Circulation 123:1873-1880, 2011. 69d.  Chaudhry SI, Mattera JA, Curtis JP, et al: Telemonitoring in patients with heart failure. N Engl J Med 363:2301-2309, 2010. 70. Zrenner B, Ndrepepa G, Müssig D, et al: The recording of monophasic action potentials with fractal-coated iridium electrodes in humans. Pacing Clin Electrophysiol 23:54-62, 2000. 71. Wilensky RL, Yudelman P, Cohen AI, et al: Serial electrocardiographic changes in idiopathic dilated cardiomyopathy confirmed at necropsy. Am J Cardiol 62:276-283, 1988. 72. Callaghan F, Vollmann W, Livingston A, et al: The ventricular depolarization gradient: effects of exercise, pacing rate, epinephrine, and intrinsic heart rate control on the right ventricular evoked response. Pacing Clin Electrophysiol 12:1115-1130, 1989. 73. Cohn J, Levine TB, Olivari MT, et al: Plasma norepinephrine as a guide to prognosis in patients with chronic congestive heart failure. N Engl J Med 311:819-823, 1984. 74. Kadhiresan VA, Pastore J, Auricchio A, et al: Pacing therapies in congestive heart failure: a novel method—the activity log index—for monitoring physical activity of patients with heart failure. PATH-CHF Study Group. Am J Cardiol 89:1435-1437, 2002. 75. Braunschweig F, Mortensen PT, Gras D, et al: Monitoring of physical activity and heart rate variability in patients with chronic heart failure using cardiac resynchronization devices. InSync III Study Investigators. Am J Cardiol 95:1104-1107, 2005. 76. Bonaduce D, Petretta M, Marciano F, et al: Independent and incremental prognostic value of heart rate variability in patients with chronic heart failure. Am Heart J 138:273-284, 1999. 77. La Rovere MT, Pinna GD, Maestri R, et al: Short-term heart rate variability strongly predicts sudden cardiac death in chronic heart failure patients. Circulation 107:565-570, 2003. 78. Fantoni C, Raffa S, Regoli F, et al: Cardiac resynchronization therapy improves heart rate profile and heart rate variability of patients with moderate to severe heart failure. J Am Coll Cardiol 46:1875-1882, 2005. 79. Carbson G, Girouard S, Schlegl M, Butter C: Three-dimensional heart rate variability diagnostic for monitoring heart failure through an implantable device. J Cardiovasc Electrophysiol 15: 506, 2004. 80. Roosevelt G, Singh JP, Mullin CM, et al: Prognostic value of heart rate variability footprint and standard deviation of average 5-minute intrinsic R-R intervals for mortality in cardiac resynchronization therapy patients. J Electrocardiol 40:336-342, 2007. 81. Singh JP, Rosenthal LS, Hranitzky PM, et al: Device diagnostics and long-term clinical outcome in patients receiving cardiac resynchronization therapy. Europace 11:1647-1653, 2009.

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82. Adamson PB, Smith AL, Abraham WT, et al: Continuous autonomic assessment in patients with symptomatic heart failure: prognostic value of heart rate variability measured by an implanted cardiac resynchronization device. InSync III Model 8042 and Attain OTW Lead Model 4193 Clinical Trial Investigators. Circulation 110:2389-2394, 2004. 82a.  Koplan BA, Kaplan AJ, Weiner S, et al: Heart failure decompensation and all-cause mortality in relation to percent biventricular pacing in patients with heart failure: is a goal of 100% biventricular pacing necessary? J Am Coll Cardiol 27:53:355-360, 2009. 83. Kamath GS, Cotiga D, Koneru JN, et al: The utility of 12-lead Holter monitoring in patients with permanent atrial fibrillation for the identification of nonresponders after cardiac resynchronization therapy. J Am Coll Cardiol 53:1050-1055, 2009. 84. Steimle AE, Stevenson LW, Chelimsky-Fallick C, et al: Sustained hemodynamic efficacy of therapy tailored to reduce filling pressures in survivors with advanced heart failure. Circulation 96:1165-1172, 1997. 85. Gibbs JS, Keegan J, Wright C, et al: Pulmonary artery pressure changes during exercise and daily activities in chronic heart failure. J Am Coll Cardiol 15:52-61, 1990. 86. Nathan AW, Perry SG, Cochrane T, et al: Ambulatory monitoring of pulmonary artery pressure: a preliminary clinical evaluation. Br Heart J 49:33-37, 1983. 87. Lau CP, Butrous GS, Ward DE, Camm AJ: Comparison of exercise performance of six rate-adaptive right ventricular cardiac pacemakers. Am J Cardiol 63:833-838, 1989. 88. Ohlsson A, Nordlander R, Bennett T, et al: Continuous ambulatory haemodynamic monitoring with an implantable system: the feasibility of a new technique. Eur Heart J 19:174-184, 1998. 89. Steinhaus DM, Lemery R, Bresnahan DR, Jr, et al: Initial experience with an implantable hemodynamic monitor. Circulation 93:745-752, 1996. 90. Reynolds DW, Bartelt N, Taepke R, Bennett TD: Measurement of pulmonary artery diastolic pressure from the right ventricle. J Am Coll Cardiol 25:1176-1182, 1995. 91. Magalski A, Adamson P, Gadler F, et al: Continuous ambulatory right heart pressure measurements with an implantable hemodynamic monitor: a multicenter, 12-month follow-up study of patients with chronic heart failure. J Card Fail 8:63-70, 2002. 92. Zile MR, Bourge RC, Bennett TD, et al: Application of implantable hemodynamic monitoring in the management of patients with diastolic heart failure: a subgroup analysis of the COMPASS-HF trial. J Card Fail 14:816-823, 2008. 93. Bourge RC, Abraham WT, Adamson PB, et al: Randomized controlled trial of an implantable continuous hemodynamic monitor in patients with advanced heart failure: the COMPASS-HF study. J Am Coll Cardiol 51:1073-1079, 2008. 94. Zile MR, Bennett TD, St John Sutton M, et al: Transition from chronic compensated to acute decompensated heart failure: pathophysiological insights obtained from continuous monitoring of intracardiac pressures. Circulation 118:1433-1441, 2008. 94a.  Adamson PB, Conti JB, Smith AL, et al: Reducing events in patients with chronic heart failure (REDUCEhf) study design: continuous hemodynamic monitoring with an implantable defibrillator. Clin Cardiol 30:567-575, 2007. 95. Ritzema J, Melton IC, Richards AM, et al: Direct left atrial pressure monitoring in ambulatory heart failure patients: initial experience with a new permanent implantable device. Circulation 116:2952-2959, 2007. 96. Ritzema J, Troughton R, Melton I, et al: Physician-directed patient self-management of left atrial pressure in advanced chronic heart failure. Hemodynamically Guided Home SelfTherapy in Severe Heart Failure Patients (HOMEOSTASIS) Study Group. Circulation 121:1086-1095, 2010. 97. Verdejo HE, Castro PF, Concepción R, et al: Comparison of a radiofrequency-based wireless pressure sensor to Swan-Ganz catheter and echocardiography for ambulatory assessment of pulmonary artery pressure in heart failure. J Am Coll Cardiol 50:2375-2382, 2007. 97a.  Adamson PB, Abraham WT, Aaron M, et al: CHAMPION trial rationale and design: the long-term safety and clinical efficacy of a wireless pulmonary artery pressure monitoring system. J Card Fail 17:3-10, 2011. 97b.  Abraham WT, Adamson PB, Bourge RC, et al: Wireless pulmonary artery haemodynamic monitoring in chronic heart failure: a randomised controlled trial. Lancet 377:658-666, 2011. 97c.  Hoppe UC, Vanderheyden M, Sievert H, et al: Chronic monitoring of pulmonary artery pressure in patients with severe heart failure: multicentre experience of the monitoring Pulmonary Artery Pressure by Implantable device Responding to Ultrasonic Signal (PAPIRUS) II study. Heart 95:1091-1097, 2009. 98. Wang L, Lahtinen S, Lentz L, et al: Feasibility of using an implantable system to measure thoracic congestion in an ambulatory chronic heart failure canine model. Pacing Clin Electrophysiol 28:404-411, 2005. 99. Catanzariti D, Lunati M, Landolina M, et al: Monitoring intrathoracic impedance with an implantable defibrillator reduces hospitalizations in patients with heart failure. Italian Clinical Service OptiVol-CRT Group. Pacing Clin Electrophysiol 32:363370, 2009. 100. Maines M, Landolina M, Lunati M, et al: Intrathoracic and ventricular impedances are associated with changes in ventricular volume in patients receiving defibrillators for CRT. Italian Clinical Service OptiVol-CRT Group. Pacing Clin Electrophysiol 33:64-73, 2010.

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101. Perego GB, Landolina M, Vergara G, et al: Implantable CRT device diagnostics identify patients with increased risk for heart failure hospitalization. OptiVol-CRT Clinical Service Observational Group. J Interv Card Electrophysiol 23:235-242, 2008. 102. Lüthje L, Vollmann D, Drescher T, et al: Intrathoracic impedance monitoring to detect chronic heart failure deterioration: relationship to changes in NT-proBNP. Eur J Heart Fail 9:716-722, 2007. 102a.  Small RS, Wickemeyer W, Germany R, et al: Changes in intrathoracic impedance are associated with subsequent risk of hospitalizations for acute decompensated heart failure: clinical utility of implanted device monitoring without a patient alert. J Card Fail 15:475-481, 2009. 103. Maines M, Catanzariti D, Cemin C, et al: Usefulness of intrathoracic fluids accumulation monitoring with an implantable biventricular defibrillator in reducing hospitalizations in patients with heart failure: a case-control study. J Interv Card Electrophysiol 19:201-207, 2007. 104. Santini M, Ricci RP, Lunati M, et al: Remote monitoring of patients with biventricular defibrillators through the CareLink system improves clinical management of arrhythmias and heart failure episodes. J Interv Card Electrophysiol 24:53-61, 2009. 105. Jhanjee R, Templeton GA, Sattiraju S, et al: Relationship of paroxysmal atrial tachyarrhythmias to volume overload: assessment by implanted transpulmonary impedance monitoring. Circ Arrhythm Electrophysiol 2:488-494, 2009. 106. Ip JE, Cheung JW, Park D, et al: Temporal associations between thoracic volume overload and malignant ventricular arrhythmias: a study of intrathoracic impedance. J Cardiovasc Electrophysiol 22:293-299, 2011. 107. Ypenburg C, Bax JJ, van der Wall EE, et al: Intrathoracic impedance monitoring to predict decompensated heart failure. Am J Cardiol 99:554-557, 2007.

108. Vollmann D, Nägele H, Schauerte P, et al: Clinical utility of intrathoracic impedance monitoring to alert patients with an implanted device of deteriorating chronic heart failure. European InSync Sentry Observational Study Investigators. Eur Heart J 28:1835-1840, 2007. 108a.  Abraham WT, Compton S, Haas G, et al: Intrathoracic impedance vs daily weight monitoring for predicting worsening heart failure events: results of the Fluid Accumulation Status Trial (FAST). Congest Heart Fail 17:51-55, 2011. 109. Timperley J, Mitchell AR, Brown P, Betts TR: Changes in intrathoracic impedance from a pneumothorax: insights from an implanted monitoring system. Pacing Clin Electrophysiol 28:1109-1111, 2005. 110. Salo RW, Pederson BD, Olive AL, et al: Continuous ventricular volume assessment for diagnosis and pacemaker control. Pacing Clin Electrophysiol 7:1267-1272, 1984. 111. Kaye G, Arthur W, Edgar D, et al: The use of unipolar intracardiac impedance for discrimination of haemodynamically stable and unstable arrhythmias in man. Europace 8:988-993. 2006. 112. Panescu D, Naware M, Siou J, et al: Usefulness of monitoring congestive heart failure by multiple impedance vectors. Proc IEEE Eng Med Biol Soc 2008:5668-5670, 2008. 113. Zima E, Lippert M, Czygan G, Merkely B: Determination of left ventricular volume changes by intracardiac conductance using a biventricular electrode configuration. Europace 8:537-544, 2006. 114. Stahl C, Beierlein W, Walker T, et al: Intracardiac impedance monitors hemodynamic deterioration in a chronic heart failure pig model. J Cardiovasc Electrophysiol 18:985-990, 2007. 115. Bocchiardo M, Meyer zu Vilsendorf D, Militello C, et al: Intracardiac impedance monitors stroke volume in resynchronization therapy patients. Europace 12:702-707, 2010.

116. Page E, Cazeau S, Ritter P, et al: Physiological approach to monitor patients in congestive heart failure: application of a new implantable device-based system to monitor daily life activity and ventilation. Europace 9:687-693, 2007. 117. Faxon D, Lenfant C: Timing is everything: motivating patients to call 9-1-1 at onset of acute myocardial infarction. Circulation 104:1210-1211, 2001. 118. Fischell TA, Fischell DR, Fischell RE, et al: Real-time detection and alerting for acute ST-segment elevation myocardial ischemia using an implantable, high-fidelity, intracardiac electrogram monitoring system with long-range telemetry in an ambulatory porcine model. J Am Coll Cardiol 48:2306-2314, 2006. 119. Hopenfeld B, John MS, Fischell DR, et al: The Guardian: an implantable system for chronic ambulatory monitoring of acute myocardial infarction. J Electrocardiol 42:481-486, 2009. 120. Leung SK, Lau CP, Lam CT, et al: Automatic optimization of resting and exercise atrioventricular interval using a peak endocardial acceleration sensor: validation with Doppler echocardiography and direct cardiac output measurements. Pacing Clin Electrophysiol 23:1762-1766, 2000. 121. Whellan DJ, Ousdigian KT, Al-Khatib SM, et al; PARTNERS Study Investigators: Combined heart failure device diagnostics identify patients at higher risk of subsequent heart failure hospitalizations: results from PARTNERS HF (Program to Access and Review Trending Information and Evaluate Correlation to Symptoms in Patients with Heart Failure) study. J Am Coll Cardiol 55:1803-1810, 2010.

6 

Power Systems for Implantable Pacemakers, Cardioverters, and Defibrillators DARREL F. UNTEREKER  |  ANN M. CRESPI  |  ANTHONY RORVICK  |  CRAIG L. SCHMIDT  |  PAUL M. SKARSTAD

This chapter discusses batteries and capacitors used to power pace-

makers, defibrillators, and similar implantable devices. Batteries are active components that convert chemical energy into electrical energy. Capacitors are passive and temporarily store energy, often to increase the available power (rate of energy delivery) in an electric circuit. This chapter provides practical information to help physicians manage patients who have implanted medical devices that require electrical power. The battery is conceptually different from the other components of an implantable medical device. In principle, the other components are designed to last indefinitely. However, for current cardiac devices such as pacemakers and defibrillators, the available chemical energy of the battery is consumed during its use. Eventually, the output of the battery becomes insufficient to operate the electronics and must be replaced. At present, batteries for implantable cardiac devices are part of the device, and the entire pulse generator must be replaced to renew the battery. However, this is not true of all implantable pulse generators. Many neurologic stimulators now use rechargeable batteries. In the future, more implantable devices may also use rechargeable batteries so that the energy powering the device can be renewed. While batteries transform chemical energy into electrical energy, capacitors act as energy storage devices. Capacitors that intermittently boost the power capability of electronic circuits are of principal interest here. The large capacitors in an implantable cardioverterdefibrillator (ICD) allow the device to deliver a therapeutic, highvoltage, high-energy shock to the heart over a few milliseconds, which the battery itself could not do. Understanding the properties and limitations of the components used to power implantable devices will help physicians not only better understand the devices, but also deliver the best care to their patients.

Batteries BASIC FUNCTION AND ELECTROCHEMISTRY Energy Storage A battery converts chemical energy into electrical energy. The source of this energy is the electrochemical reactions that occur within the battery. The type and amount of the materials participating in these reactions determine the deliverable energy of a battery. Chemical Reactions During a spontaneous chemical reaction, substances react to form reaction products. For example, in combustion, a fuel combines with oxygen (O2) from air to form water (H2O), carbon dioxide (CO2), and other species. The fuel is oxidized (loses electrons) during this process. Oxygen from air is reduced (gains electrons). The exact amounts of each that react are determined by the stoichiometry of the reaction. Equations 6-1 to 6-4 are examples of different types of reactions in which oxidation and reduction occur, called “redox” reactions, because electrons are transferred from one reactant to another.

2H2 + O2 → 2H2O (formation of water)



(6-1)

CH4 + 2O2 → CO2 + 2H2O (combustion of methane) (6-2)



4Fe + 3O2 → 2Fe2O3 (rusting of iron)

(6-3)

2Li + I2 → 2LiI (discharge of lithium/iodine pacemaker battery) (6-4) Such chemical reactions occur spontaneously because the products are in a lower energy state than the reactants. The difference in energy appears as heat in the case of combustion, or rusting. A battery or galvanic cell is designed to convert much of this energy difference into electrical energy rather than heat. A battery operates because the electrons transferred during a redox reaction are channeled from one terminal of the battery through the external circuit and back to the battery through its other terminal. The maximum work these electrons can do outside the battery is related to the difference in a thermodynamic quantity called the free energy of the reaction. MAJOR COMPONENTS OF BATTERIES Figure 6-1 shows a simple battery; the major parts are the anode, cathode, and electrolyte. The anode and cathode must be physically separated, and both must be in contact with the electrolyte. Anode and Cathode The two battery components involved in the electrochemical reaction during discharge are the anode and the cathode. The anode, lithium (Li) in most medical batteries, furnishes electrons to the external circuit, while the cathode, usually an oxide or halogen-rich compound, receives the electrons. The anode is defined as the electrode at which electrochemical oxidation occurs, and the cathode is the electrode at which electrochemical reduction occurs. Oxidation is a process in which electrons are removed or lost, and reduction is a process in which electrons are gained. This is true in all electrochemical cells. In a spontaneous galvanic cell such as a battery, the anode is electrically negative and the cathode electrically positive. This creates a seeming inconsistency. In nonspontaneous electrochemical reactions, such as those that occur at pacing electrodes, electrochemical reactions are driven by an externally applied voltage. In this case, oxidation (loss of electrons) will occur at the positive electrode and reduction (gain of electrons) at the negative electrode. Thus, for pacing electrodes, the anode is positive and the cathode is negative. The polarities are reversed, but the underlying electrochemical processes defining anode and cathode—oxidation and reduction, respectively—are the same. The terminology is entirely consistent. Electrolyte The anode and the cathode must be separated so that they cannot directly react with each other, but they also need to be connected by an ionically conducting medium called the electrolyte. As the battery discharges, it furnishes electrons at one terminal, pushes them through an external circuit, and receives them back at a second terminal. The

175

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of the battery for a test gas, usually helium, is less than 1 × 10−7 cm3/ sec at 1-atm pressure difference between the inside and outside of the battery.

e− External load

CLASSIFICATION OF BATTERIES A+ + C− e−

−

A+

C−

e−

+

AC

Batteries may be classified in many ways. For example, application, functional characteristics, chemistry, or the physical state of some component are often used to categorize batteries. One fundamental distinction is between primary and secondary batteries. Primary Batteries

Anode

Anode current collector

Electrolyte

Cathode current collector

Cathode

Figure 6-1  Schematic representation of sealed battery. The anode material is on the left, the cathode material on the right, and the electrolyte in the middle of the cell. The discharge reaction is A + C → AC. In this example, both A+ and C− are mobile ions, and the discharge product, AC, is insoluble and precipitates in the electrolyte solution.  In this figure, connections to the interior of the cell are made via feedthroughs.

electrolyte allows electric charge, as ions, to flow within the battery, completing the circuit. The electrolyte must conduct ions but not electrons. If the electrolyte conducted electrons, the battery would be shorted internally just as if the terminals were connected by a metal wire. Batteries with lithium anodes, such as those used to power pacemakers and ICDs, must use nonaqueous electrolytes because water readily reacts with lithium. Most Li-based batteries employ a mixture of organic ethers and esters as solvents for the electrolyte. A dissolved Li salt is used to make the electrolyte conductive. For example, lithium/ manganese dioxide batteries, widely used to power automatic cameras, often use electrolytes containing a lithium salt dissolved in dimethoxyethane ether and propylene carbonate ester. Separator The separator is a structural member of the battery that keeps the anode and cathode materials physically apart, thus preventing shorting of the battery. In batteries with liquid electrolytes, the separator is usually a porous polymer film that is immersed in and permeated by the electrolyte. In the case of lithium/iodine (Li/I2) batteries, traditionally used for pacemakers, the separator and electrolyte are the same (i.e., growing layer of solid, ionically conductive LiI discharge product). Current Collector The current collector makes the connection between the positive or negative terminal of the battery and its respective active electrode material inside the cell. A current collector is usually a wire connected to a screen, or grid, which is embedded in the anode or cathode material. The current collector may also serve as a structural member of the battery to provide physical integrity and strength to that electrode. Some medical batteries use the case of the cell as the current collector for one of the electrodes. Batteries are termed “case negative” if the anode is in contact with the case and “case positive” if the cathode is in contact with the case. SEALING OF BATTERIES Batteries used to power implantable pacemakers and ICDs are hermetically sealed, most often in welded containers with glass electrical feedthroughs to make electrical connections between the inside and the outside of the battery. This is necessary to prevent slow interchanges of materials between the battery and its surroundings. Medical batteries are typically considered hermetically sealed if the leak rate out

Primary batteries can only be used once. They are not designed to be recharged; in fact, it is dangerous to attempt to recharge them. A familiar example of a primary battery is the alkaline zinc/manganese dioxide flashlight battery. Most batteries used to power modern implantable medical devices are primary batteries that use lithium anodes because such batteries have very high energy densities. Secondary Batteries Secondary batteries are designed to be repetitively discharged and recharged. Familiar examples include the lead-acid batteries used in automobiles and lithium-ion batteries, widely used in portable electronic devices, home appliances, power tools, and electric vehicles. Lithium-ion batteries are also the most relevant secondary medical battery, used as supplemental power sources in some left ventricular assist devices and widely used to power implantable neurologic stimulators for control of chronic pain. Although not currently in development for pacemakers and ICDs, features such as more frequent telemetry could lead manufacturers to reconsider lithium-ion battery use in cardiac applications in the future. FUNCTIONAL CHARACTERISTICS OF BATTERIES Capacity The fundamental unit of battery capacity is the coulomb (C), or ampere-second (A-sec). This is the amount of charge delivered by one ampere (1 A) of current in 1 second. In the context of implantable devices, it is customary to express battery capacity in terms of amperehours (A-hr), which represents the charge carried by a current of 1 A flowing for 1 hour; 1 A-hr = 3600 C. A battery for an implantable medical device usually has a capacity rating of 1 2 to 2 A-hr. The methods for determining and specifying the capacity of a medical battery produce numbers ranging from upper-limit theoretical values that can never be achieved in the field to realistic values based on detailed models or accelerated testing.1-3 More sophisticated methods account for limited availability of the components within the battery, kinetic limitations on the chemical reaction, and losses to side-reactions over time. Estimating the deliverable capacity of implantable medical batteries is especially difficult because of their long service life. The time frame for the operation of most implantable medical devices is so long (5-10 years) that real-time measurements of capacity are not practical. Therefore, accelerated tests and models are typically used to estimate the amount of deliverable capacity in these batteries. Technology in this area is now highly developed, and it is possible to make highly accurate projections of deliverable battery capacity under a range of usage conditions.4-6 Energy and Energy Density The fundamental unit of energy is the joule (J). This is the energy given to one coulomb (1 C) of charge that is accelerated by a difference in potential of one volt (1 V). One joule is also the energy transferred by one watt (1 W) of power in 1 second. Just as battery capacity is often measured in ampere-hours, battery energy is often expressed in watthours (W-hr) instead of joules; 1 W-hr = 3600 J. An important battery parameter in the design of an implantable device is energy density, a quantity that can be expressed on either a

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6  Power Systems for Implantable Pacemakers, Cardioverters, and Defibrillators

mass or a volume basis. For medical applications, volume is usually more important than mass, so ratings based on volumetric energy density are most often used. The time integral of the product of voltage and current divided by the total volume of the battery is its deliverable energy density. Modern batteries for implantable devices have energy densities as high as 1 W-hr/cm3, including the case. Stoichiometry and Cell Balance The specific amount of anode and cathode materials that will react is determined by the stoichiometry of the battery reaction. A cell that contains exactly the required ratio of anode and cathode materials is said to be a “balanced” cell. However, most medical batteries are not designed with exactly the stoichiometric ratio of the active cathode and anode to provide predictable end-of-service characteristics. In many cases the anode (lithium) is in excess so that the voltage decrease near the end-of-service life is not too abrupt. The open-circuit voltage of a battery can be calculated from the thermodynamic free energy for the discharge reaction. This is the voltage that will be measured when there are no kinetic limitations, a condition that occurs only when an insignificant amount of current is being drawn from the battery. With the onset of current flow, the voltage at the battery terminals will be smaller than the open-circuit value. Both chemistry and battery design determine the relationship between voltage and current drawn from the battery. For example, a lead-acid battery for automotive use is constructed of very conductive materials and is designed with large-surface-area electrodes so that extremely high currents can be drawn from it to run an engine’s starter motor. On the other hand, a transistor radio battery is designed with small electrodes because relatively low currents are typically needed to power small portable electronic devices. Figure 6-2 shows a typical currentvoltage relationship in which the load voltage approaches the opencircuit voltage (OCV) as the current approaches zero. At the other extreme, the maximum (short-circuit) current is observed when the load voltage approaches zero. Internal Resistance and Impedance Electrical impedance and resistance are important battery properties that play a significant role in the clinical performance of many implantable devices. The terms impedance and resistance are often used interchangeably but are not the same. Both are terms for the change in voltage per unit change in current in an electric circuit, but they are measured under different conditions. Impedance is the more general term, encompassing effects of resistance, capacitance, inductance, and

OCV

Voltage

Increasing current density

2.5 2 A

1.5

B

C

D

1 0.5 0

0

0.2

0.4

0.6

0.8

1

1.2

1.4

Discharged capacity (ampere-hours)

Cell Voltage and Current

Load voltage

3

V load (V)



Polarization

Current drain

Short-circuit current

Figure 6-2  Typical load voltage versus current drain plot for a battery. The load voltage is always less than the open-circuit voltage (OCV). Current drain increases toward the right. The maximum cell voltage is obtained at zero current, and the maximum current drain occurs at zero load voltage. The exact shape of the plot depends on the chemistry and design of the battery.

Figure 6-3  Load voltage versus capacity plot. For a lithium/iodine battery discharged at four magnitudes of constant current. Curve A = 40 µA/cm2, curve B = 20 µA/cm2, curve C = 10 µA/cm2, and curve D = 2 µA/cm2. The deliverable capacity decreases with increasing current in this current range because of increasing effects of polarization.

other circuit elements on the relationship between voltage and current. The resistive component of impedance is measured using directcurrent methods. Alternating-current and transient methods are used to measure the additional components of impedance besides resistance. For simple, resistive electric circuit elements, Ohm’s law, V = IR, accurately describes a linear relationship between voltage drop, V, and the corresponding change in current, I, with resistance, R, as the proportionality constant. However, a battery is a complex electrochemical device with several nonlinear processes operating in series/parallel combinations. Different processes may dominate at different current levels, depths of discharge, and time. Consequently, the relationship between current and voltage for a battery is, in general, nonlinear, even at very low currents. Although this relationship is sometimes characterized by a quantity called RDC, Ohm’s law should only be applied to batteries with caution. Non-Ideal Battery Behavior In addition to the principles and the nomenclature of battery operation, it is also important to understand factors that limit a battery’s ability to power an implantable device. Polarization.  Polarization is any process that causes the voltage at the terminals of a battery to drop below its open-circuit value when it is providing current. The internal resistance of the battery is one important cause. This is well illustrated in Figure 6-3 for the lithium/iodine battery, but the same is true for all batteries to some degree. The differences in the curves for discharge voltage versus capacity at four rates of constant current discharge are mainly caused by the voltage drop associated with internal resistance of the battery. Other contributing elements of voltage loss when a battery provides current are concentration polarization, which is associated with concentration gradients that may develop in the electrolyte or the active electrode materials, and activation polarization, which is associated with the kinetics of the electron-transfer reactions at the electrode/electrolyte interfaces. When current is drawn from a battery, all these processes occur to some extent. The net effect of these kinetic limitations is always observed as a decrease in the voltage at the terminals of the battery. In general, neither concentration polarization nor electron-transfer polarization conforms to Ohm’s law. Self-Discharge.  Self-discharge is the spontaneous discharge of a cell or battery by an internal chemical reaction rather than through useful electrochemical discharge. We have all seen the effects of self-discharge,

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SECTION 1  Basic Principles of Device Therapy

such as the flashlight that does not work when needed after prolonged storage. One mechanism of self-discharge involves a slow, direct reaction between the anode and cathode; for example, if one or both of the active electrode materials are slightly soluble in the electrolyte. Other self-discharge processes may involve reactions between anode or cathode and another substance in the battery, such as the solvent in the electrolyte; for example, the typical reaction between the anode and the electrolyte solvent to form a passive film on the lithium anode or a gas that pressurizes the battery case. These parasitic reactions are usually very slow, but because medical batteries are expected to operate for many years, their accumulated effects can be appreciable. Although it is difficult to measure the very slow rate of self-discharge reactions, techniques such as microcalorimetry can detect the small amounts of heat involved.7,8 The heat generated can be used to calculate the rate of self-discharge by applying thermodynamic principles. Accurate assessment of self-discharge is an important element in predictive models of battery performance used in both device design and longevity estimation. BATTERIES IN IMPLANTABLE CARDIAC RHYTHM MANAGEMENT DEVICES Battery Design Requirements Implantable medical devices must be viewed as a system. The main requirement in battery selection for implantable devices is high reliability. Other factors include the desired longevity of the device (directly related to battery energy density, circuit design, and overall size) and an appropriate indication of impending battery depletion (end-of-service warning). The basic considerations when designing a battery for an implantable medical device include the current variations that will be imposed by the circuit as the device provides its service for the individual patient. Once the span of these application requirements is formulated, the current, voltage, and capacity requirements of the battery can be determined. An important limitation is the physical space, both volume and shape, within the device that can be allotted to the battery. Once this is determined, the required energy density can be calculated, and various battery systems and designs can be considered and evaluated. To put these decisions in perspective, one needs to realize that the battery occupies about a third of the volume of an ICD. The circuitry occupies another third, and the capacitors fill the remaining volume. In a pacemaker the battery and circuitry each occupy about half the volume. Power Requirements An important parameter for a device is the peak power requirement. For example, the rate of energy consumption differs greatly for pacemakers and defibrillators. Pacemakers use very small amounts of energy when they stimulate the heart, on the order of 15 µJ. Defibrillators, on the other hand, deliver as much 40 J for a defibrillation shock. A battery optimized for a pacemaker could never come close to supplying energy at the rate required to power a defibrillator. Likewise, a defibrillator battery is not an optimum choice to power a pacemaker, although it could easily supply the current needed. The high-power design of a defibrillator battery has a significantly lower energy density than that of a pacemaker battery, by a factor of as much as two. Thus, if a defibrillator battery was used primarily for pacing, and everything else were equal, it would need to be twice as large as an optimized pacemaker battery to obtain the same longevity. Optimizing a battery for longevity and power becomes more complicated when a device performs multiple functions, such as both bradycardia pacing and defibrillation. For example, up to now, the lithium/ iodine battery has been the dominant power source for implantable cardiac pacemakers, which typically have peak power demands of 100 to 200 µW. Under these conditions, the lithium/iodine battery, which is still used in many pacemakers, can maintain an adequate voltage, even when its internal resistance reaches several thousand ohms. On the other hand, an implantable cardiac defibrillator may have peak

power requirements approximately 10,000 times greater than those of a pacemaker. Under such a high power demand, the voltage of an Li/I2 battery would drop to almost zero, and the power delivered to the device would be almost nil. In recent years the distinction between a need for high-rate and low-rate batteries has become more blurred because features such as distance (“wireless”) telemetry and multisite pacing need both more current and more capacity to operate. The result is that battery designers have been challenged to develop medium-rate batteries that can deliver more power than pacemaker batteries of the past while still having a high energy density. Average vs. Instantaneous Current Drain Implantable medical devices are often characterized by their average current drain, although events such as therapy delivery and telemetry may require temporary current excursions that can differ greatly from the average value. The effect of the instantaneous current demand can be mitigated by using a capacitor that buffers the battery to allow short bursts of power that may be more than two orders of magnitude higher than it could deliver directly. This allows the use of battery designs with reduced anode and cathode surface areas and improved volumetric efficiency. The same principle involving a battery plus a capacitor is used to deliver the defibrillation therapy to the heart from an ICD, even though the magnitude of the current and voltage is much greater. Shape, Size, and Mass Constraints All device requirements (e.g., longevity, end-of-service indication, peak power) must be balanced against device volume, shape, and mass. Volume and thickness are particularly important for safe implantation in children and for esthetic reasons in many adults. The ratio of volume to surface area in the battery is an important parameter. The performance of a battery of fixed volume will vary substantially depending on its electrode surface area/volume ratios. The operating current and longevity demanded of the battery determine both the minimum areas and the amounts of anode and cathode needed. Increasing the electrode area/volume ratio increases current capability, but this ratio must not be too large, or the battery will be too costly and will have a diminished energy density and most likely greater complexity, raising reliability concerns. Size, Energy Density, and Current Drain The relationship between battery size and average current is not one of direct proportionality. For example, decreasing the average current by 50% will not permit a 50% reduction in battery size without compromising longevity because of the inactive materials in a battery (e.g., case, electrolyte, current collectors). Likewise, the usable energy density is also a function of the current demand on the cell. As the current from the cell is increased, the resulting voltage drops significantly (see Fig. 6-3), which reduces the time during which the cell can provide current at or above the minimum voltage necessary to operate the electronic circuits. Thus, usable energy density, which is directly proportional to the area under the discharge (voltage vs. capacity) curve, is also reduced. For very-high-rate batteries such as those used to power ICDs, this is not such an issue because their internal resistance is extremely low. THE BATTERY AND LONGEVITY OF PULSE GENERATOR Longevity is typically defined as the interval between device implantation and detection of the end-of-service indicator. Because therapy can vary substantially from patient to patient, the longevity requirement is typically linked to a specified set of nominal conditions and programmed parameters. The minimum battery capacity required to achieve the specified longevity can be calculated from the average current needed for this nominal set of conditions. The following equation relates the longevity of the pulse generator, L, to the deliverable capacity of the battery, Qdel, and the average pacing current, I:



6  Power Systems for Implantable Pacemakers, Cardioverters, and Defibrillators L = Qdel /8766 I



(6-5)

The unit of L is years, Qdel is in milliampere-hours (mA-hr), and I in milliamperes (mA). The conversion factor 8766 (365.25 days/yr × 24 hours/day) is needed because longevity is expressed in years, not hours. The actual capacity that is built into the battery must be larger than Qdel, because additional capacity is needed to account for self-discharge and other parasitic losses of capacity (Qsd). More capacity must also be included to allow for an interval between the end-of-service indicator and the time when the battery can no longer power the device (QEOL). The total capacity (QTotal) is defined as follows: Q Total = Qdel + Qsd + Q EOL



(6-6)

The average current drain in Equation 6-5 depends on the characteristic of the pulse generator circuitry and the requirements for therapy. It has two main components: the static current drain, which powers the electronic components even when no therapy is delivered, and the therapeutic current. The trend throughout the evolution of implantable devices has been that current demands decrease as technology is improved, which leads to smaller batteries and pulse generators while maintaining relatively constant longevity. Some expect that this trend will continue, but the path to lower current often has a sawtooth profile because new features and therapeutic modalities temporarily increase the required current. For ICDs, the situation is more complicated because of the unpredictable mixture of bradycardia and tachyarrhythmia therapy delivery and the constant need to have a very high power capability in any device that may be required to deliver a defibrillation shock quickly. Effect of Pulse Width on Pacing Current Increasing the pacing rate, pulse width, or pulse amplitude increases the average pacing current. The average pacing current, excluding static current, is directly proportional to the pacing rate. Recall that the pacing pulse results from the discharge of a capacitor through the electrode-heart interface. This capacitor produces a pulse in which the current decays exponentially with time, as shown for two different pulse widths in Figure 6-4. Thus, the time-dependent behavior of the current during the pacing pulse is given by the following equation: I = (VA /R H ) e − t/R HC



(6-7)

where VA is the amplitude at the beginning of the pacing pulse, RH is the impedance of the lead plus the heart (discussed later), C is the value of the capacitor that delivers the pacing pulse, and t is the time since the beginning of the pacing pulse. In Figure 6-4, A and B, the pulse width is tw and tw/2, respectively. The area under each current-time curve gives the total charge delivered during the pulse. Although the width of the pulse in B is half that of the pulse in A, the charge delivered by this pulse is considerably more than half that of the longer pulse. The exact ratio of the charge delivered in the two cases depends on the

179

values of RH and C. Nevertheless, reducing the pulse width by a given fraction will always reduce the average pacing current by a substantially smaller fraction because of the exponentially decaying shape of the pacing stimulus current curve. Effect of Pulse Amplitude on Pacing Current The definition of pacing pulse amplitude may vary somewhat between manufacturers of implantable pulse generators. For our purposes, pulse amplitude is defined as the voltage delivered to the heart at the beginning of the pacing pulse (“leading edge” voltage). As stated earlier, the area under the current-time curve gives the charge delivered per pulse. Thus, doubling the amplitude doubles the current and the total charge delivered to the heart. Also, because the charge per pulse is doubled, it might seem that the average pacing current drawn from the battery would also be doubled. However, the impact on the pacing current is much larger than that, as seen from the following argument. The energy per pacing pulse is defined as follows:

E = VA × I A × t w

(6-8)

In Equation 6-8, VA is the average pacing stimulus output voltage of the pulse generator, IA is the average pacing current delivered to the heart, and tw is the pulse width. If we consider the lead-electrode-heart interface to be mainly resistive, Ohm’s law, I = V/R, can be substituted in Equation 6-8, which becomes the following:

E = (VA 2 /R H ) × t w

(6-9)

In Equation 6-9, RH is the effective ohmic load of the heart and lead. From this equation it is readily apparent that energy consumption increases with the square of the output voltage. The effect of this energy increase on the current drawn from the battery may not be intuitive. Because the battery supplies all the energy delivered to the heart at a relatively constant voltage, any increase in energy is accompanied by a proportional increase in current drawn from the battery. Thus the average pacing current is increased by a factor of four if VA is doubled. In fact, this is the best situation; additional energy losses occur when the stimulus voltage is programmed to a higher level because the electronic circuit for increasing the stimulus voltage is not 100% efficient. EFFECT OF LEAD IMPEDANCE ON PACING CURRENT It is also important to consider the effect of RH on the average pacing current. RH is sometimes called “lead impedance,” but this is a misnomer. The impedance of the lead, itself, is mainly a resistance and it is relatively small (50-100 ohms). Most of the “lead” impedance actually arises at the electrode-tissue interface (500-1000 ohms or more). The factors affecting the impedance of this interface are discussed in Chapter 1. In general, the average pacing current is approximately inversely proportional to the sum of the actual lead and tissue interface impedances. Thus, there is a substantial interest in technologies that can increase RH at the electrode-tissue interface (without increasing the lead conductor resistance), thereby decreasing the pacing current while maintaining a constant or even reduced pacing threshold voltage.

Current

Current

Programming Effects on Longevity of Bradycardia Pulse Generators

tw/2

tw

A

Time

B

Time

Figure 6-4  Comparison of charge delivered to heart as pacing pulse width is shortened. The two cases (A and B) assume a discharge into the same patient. The pulse width, tw, is typically between 0.2 and 1.5 msec, and the maximum current is between 2 and 15 mA, depending on the pulse generator output voltage and the lead-heart interface resistance.

In summary, the wide range of available pacing parameters can have a dramatic effect on the current drain from the battery in an implanted pulse generator. For example, in the same patient, a bradycardia pulse generator with 6 years of longevity under nominal pacing parameters may reach its replacement time in 2 years at one extreme or more than 10 years at the other extreme. Therefore, it is important for clinicians to optimize parameters for the individual patient and to consider each device’s settings when predicting remaining device longevity for a specific patient. Considerations for Longevity of ICDs Most of the arguments concerning bradycardia pacemakers also apply to ICDs. However, ICDs are much more complicated than typical

SECTION 1  Basic Principles of Device Therapy

BATTERY END-OF-SERVICE INDICATION End-of-service requirements result from the need to indicate impending battery depletion in both pacemakers and defibrillators in a manner that allows adequate time to replace the device. In general, this requires a battery to have some measurable characteristic, such as voltage or impedance, which can be related to its state of discharge. The pulse generator end-of-service indication must occur well before the battery loses so much voltage that it cannot sustain cardiac pacing or perform defibrillation. Because longevity is a strong function of device settings, the longevity requirement is typically linked to a specific set of nominal pacing or defibrillation parameters. A detailed knowledge of the variations in battery performance, the changes in load current with pulse generator settings, and the accuracy of the end-of-service measurement circuitry is necessary to ensure that these requirements will be met. Elective Replacement Indicator All modern implantable pulse generators have an elective replacement indicator (ERI) that alerts the clinician to impending battery depletion and allows adequate time for replacement of the device. Typically, the ERI is designed to alert at least 3 months before the battery voltage drops to a level at which erratic pacing, loss of capture results, or loss of other critical features occur. Future devices will conform to CENELEC standards (European Committee for Electrotechnical Standardization). For pacemakers, the standard requires a recommended replacement time (RRT) such that at least 95% of the devices will have a prolonged service period (PSP) of at least 180 days. For ICD and CRT devices, the requirement is that at least 95% of the devices have a PSP of at least 90 days. Device manufacturers may choose to retain the ERI nomenclature and may set ERI to be coincident with RRT or between RRT and end of service. Methods for Monitoring State of Battery Discharge Battery Voltage.  The most common method to indicate impending battery depletion is to measure the battery voltage. Most modern devices incorporate a voltage measurement circuit in the form of an analog-to-digital converter. The digitized battery voltage can be compared with a value stored in a nonvolatile memory to trigger the ERI, or it can be telemetered to the programmer where the comparison is made. For lithium/iodine batteries, the battery voltage remains relatively constant throughout most of its discharge under low load conditions, as shown in the voltage versus capacity curve in Figure 6-5. This figure also shows the resistance of this battery as it discharges. Note that the resistance changes from a modest value at the beginning of service to a quite large value when it is almost depleted. Because the battery voltage is relatively constant for much of the pulse generator’s useful life, the telemetered voltage of this battery may not be particularly useful for estimating remaining service. Voltage characteristics during discharge differ for different battery chemistries, so clinicians should not assume familiarity with one model or that one type of battery can be applied to another. For example, the lithium/ iodine battery has a fairly flat discharge curve throughout most of its life at pacing currents. However, other battery chemistries have voltage characteristics much different than Li/I2. The lithium/silver vanadium oxide (Li/SVO) battery used in many ICDs has two distinct voltage regions before the typical ERI voltage. The region following the second plateau is sloped. The switch to the sloped discharge region can be used as the ERI trigger for this battery chemistry. This voltage is much less sensitive to current variances than the voltage chosen for an Li/I2

3

10 8

2.5

8

Voltage

7

2

6 5

1.5

4

1

3

Resistance

2

0.5 0

Resistance (kohms)

pacemakers, so it is not nearly as easy to generalize about the effect of different factors on their longevity. In general, the two main factors to consider are the expected frequency of tachyarrhythmia therapies and the percentage of time spent pacing the heart. The longevity may vary by a factor of two to three because of these issues alone, as discussed later.

V load (V)

180

1 0

0.2

0.4

0.6

0.8

1.0

1.2

1.4

0

Discharged capacity (ampere-hours) Figure 6-5  Relationships among voltage, resistance, and delivered capacity. Voltage remains reasonably constant through the service life of lithium/iodine battery, whereas the resistance steadily builds up with discharge. Note the rapid falloff in voltage as the resistance rapidly increases near the end of service life.

battery because the internal resistance of the Li/SVO battery is much lower than the Li/I2, so for the Li/SVO system, voltage is a good indicator of remaining service life. Battery Impedance.  Battery impedance is another parameter used to signal the elective replacement point. Battery impedance is generally much less dependent on current than battery voltage and may convey more information about the battery’s state of discharge. For example, although the voltage of an Li/I2 battery remains relatively constant through most of its discharge, its impedance increases continuously, and especially rapidly as the battery approaches depletion (see Fig. 6-5). At depletion, battery impedance is not only useful for signaling the elective replacement point but may also provide an estimate of remaining service life. This feature has been incorporated into some pacemaker designs. Battery impedance is usually determined from the open-circuit and load-circuit voltages (or two different load-circuit voltages) of the battery by measuring the voltage drop across a known resistor within the pulse generator, then assuming Ohm’s law can be applied. Consumed Charge.  A final method used to indicate remaining battery life has been to measure the cumulative sum of the charge removed from the battery. This is accomplished by monitoring the current drawn from the battery or the current and voltage (i.e., energy) drawn from the battery. This method requires an accurate knowledge of the original deliverable capacity of the battery because the technique actually measures the capacity already used, and the amount left must be calculated by subtracting this from the initial value. False Triggering Any large increase in average battery current, even if temporary, in pulse generators that utilize battery voltage to indicate the elective replacement point may drop the battery voltage below the ERI trigger value before the true ERI point is reached; this occurs particularly with the Li/I2 battery because of its high resistance. For example, changes in current can result from electrophysiologic studies with noninvasively programmed stimulation (NIPS) in which the pacemaker is used to interrupt an episode of tachycardia by delivering short bursts of veryhigh-rate pacing, or from high pacing rates or telemetry sessions. Because ERI is usually latched (stored) when it is triggered, these temporary increases in current can cause a false, premature appearance of the ERI. In most cases, devices are designed to allow an ERI prematurely triggered to be reset. In some devices, ERI is inhibited during temporary high-current events to prevent false triggers.



6  Power Systems for Implantable Pacemakers, Cardioverters, and Defibrillators

In ICDs, ERI is usually based on the battery voltage or the time required to charge the capacitor. The battery voltage is usually measured during normal sensing/pacing operation and not during a defibrillation therapy, when the battery voltage is depressed. Because of the very low internal resistance of ICD batteries, pacing parameters have a much smaller effect on triggering ERI in these devices, so false triggering is less of a problem. Alternatively, some ICDs base ERI determination on the time required to charge the output capacitor. This approach is possible because most battery designs have a reduced power capability as the battery approaches depletion. This method does not usually cause false ERI triggering.

181

Negative terminal Welded seal

Fill port Lithium anode P2VP film Cathode material

CLINICAL INDICATORS OF BATTERY REPLACEMENT TIME The clinical indicators of battery depletion vary widely among manufacturers of pulse generators. In fact, a single manufacturer may provide significant variations between various models. Of the following common indicators, the first three mainly pertain to older pacemakers that may still be implanted in some patients: 1. Stepwise change in pacing rate. Elective replacement time is indicated by a change in the pacing rate to a predetermined fixed rate (e.g., 65 bpm) or a fractional change in rate (e.g., 10% decrease from programmed rate). 2. Stepwise change in magnet rate. The magnet-pacing rate decreases in a stepwise fashion related to remaining battery life. 3. Pacing mode change. DDD and DDDR pulse generators may automatically revert to another mode such as VVI or VOO to reduce current drain and extend battery life. 4. Telemetered battery voltage or impedance. In modern pacemakers the battery voltage or the battery impedance can be telemetered to the programming device. This information is used in estimating remaining battery service life or in indicating an imminent ERI by algorithms performed outside the pulse generator. All manufacturers provide technical manuals containing tables or graphs that indicate the relationship between battery voltage or impedance and the estimated remaining service life of the device. BATTERY CHEMISTRIES USED IN PACEMAKERS Lithium/Iodine Battery The lithium/iodine battery is probably the most well known implantable battery because it has been used in the vast majority of cardiac pacemakers. The first implant of a pacemaker powered by an Li/I2 occurred in 1972.5,6,9 At least 13 million Li/I2-powered pacemakers have now been implanted. Many factors favor the use of this battery system. When the current demand is low, it is difficult to improve on the performance of Li/I2 batteries; their high energy density and low rate of self-discharge result in good longevity and small size. The inherently high impedance of the Li/I2 battery has not been a major disadvantage up to now because the current required by modern pacemaker circuits is low, typically about 10 µA. Note that the much larger current delivered during a (short duration) pacing pulse is drawn from a capacitor, which can charge between pacing pulses. The voltage and impedance characteristics of the Li/I2 cell also allow the clinician to monitor the approaching end-of-service indication. This battery system is simple, elegant in concept and is inherently resistant to many common modes of failure, as will be discussed below. As a result, Li/I2 batteries have attained a record of reliability unsurpassed among electrochemical power sources. Cell Structure.  Figure 6-6 shows a cutaway view of a typical Li/I2 battery. In general, Li/I2 batteries have a single, central lithium anode that is surrounded by cathode material, which is at least 96% iodine and has been thermally reacted with a polymer material to render a conductive mixture. The central anode is seen with an embedded current collector wire, and the iodine cathode fills much of the volume

Cathode P2VP film Anode Figure 6-6  Cutaway view of typical lithium/iodine battery. Easily identifiable interior parts include the anode feedthrough, anode current collector (wire), cathode, poly-2-vinylpyridine (P2VP) film on the anode, and port where the cathode material is poured into the cell. The connection to the cathode is through the battery case.

inside the battery. Figure 6-6 also shows several other important structures, including the electrical feedthrough that connects the anode to the outside of the cell. The case serves as the site of the electrical connection to the cathode, which is in direct contact with the inside of the container. Another visible feature is the fill port, the means by which the cathode mixture is introduced into the cell, after which the fill port is hermetically sealed. Discharge Curve.  Figure 6-5 shows the characteristic shapes of the voltage and resistance curves as a function of discharge for a typical Li/I2 battery. During most of the battery life, the voltage is stable near 2.8 V, and the resistance change is gradual and dominated by the growing electrolyte (LiI) thickness. Near the end of battery life, the cathode resistance rapidly dominates the resistance of the cell, as the cathode becomes lower in iodine content and less conductive. The region of discharge dominated by cathode resistance and declining cell voltage is used to signal the approaching “end of service” for most pacemakers powered by Li/I2 batteries.10 Effects of Current Drain on Deliverable Capacity.  Unfortunately, the high energy density of the Li/I2 battery may be negated if the application requires frequent periods of high current drain. This is caused by the high internal resistance of the battery, which results in voltage losses due to polarization (see Figs. 6-2 and 6-3). The Li/I2 battery system provided small, simple, highly reliable power sources with power characteristics almost ideally suited to the requirements of cardiac pacing for more than 3 decades. In the late 1990s, however, several device features that required higher power began to emerge. In many cases, these peak power requirements are beyond the capability of Li/I2 batteries. These new features include increased use of addressable memory to capture and store information about the electrical activity of the heart, the need for faster and longer-range telemetry to transmit this information outside the body, new physiologic sensors, and new therapies with higher power requirements, such as cardiac resynchronization and treatment of atrial fibrillation. These higherpower features are now being supported by new battery chemistries with similar energy density to lithium/iodine, but a much higher power capability. Lithium/Carbon Monofluoride Battery Lithium/carbon monofluoride (Li/CFx) batteries were introduced in commercial coin-type batteries in 1976. Carbon monofluoride (also

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SECTION 1  Basic Principles of Device Therapy

known as CFx) is a cathode material with very high capacity and moderate power capability. The CFx material is a powder that is mixed with carbon as a conductivity enhancer plus a polymer binder and then pressed into a porous pellet. The battery comprises of a lithium anode, porous polymer separator material, and porous pressed-powder cathode pellet that are inserted into a case and hermetically sealed, similar to an Li/I2 battery. In contrast to the Li/I2 battery, however, the Li/CFx battery uses a liquid electrolyte composed of a lithium salt dissolved in an organic solvent (or solvent mixture). The energy density of this battery is similar to that of the Li/I2 battery, but it delivers significantly higher power. A typical discharge curve is shown in Figure 6-7, A. One feature of the Li/CFx system is that the voltage decline near end of battery life tends to be fairly abrupt, making it challenging for device manufacturers to engineer means of maintaining adequate replacement-time warning.

3.2 3.0 2.8 2.6 2.4 2.2 2.0 0%

20%

A

40%

60%

80%

100%

80%

100%

80%

100%

Cathode utilization

Lithium/Hybrid Cathode Battery New lithium battery chemistry with a cathode consisting of silver vanadium oxide (SVO) plus CFx has been developed to meet the needs of implantable devices with higher rate therapies and features.11,12 SVO has high power capability and is the same cathode material used to power ICDs. As previously described, CFx has modest power capability and more abrupt end-of-service characteristics, but a very high capacity. The blend of the two cathode materials, called a hybrid cathode, yields a primary battery that has an energy density equal to that of a lithium/iodine battery with approximately 100 times more power capability. It can support the new device features and also has good end-of-service detection because of the SVO. Hybrid cathode batteries have been used in implantable cardiac devices since 1999. Cumulative implants through 2009 total about 300,000. Plots of voltage versus percent of discharged capacity for SVO and CFx batteries are shown in Figure 6-7, A and B. The voltage curve of hybrid cathode behaves like a superposition of the two in proportion to the starting composition of the mixture (Fig. 6-7, C). The composition of the hybrid can be chosen based on the application, enhancing the capacity with a greater proportion of CFx or enhancing the power capability and end-of-service characteristics with more SVO. For low- and medium-rate applications, the composition can be chosen such that 85% to 90% of the battery capacity comes from CFx. In this composition range, efficiently designed medium-rate hybrid cathode batteries can match lithium/iodine in energy density at about 1 W-hr/cm3. Two basic hybrid cathode battery designs have been used in implantable pacemakers. Both types use lithium anodes, a porous polymer separator, a porous pressed-powder cathode pellet, and a liquid electrolyte of lithium salt dissolved in an organic solvent blend. The differences in the two designs are related to the construction of the cathode pellet, either a mixture of the SVO and CFx active materials or the CFx and SVO in separate layers, with the SVO layer adjacent to the lithium anode. In both cases the cathodes also contain a polymer binder and carbon as a conductivity enhancer. Although the methods of construction are different, the discharge behavior of either design is similar to that in Figure 6-7, C. BATTERY CHEMISTRIES USED IN DEFIBRILLATORS

3.2 3.0 2.8 2.6 2.4 2.2 2.0 0%

20%

B

40%

60%

Cathode utilization

3.2 3.0 2.8

Primarily CFx capacity

2.6 2.4

Primarily SVO capacity

2.2 2.0 0%

C

20%

40%

60%

Cathode utilization

Figure 6-7  Voltage versus percent discharged capacity. A, Typical discharge curve of a lithium/carbon monofluoride (Li/CFx) battery. B, Typical discharge of a lithium/silver vanadium oxide (Li/SVO) battery. C, Typical discharge curve of a CFx-SVO hybrid cathode battery.

Comparison of Pacemaker and Defibrillator Batteries Implantable defibrillators deliver up to 40 J of energy to the heart within a few milliseconds. By contrast, the energy required to stimulate the heart in bradycardia pacing is 1 to 10 µJ over 1 msec or less. Clearly, the demands placed on the battery that powers an implantable defibrillator are much different from those of a battery that powers a bradycardia pacemaker. Implantable defibrillators are designed to deliver a shock within about 10 seconds after the fibrillation or tachycardia is detected. When the ICD determines that a shock may be required, it begins to charge a high-voltage capacitor (discussed later). The time needed to charge the capacitors before the shock is delivered depends

on the ability of the battery to sustain a very high current during this period. Thus the implantable defibrillator requires a battery with high peak power, where power is the product of current and voltage (P = I × V). The high power required of defibrillator batteries dictates a different design than used for pacemaker batteries. Figure 6-6 shows a cutaway view of a typical battery used to power a pacemaker, with small and rather thick electrodes. Contrast this with Figure 6-8, which shows an analogous cutaway for a typical defibrillator battery, with very different construction. The long, thin anode and cathode, with two layers of



6  Power Systems for Implantable Pacemakers, Cardioverters, and Defibrillators Positive terminal

Fill port Anode connection to case Electrolyte Case liner Separator Cathode

Separator Separator Anode Figure 6-8  Cutaway view of typical high-rate defibrillator battery. In wound construction a sandwich of anode, separator, and cathode is rolled, flattened, and fitted inside the battery case. This battery uses a liquid electrolyte to ensure high internal conductivity. In this battery the anode is connected to the case and the cathode is connected to the feedthrough. Note the multiple connections to the cathode to ensure low internal resistance.

183

rapidly the battery is depleted. When the battery is depleted over a longer period, its internal impedance is higher in the latter half of the discharge curve. Figure 6-11 illustrates this phenomenon, with background voltage and internal impedance of Li/SVO batteries discharged over 3, 5, and 7 years. The internal impedance is unaffected by the discharge rate until midway through discharge. After that point the internal impedance always increases, but to a larger degree for slower discharge rates (i.e., longer discharge times). The higher internal resistance results in longer charge times, especially for the longest discharge periods. Charge Time–Optimized Li/SVO ICD Batteries.  One manufacturer has addressed the problem of the time-dependent increase in charge time by changing the ratio of anode to cathode material in the battery. The first Li/SVO batteries were made with an excess of lithium anode material so that the voltage only dropped slowly at the end of service life. “Cathode limitation” is the most efficient battery design if the goal is to optimize energy density. However, if the goal is to maintain a low charge time, it makes sense to use only part of the cathode capacity, ending the discharge partway onto the 2.6-V plateau where the charge time starts to increase. This is accomplished by reducing the amount of anode and increasing the amount of SVO cathode. As a result, the battery always remains in the region where internal impedance is low and charge times are short. 3.4

TYPES OF BATTERIES USED IN ICDS Three different types of battery chemistries are in use currently in ICDs: lithium/silver vanadium oxide, lithium/manganese dioxide, and lithium/layered SVO-CFx (also called “dual cathode”). The shapes of the discharge curves, resistance of the battery, and effects of time since implant vary among battery types. Regardless of the battery chemistry, all ICD batteries are used in much the same way. Most of the time the battery delivers a low current drain as required for sensing and pacing, which has negligible effect on the battery voltage. When it is necessary to charge the high-voltage capacitors, whether for capacitor formation (see Capacitors) or to deliver a therapeutic shock, the voltage drops quickly by about 1 V, then, after the capacitors are charged, recovers quickly to near the previous value (Fig. 6-9). Lithium/Silver Vanadium Oxide The first ICD battery chemistry put into widespread use was Li/SVO.11-17 These batteries have a distinctive discharge curve, with a slightly sloping voltage at about 3 V in the first third of discharge, followed by a decline in voltage to a flat region with about 2.6 V, and ending in a short region of declining voltage, as seen in Figure 6-10. The uppermost curve represents the battery voltage during normal sensing and bradycardia pacing operations (also called the “background” voltage). It is this background voltage that is typically monitored by the ICD and reported via telemetry. The “load” voltage is the voltage measured during charging of the defibrillation capacitors, and it is typically much lower than the background voltage. After many years of laboratory testing and experience in the field, it is apparent that the performance of Li/SVO batteries depends on how

Detection of arrhythmia

3.2

separator film in between, are rolled into a flattened coil. Alternatively, either or both of the electrodes of ICD batteries can be individual flatplate electrodes electrically connected to each other. Either method provides the high-power electrode area necessary for ICD batteries. The electrolyte in the defibrillator battery is a highly conductive solution of a lithium salt in organic solvents. The design with large, thin electrodes and liquid electrolyte gives the defibrillator battery the high power needed for quickly charging the high-voltage capacitors in the defibrillator. As mentioned, however, these same characteristics also reduce the energy density of the battery.

3 Voltage recovery

2.8 2.6

Shock delivered

2.4

Charge time

2.2 –5

2

0

5

10

15

20

Time (seconds) Figure 6-9  Typical voltage-time curve showing defibrillator battery voltage during delivery of a defibrillator shock. Left, Voltage before and during detection; middle, during charging of high-voltage capacitor; and right, voltage recovery after delivery of defibrillation shock.

3.4

20

Background voltage

3.2

18

Charge time

3.0

16

Load voltage

2.8

14

2.6

12

2.4

10

2.2

8

2.0

6

1.8

4

1.6

2

1.4

0 900

0

100

200

300

400

500

600

700

800

Capacity (ampere-hours) Figure 6-10  Typical plot of discharge characteristics of the lithium/ silver vanadium oxide ICD battery.

SECTION 1  Basic Principles of Device Therapy

Background voltage (V)

3.5 3.0

Background voltage

70

40

60

35

2.5

50

2.0

40

1.5 1.0

Area-normalized resistance

0.5 0.0 0%

30

7 years 5 years 3 years Zero-time

50%

20 10 100%

Area-normalized resistance (ohm cm2)

184

0

Percent depth of discharge Figure 6-11  Comparison of area-normalized resistance of lithium/ silver vanadium oxide battery for various battery longevities. The areanormalized resistance is the internal impedance of the battery, calculated from the drop in voltage during the pulse, the current, and  the area of the cathode. The area-normalized resistance = Acathode • (Vbackground − Vpulse)/Ipulse.

Lithium/Manganese Dioxide Another ICD manufacturer uses manganese dioxide batteries for use in ICDs. Lithium/manganese dioxide (Li/MnO2) batteries have been used in consumer and military batteries since the 1980s. For example, high-power Li/MnO2 batteries used for photoflash applications have similar electrical characteristics to those used in ICDs. Figure 6-13 provides a graph of Li/MnO2 battery performance. During the first half of battery life, the voltage remains almost constant at 3.1 V. In the second half of battery life, the voltage gradually slopes downward. This characteristic of the battery voltage, along with charge time and the amount of battery energy consumed, allows the clinician to judge how much longer the battery may last. The internal resistance and thus the charge time of Li/MnO2 batteries remain relatively constant until battery replacement is indicated by the ICD.19 Lithium/Layered Silver Vanadium Oxide–Carbon Monofluoride Another manufacturer of ICD batteries has combined the properties of two cathode materials, silver vanadium oxide and carbon monofluoride, to create a novel ICD battery that exploits the best features of those two materials. SVO provides high power while CFx contributes high capacity per unit volume. The construction of the cathode portion of the battery is rather complex, with a central CFx cathode sandwiched between two metallic current collectors, with SVO electrodes on each side. This arrangement of electrodes allows the SVO, which can discharge at very high rates, to be preferentially discharged when high current is needed to charge the capacitors. Over time the extra lithium ions that accumulate in the SVO part of the electrode equalize with the CFx part, so that the lithium ions become distributed equally within both parts of the cathode. The voltage curve starts at about 3.2 V, stays almost flat from about 20% to 50%, and then gradually declines. Figure 6-14 shows a typical discharge curve.20

3.00 2.75

25

2.50

20

2.25

15

2.00

Conventional SVO charge time

10

0

0

1

2

3

A

4

5

6

7

8

1.50 9

1.25

Longevity (yr)

40

3.25

Charge time– optimized battery voltage

35 30

1.75

Charge time– optimized charge time

5

25

Figure 6-12 compares the voltage and charge time of conventional SVO ICD batteries with charge time–optimized batteries that occupy the same volume in an ICD. For example, when discharged over 7 years, the charge time is considerably lower starting after 4 years for the charge time–optimized battery than the conventionally balanced battery (Fig. 6-12, A). The charge time–optimized battery maintains a much lower charge time in the latter half of ICD life, without substantially sacrificing longevity.18 The difference in charge time is much less when the battery is depleted over a shorter time, such as 3 years (Fig. 6-12, B).

Conventional battery voltage

30

3.25

Charge time– optimized battery voltage

3.00 2.75

Conventional battery voltage

2.50

20

2.25 Conventional SVO charge time

15 10

2.00

1.75 Charge time–optimized charge time 1.50

5 0 0.0

0.5

1.0

B

1.5

2.0

2.5

3.0

3.5

4.0

1.25

Longevity (yr)

Figure 6-12  Comparison of the performance of charge time– optimized batteries with conventionally balanced silver vanadium oxide (SVO) batteries as a function of time for battery longevity of 7 years (A) and 3 years (B). Both batteries occupy the same volume in the ICD.

SAFETY Most ICD batteries are designed to provide the power needed to deliver a therapeutic shock within 5 to 10 seconds after an arrhythmia is detected. With efficiency losses, this corresponds to a power capability of approximately 10-20 W. This is in marked contrast to that of 3.2

8

2.8

6 Battery voltage Internal resistance

2.4

4

2

2.0

1.6 0%

20%

40%

60%

80%

100%

0 120%

Extent of battery depletion Figure 6-13  Voltage and internal resistance with extent of discharge for a lithium/manganese dioxide ICD battery. Charge times remain fast and relatively steady as the internal resistance remains low until near the end of battery life. (Courtesy Boston Scientific.)



6  Power Systems for Implantable Pacemakers, Cardioverters, and Defibrillators 3500

Voltage (mV)

3000 2500 2000 1500 1000 500 0

0

10

20

30

40

50

60

70

80

90 100

% DOD Figure 6-14  Discharge curve for lithium/layered silver vanadium oxide–carbon monofluoride battery as a function of percent depth of discharge (DOD). This test lasted 6 months. A 3.15-A pulse was applied about every 15 days. Upper curve, Battery voltage on a low current drain. Lower curve, Battery voltage at the end of the 3.15-A pulse. Batteries of two different sizes were compared, 6.5 cm3 (blue) and 9.0 cm2 (orange), with minimal differences. (Courtesy Dr. Esther Takeuchi, State University of New York, Buffalo.)

most pacemaker batteries, which have power capability of less than 0.001 W. Typical ICD batteries have power capability comparable to AA Li/ MnO2 batteries widely used in consumer applications, and far less power capability than many other batteries used in familiar consumer products. Nevertheless, higher power by itself carries an increased risk of overheating in the event of a short circuit. To avoid this, ICD batteries are carefully designed and rigorously manufactured to prevent short circuits. Moreover, special separator materials are used that limit currents to safe levels in the event of a short circuit. Finally, all ICD battery designs are extensively tested under rigorous abuse conditions to verify the safety of the design.

Emerging Power Sources

185

6. Long calendar life. The best lithium-ion batteries will maintain more than 80% of initial capacity after 10 years of use. The only effect of capacity loss to the patient is reduced time between required recharges. It is not yet clear what future applications will emerge for rechargeable batteries in the cardiac rhythm management arena, but some applications are likely, and substantial patent activity is already occurring. In fact, neurologic spinal cord stimulators powered by rechargeable lithium-ion batteries have been developed by three different device manufacturers and make up the majority of current implants. Lithiumion batteries are also used to provide temporary power for some left ventricular assist devices (e.g., while patient is bathing). In many applications the time between recharges could be as long as weeks or months. Possible modes of lithium-ion battery application include the following: 1. Supporting relatively high-power, non–life-support features of a device (e.g., long-distance telemetry, physiologic sensors). In this case a primary battery would still be used to provide life-sustaining therapy. 2. Using a high-power rechargeable battery to charge the capacitor of an ICD while incorporating a low-power, high–energy density primary battery to provide the recharge energy. In some cases, this may be accomplished by placing the two cells in parallel. 3. Providing excellent longevity (up to 20 years or more) with small devices. In this case the rechargeable lithium-ion battery would serve as the sole power source. Presumably, such devices would be limited to patients who are physically and mentally capable of managing the recharge requirements. Principles of Operation The defining feature of a lithium-ion battery is that it contains no metallic lithium. Instead, lithium ions (Li+) are shuttled back and forth between the positive and negative electrodes during charge and discharge, as shown in Figure 6-15. The most common electrode materials are lithium cobalt oxide (LiCoO2) for the positive cathode material and a graphitic carbon to contain the intercalated lithium for the negative anode material. Many alternative materials are being developed and introduced in commercial batteries and will eventually migrate to batteries for implantable medical devices. The general cell construction is

RECHARGEABLE LITHIUM-ION BATTERIES Rechargeable lithium-ion cells were introduced to the consumer market by the Sony Corporation in 1992. Since that time, lithium-ion technology has been the major focus of research and development investment by many leading battery manufacturers. Lithium-ion batteries exhibit several characteristics that make them highly suitable for implantable medical applications, as follows: 1. High energy density. Although lithium-ion batteries have only about half the energy density of a typical ICD battery, they have double the energy density of older rechargeable batteries (e.g., NiCad). For applications that have a high average power requirement, lithium-ion batteries can provide both small size and excellent longevity. 2. No “memory effect.” Lithium-ion batteries can be partially charged or discharged with no deleterious effects. Thus the patient is free to develop a recharge schedule that is convenient. 3. High voltage. Lithium-ion batteries typically operate in the range of 3 to 4 V and are generally compatible with electronics for implantable devices that have historically used 3-V primary lithium batteries. 4. Low self-discharge. Lithium-ion batteries can be designed to have self-discharge rates of less than 1% per month. Furthermore, the self-discharged capacity is almost fully recoverable on subsequent recharge. 5. High cycle life. Several thousand charge/discharge cycles can be completed with much more than 50% capacity retention. This greatly exceeds the demands of most anticipated implantable applications.

e

e V

Charge

Li+

Discharge

Li+ discharge

Li+

Li+

Li+

Negative electrode (typically LixC)

Li+

Li+

Li+ charge

(Electrolyte)

Li+

Li+

Positive electrode (typically Li(1-x)CoO2)

Figure 6-15  Schematic diagram of a lithium-ion (Li+) battery showing lithium ions shuttling between the positive and negative electrodes as the battery is charged and discharged.

186

SECTION 1  Basic Principles of Device Therapy

4.2

similar to that shown for the Li/SVO battery. However, lithium-ion batteries are manufactured in their discharged state (e.g., using LiCoO2 and graphite) and then charged (or “formed”) after the cell is fully assembled.

4 3.8

Method of Recharge Cell voltage

To date, all of the implantable applications for rechargeable batteries involve recharge by transcutaneous electromagnetic induction. The implanted device incorporates a wound-wire coil that acts as the secondary to an externally located primary coil that is placed adjacent to the implanted device. Typical lithium-ion batteries are capable of accepting a full charge in 2 hours or less. Recharge times in future batteries with improved chemistries could become as short as 10 to 30 minutes. In addition to battery capacity, actual recharge time in ICDs depends on the relative positions of the primary and secondary coils, the depth of the implant, and the maximum power that can be transmitted while limiting heating of the primary coil or the implanted device.

3.4 3.2 3 2.8 2.6 0%

End of Service Life

20%

40%

60%

80%

100%

Depth of discharge

Lithium-ion batteries typically exhibit a gradually declining voltage during discharge that can be used as a “gas gauge” to indicate the state of discharge at a given time (Fig. 6-16). Information about the voltage can be telemetered to a patient controller or a monitor to alert the patient of the need to recharge the battery. However, the end of service for a device powered by a lithium-ion battery is not as readily apparent as for one powered by a primary battery. Because the lithium-ion battery slowly loses capacity as a function of both time and the total number of charge/discharge cycles, the patient may eventually experience a reduced interval between recharge sessions becomes too short to be acceptable to the patient. However, because rechargeable batteries can, in principle, provide unlimited energy, the eventual end of service may be determined by the lifetime of some other component of the implanted device. CAPACITORS The ICDs defibrillate the heart by delivering one or more high-energy, high-voltage shocks. The power capability and voltage of the battery alone are not sufficient to accomplish this. High-voltage capacitors are necessary to store and deliver the energy required for these shocks. The use of capacitors has important clinical implications, including delivered energy, device size, charge time, and longevity. Capacitance In its simplest form, a capacitor consists of two conductors, separated in space and electrically insulated from each other. When the conductors are charged, each stores an equal and opposite amount of charge. In an ideal capacitor, the amount of charge (Q) on each conductor is proportional to the difference in voltage (V) between the two conductors. The proportionality constant, C, is called the capacitance. Q=C×V



(6-10)

Figure 6-16  Monotonic discharge curve of the lithium-ion battery shows how it could be used as a “gas gauge” for determining the remaining capacity in a battery.

The simplest model of a capacitor is shown in Figure 6-17, A, in which the plates are separated in a vacuum. The conductors (or electrodes) consist of two parallel plates of area (A) separated by a distance (d). The capacitance, or ability to store charge as a function of voltage, is equal to A/d multiplied by a fundamental constant, ε0, the permittivity of free space, which is related to the energy required to separate charge in a vacuum. The value of the capacitance can be increased by inserting insulating materials, known as dielectrics, between the electrodes. The factor by which these materials increase the capacitance is called the dielectric constant, k, which is the ratio of the permittivity of the material to the permittivity of free space, ε/ε0. A simple capacitor containing a dielectric material is shown in Figure 6-17, B. Various materials are used as dielectrics the properties of which largely determine the properties of the capacitor. Energy Storage In an ICD, several capacitors are required to deliver a high-voltage, high-energy electric shock through the lead system to the heart. Capacitors are passive devices and therefore must store energy before it can be delivered. The process of storing energy in a capacitor is called “charging.” For a capacitor to be charged, both terminals must be connected to a charging circuit that includes a transformer and to a voltage source such as a battery. The transformer converts the low-voltage energy supplied by the battery into the high-voltage energy stored by the capacitor. The amount of energy stored on the capacitor is proportional to the square of the voltage on the capacitor, as follows: E = 12 C × V2 (6-11)

Parallel-plate capacitor

Parallel-plate capacitor with dielectric

d

d

A = Area

A

3.6

Dielectric constant = κ = ε/εo

A = Area C0 = Capacitance =

Charge Q ε A = = o Voltage V d

B

C = Capacitance =

Charge Q κεoA = = Voltage V d

Figure 6-17  Schematic representation of a parallel-plate capacitor with A, a vacuum, and B, a dielectric, between the capacitor plates. The capacitance is directly proportional to the area of the plates (A), and inversely proportional to the distance between the plates (d). The insertion of a dielectric material between the plates increases the capacitance by a factor of k, the dielectric constant, which is a characteristic of the particular insulating material.



6  Power Systems for Implantable Pacemakers, Cardioverters, and Defibrillators 800

Voltage (volts)

700 600 500 400

Voltage at time = one time constant = 750 (1 – 0.63) = 278 V

300 200

One time constant = RC = 5 ms

100 0

0

5

10

15

20

Time (milliseconds) Figure 6-18  Waveform for the discharge of an ideal 100-microfarad capacitor, charged to 750 V, into a 50-ohm load. These parameters approximate those for an implantable defibrillator in use.

Energy Delivery During defibrillation, the energy stored in the capacitor is delivered to the heart using an electrical path between the opposing electrodes of the capacitor. This connection allows the separated charge to recombine by traveling through the heart. When an ideal capacitor is discharged through a constant resistance, R, the voltage decays with time in an exponential manner, as seen in Figure 6-18. For an ideal capacitor discharged to 0 volts, the delivered energy would equal the stored energy. In ICDs the amount of energy delivered from the capacitor is less than the amount stored. One reason is that the therapy delivery is truncated before the capacitor is completely discharged, and the remaining energy is discarded. A truncated waveform is used because it defibrillates the heart more effectively than a waveform that is allowed to decay to 0 V, despite less energy being delivered to the heart. Other inefficiencies associated with the capacitor, device circuit, and delivery system also reduce the amount of stored energy delivered to the heart, as discussed later. Energy Density Energy density is an important parameter for capacitors as well as batteries. Capacitor energy density is typically expressed in terms of volume and has units of joules per cubic centimeter (J/cm3). The most energy-dense modern defibrillation capacitors deliver energies in excess of 5 J/cm3. As previously discussed, the highest energy density seen for implantable batteries is about 1 W-hr/cm3, which is the same as 3600 J/cm3, or more than 500 times as high as the best defibrillation capacitor. Why, then, use capacitors at all? The reasons are voltage and power. Defibrillation therapy requires the energy to be delivered rapidly at about 1000 W of power and starting at a voltage of 650 to 850 V. ICD batteries are designed to deliver the energy to the circuit at about 10 W of power with voltage between 2.5 and 3.0 V. Consequently, the ICD is designed with the battery as a large energy reservoir, analogous to the gas tank of an automobile, and the capacitor as the fuel injector to the engine. Nevertheless, it is important for the capacitors to have the highest practicable energy density in order to minimize the overall volume of the ICD.

187

to make the metal particles bond to one another (called sintering) while still remaining porous. For either material, an oxide film is electrochemically grown (called forming) on the exposed surfaces of the metal electrode. The oxide serves as the capacitor dielectric. Aluminum anodes are typically formed to support charging to about 400 V. Two capacitors are connected in series in the ICD to provide sufficient voltage for defibrillation. Tantalum anodes are typically formed to support charging to 175 to 250 V with three or four capacitors connected in series. The high capacitance, high voltage, and relatively low resistance of aluminum and tantalum electrolytic capacitors provide their high energy density (3-5 J/cc) and quick, efficient energy delivery that make them the capacitors of choice for ICDs. ICD Capacitor Construction Early ICD devices used commercially produced, cylindrically wound, aluminum electrolytic capacitors developed for flash photography applications. The need for smaller, thinner capacitors and specialized shapes led ICD manufacturers to develop stacked-plate aluminum electrolytic capacitors and custom tantalum electrolytic capacitors.21,22 Aluminum Electrolytic Capacitor Components.  All aluminum electrolytic capacitors are produced using the same basic materials of construction. Cylindrical and stacked-plate designs both incorporate an anode, a cathode, an electrolyte, and a separator material placed between the anode and cathode. The anode and cathode are aluminum foils that have been etched and formed, as previously described. Cylindrical Aluminum Electrolytic Construction.  Most commercially produced aluminum electrolytic capacitors employ wound electrodes in a cylindrical shape. Strips of anode and cathode foil, spaced apart by a paper separator, are wound into a coil. Figure 6-20 shows a cutaway view of an aluminum electrolytic photoflash capacitor of the type typically used in the first few generations of ICDs. Stacked-plate Aluminum Electrolytic Construction.  ICD manufacturers have developed stacked-plate aluminum electrolytic capacitor technology to improve the energy density, minimize the thickness, and allow a variety of shapes compared with cylindrically wound capacitors. The details of construction vary among manufacturers, but several design themes are common. Foils and separators are cut to a specific geometry and layered such that multiple anode plates are

ICD Capacitor Types The defibrillation capacitors used in past and current ICDs are called electrolytic capacitors, most often based on aluminum or tantalum electrode materials. The metal electrode materials are processed so they have a very large surface area, which allows for a high capacitance. For aluminum electrolytic capacitors, the positive electrode (anode in a capacitor) starts as a thin foil that has been etched to create a large number of tunnels. Figure 6-19 shows a scanning electron micrograph of a cross section of a highly etched anode foil. For tantalum electrolytic capacitors, the anode is made from tantalum metal powder pressed into a porous pellet and heated to a high enough temperature

Figure 6-19  Photomicrograph of a replica made from the etched anode foil used in aluminum electrolytic capacitors. The long, linear structures are tunnels etched into the aluminum foil. In this example some of the tunnels are etched completely through the foil to provide an electrical connection between the two sides. Magnification is about 800×.

188

SECTION 1  Basic Principles of Device Therapy

Terminals Crimp seal

Wound stack

Separator Anode Anode Separator Cathode Figure 6-20  Cutaway view of wound aluminum electrolytic capacitor used in early implantable defibrillators. The anode and cathode connections are brought out of the capacitor using feedthroughs. The capacitor is closed with a crimp seal. Notice that two layers of anode material are used for each layer of cathode material. The entire wound capacitor coil is saturated with electrolyte (not shown).

Deformation.  The non-ideal process that is most apparent is called deformation. As previously explained, the dielectric material in electrolytic capacitors is created by “forming,” immediately after which the dielectric is almost free of defects. Over time, the chemical environment within the capacitor or relaxation of residual stress will cause microscopic imperfections in the oxide film on the anode. When the capacitor is charged to a high voltage after a long period of disuse, additional energy is required to regrow oxide in these areas to heal the defects.23 This process is known as re-formation. The additional energy needed to re-form the dielectric results in longer charge times in ICDs. In a typical aluminum capacitor, 25% or more additional energy (and time) may be required to charge the capacitor if it has been many months since the last charge. The amount of additional energy drops to about 15% for capacitors not charged for about a month. Tantalum capacitors perform similarly, although the rate of deformation may be somewhat slower. Figure 6-23 shows capacitor deformation in typical aluminum capacitors for various charging intervals. To minimize the clinical effect of deformation, current ICDs incorporate a programmable maintenance routine that automatically charges the capacitor to high voltage to heal the dielectric at intervals of 1 month, 3 months, or 6 months. Leakage Current.  Leakage current is another non-ideal process that occurs in all capacitors to some extent. After a capacitor is charged, some of the energy “leaks” off the capacitor, and the voltage decreases with time. This happens quickly enough to affect how ICDs are designed; leakage current is the major reason why ICDs are only charged just before a shock is to be delivered. Internal Resistance.  The final important non-ideality is internal resistance. All capacitor types have some level of internal resistance, as do batteries. Electrolytic capacitors used in ICDs typically have a

Connector

sandwiched between cathodes and layers of porous separator paper. The geometry of the plates is selected so that the finished capacitor fits efficiently in the ICD. The energy density of stacked-plate capacitors can approach 4 J/cc versus the 2.5 J/cc typical of the coiled configuration. Figure 6-21 shows a cutaway view of a stacked-plate aluminum electrolytic capacitor similar to those used in current ICDs. Tantalum Electrolytic Capacitor Construction.  Tantalum electrolytic capacitors have the same basic elements as aluminum electrolytic capacitors. In tantalum capacitors the anode is made using pressed, sintered, and formed tantalum pellet. The cathode in a tantalum capacitor is typically deposited directly on the inside of the capacitor case, making it the negative terminal of the capacitor. Tantalum capacitors have similar benefits to stacked-plate aluminum capacitors but provide higher energy and simpler construction than aluminum capacitors. Rather than layered anode foils, tantalum capacitors have a single anode pellet. The simple and efficient construction minimizes the volume of materials that do not contribute to energy storage (cathode, separator, encasement), thus maximizing energy density. Tantalum capacitors used in current ICDs have an energy density of 5 J/cc or more. The primary drawback of tantalum capacitors is their mass. Because tantalum is very dense, tantalum capacitors are approximately 50% heavier than aluminum capacitors with equivalent energy. Figure 6-22 shows a cutaway view of a tantalum electrolytic capacitor similar to those used in some current ICDs. Non-Ideal Behavior in Capacitors Although the electrolytic capacitors used in ICDs have performance characteristics well suited to the ICD application, their performance does deviate from ideal in ways that can be clinically significant.

Feedthroughs

Anode/cathode stack

Cathodes

Paper wrap

Papers Anodes

Fill port Figure 6-21  Cutaway view of stacked-plate aluminum electrolytic capacitor used in current implantable defibrillators. Anode and cathode layers are cut to a specific geometry and are stacked upon one another. Anode and cathode terminals exit the welded capacitor housing through feedthroughs. The capacitor encasement is laser welded. Note that four layers of anode material are used for each layer of cathode material (some designs use up to five layers). The entire capacitor stack is saturated with electrolyte (not shown).



6  Power Systems for Implantable Pacemakers, Cardioverters, and Defibrillators

189

Capacitor Failure Modes

Encasement Cathode

Capacitor failure modes depend greatly on the specific conditions of device use. The failure modes most relevant to the ICD application are discussed here.

Separator

Short Circuit.  Short circuit occurs when the electrodes at opposite polarity come into direct electrical contact. In electrolytic capacitors, a short circuit occurs when the anode and cathode touch, such as when the separator is punctured, torn, or improperly aligned. Short circuit may result in the total failure of the capacitor or only partial loss of function. If the capacitor short-circuits while it is charged, its energy can quickly be discharged, which can irreversibly harm the device. If the short circuit occurs when the capacitor is uncharged, it will not be able to be charged, and the full energy of the therapy will not be delivered.

Encasement Cathode

Anode

Dielectric Oxide Degradation.  As previously discussed, capacitor deformation is an important ICD capacitor performance characteristic. The increases in charge time resulting from deformation are minimized by charging the device periodically to repair the microscopic oxide defects that form over time. In rare cases where the oxide is not properly formed or contains contaminants, normal oxide degradation can be accelerated. Depending on the severity of the issue, periodic maintenance charging may not be sufficient to keep device charge time at specified levels, and in extreme cases the device may fail to charge. Manufacturers thoroughly test incoming materials and assemble products in clean environments to ensure robust and predictable performance. Future Developments

Figure 6-22  Cutaway view of a tantalum electrolytic capacitor used in current implantable defibrillators. Design incorporates a single tantalum anode pellet pressed from tantalum powder. A tantalum lead wire exits the welded housing by a feedthrough. Highcapacitance cathode material is deposited directly on the case. The entire electrode assembly is saturated with electrolyte (not shown).

Electrolytic capacitors are expected to remain the technology of choice for many future generations of ICD. Research continues on materials and processing methods to continue to improve energy density and performance. There is no clear path to integration of alternate capacitor technologies, such as ceramic, polymer, or double layer, but new materials and processes are being investigated that might provide smaller size and improved performance characteristics.

resistance of about 2 to 4 ohms, which results in about 5% energy loss when a capacitor is discharged into a typical 50-ohm therapeutic load. Resistance is accounted for in the design of the capacitor and ICD so that the desired energy is actually delivered during a defibrillation therapy.

Effects of Batteries and Capacitors on Defibrillation Performance

60% Photoflash capacitor performance

Percent deformation

50% 40% 30% 20%

Stacked-plate capacitor performance

10% 0%

0

2

4

6

8

Time since last charge (months) Figure 6-23  Percent deformation. For typical photoflash capacitor used in early ICDs and typical stacked-plate performance. Percent deformation is percentage of additional energy required to charge a capacitor after some period at open-circuit voltage.

ENERGY LOSSES IN DEFIBRILLATORS The processes of charging the capacitors and discharging the energy through the leads of a cardiac defibrillator have numerous inefficiencies, which result in a substantial disparity between the chemical energy consumed in the battery and the electrical energy delivered to the heart. Besides capacitor deformation, leakage current, and capacitor resistance, as previously discussed, other losses are inherent in the design of the device-charging circuit and leads. The most significant loss of energy is associated with the battery. Batteries in ICDs are designed to be as small as possible, which means that they tend to be operated near their maximum power capability. An ideal power source operates at maximum power when the load placed on the power source matches the internal resistance of the power source. This well-known concept is often referred to as “impedance matching.” In terms of energy efficiency, it means that when a power source is operating at maximum power, only half of the total energy consumed is delivered to the external load. The remainder of the energy is dissipated as heat inside the ICD. Considering all these contributions, the overall efficiency of the battery/charging-circuit/ capacitor/lead system for current ICDs is less than 50%. To put this energy consumption in clinical perspective, the energy content of a typical ICD battery is on the order of 15,000 J. A system operating at 50% efficiency would consume 70 J from the battery, or about 0.5% of the total energy. This equates to about 10 to 15 days of longevity reduction for each 35-J shock delivered.

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SECTION 1  Basic Principles of Device Therapy

CLINICAL IMPLICATIONS OF DESIGN When comparing the performance of ICDs, many principal parameters are interrelated, and optimizing one may degrade another. The clinician should be aware of these issues to develop realistic expectations for device performance. Four principal characteristics of an ICD interact, primarily because of the battery and the defibrillation capacitors: size, longevity, charge time, and defibrillation energy. Smaller ICDs may be desirable, but all else being equal, small size trades off longevity, charge time, and maximum shock energy. Higher maximum shock energy will extend charge time and reduce longevity. The ICD manufacturer must strike an appropriate balance between ICD volume and clinically important parameters such as longevity, charge time, and maximum shock energy. The clinician should also be aware that charge times are generally specified assuming formed capacitors. Since deformed capacitors

consume additional energy in charging, charge times for devices that have not been recently charged can be up to 25% longer for the first defibrillation therapy delivered. Specific device use conditions also have a significant impact on observed device performance. As discussed earlier, the energy removed from the battery during each charge of the output capacitor represents a significant fraction of its total energy, on the order of 0.5% to 1.0%. Thus the longevity of a defibrillator depends greatly on the frequency of the shock therapy. This is further complicated by single-, dual-, or triple-chamber pacing conditions. For example, a device used only as a defibrillator that delivers few therapeutic shocks may last more than 13 years, whereas the same device may only last 3 to 5 years if implanted in a patient who requires significant triplechamber pacing or if many shocks are delivered. Longevity differences can be a factor of three or greater depending on how the device is used.

REFERENCES 1. Brennen KR, Fester KE, Owens BB, Untereker DF: Pacemaker battery capacity: a consideration of the manufacturer’s problem. Proceedings of the VIth World Symposium on Cardiac Pacing, Montreal, 1979, Pacesymp. 2. Brennen KR, Fester KE, Owens, BB, Untereker DF: A capacity rating system for cardiac pacemaker batteries. J Power Sources 5:25, 1980. 3. Broadhead J: Electrochemical principles and reactions. In Linden D, editor: Handbook of batteries and fuel cells, ed 2, New York, 1994, McGraw-Hill. 4. Holmes CF: An approach to the reliability of implantable lithium batteries. J Power Sources 26:185, 1989. 5. Schmidt CL, Skarstad PM: Impedance behavior in lithium iodine batteries. In Abraham KM, Solomon M, editors: Proceedings of the Symposium on Primary and Secondary Lithium Batteries, PV91 3, 1991, Electrochemical Society, pp 75-85. 6. Schmidt CL, Skarstad PM: Development of a physically based model for the lithium iodine battery. In Keily T, Baxter BW, editors: Power sources 13, Leatherhead, England, 1991, International Power Sources Committee, pp 347-361. 7. Untereker DF, Owens BB: Microcalorimetry: a tool for the assessment of self-discharge processes in batteries. Reliability Technology for Cardiac Pacemakers II, Gaithersburg, Md, 1977, National Bureau of Standards Workshop.

8. Untereker DF: The use of a microcalorimeter for analysis of load dependent processes occurring in a primary battery. J Electrochem Soc 125:1907, 1978. 9. Owens BB, editor: Batteries for implantable biomedical devices. New York, 1986, Plenum. 10. Antonioli G, Baggioni F, Consiglio F, et al: Stimulatore cardiaco implantible con nuova battaria a sato solido al litio. Minerva Med 64:2298, 1973. 11. Weiss DJ, Cretzmeyer JW, Crespi AM, et al: Electrochemical cells with end-of-service indicator. US Patent 5:180, 642, 1993. 12. Schmidt CL, Skarstad PM: The future of lithium and lithium-ion batteries in implantable medical devices. J Power Sources 97/98:742-746, 2001. 13. Takeuchi ES, Quattrini JP: Batteries for implantable defibrillators. Med Electronics 119:114-117, 1989. 14. Liang CC, Bolster E, Murphy RM: Metal oxide cathode material for high energy density batteries. US Patents 4:310, 609_1982 and 4,391,729_1983. 15. Holmes CF, Keister P, Takeuchi E: High rate lithium solid cathode battery for implantable medical devices. Prog Batteries Solar Cells 6:64, 1987. 16. Crespi AM: Silver vanadium oxide cathode material and method of preparation. US Patent 5:221, 453, 1993.

17. Crespi A, Schmidt C, Norton J, et al: Modeling and characteristics of the resistance of lithium/SVO batteries for implantable cardioverter defibrillators. J Electrochem Soc. 148:A30-A37, 2001. 18. Medtronic. Tachyarrhythmia technical concept paper. Vol V, No 1. Marquis ICD Battery Technology. 19. Root MJ: Lithium-manganese dioxide cells for implantable defibrillator devices: discharge voltage models. J Power Sources 2008. doi:10.1016/j.jpowsour.12:083, 2009. 20. Gan H, Rubino RS, Takeuchi ES: Dual-chemistry cathode system for high-rate pulse applications. J Power Sources 146:101-106, 2005. 21. Strange TF, Graham TV: Beyond photoflash; development of a polyanode capacitor. In Proceedings of the 20th Capacitor and Resistor Technology Symposium, CARTS 2000, Huntington Beach, Calif, 2000, pp 16-21. 22. Breyen MD, Rorvick AW, Skarstad PM: Efficiently packaged aluminum electrolytic capacitors for implantable cardioverterdefibrillators. In Proceedings of the 22nd Capacitor and Resistor Technology Symposium, CARTS 2002, New Orleans, 2002, pp 197-202. 23. Norton J, Anderson C: Energy management in aluminum electrolytic capacitors: In Proceedings of the 23rd Capacitor and Resistor Technology Symposium, CARTS 2003, Scottsdale, Ariz, 2003, pp 269-277.

7 

The Biologic Pacemaker J. KEVIN DONAHUE

In October 1958, the first fully internalized pacemaker was implanted

at the Karolinska Institute in Sweden. Within hours, the unit ceased to function, or as Senning wrote in a retrospective account, “At 2 am the pacemaker became silent.”1 In a Medline search with the keywords “implantable pacemaker,” the first five citations (starting in 1960) are case reports of various implants, and the sixth (1962) is entitled “Complications of an implantable cardiac pacemaker.”2 Since that initial implant, millions of pacemakers have been implanted, saving lives and reducing or even eliminating symptoms. Along with these benefits of electronic pacemakers, however, complications occur in 5% to 10% of patients undergoing implant procedures.3,4 Acute complications range from minor hematomas to more severe infections, pneumothoraces, and cardiac perforations. Long-term management of pacemaker patients is complicated by device and lead failures and a requirement to change out the device when the battery fails. These limitations of device therapy have motivated the search for alternatives. One widely publicized approach over the last few years is to create a “biologic pacemaker” by either gene transfer or cell transplantation methods (Table 7-1). This chapter reviews the literature on this approach.

Gene Therapy Approaches BETA-2 ADRENERGIC RECEPTOR GENE TRANSFER In the first attempt at genetic modulation of heart rate, Edelberg et al.5 injected plasmids containing the beta-2 adrenergic receptor (β2-AR) into mouse right atria. They found expression of the vector in 81% of myocytes at the injection site and a heart rate of 550 beats per minute (bpm) for the active-treatment group, significantly higher than the 370 bpm rate in controls. In a subsequent pig study, these investigators did not quantify gene transfer but reported a heart rate of 163 bpm for the β2-AR–receiving pigs relative to 127 bpm for controls.6 These studies showed promising results but also raised questions. The investigators did not report the time course of heart rate increase for the mice, but in the pigs they found a statistically significant increase in heart rate only on the second day after injection. The usual experience with plasmid-mediated gene transfer is that the effect persists for several weeks or even months after gene transfer.7-9 The transient effect for the β2-AR experiments suggests that some toxic effect from expression of this gene may be a limiting factor. Additional concerns involve possible mechanisms for the increase in heart rate. The mouse study did not target the specialized conducting system, and under ordinary circumstances, atrial myocytes should not display automaticity. A potential mechanism for the heart rate increase from β2-AR expression in atrial myocytes is Gs-mediated activation of adenylate cyclase, causing an intracellular cascade that ultimately results in phosphorylation of the type 1 calcium channel and the ryanodine receptor. Increased activity of these calcium-handling proteins could cause automaticity by the calcium clock mechanism, but sustained increase in intracellular calcium would likely be toxic. The porcine study did target sinus node by injecting in the area electrically mapped as the earliest site of activation during sinus rhythm, but the extremely transient functional change in pigs suggests that toxicity may be a concern even in sinus nodal tissues. An additional concern from the figures demonstrating gene expression is that the green from the green fluorescent protein in controls and the red from β2-AR immunostaining in the active-treatment group

are apparently extracellular. These concerns should probably be addressed before considering translation of the β2-AR findings. IK1 KNOCKOUT In the next illustration of heart rate modulation, investigators used a dominant negative mutation of the gene encoding the IK1 channel (KCNJ2 GYG144-146AAA) to create automaticity.10 IK1 (also called the inward rectifier current) is the principal current responsible for maintaining the cardiac myocyte’s resting membrane potential. The logic behind this strategy was that differences in IK1 between atrial and sinus nodal myocytes account for the relative instability of sinus nodal membrane potential, so that a deficit of IK1 would increase the probability of phase 4 depolarization and automaticity. The investigators admitted that their primary interest was in the electrophysiologic effects of IK1 on the action potential, and that they did not intentionally create pacemaking nodes. Their delivery method essentially scattered the transgene around the ventricular myocardium. Nonetheless, they found evidence of ventricular automaticity in 40% of animals expressing the KCNJ2-AAA mutation. Further investigation found that at least 80% of endogenous IK1 needed to be eliminated for automaticity to occur.11 There have been no subsequent reports applying this strategy to other models. No attempt was made to quantify rate or stability of the idioventricular rhythm in this report.10 Other concerns rising from the KCNJ2AAA strategy include the repolarization delaying effects of gene expression and the possibility of atrioventricular (AV) nodal suppressing effects. The investigators reported some level of action potential prolongation and increased QT interval at all expression levels of the transgene, as opposed to the pacing phenotype that was only visible with 80% suppression of IK1. Inspection of the limited reported data also shows no apparent AV nodal conduction, even in areas where the atrial beat is timed far enough away from the proceeding ventricular beat that AV nodal conduction should reasonably be expected. This finding is extrapolated from extremely limited data and should be assessed with more care if IK1 knockout becomes part of a biopacing strategy. Overall, the findings with IK1 knockout suggest that this may be a component of an overall biopacing strategy (with the previous caveats), but that the limited rate in pigs suggests that it might need to be applied with other interventions to increase automaticity. IF MANIPULATIONS Since the early attempts at pacemaker induction, almost all subsequent gene transfer approaches have involved manipulation of If, the putative pacemaker (funny) current. The first attempts consisted of intramyocardial injections of adenovirus vectors encoding wild-type HCN2. The initial report targeted the left atrium,12 and a subsequent report showed gene transfer to the left bundle branch (LBB).13 In both reports, transgene function was not apparent during normal sinus rhythm. During vagal stimulation–induced sinus arrest, the transgene-expressing animals had a significantly higher escape rhythm than controls, and this rhythm mapped to the target area. Overall, the escape rate was inadequate to be a viable pacemaker replacement, and the transgene-driven rhythm could not be seen outside of vagal stimulation, so the investigators could not comment on stability of this rhythm over time. An additional report with the HCN4 isoform suggested that automaticity was not limited to HCN2. Cai et al.14 injected adenoviruses

191

192

TABLE

7-1 

SECTION 1  Basic Principles of Device Therapy

Results of Molecular Therapy Experiments to Induce Cardiac Automaticity

Gene Therapy Transgene [Reference] β2-Adrenergic receptor [6]

Vector Plasmid

Target Right atrium

Results 50% increase in HR

KCNJ2 GYG144-146AAA [10] HCN1 EVY235-237ΔΔΔ [16]

Adenovirus

Whole heart

Adenovirus

Left atrium

Observed ventricular automaticity Average HR 64 bpm in sick sinus syndrome model

HCN2 wild type [12]

Adenovirus

Left atrium

[13, 15]

Adenovirus

LV left bundle branch

HCN2 E324A [15]

Adenovirus

HCN212 [17]

Adenovirus

HCN4 wild type [14]

Adenovirus

LV left bundle branch LV left bundle branch Right ventricle

Cell Therapy Cell Type [Reference] Fetal atrial myocytes [18]

Manipulation None

Human embryonic stem cells [19] Human mesenchymal stem cells [20, 21] Guinea pig lung fibroblasts [22]

Problems Effect only seen for 1 day; rhythm not mapped to gene transfer site QT prolongation, no attempt to quantify rate or stability Only 2 animals per group; competing electronic pacemaker DDD 60 bpm limits analysis of biopacing rhythm; limited data Long pause before onset of escape rhythm; no visible phenotype during sinus rhythm; no data on durability Long pause before onset of escape rhythm; no visible phenotype during sinus rhythm (visible with AV block); required backup pacing 26% of the time Required VVI 45 bpm pacing 36% of the time

Escape rhythm during vagal arrest mapped to target area 54 bpm LV escape rhythm during vagal arrest, average HR 57 bpm after AV node ablation Average HR 53 bpm after AV node ablation All animals developed VT

All animals developed VT

Average HR 69 bpm after AV node ablation

No data on rate stability; animals only followed for 24 hours

Target Left ventricle

Results Average HR 70 bpm after AV node ablation

CM derivation and isolation of beating cells HCN2 gene transfer

Left ventricle

Average HR 59 bpm after AV node ablation

Problems Only 2 animals per group; long pause before stable biopacing; no long-term rhythm stability data; questionable feasibility for translation Stable rhythm seen in only half of study animals; no long-term rhythm stability data

Left ventricle

Average HR 52-61 bpm after AV node ablation

HCN1 gene transfer, co-infusion with PEG5000 for cell fusion

Left ventricle

Observed ventricular escape beats during methacholineinduced sinus arrest

Long pause before stable biopacing; after AV node ablation, animals still had 25% of beats from VVI 35 electronic pacemaker No stable biopacing

AV, Atrioventricular; HR, heart rate; LV, left ventricular; VT, ventricular tachycardia.

encoding either wild-type HCN4 with green fluorescent protein (GFP) or GFP alone into porcine ventricles.14 They performed AV nodal ablation 3 to 4 days after gene transfer and measured the residual ventricular rate. The HCN4-expressing animals had a ventricular escape rate of 69 bpm, whereas the rate was only 41 bpm in the GFP controls. Pace mapping of the rhythm in the HCN4-expressing animals showed that the rhythm originated from the target region. The heart rate measurement was essentially an acute study, so no data on rate stability were provided. Bucchi et al.15 tested the hypothesis that mutations of the HCN channel could improve automaticity. They compared wild-type HCN2 to HCN2-E324A. This mutation was chosen because biophysical analysis suggested it would increase channel activation. The mutation alters the voltage dependence of activation so that the channel becomes active at less negative membrane potential, thus increasing the rate of channel opening during cellular repolarization. The authors hypothesized that this increase in If would lead to faster phase 4 depolarization and thus a faster heart rate. Events did not proceed as planned. The investigators injected adenoviruses encoding the wild-type HCN2, the HCN2-E324A mutant, or GFP into the LBB of dogs. They implanted electronic pacemakers, ablated the AV node, and watched. They saw no significant increase in heart rate when comparing the wild-type and mutant channels (both were better than GFP control), and they noted a greater need for the electronic pacemaker backup in animals expressing the HCN2 mutant than the wildtype HCN2. To investigate this discrepancy with expectations, the investigators found that the mutant channel did, indeed, have the predicted kinetics, but it also had a decrease in peak current caused by a decrease in expression compared with the wild-type channel. The investigators did not define the cause for the decreased expression, but presumably it was cellular toxicity to the mutant channel because the virus vector and DNA control elements were the same for the wild-type and mutant HCN2 genes.

In addition to finding that behavior of mutant channels involves factors other than the predicted biophysical profile, the AV node ablation allowed Bucchi et al.15 to comment on the long-term behavior of the biopacing node. Two important findings were the inadequacy of both pacing rate and rhythm stability after gene transfer. As noted, the average rate of animals receiving wild-type HCN2 was only 57 bpm, which is considerably lower than the normal canine heart rate. Even with this average heart rate of 57 bpm, the wild-type HCN2–expressing animals had frequent episodes when biopacemaking activity was lost, requiring the backup electronic pacemaker 26% of the time. The electronic pacemaker was programmed VVI 45 bpm, so drop-out of the biopacemaker was not caused by overdrive suppression from the electronic pacemaker. The cause was certainly failure of the biopacing node, but the exact mechanism was not reported; loss of automaticity and exit block were likely possibilities. Tse et al.16 reported another strategy to improve pacing rate with a mutation of the HCN1 isoform. HCN1-EVY235-237ΔΔΔ was a deletion mutation that shortened the S3-4 linker of the channel, with a predicted biophysical effect of increased probability of channel opening. Using a porcine model, the investigators ablated enough of the sinus node to cause sick sinus syndrome, implanted an electronic dual-chamber pacemaker programmed to DDD mode with a lower rate limit of 60 bpm, and then injected either an adenovirus vector encoding the HCN1 mutant or saline into the left atrial appendage. There were only two animals in each group. After gene transfer, the active-treatment group had a spontaneous rate of 64 bpm but still required the electronic pacemaker 14% of the time, which is difficult to interpret given the programmed lower rate limit of 60 bpm and the observed average spontaneous rate of 64 bpm. There is little room for movement below the average rate, although the only 14% use of electronic pacing suggests that the heart rate could not have varied far from the average, or one would have expected a higher percentage of pacing. No rate or rhythm data from the two saline-injected animals were



7  The Biologic Pacemaker

provided. Given the limitations of this report, it is difficult to interpret it in the overall context of biopacing strategies. The danger of HCN mutation was illustrated by Plotnikov et al.,17 who reported ventricular tachycardia in dogs after injection of a chimeric HCN channel that had components of both HCN1 and HCN2 isoforms. One day after injection of adenoviruses encoding this chimeric channel into LBB, they noted ventricular tachycardia (VT) at rates in excess of 200 bpm in all animals. They further connected the arrhythmia with the gene transfer by showing VT elimination after exposure to ivabradine, an If-blocker with relative HCN1 specificity. Overall, gene transfer with HCN-family genes suggests that automaticity is achievable, but concerns about baseline rate, stability of heart rate over time, and most importantly, assessment for proarrhythmia need to be addressed before this strategy becomes viable.

Cell Therapy Strategies The first report of cell transplantation for cardiac pacing used a strategy of first isolating canine fetal atrial myocytes, “including sinus node,” and then injecting these cells into the ventricles of adult dogs.18 The investigators did not attempt to identify or isolate a spontaneously beating subpopulation, choosing instead to inject the mixed population from cell isolation. Two dogs received the cardiomyocytes, and two control dogs received fibroblast injections. Three weeks after cell injections, the investigators performed AV node ablation and observed the ventricular rate. Initially after ablation, all animals had a nodal escape of approximately 35 bpm. After 5 to 30 minutes, the two dogs receiving atrial myocytes developed a ventricular escape rhythm at 70 bpm that mapped to the cell injection site. The two control animals had no alteration in rhythm beyond the initial 35-bpm nodal escape. Further study showed that the ventricular rhythm had an escape recovery time of 710 msec after 150-bpm burst pacing, which suggested some level of stability for the rhythm. As with many of the original gene therapy pacing studies, this phenotyping study was acute, and long-term rhythm stability was not assessed. Several practical problems limit translation of these findings, primarily the need to isolate cells from third-trimester fetuses and the need for information on long-term stability of the cells and the rhythm. Immunologic reaction to the cell transplant is another consideration. Overall, however, this report established the feasibility of cell transplant pacemakers. Similar work showed viability of embryonic stem cell (ESC)–derived cardiomyocytes for pacing. Kehat et al.19 used methods frequently reported in the field to cultivate spontaneously beating cardiomyocytes from undifferentiated ESCs. They first allowed the human cells to grow with specialized culture medium overlying a “feeder layer” of neonatal mouse fibroblasts. The resulting cell population developed into “embryoid bodies,” with suspension culture again using specialized medium. Spontaneously beating cells were mechanically dissected away from the embryoid bodies and injected into pigs after AV node ablation. The 13 pigs that received cell therapy had backup electronic pacemakers in this uncontrolled study. Shortly after AV node ablation, the animals had only a nodal escape rhythm. At electrophysiologic study 1 to 3 weeks after cell injection/ablation, two of the animals still only had a nodal rhythm, five had isolated monomorphic premature ventricular contractions (PVCs) or short runs of ventricular rhythm, and six had sustained ventricular rhythm that mapped to the cell injection site. The average rate of the ventricular rhythm (59 bpm) was competitive with the nodal escape rhythm (61 bpm), so the investigators were unable to comment on rhythm stability or long-term viability of this approach. These issues are relevant when contemplating translation of the ESC results, in addition to ethical concerns voiced by some about the use of human ESCs, and the need for extensive manipulation of the cells, which would not be readily scalable. Potapova et al.20 reported a novel hybrid cell and gene therapy approach. They transduced human mesenchymal stem cells (MSCs) with the HCN2 gene, then transplanted these cells into canine left ventricle. Their hypothesis was that the HCN2-expressing cell would couple with nearby ventricular myocytes and essentially act as a driver

193

for phase 4 depolarization in the myocytes. With sufficient connectivity, the myocyte would transmit its membrane potential to the MSC during repolarization; this would activate If in the MSC, and that current would be transmitted across the gap junctions to the ventricular myocyte, causing phase 4 depolarization and thus automaticity in the two-cell unit. The investigators were motivated to use the cell therapy approach by concerns about dangers in long-term expression gene transfer vectors that would be needed to transduce the ventricular myocytes directly in vivo, although they admitted that these concerns had not yet been demonstrated in animal or human studies using adeno-associated virus. Potapova et al.20 compared four control animals to six animals receiving the HCN2-transduced MSCs. As in their initial HCN2 gene transfer studies, they saw no spontaneous rhythms, so they used strong vagal stimulation to induce sinus arrest. All animals had nodal or ventricular escape rhythms during the sinus arrest period. Of the four control animals, two had right ventricular (RV) escape rhythms and 2 had left ventricular (LV) escape rhythms. The average escape rate for the controls was 45 bpm. Five of the six HCN2-MSC animals had LV escape rhythms, and the average rate of 61 bpm was significantly higher than in the controls. The investigators mapped the escape to the target site in the HCN2-MSC animals. In a subsequent study, Plotnikov et al.21 tested the dose-response for HCN2-transduced MSC injections.21 They evaluated animals receiving 155,000 to 850,000 cells. In this study the researchers performed AV nodal ablation with electronic pacemaker implantation (VVI 35 bpm) for long-term follow-up of the escape rhythm. In all animals, spontaneous rhythms became apparent after 2 to 3 days, but this rhythm remained unstable for at least 2 weeks. For animals receiving less than 600,000 cells, the rhythm never stabilized, and the average rate was 43 bpm. After 2 weeks the animals receiving 600,000 to 850,000 cells had a relatively stable rhythm at an average rate of 52 bpm for the 42-day duration of the study. With linear regression analysis, the investigators showed a correlation between rate and number of injected cells for the higher-dose animals. Even with this level of biopacing, however, 25% of observed beats originated from the VVI 35 bpm electronic pacemaker. A further concern was that only 70% of the ventricular beats pace-mapped to the injection site in the high-dose group, although this was an improvement over the 10% mapped to the target site in the low-dose group. Cho et al.22 reported a variation on the HCN-transduced cell transplant approach. They transduced fibroblasts with the HCN1 isoform and then used polyethylene glycol 5000 (PEG5000) to induce fusion of the HCN1-expressing fibroblasts with ventricular myocytes. This largely in vitro study showed viability of the cell fusion concept and automaticity of the resulting fibroblast-myocyte complex. Very limited in vivo data in guinea pigs injected with the PEG5000/transduced fibroblast mixture showed occasional ventricular escape beats coming from the target region after methacholine-induced sinus arrest. The investigators hypothesized that an additional factor beyond the HCN1 expression was the increased cell surface area of the fibroblast/myocyte complex that diluted IK1 from the myocyte by approximately 20%. An important in vitro investigation by Fahrenbach et al.23 showed improved pacing with suppression of connexin 43 (Cx43) expression. These investigators used ESCs isolated from Cx43 knockout transgenic mice. The cells still expressed connexins 40 and 45, so some cellular connectivity was possible, but the overall level of connectivity was reduced. The investigators cultured spontaneously beating ESCderived cardiomyocytes from either wild type of Cx43−/− mice with HL-1 cells (an immortalized mouse atrial tumor cell line). The Cx43 knockout cells required a longer time to connect to the underlying HL-1 cells, but once connectivity was established, the Cx43−/− cells had a much higher probability of becoming the dominant pacing node for the culture dish. These results are consistent with the observations of reduced cell connectivity in histologic studies of sinus node tissues. The concept from sinus node studies is that limited cellular connectivity allows the rhythm to arise gradually in an isolated node, without suppression from the strong polarizing influence of the surrounding myocytes.

194

SECTION 1  Basic Principles of Device Therapy

Based on the Fahrenbach et al.23 results, some rhythm stability problems in the previously noted gene and cell therapy examples could be caused by too much connectivity between the automatic cells, and the surrounding myocytes could be suppressing phase 4 depolarization in the biopacing node.

Conclusion Given the overwhelming success of electronic pacemakers, the efficacy and safety of biologic pacemakers will need to be well established before

translation to clinical practice is contemplated. Currently, both gene transfer and cell transplantation strategies indicate that automaticity can be induced after a variety of manipulations. Problems common to all investigated strategies (where data are available to evaluate) include drop-off of automaticity resulting in either asystole or activation of a backup electronic pacemaker, inadequacy of baseline rate, and demonstration of long-term rhythm stability. Still, these are very early studies; the challenges are there, but none of the identified problems is insurmountable. Over the next several years, we can anticipate continued research in this field and likely solutions to these problems.

REFERENCES 1. Senning A: Cardiac pacing in retrospect. Am J Surg 145: 733-739, 1983. 2. Appelbaum I, Parsonnet V, Sonnet V, et al: Complications of an implantable cardiac pacemaker. J Newark Beth Isr Hosp 13:166174, 1962. 3. Pakarinen S, Oikarinen L, Toivonen L: Short-term implantationrelated complications of cardiac rhythm management device therapy: a retrospective single-centre 1-year survey. Europace 12:103-108, 2010. 4. Parsonnet V, Bernstein A, Lindsay B: Pacemaker-implantation complication rates: an analysis of some contributing factors. J Am Coll Cardiol 13:917-921, 1989. 5. Edelberg J, Aird W, Rosenberg R: Enhancement of murine cardiac chronotropy by the molecular transfer of the human beta-2 adrenergic receptor cDNA. J Clin Invest 101:337-343, 1998. 6. Edelberg J, Huang D, Josephson M, et al: Molecular enhancement of porcine cardiac chronotropy. Heart 86(5):559-562, 2001. 7. Perales JC, Ferkol T, Beegen H, et al: Gene transfer in vivo: sustained expression and regulation of genes introduced into the liver by receptor-targeted uptake. Proc Natl Acad Sci USA 91:40864090, 1994. 8. Muramatsu T, Arakawa S, Fukazawa K, et al: In vivo gene electroporation in skeletal muscle with special reference to the duration of gene expression. Int J Mol Med 7:37-42, 2001.

9. Ziady AG, Kim J, Colla J, et al: Defining strategies to extend duration of gene expression from targeted compacted DNA vectors. Gene Ther 11:1378-1390, 2004. 10. Miake J, Marban E, Nuss H: Biological pacemaker created by gene transfer. Nature 419:132-133, 2002. 11. Miake J, Marban E, Nuss H: Functional role of inward rectifier current in heart probed by Kir2.1 overexpression and dominantnegative suppression. J Clin Invest 111:1529-1536, 2003. 12. Qu J, Plotnikov A, Danilo PJ, et al: Expression and function of a biological pacemaker in canine heart. Circulation 107: 1106-1109, 2003. 13. Plotnikov A, Sosunov E, Qu J, et al: Biological pacemaker implanted in canine left bundle branch provides ventricular escape rhythms that have physiologically acceptable rates. Circulation 109:r31-r37, 2004. 14. Cai J, Yi FF, Li YH, et al: Adenoviral gene transfer of HCN4 creates a genetic pacemaker in pigs with complete atrioventricular block. Life Sci 80:1746-1753, 2007. 15. Bucchi A, Plotnikov A, Shlapakova I, et al: Wild-type and mutant HCN channels in a tandem biological-electronic cardiac pacemaker. Circulation 114:992-999, 2006. 16. Tse H, Xue T, Lau C, et al: Bioartificial sinus node constructed via in vivo gene transfer of an engineered pacemaker HCN channel reduces the dependence on electronic pacemaker in a sick-sinus syndrome model. Circulation 114:1000-1011, 2006.

17. Plotnikov AN, Bucchi A, Shlapakova I, et al: HCN212-channel biological pacemakers manifesting ventricular tachyarrhythmias are responsive to treatment with I(f) blockade. Heart Rhythm 5:282-288, 2008. 18. Ruhparwar A, Tebbenjohanns J, Niehaus M, et al: Transplanted fetal cardiomyocytes as cardiac pacemaker. Eur J Cardiothorac Surg 21:853-857, 2002. 19. Kehat I, Khimovich L, Caspi O, et al: Electromechanical integration of cardiomyocytes derived from human embryonic stem cells. Nat Biotechnol 22:1282-1289, 2004. 20. Potapova I, Plotnikov A, Lu Z, et al: Human mesenchymal stem cells as a gene delivery system to create cardiac pacemakers. Circ Res 94:952-959, 2004. 21. Plotnikov AN, Shlapakova I, Szabolcs MJ, et al: Xenografted adult human mesenchymal stem cells provide a platform for sustained biological pacemaker function in canine heart. Circulation 116:706-713, 2007. 22. Cho HC, Kashiwakura Y, Marban E: Creation of a biological pacemaker by cell fusion. Circ Res 100:1112-1115, 2007. 23. Fahrenbach JP, Ai X, Banach K: Decreased intercellular coupling improves the function of cardiac pacemakers derived from mouse embryonic stem cells. J Mol Cell Cardiol 45:642-649, 2008.

8 

Pacemaker, Defibrillator, and Lead Codes and Headers ROBERT ANDREW PICKETT  |  GEORGE H. CROSSLEY III

In the continuing pursuit of therapy for the arrhythmia patient,

tremendous strides have been made in pacing and defibrillation platforms. Along with these advances, much standardization has been introduced to limit differences and disparities among device types and to allow for compatibility. In the 1960s, pacemakers served to stimulate one chamber at a fixed rate. By the 1980s, implantable cardioverter-defibrillators (ICDs) offered lifesaving therapy by way of shock only, without supportive pacing for bradycardia. However, many patients require both antibradycardia and antitachycardia therapies. Later, cardiac resynchronization therapy (CRT) for congestive heart failure was developed and has become the standard of care for these patients. With each of these advances in device therapy, the traditional means of describing device function have become inadequate, requiring new standardized terminology. Devices are used for more indications in the treatment of patients with bradyarrhythmias, tachyarrhythmias, and congestive heart failure, and more algorithms are developed either to stimulate multiple chambers (cardiac resynchronization) or to minimize unnecessary pacing when possible.1,2 It is noteworthy that, unlike some therapies, the generic pacemaker codes have withstood the test of time.

Pacemaker Codes THREE-POSITION ICHD CODE (1974) In 1974 the Inter-Society Commission for Heart Disease Resources (ICHD) proposed a three-position code to allow a uniform system to describe the functionality of a pacemaker platform.3 The goal was to create a generic code that would work among all platforms; it would also allow for concise description of what was becoming a more complex issue. The first letter of the code was to distinguish the chamber(s) being paced. The second letter was to determine the chamber(s) that were sensed for intrinsic electrical depolarizations and to start the timer for determining the next paced beat delivery. The third letter of the code was to describe how the pacing was to be affected by a sensed event (i.e., inhibiting pacing or triggering pacing in response to sensing). Position one can be the atrium (A), the ventricle (V), or both atrium and ventricle (D) for location of pacing. Position two again represents atrium (A), the ventricle (V), or both (D), meaning sensing can occur for spontaneous electrical depolarizations in either chamber. The third position describes whether sensing leads to inhibition of pacing (I), triggering the delivery of a pacing stimulus (T), or both inhibition and triggering (D). Tables 8-1 and 8-2 outline the options and different pacing configurations. The outstanding feature of this code is its utility in describing pacemaker function despite all the therapeutic advances since its introduction. Changes have been proposed, but this code endures with the addition of a letter to indicate rate-responsive pacing. FIVE-POSITION ICHD CODE (1981) In 1981, in response to changes in the technology of pacing, two more positions were added. The first three positions remained unchanged, and position four was developed to convey the ability to program the pacemaker. With advancements in pacing, it became important to determine if a pulse generator could be programmed. As the ability to

reprogram the pacing rate, pacing output, and mode were introduced into pacemakers, the fourth position could be listed as programmable (P), which generally implied one feature, or multiprogrammable (M), which implied multiple parameters, or none (O). A fifth letter was included to determine if the device had the capability of antitachycardia pacing. Interestingly, this fifth letter was used in the pacing platform and was not for ICD platforms, which were not available at the time this code was developed. A second modification of the ICHD nomenclature occurred in 1983 to distinguish, in the fourth letter, if a device had the ability to internally store data such as battery voltage, serial numbers, or electrograms. This was designated communicating (C). Table 8-3 shows the different options for all five positions, but as one would surmise, it is now assumed all devices can store information and can be programmed, thus obviating the need for that nomenclature. Currently, pacing platforms that provide antitachycardia pacing are typically identified based on the specifics of the device platform rather than an additional letter in the code (e.g., Medtronic EnRhythm pacing platform allows atrial antitachycardia pacing therapy). Using this convention, a single-chamber pacing system today would be described as VVIM to designate multiprogrammability or VVIC to designate the presence of telemetry (communication). In reality, this fourth letter is implied due to the presence of telemetry and multiprogrammability in all of our current platforms. Table 8-4 provides examples of the many different ways to describe pacemaker function that are highly descriptive of the available features on pacing systems, but not readily used in day-to-day management of pacemakers. NASPE/BPEG GENERIC (NBG) PACEMAKER CODE The Mode Code Committee of the North American Society of Pacing and Electrophysiology (NASPE), now known as the Heart Rhythm Society, in cooperation with the British Pacing and Electrophysiology Group (BPEG), developed a code in 1987 (Table 8-5).4 The one change still used today is the letter R in the fourth position to represent ratemodulated pacing to treat chronotropic incompetence. The R in the fourth position was developed in response to the many sensors available to help the pacemaker mimic sinus nodal function in response to external or internal stimuli. The other development in the nomenclature is the use of the letter S for shock delivery for tachyarrhythmias with the fifth position in the code. The NASPE/BPEG Generic (NBG) Pacemaker Code retains all the characteristics of the 1974 ICHD Code and some of those of the later five-position codes. To deal with the increasing complexity of later devices, however, it was considered necessary to clarify several features and to define more explicitly how the code was to be used. Positions I, II, and III were still used exclusively for antibradycardia pacing, but position IV had the option of describing either programmability or rate-responsive pacing. With establishment of programmability in almost all devices, position IV evolved to denote the presence or absence of rate-response pacing. Position V is for describing the ability to deliver antitachycardia therapy, as either antitachycardia pacing or a shock. As of 1987, shocks were described as either low-energy cardioversions or high-energy defibrillation. No distinction was made between low-energy and high-energy shocks in the nomenclature. The original code of 1981 allowed for different letters to describe different

195

196

TABLE

8-1 

SECTION 1  Basic Principles of Device Therapy

The 1974 Three-Position ICHD Code or Three-Letter Identification Code*

First Letter Chambers paced

Second Letter Chambers sensed

Third Letter Mode of response

*Letters used: A, atrium; D, double-chamber; I, inhibited; O, not applicable; T, triggered; V, ventricle. Adapted from Parsonnet V, Furman S, Smyth NPD: Implantable cardiac pacemakers: Status report and resource guideline. Pacemaker Study Group, Inter-Society Commission for Heart Disease Resources (ICHD). Circulation 50:A-21, 1974. Copyright 1974 American Heart Association.

forms of pacing trains used in tachycardia termination, such as burst (B), scanning (S), or external (E). The update of 1987 simplified the descriptors to pacing (P), shock (S), or pacing and shocks (D).4,5 REVISED NBG CODE In light of evolving pacemaker technology and increasing interest in multisite pacing, the 1987 NBG Code was modified by a multinational task force under the chairmanship of David Hayes, MD. Three major changes were made that simplified the code. It was assumed that all pacemakers were communicative and programmable, eliminating the need for those codes. Further, this revised code established a system to describe the function of ICDs. The structure of the revised NBG Code is summarized in Table 8-6, with examples of its use in Table 8-7.6 Note that with each modification, the first three positions continue to describe antibradycardia pacing, but in this update, no mention of antitachycardia pacing is made.7 Positions I, II, and III indicate the chambers in which pacing and sensing occur and the effect of sensing on the triggering or inhibition of subsequent pacing stimuli. In this context, “sensing” refers specifically to the detection of spontaneous cardiac depolarizations (or spurious interference signals interpreted as spontaneous cardiac depolarizations) outside the refractory periods of the pulse generator. Position IV is now used only to indicate the presence (R) or absence (O) of an adaptive-rate mechanism (rate modulation). This position is also used to denote that changes can occur in pacing parameters, such as changes in the atrioventricular (AV) delay in response to increased heart rates. Position V indicates whether multisite pacing, as described earlier, is present in none of the cardiac chambers (O); in one or both of the atria (A), with stimulation sites in each atrium or more than one stimulation site in either atrium; in one or both of the ventricles (V), with stimulation sites in both ventricles or more than one stimulation site in either ventricle; or in dual chambers (D), in one or both of the atria and in one or both of the ventricles. Resynchronization pacing is often used to treat congestive heart failure in patients with a wide QRS, but the device is rarely described using a fifth letter to denote ventricle (V) as the site of pacing in both ventricles. In reality, a biventricular pacemaker with rate-responsive pacing “on” would be categorized as DDDRV pacing in both atrium and ventricle, sensing in atrium and ventricle, inhibition or triggered activity dependent on intrinsic activation rates, with rate-responsive pacing and the presence of pacing leads in the right and left ventricle. TABLE

8-3 

TABLE

8-2 

Mode VOO AOO DOO VVI VVT AAI AAT VAT DVI

Pacing Modes Described by the Three-Position ICHD Code Description Asynchronous ventricular pacing; no sensing function Asynchronous atrial pacing; no sensing function Dual-chamber (AV-sequential) asynchronous pacing; no sensing function Ventricular pacing inhibited by ventricular sensing Ventricular pacing triggered instantaneously by ventricular sensing Atrial pacing inhibited by atrial sensing Atrial pacing triggered instantaneously by atrial sensing Ventricular pacing triggered after a delay by atrial sensing Dual-chamber (AV-sequential) pacing inhibited by ventricular sensing

Modified from Parsonnet V, Furman S, Smyth NPD: Implantable cardiac pacemakers: status report and resource guidelines. Pacemaker Study Group, Inter-Society Commission for Heart Disease Resources (ICHD). Circulation 50:A21, 1974. Copyright 1974 American Heart Association. AV, Atrioventricular.

As multisite pacing and physiologic pacing with leads in specific positions (e.g., Bachman bundle) are further investigated, use of the fifth letter may become more prevalent.8,9

Defibrillator Codes NASPE/BPEG DEFIBRILLATOR (NBD) CODE On January 23, 1993, the NASPE (now Heart Rhythm Society) Board of Trustees approved the adoption of the NASPE/BPEG Defibrillator (NBD) Code (Table 8-8).10 It was developed by the NASPE Mode Code Committee, composed of members of NASPE and BPEG, and was intended for describing cardiac-defibrillator capabilities and operation. The resultant code for ICDs is similar in form to the longestablished pacing code. Although the pacing code described antibradycardia pacing in detail and then acknowledged the presence of antitachycardia therapy, the ICD code describes the antitachycardia therapy in detail and then acknowledges the presence of antibradycardia pacing. No indication is made to distinguish between low-energy cardioversion and high-energy defibrillation. Position (potential) I denotes the chamber where the shock is delivered; position II indicates antitachycardia pacing function; position III reports how the tachycardia is detected, i.e. based on electrograms or hemodynamic changes, and position IV reports the presence of antibradycardia pacing. In current practice, given the lack of a selective atrial defibrillator, the first letter will always be V for ventricle as the location of the shock and the third letter will always be E for electrogram because all ICDs are using electrograms for determination of arrhythmia. The second position will be rarely used because antitachycardia pacing is so widely prevalent on ICDs, and with so many combinations of therapies, a single letter cannot reliably describe them. Given the state of current practice, an ICD is still described most often using the same four letters (including rate-responsive pacing) as in the pacing realm.

The 1981 Five-Position ICHD Code

Position I (Chambers Paced)

Position II (Chambers Sensed)

Position III (Modes of Response)

V = ventricle A = atrium D = double

V = ventricle A = atrium D = double O = none

S = single chamber†

S = single chamber†

T = triggered I = inhibited D = double* O = none R = reverse Comma optional here†

Position IV (Programmable Functions) P = programmable (rate and/or output) N = normal-rate competition M = multiprogrammable C = communicating O = none

Modified from Parsonnet V, Furman S, Smyth NPD. Revised code for pacemaker identification. PACE 4:400, 1981. *Triggered and inhibited response. †Manufacturer’s designation only.

Position V (Special Antitachyarrhythmia Functions) B = bursts S = scanning E = external



8  Pacemaker, Defibrillator, and Lead Codes and Headers

TABLE

8-4 

Pacing Modes Described by the Five-Position ICHD Code

Mode VDD,M (VDDM)

DDD,M (DDDM)

VVI,MB (VVIMB)

AAR,ON (AARON) AOO,OE (AOOOE)

Description Ventricular antibradycardia pacing inhibited by ventricular sensing, triggered after delay by atrial sensing; multiprogrammable device; no antitachyarrhythmia (antitachycardia) function. Dual-chamber (AV-sequential) antibradycardia pacing inhibited by sensing in either chamber, with ventricular pacing triggered after delay by sensing in atrium after ventricular event; multiprogrammable device; no antitachycardia function. Ventricular antibradycardia pacing inhibited by ventricular sensing; multiprogrammable device; pacing bursts for ventricular tachycardia, means of activation unspecified. No antibradycardia function; nonprogrammable device; normal-rate competition for termination of atrial tachycardia, activated by atrial sensing. Asynchronous atrial antibradycardia pacing; nonprogrammable device; externally activated atrial antitachycardia pacing, nature unspecified.

Modified from Parsonnet V, Furman S, Smyth NPD: Revised code for pacemaker identification. PACE 4:400, 1981. AV, Atrioventricular.

NBD CODE, SHORT FORM As an additional means of concisely distinguishing among devices limited to cardioversion or defibrillation and those that incorporate antitachycardia and antibradycardia pacing, a short-form code was defined (Table 8-9). It is intended for use only in conversation and again rarely used given the widespread prevalence of both antitachycardia pacing and shocking across platforms and manufacturers.

Lead Code NASPE/BPEG PACEMAKER-LEAD (NBL) CODE In 1996 the NASPE Board of Trustees voted to adopt a generic code for pacemaker leads, to be known as the NASPE/BPEG PacemakerLead (NBL) Code. The code was approved subsequently by BPEG, and its definition and usage conventions were published as a NASPE Policy TABLE

8-5 

197

Statement. The NBL Code is a four-position generic code intended for use in conversation, record keeping, writing, and labeling. All four positions (potentials), as defined in Table 8-10, are used in every circumstance, unlike the NBG Code, for which only three or four of the five positions are often used. Table 8-11 provides examples of the use of the simple NBL Code.11 The NBL Code is designed so that it cannot be confused with the NBG or NBD Codes. As with those earlier codes, the characteristics were chosen in terms of clinical priority to describe characteristics with significant influence on the behavior of the device under consideration. For this reason, the NBL Code does not address such features as connector design and electrode location. Moreover, until there was emergence of clearer patterns in the electrode configurations used for cardioversion and defibrillation, with or without the pulse generator housing serving as part of the electrode system, it was believed that attempts to design a practical code encompassing leads used for cardioversion and defibrillation as well as pacing would be premature.12

Headers In the early history of pacing systems, differences among lead sizes, pacing configurations and polarity, and manufacturer variability led to a wide variety of pacemaker/ICD headers (connections of the lead[s] to the generator). With the proliferation of international standards (IS-1) technology, many of the connection issues of the past have disappeared. IS-1 lead technology and defibrillation connections (DF-1) header standards have resolved many of the differences in compatibility between manufacturers and platforms. Physicians now have a wider array of generators and systems from which to choose depending on patient need. The new international standard connections for leads and defibrillation leads (IS-4 and DF-4) will likely replace the bulkier IS-1/DF-1 technology (Figs. 8-1 and 8-2). This will further streamline connections and decrease the complexity and frequency of connections that could be a failure point in the system function. This should reduce the size of the junction for connecting the lead(s) to the generators, commonly referred to as headers (Figs. 8-3 and 8-4). With IS-4/DF-4 technology, rather than making three separate connections between ICD pace/sense, right ventricular (RV) defibrillation coil, and superior vena cava (SVC) coil with the

The NASPE/BPEG Generic (NBG) Pacemaker Code*

Position I (Chambers Paced)

Position II (Chambers Sensed)

Position III (Response to Sensing)

Position IV (Programmability, Rate Modulation)

Position V (Antitachyarrhythmia Functions)

O = none A = atrium V = ventricle D = dual (A + V)

O = none A = atrium V = ventricle D = dual (A + V)

O = none T = triggered I = inhibited D = dual (T + I)

O = none P = simple programmable M = multiprogrammable C = communicating R = rate modulation

O = none P = pacing (antitachyarrhythmia)

S = single (A or V)†

S = single (A or V)†

S = shock D = dual (P + S)

Modified from Bernstein AD, Camm AJ, Fletcher RD, et al: The NASPE/BPEG Generic Pacemaker Code for antibradyarrhythmia and adaptive-rate pacing and antitachyarrhythmia devices. PACE 10:794, 1987. *Positions I to III are used exclusively for antibradyarrhythmia (antibradycardia) function. †Manufacturer’s designation only

TABLE

8-6 

Revised NASPE/BPEG Generic Code for Antibradycardia Pacing

Position I (Chambers Paced)

Position II (Chambers Sensed)

Position III (Response to Sensing)

Position IV (Rate Modulation)

Position V (Multisite Pacing)

O = none A = atrium V = ventricle D = dual (A + V) S = single (A or V)*

O = none A = atrium V = ventricle D = dual (A + V) S = single (A or V)*

O = none T = triggered I = inhibited D = dual (T + I)

O = none R = rate modulation

O = none A = atrium V = ventricle D = dual (A + V)

Modified from Bernstein AD, Daubert JC, Fletcher RD, et al. The Revised NASPE/BPEG Generic Code for antibradycardia, adaptive-rate, and multisite pacing. PACE 25:260-264, 2000. *Manufacturer’s designation only.

198

TABLE

8-7 

SECTION 1  Basic Principles of Device Therapy

TABLE

Examples of Revised NBG Code

Code VOO, VOOO, or VOOOO VVIRV

AAI, AAIO, or AAIOO AAT, AATO, or AATOO

AATOA

DDD, DDDO, or DDDOO

DDI, DDIO, or DDIOO DDDR or DDDRO DDDRA DDDOV DDDRD

8-9 

Meaning Asynchronous ventricular pacing; no sensing, rate modulation, or multisite pacing. Ventricular inhibitory pacing with rate modulation and multisite ventricular pacing (i.e., biventricular pacing or more than one pacing site in one ventricle). This mode is often used in patients with heart failure, chronic atrial fibrillation, and intraventricular conduction delay; assessed by atrial fibrillation group in MUSTIC study.* Atrial pacing inhibited by sensed spontaneous atrial depolarizations; no rate modulation or multisite pacing. Atrial pacing with atrial outputs elicited without delay on atrial sensing during alert period outside pulse generator’s refractory period (used primarily as diagnostic mode to determine exactly when atrial depolarizations are sensed); no rate modulation or multisite pacing. Atrial pacing with atrial outputs elicited without delay on atrial sensing during alert period outside pulse generator’s refractory period, without rate modulation but with multisite atrial pacing (i.e., biatrial pacing or more than one pacing site in one atrium).† Dual-chamber pacing (normally inhibited by atrial or ventricular sensing during alert portion of VA interval or by ventricular sensing during alert portion of AV interval, and with ventricular pacing triggered after programmed PV interval by atrial sensing during alert portion of VA interval); no rate modulation or multisite pacing. Dual-chamber pacing without atrium-synchronous ventricular pacing (atrial sensing merely cancels pending atrial output without affecting escape timing); no rate modulation or multisite pacing. Dual-chamber, adaptive-rate pacing; no multisite pacing. Dual-chamber, adaptive-rate pacing with multisite atrial pacing (i.e., biatrial pacing or more than one pacing site in one atrium); assessed in multicenter DAPPAF study.‡ Dual-chamber pacing without rate modulation but with multisite pacing (i.e., biventricular pacing or more than one pacing site in one ventricle).§ Dual-chamber pacing with rate modulation and multisite pacing both in the atrium (i.e., biatrial pacing or pacing in more than one site in one atrium) and in the ventricle (i.e., biventricular pacing or pacing in more than one site in one ventricle).

Modified from Bernstein AD, Daubert J-C, Fletcher RD, et al. The Revised NASPE/ BPEG Generic Code for antibradycardia, adaptive-rate, and multisite pacing. PACE 25:260264, 2000. *Leclerq C, Walker S, Linde C, et al. Comparative effects of permanent biventricular and right-univentricular pacing in heart failure patients with chronic atrial fibrillation. Eur Heart J 23:1780-1787, 2002. †Revault d’Allonnes G, Pavin D, Leclerq C, et al. Long-term effects of biatrial synchronous pacing to prevent drug-refractory atrial tachyarrhythmias: a nine-year experience. J Cardiovasc Electrophysiol 11:1081-1091, 2000. ‡Fitts SM, Hill MR, Meral R, et al. Design and implementation of the Dual Site Atrial Pacing to Prevent Atrial Fibrillation (DAPPAF) clinical trial. J Interv Card Electrophysiol 2:139-144, 1998. §Cazeau S, Leclerq C, Lavergne T, et al. Effects of multisite biventricular pacing in patients with heart failure and intraventricular conduction delay. N Engl J Med 344:873880, 2001. MUSTIC, Multisite Stimulation in Cardiomyopathy.

TABLE

8-8 

The NASPE/BPEG Defibrillator Code, Short Form

Code ICD-S ICD-B ICD-T

Meaning ICD with shock capability only ICD with bradycardia pacing as well as shock ICD with tachycardia (and bradycardia) pacing as well as shock

Modified from Bernstein AD, Camm AJ, Fisher JD, et al. The NASPE/BPEG Defibrillator Code. PACE 16:1776, 1993. ICD, Implantable cardioverter-defibrillator.

TABLE

8-10 

The NASPE/BPEG Pacemaker-Lead (NBL) Code

Potential I (Electrode Configuration)

Potential II (Fixation Mechanism)

Potential III (Insulation Material)

Potential IV (Drug Elution)

U = unipolar B = bipolar M = multipolar

A = active P = passive O = none

P = polyurethane S = silicone rubber D = dual (P + S)

S = steroid N = nonsteroid O = none

Modified from Bernstein AD, Parsonnet V. The NASPE/BPEG Pacemaker-Lead Code. PACE 19:1535, 1996.

TABLE

8-11  Code UPSO BAPS

Examples of the NBL Code Meaning Unipolar passive-fixation lead with silicone-rubber insulation but without elution of an anti-inflammatory agent Bipolar active-fixation lead with polyurethane insulation and steroid elution

Modified from Bernstein AD, Parsonnet V. The NASPE/BPEG Pacemaker-Lead Code. PACE 19:1535, 1996.

header, an IS-4 lead will have one connection with a DF-4 header, minimizing the risk of poor contact between header set screw and lead pin.

Conclusion The practice of electrophysiology has benefited greatly from the standardization of lead-header connections and the use of standard tools of communication regarding the configuration of pacemakers, ICDs, and defibrillators.

The NASPE/BPEG Defibrillator (NBD) Code

Potential I (Shock Chamber)

Potential II (AntitachycardiaPacing Chamber)

Potential III (Tachycardia Detection)

Potential IV (AntibradycardiaPacing Chamber)

O = none A = atrium

O = none A = atrium V = ventricle D = dual (A + V)

E = electrogram H = hemodynamic V = ventricle D = dual (A + V)

O = none A = atrium V = ventricle D = dual (A + V)

Modified from Bernstein AD, Camm AJ, Fisher JD, et al. The NASPE/BPEG Defibrillator Code. PACE 16:1776, 1993.

Figure 8-1  Standard dual-coil DF4 ICD lead. (Courtesy Medtronic, Minneapolis.)



8  Pacemaker, Defibrillator, and Lead Codes and Headers

199

Figure 8-2  Standard dual-coil DF1/IS1 ICD lead. (Courtesy Medtronic, Minneapolis.)

Figure 8-3  Standard IS-1/DF-1 dual-coil ICD header. (Courtesy Medtronic, Minneapolis.)

Figure 8-4  DF-4 dual-coil ICD header. (Courtesy Medtronic, Minneapolis.)

REFERENCES 1. Bristow MR, Saxon LA, Boehmer J, et al: Cardiac-resynchronization therapy with or without an implantable defibrillator in advanced chronic heart failure. N Engl J Med 350(21):2140-2150, 2004. 2. Mirowski M: The automatic implantable cardioverterdefibrillator: an overview. J Am Coll Cardiol 6:461-466, 1985. 3. Parsonnet V, Furman S, Smyth NPD: Implantable cardiac pacemakers: status report and resource guidelines. Pacemaker Study Group, Inter-Society Commission for Heart Disease Resources (ICHD). Circulation 50:A21, 1974. 4. Bernstein AD, Camm AJ, Fletcher RD, et al: The NASPE/BPEG Generic Pacemaker Code for antibradyarrhythmia and adaptiverate pacing and antitachyarrhythmia devices. Pacing Clin Electrophysiol 10(4 Pt 1):794-799, 1987. 5. Bernstein AD, Brownlee RR, Fletcher RD, et al: Report of the NASPE Mode Code Committee. PACE 7:395, 1984.

6. Bernstein AD, Daubert JC, Fletcher RD, et al: The Revised NASPE/ BPEG Generic Code for antibradycardia, adaptive-rate, and multisite pacing. PACE 25:260-264, 2000. 7. Bernstein AD, Camm AJ, Furman S, Parsonnet V: The NASPE/ BPEG codes: use misuse, and evolution. PACE 24:787-788, 2001. 8. Epstein AE, Dimarco JP, Ellenbogen KA, et al: ACC/AHA/HRS 2008 guidelines for Device-Based Therapy of Cardiac Rhythm Abnormalities: executive summary. Heart Rhythm 5(6):934-955, 2008. 9. Epstein AE, DiMarco JP, Ellenbogen KA, et al: ACC/AHA/HRS 2008 Guidelines for Device-Based Therapy of Cardiac Rhythm Abnormalities: a report of the American College of Cardiology/ American Heart Association Task Force on Practice Guidelines (Writing Committee to Revise the ACC/AHA/NASPE 2002 Guideline Update for Implantation of Cardiac Pacemakers and

Antiarrhythmia Devices), with the American Association for Thoracic Surgery and Society of Thoracic Surgeons. J Am Coll Cardiol 51:e1-e62, 2008. 10. Bernstein AD, Camm AJ, Fisher JD, et al: The NASPE/BPEG Defibrillator Code. PACE 16:1776, 1993. 11. Bernstein AD, Parsonnet V: The NASPE/BPEG Pacemaker-Lead Code. PACE 19:1535, 1996. 12. Gregoratos G, Cheitlin MD, Conill A, et al: ACC/AHA Guidelines for Implantation of Cardiac Pacemakers and Antiarrhythmia Devices: a report of the American College of Cardiology/ American Heart Association Task Force on Practice Guidelines (Committee on Pacemaker Implantation). J Am Coll Cardiol 31:1175-1209, 1998.

9 

Basic Physiology and Hemodynamics of Cardiac Pacing FRITS W. PRINZEN  |  MARC STRIK  |  FRANÇOIS REGOLI  |  ANGELO AURICCHIO

C

ardiac pacing has significantly improved the survival and quality of life in patients with bradycardia. When first introduced, pacemakers were simply lifesaving devices that provided a fixed pacing rate during bradycardia. With advances in technology and understanding of cardiac physiology during artificial pacing, devices have been developed that mimic the normal cardiac automaticity and atrioventricular (AV) activation sequence. Accordingly, pacemakers are capable of raising the pacing rate with exercise without the need to rely on sinus activity, so-called rate-adaptive pacing. Moreover, a novel target of pacing was introduced to improve electromechanical synchrony rather than cardiac rhythm: cardiac resynchronization therapy (CRT). The use of CRT to treat patients with heart failure has given new impetus to evaluation of the hemodynamic effect of spontaneously occurring or pacing-induced ventricular conduction delay, as well as abnormal AV timing. In addition, inappropriate right ventricular pacing not only may have deleterious effects on ventricular mechanics, but ultimately may increase morbidity and mortality. It is now clear that a truly physiologic pacemaker maintains the normal sequence and timing of atrial and ventricular activation over a wide range of heart rates, varies the heart rate in response to metabolic demands, and preserves the normal, rapid, synchronous sequence of ventricular activation when required.

Physiology of Electrical and Mechanical Activation ELECTRICAL ACTIVATION DURING SINUS RHYTHM The cardiac action potential originates from the sinus node, located high in the right atrium (Fig. 9-1). Its cells depolarize spontaneously and initiate the spontaneous depolarization of action potentials at a regular rate from the sinus node. This rate depends on various conditions, such as atrial stretch and sympathetic activation, but is usually between 60 and 100 beats per minute (bpm) at rest. Myocytes are electrically coupled to each other through gap junctions. These structures consist of connexin molecules and allow direct intercellular communication.1 Gap junctions do not have a preferential direction of conduction, but because the action potential starts in the sinus node, it spreads from there through the atria. Evidence suggests specialized conduction pathways in the atrium, but their (patho) physiologic relevance is still disputed. In the human heart, spreading of the action potential through the atria takes approximately 100 msec, after which the impulse reaches the AV node (see Fig. 9-1). In the normal heart the AV node is the only electrical connection between the atria and ventricles, because a fibrous ring (anulus fibrosus) is present between the remaining parts of the atria and ventricles. The AV nodal tissue conducts the electrical impulses very slowly; indeed, it takes approximately 80 msec for these impulses to travel from the atrial side to the ventricular side of the AV node. This delay between atrial activation and ventricular activation has functional importance because it allows optimal ventricular filling. Its slow conduction renders the AV node sensitive to impaired conduction and even complete conduction block, an important indication for ventricular pacing. As with other locations in the heart, conduction in

the AV node has no preferential direction. Consequently, impulses can also be conducted retrogradely through the AV node, a condition that can occur when the ventricles are electrically stimulated. From the AV node, the electrical impulse reaches the His bundle, the first part of the specialized conduction system of the ventricles called the Purkinje system. Within this system the electrical impulse is conducted approximately four times faster (3-4 m/sec) than in the working myocardium (0.3-1 m/sec). This difference results because Purkinje cells are longer and have a higher content of gap junctions.1-5 The intraventricular conduction system can be regarded as a trifascicular conduction system consisting of the right bundle branch (RBB) and two divisions of the left bundle branch (LBB). The RBB proceeds subendocardially along the right side of the interventricular septum until it terminates in the Purkinje plexuses of the right ventricle; the LBB also has a short subendocardial route. It seems that the bifascicular vision of structure of the LBB is oversimplified.6 The general picture that now emerges is that the left ventricular Purkinje network is composed of three main, widely interconnected parts, depending on the anterior subdivision, the posterior subdivision, and a centroseptal subdivision of the left main bundle. This third medial or centroseptal division supplies the midseptal area of the left ventricle and arises from the LBB, from its anterior or posterior subdivision, or both. The fascicles continue in a network of subendocardially located Purkinje fibers. In the left ventricle, Purkinje fibers form a network in the lower third of the septum and free wall, which also covers the papillary muscles.4,7 In humans the bundles are present only underneath the endocardium, whereas other species (e.g., ox, sheep, goat) have networks of Purkinje fibers across the entire ventricular wall.8,9 It is important to note that the His bundle as well as the right and left bundle branches and their major tributaries are electrically isolated from the adjacent myocardium. The only sites where the Purkinje system and the normal working myocardium are electrically coupled are the Purkinje-myocardial junctions. The exits of the Purkinje system are located in the subendocardium of the anterolateral wall of the right ventricular (RV) and the inferolateral left ventricular (LV) wall4,10 (see Fig. 9-1). This area of exit corresponds with the area of the ventricular muscle that is activated earliest.2-5,11,12 The distribution of the Purkinjemyocardial junctions is spatially inhomogeneous, and the junctions themselves have varying degrees of electrical coupling.13 Endocardial activation of the right ventricle starts near the insertion of the anterior papillary muscle 10 msec after onset of LV activation11 (Fig. 9-2). After activation of the more apical regions, the activation of the ventricular working myocardium occurs predominantly from apex to base, both in the septum and in the LV and RV free wall. Further depolarization occurs centrifugally from endocardium to epicardium as well as tangentially.11,12 The earliest epicardial breakthrough occurs in the pretrabecular area in the right ventricle, from which there is an overall radial spread toward the apex and base. The last part of the right ventricle that becomes activated is the AV sulcus and pulmonary conus. Overall, the posterobasal area of the left ventricle (or an area more lateral) is the last part of the heart to be depolarized (see Fig. 9-2). The time between arrival of the impulse in the His bundle and the first ventricular muscle activation is approximately 20 msec,4 whereas

203

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SECTION 2  Clinical Concepts

AV node His bundle (4)

Sinus node 1

3

2

8

9 Purkinjemyocardial junctions (7)

Bundle branch (5)

Purkinje fibers (6) Figure 9-1  Conduction of the impulse in the heart. The impulse originates in the sinus node (1), continues in the atrial wall (2), and is delayed in the atrioventricular (AV) node (3). Conduction within the ventricles is initially rapid within the rapid conduction system: His bundle (4), right and left bundle branches (5), and Purkinje fibers (6). The impulse is transferred from the rapid conduction system to the working myocardium in the Purkinje-myocardial junctions (7), which are located in the endocardium. Within the slowly conducting working myocardium, the impulse is conducted from endocardium to epicardium.

5 10 15 20 25 30 35 40 45 50 55 60 65 70 75 80 85 90 95 100

Figure 9-2  Three-dimensional isochronic representation of the electrical activation in an isolated human heart. Shaded scale indicates activation time in milliseconds. (Modified from Durrer D, van Dam RT, Freud GE, et al: Total excitation of the isolated human heart. Circulation 41:899-912, 1970.)

(Ca2+-induced Ca2+ release). These processes increase intracellular Ca2+ from approximately 10−7 M to 2 × 10−6 M. This high calcium concentration catalyzes the interaction between myosin and actin filaments, leading to contraction. After repolarization, relaxation occurs as Ca2+ dissociates from the contractile apparatus and is taken up again by the sarcoplasmic reticulum through the action of sarcoplasmic reticular Ca2+–adenosine triphosphatase (SERCA). Intracellular Ca2+ homeostasis is also maintained through action of the sodium-calcium (Na+/ Ca2+) exchanger (NCX). Normally, this exchanger removes Ca2+ from the cell (forward mode). Figure 9-3 also shows that the intracellular Ca2+ concentration rises rapidly after the upstroke of the action potential, but that there is some delay between the Ca2+ increase and the development of force. This delay is the main determinant of the delay between electrical and mechanical activation. The electromechanical delay, that is, the delay between the depolarization and the onset of force development, amounts to approximately 30 msec.15,16 On a global basis, this delay can be observed as the delay between the R wave of the electrocardiogram (ECG) and the rise in LV pressure (Fig. 9-4). Electromechanical coupling in failing hearts is

total ventricular activation lasts 60 to 80 msec, corresponding to a QRS duration of 70 to 80 msec.11 These numbers illustrate the important role of the Purkinje fiber system in the synchronization of myocardial activity. This role results from the system’s unique propagation properties and its geometrically widespread distribution. During normal orthodromic excitation, fast propagation over long fibers, together with wide distribution of Purkinje-myocardial junctions, induces a high degree of coordination between distant regions of the myocardium. MECHANICAL ACTIVATION DURING SINUS RHYTHM As in many other muscle cells, electrical activation leads to contraction in cardiac myocytes, a process referred to as excitation-contraction (E-C) coupling. In cardiac E-C coupling, the calcium ion (Ca2+) plays a central role14 (Fig. 9-3). The contraction-relaxation cycle starts when depolarization leads to entry of Ca2+ into the cell through voltagedependent L-type Ca2+ channels. This Ca2+ entry triggers a much larger amount of Ca2+ to be released from the sarcoplasmic reticulum K

Ca2

Na NCX Ca2

Na a

Na

1 2

Ca2 b c

A

NCX f

Sarcoplasmic reticulum

Ca2

0

Ca2

d

Ca2

e

3 4

SERCA

Action potential Ca2 transient

B

200 ms

Force

Figure 9-3  Excitation-contraction (E-C) coupling in myocardium. A, Schematic representation of the major pathways involved in calcium ion (Ca2+) homeostasis in the cardiomyocyte. a, Ca2+entry via L-type Ca2+ channel; b, Ca2+-induced Ca2+ release from the sarcoplasmic reticulum; c, Ca2+induced contraction; d, dissociation of Ca2+ from the contractile apparatus; e, reuptake of Ca2+ via sarcoplasmic reticular Ca2+–adenosine triphosphatase (SERCA); f, Ca2+ efflux via the Na+/Ca2+ exchanger (NCX). B, Schematic tracing of a Ca2+ transient, an action potential, and contractile force. The numbers indicate the different phases of the action potential: 0, upstroke; 1, fast early depolarization; 2, plateau phase; 3, repolarization phase; 4, resting membrane potential.

9  Basic Physiology and Hemodynamics of Cardiac Pacing

Electrical events



R

R T

P Q

S Ventricular systole

P Q Ventricular diastole

S

AV closure AV opening

Pressure

Aortic pressure

Left ventricular pressure

MV closure

MV opening

Ventricular volume

Left atrial pressure

IC Ejection Systole

IR RVF Diastasis AC Diastole

Figure 9-4  Events of the cardiac cycle. Left atrial, aortic, and left ventricular pressures are correlated with electrical events and left ventricular volume. AC, Atrial contraction; AV, aortic valve; IC, isovolumic contraction; IR, isovolumic relaxation; MV, mitral valve; RVF, rapid ventricular filling. See text for description.

different from that in normal hearts. A prominent feature of heart failure is that the relation between SERCA and the NCX changes. Often, SERCA is downregulated17 and NCX upregulated in the heart failure state.18 On the one hand, the reduced SERCA activity leads to less Ca2+ loading of the sarcoplasmic reticulum, resulting in less Ca2+ release during the subsequent activation, and thus a weaker contraction (systolic dysfunction). On the other hand, during diastole, Ca2+ removal from the cytosol is slower and incomplete, leading to a slower relaxation and greater diastolic stiffness (diastolic dysfunction). The upregulation of the NCX and the elevated intracellular Na+ concentrations in failing myocardium19 may compensate for the SERCA downregulation.20,21 The high Na+ concentration facilitates the NCX to work in the “reverse mode,” that is, to remove intracellular Na+ and exchange it for extracellular Ca2+. This additional Ca2+ enhances systolic function at least to some extent. Because the SERCA pump is much faster in pumping calcium than the NCX, insufficiency in contraction and relaxation in failing hearts is most pronounced at higher heart rates. These changes result in decreasing contractile force with increasing heart rate, a negative force-frequency relation.17 The changes in the failing myocardium lead to the altered expression of other proteins involved in E-C coupling, as well as shifts in isoforms of various contractile proteins.22 Because of the tight coupling between excitation and contraction, atrial activation is followed by atrial contraction, and ventricular activation by ventricular contraction. Consequently, atrial contraction precedes ventricular contraction, as illustrated in Figure 9-4. Because of this timing, the atrial contraction adds about 20% to the volume of the ventricles. This “atrial kick” increases the length of ventricular muscle cells and their sarcomeres. The sarcomere length is an

205

important determinant of myocardial contractile force. Over the entire physiologic range of sarcomere lengths (1.6-2.4 µm), the longer the sarcomeres, the greater is the contractile force. This effect, known as the Frank-Starling relation (Starling curve; Starling law of the heart), is only partly related to the overlap of actin and myosin, as generally mentioned in textbooks. More important to the Frank-Starling relation is the growing affinity of troponin C for calcium at increasing sarcomere length,23 although the cause of this greater sensitivity is not completely understood. Ventricular contraction starts after ventricular depolarization. After a short isovolumic contraction phase, the aortic valve opens, starting the ejection phase. The velocity of emptying of the ventricle is highest in the first half of the ejection phase because of a higher pressure gradient between the left ventricle and the aorta. In the second half of the ejection phase, LV pressure actually falls below aortic pressure, but the aortic valve stays open because of the inertia of the flowing blood. With increasingly negative LV-aortic pressure gradients, the aortic valve closes and the isovolumic relaxation phase starts, ending with the opening of the mitral valve. Because of the filling of the atrium during ventricular systole, atrial pressure is relatively high. This event, in combination with the rapid fall in LV pressure, causes a positive AV pressure gradient during the early filling phase. Thus, rapid acceleration of blood occurs, contributing to most of the LV filling in the early diastolic phase. This event is also reflected by the large Ei wave on mitral valve Doppler recordings in healthy hearts (Fig. 9-5). Atrial contraction, after the P wave on the ECG, produces a surge of blood into the ventricle at the end of diastole, causing the final and optimal filling of the ventricle. Reversal of the pressure gradient across the mitral valve, caused by atrial relaxation and ventricular contraction, pulls the valve cusps in apposition, facilitating closure of the valve. Therefore, coupling and proper timing of atrial and ventricular contractions are important determinants of ventricular pump function. The loss of atrial systole diminishes effective LV stroke volume by 25% at low heart rates and by almost 50% at higher heart rates. How different modes of pacing influence the atrial contribution to pump function and how this contribution affects pump function in different categories of patients are discussed later. Although both Frank-Starling and force-frequency relations are properties of the myocardium itself, which can regulate cardiac function to an important extent, extrinsic factors also modulate cardiac performance. Many of these factors are related to the autonomic nervous and hormonal systems. The afferent parts of these systems can be divided into low-pressure and high-pressure domains. Low-pressure

1 m/s -

.5 m/s Ei

Ai ¹⁄³ DFT

DFT

Time (ms) Figure 9-5  Pulsed Doppler recording of transmitral flow, illustrating the technique used to measure various segments of the timed velocity integral and diastolic filling period (ms, milliseconds; m/s, meters per second). The relationship between early diastolic filling and filling associated with atrial systole can be quantitated. Possible measurements include early and atrial velocities, the ratio of early to atrial velocities, the early (Ei) and atrial (Ai) flow-velocity integral (area under flow-velocity curve), the ratio of early to atrial flow-velocity integrals, and the amount of diastolic filling occurring in the first third of diastole ( 1 3 DFT [diastolic filling time]). (From Pearson AC, Janosik DL, Redd RM, et al: Doppler echocardiographic assessment of the effect of varying atrioventricular delay and pacemaker mode on left ventricular filling. Am Heart J 115:611, 1988.)

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SECTION 2  Clinical Concepts

sensors are present in the atria and pulmonary circulation. Stimulation, from increased pressures and stretch, leads to parasympathetic stimulation and sympathetic withdrawal. Moreover, atrial stretch induces release of atrial natriuretic factor (ANF). This 28–amino acid polypeptide is a potent arterial and venous vasodilator that increases urine production. The low-pressure sensors monitor the filling status of the circulation. However, erroneous readings may occur during inadequate AV synchronization, leading to elevated atrial pressures (cannon A waves). Stimulation of the reflexes and ANF production are involved in the so-called pacemaker syndrome (see later). The high-pressure sensors consist predominantly of the baroreceptors in the arterial system. These sensors feed back to the brainstem through afferent neurons in the vagal and glossopharyngeal nerves. Stimulation of these receptors through a decrease in blood pressure induces sympathetic stimulation and parasympathetic withdrawal. Although reduced parasympathetic activation is the most important factor leading to the increase in heart rate during hypotension, the greater sympathetic stimulation induces arterial and venous vasoconstriction and increases myocardial contractility. Other pressure sensors are found in the juxtaglomerular apparatus in the kidneys. Their stimulation leads to greater production of renin, angiotensin, and aldosterone, resulting in vasoconstriction as well as water and salt retention. These systems are effective in maintaining equilibrium in blood pressure and cardiac output during short-term variations. However, sustained neurohumoral stimulation is an important factor in the long-term development of hypertrophy and the rise in filling pressure. ABNORMAL ACTIVATION SEQUENCE DURING LEFT BUNDLE BRANCH BLOCK AND RIGHT VENTRICULAR PACING The normal, physiologic, and almost synchronous sequence of electrical activation is lost in diseases affecting the ventricular conduction system, such as block of the left or right bundle branch24 and the presence of an accessory pathway bypassing the AV node, as in the WolffParkinson-White syndrome.25 Additionally, ectopic impulse generation, occurring during ventricular pacing and extrasystoles,26,27 leads to abnormal impulse conduction. Under all these circumstances, the impulse is conducted primarily through the slowly conducting working myocardium rather than rapidly through the specialized conduction system. As a consequence, under conditions of abnormal activation, the time required for activation of the entire ventricular muscle, expressed as QRS duration, is at least twice as long as that during normal sinus rhythm. The extent of asynchrony (dyssynchrony) and the sequence of activation during abnormal conduction are determined by at least four myocardial properties, as follows: 1. Conduction through the myocardium is up to four times slower than conduction through the Purkinje system.4 2. Conduction is approximately two times faster along the muscle fibers than perpendicular to them.28 Therefore, in a particular layer, the wavefront around a pacing site has an elliptical shape, especially in the epicardial and midmyocardial layers.29 3. Impulses originating from the working myocardium rarely reenter into parts of the rapid conduction system. Early researchers had believed that such reentry was ubiquitous and that the amount of reentry was the main determinant of the total asynchrony of ventricular activation.30 However, later studies indicated that, although parts of the intact rapid conduction system are often present, ectopically generated impulses appear to couple poorly to the rapidly conducting specialized conduction system.12 Myerburg et al.4 elegantly showed that impulses coming from the normal myocardium can enter the Purkinje system only at the apical part of this system (presumably the Purkinje-myocardial junctions; see Fig. 9-1). This finding implies that the impulse often has traveled already for some time before reaching Purkinje-myocardial junctions. Also, to reach remote zones, the impulse must travel retrogradely through one

branch of the system all the way to the proximal part and then descend to another part. Therefore, in most cases, the sequence of activation during ventricular pacing is governed by slow conduction through the normal myocardium, away from the pacing site.16,26,27,31 4. Most endocardial fibers, even though not part of the Purkinje system, conduct impulses faster than the fibers in the rest of the LV wall.32 Moreover, the endocardial circumference is smaller than its epicardial counterpart. Therefore, total time required for electrical activation is shorter for LV endocardial pacing than for LV epicardial pacing.29,33 Detailed studies on the three-dimensional spread of activation during ventricular pacing in canine hearts have been conducted since the 1960s.12,26 The sequence of electrical activation in LBB block (LBBB) is similar to that during RV apex pacing. This sequence can be derived from the QRS configuration of the surface ECG27,34,35 and from endocardial activation maps in experimental LBBB and RV pacing.36 The similarity has lead to the use of AV sequential RV apex pacing as a model for “experimental LBBB.”34,37,38 Until recently, information on impulse conduction in patients with LBBB was limited to the data reported by Vassallo et al.24,27 These investigators mapped LV endocardial activation during RV apex pacing and in LBBB at a limited number of endocardial sites, showing that activation starts at the RV endocardium in both conditions. The first noticeable activation at the LV endocardium after right-to-left conduction of the impulse occurs at a single breakthrough site, which in almost all patients is the LV breakthrough time, 50 to 70 msec after the earliest RV activation. The impulse is conducted from the septum toward the distal free wall in a gradual manner, the site of latest activation generally being the inferoposterior wall.11 More recently, this finding has been confirmed using high-resolution mapping system39-42 (Fig. 9-6). These studies also showed that activation patterns differ among patients largely because of differences in the origin of LBBB: either a proximal conduction block43,44 or a uniform but slow conduction through the LBB.45 In approximately one third of patients with heart failure and typical LBBB QRS morphology, transseptal activation time (i.e., time between earliest RV and LV septal breakthrough point) is bimodally distributed and has a large range in both groups of patients, with transseptal conduction time of 40 msec or less and longer than 40 msec.46 This indicates great heterogeneity in the RV-to-LV activation time. Furthermore, in patients with LBBB, the total LV endocardial activation time ranges from 60 to 160 msec. Etiology does not seem to have a major impact on the total endocardial activation time. Of note, the sum of transseptal and total endocardial activation times does not account for the maximum QRS duration, which is 20 to 60 msec longer, probably because of LV endocardial-toepicardial conduction time (Fig. 9-7). These data also show that QRS duration provides a reasonable although imperfect estimate of total electrical asynchrony. Patients with LBBB show a peculiar spread of ventricular activation.40,42 Indeed, in two thirds of patients with LBBB, a functional line of block is present, resulting in a U-shaped electrical activation wavefront. The presence and location of the line of block have been also recently observed in canine model of LBBB using the same noncontact mapping system.46a The activation front propagates first around the inferior wall and then spreads onto the lateral and basal walls. The location and length of the line of block vary greatly and are related to the site and time of LV breakthrough. In patients with significant prolongation of the QRS duration (>150 msec), the line of block is consistently located in an anterior position. In contrast, in patients with QRS duration of 120 to 150 msec, the line of block is usually shorter and is located anteriorly, laterally, or posteriorly40 (Movie 9-1). ABNORMAL CONTRACTION PATTERNS DURING LEFT BUNDLE BRANCH BLOCK AND RIGHT VENTRICULAR PACING Given the tight relationship between excitation and contraction in the myocardium, it is not surprising that asynchronous electrical



9  Basic Physiology and Hemodynamics of Cardiac Pacing

207

Noncontact mapping

Contact mapping

Figure 9-6  Contact and noncontact mapping in patient with left bundle branch block and heart failure. The unipolar isopotential activation sequence (upper panel) shows a U-shaped activation front that rotated around the apex and activated the lateral wall late, whereas the bipolar propagation map (lower panel) shows a longer activation time in the anterior region. ANT, Anterior; LAT, lateral. (Modified from Auricchio A, Fantoni C, Regoli F, et al: Characterization of left ventricular activation in patients with heart failure and left bundle-branch block. Circulation 109:1133-1139, 2004.)

activation also leads to asynchronous contraction. Regions where the impulse arrives first also start to contract first.47 During LBBB or ventricular pacing, local contraction patterns differ not only in the onset of contraction, but also, and more importantly, in the pattern of contraction. These contraction patterns imply that opposing regions of the ventricular wall are out of phase and that energy generated by one region is dissipated in opposite regions. In patients with LBBB or RV pacing, the early-contracting region is most often the septum. The earliest-contracting fibers can shorten rapidly by up to 10% just during the isovolumic contraction phase, because the remaining muscle fibers are still in a relaxed state (Movie 9-2). This rapid, early shortening is followed by an additional but modest systolic shortening, eventually followed by systolic stretch (from delayed mechanical contraction of other regions, in lateral wall), and premature relaxation (Fig. 9-8). In late-activated regions, in contrast, the fibers are stretched in the early systolic phase (up to 15%) from contraction of the early-activated region. Doubling of net systolic shortening and delayed relaxation occur in late-activated regions.31,48 The discoordination between earlyand late-activated regions leads to lower output and reduced efficiency of the heart as a pump. The regional differences in contraction pattern most likely results from the local differences in myocardial fiber length during the early systolic phase. This is supported by studies using two isolated papillary muscles in series in which asynchronous stimulation caused a downward shift in the force-velocity relation in the earlier-activated muscle and an upward shift in the later-activated one.49 Furthermore, during ventricular pacing, regional systolic fiber shortening increases with greater isovolumic stretch.50 Similarly, a close correlation exists between the time of local electrical activation and the extent of systolic fiber shortening.51 Therefore, the regional differences in contraction pattern during ventricular pacing are most likely caused by regional differences in effective preload and local differences in the contraction force triggered by the Frank-Starling relation. On M-mode echocardiography, abnormal contraction patterns during RV pacing and LBBB appear as paradoxical septal motion.52

However, this motion is not actually paradoxical, because it is really the net result of different forces. The motion is caused by the asynchrony between RV and LV, which produces dynamic alterations in transseptal pressure differences,52,53 and by presystolic shortening of septal muscle fibers.54 Abnormal septal motion results in a diminished contribution of the interventricular septum to LV ejection. This has been measured in experimental LBBB using MRI tagging54 and in patients with LBBB QRS morphology using speckle tracking.55,56 The septal regions, even in absence of detectable scar, showed the lowest amount of systolic strain within LV; in contrast, the lateral wall shows the highest regional strain.55 This finding demonstrates that LBBB determines a unique and unequal LV strain distribution that may be corrected by CRT.55-57 Effect of Left Bundle Branch Block and Right Ventricular Pacing on Local Energetic Efficiency The local differences in wall motion and deformation mentioned previously reflect regional differences in myocardial work.48,58 This relationship was demonstrated by construction of local fiber stress–fiber strain diagrams and calculation of local external and total mechanical work. In regions close to the pacing site (or septum in patients with LBBB), shortening occurs at low pressure, whereas these areas are transiently being stretched at higher ventricular pressures. As a consequence, the stress-strain loops have a figure-8-like shape with a low net area, indicating the absence of external work. In regions remote from the pacing site (or lateral wall in patients with LBBB), the loops are wide, and external work can be up to twice that during synchronous ventricular activation (Fig. 9-9). Total myocardial work (sum of external work and potential energy) in LBBB and with RV pacing is reduced by 50% in early-activated regions and is increased by 50% in lateactivated regions in comparison with atrial pacing.48,54,59 Several studies reported that ventricular pacing and LBBB are associated with regional differences in myocardial blood flow,31,35,38,54,58-60 glucose uptake,38,61 and oxygen consumption.58,62 During RV apex pacing and LBBB, low values are found in the septum, the

208

SECTION 2  Clinical Concepts

300

250

Time (msec)

200

150

100

50

0 Transseptal time

A

Total endocardial activation time (by Carto mapping)

Time to end of QRS complex (QRS max)

89 msec Y+

Mitral value –58 msec

His

Total ventricular activation: 101 msec LV transseptal time: 58 msec LV RV

B

Isochronal step: 10 msec

Apex 1.48 cm

Figure 9-7  A, Distribution of transseptal time, total endocardial activation time, and total QRS duration in 140 heart failure patients with a QRS duration of 120 msec or longer. Patients are ordered according to and transseptal time. B, Three-dimensional electroanatomic mapping of the right ventricle (RV) and left ventricle (LV). A color-coded activation sequence indicated the earliest (red) and the latest (blue) endocardial activation region. The earliest activated region is the anterolateral region of the RV, and the latest part is the posterobasal region of the LV.

early-activated part of the left ventricle. Compared with sinus rhythm, myocardial blood flow and oxygen consumption are 30% lower in early-activated regions and 30% higher in late-activated regions.31,58,59 Controversy surrounds the cause of the reduced septal blood flow and glucose uptake in LBBB and RV apex pacing. Various observations, such as close correlations among local myocardial oxygen consumption, work, and blood flow, suggest that the lower blood flow is a physiologic adaptation to the lower demand. However, at higher heart rates, the asynchronous contraction pattern may also impede myocardial perfusion.38,63 In patients with RV pacemakers, false-positive findings of perfusion defects may result from this redistribution.64

The nonuniform contraction patterns translate into a lower efficiency of conversion of metabolic energy into mechanical pump function. This was observed in isolated, isovolumically beating hearts,65 as well as in anesthetized open-chest,66,67 closed-chest,59 and conscious dogs.68 Mills et al.59 also showed that efficiency remained low after 4 months of RV pacing. In all these preparations, RV apex pacing reduced mechanical output, whereas myocardial oxygen consumption was unchanged or even increased compared with atrial pacing; thus efficiency decreased by 20% to 30% in these studies. The opposite change was found when ventricular activation in patients with preexisting dyssynchrony caused by LBBB-like activation was improved by biventricular (BiV) or LV pacing alone. In these patients, BiV pacing



9  Basic Physiology and Hemodynamics of Cardiac Pacing Midseptum

Anterior

Midlateral

Normal

RV apex pacing mJ/g 8

LV free wall pacing

*

0 Fiber stress Fiber length Apex 0

200

Time (msec)

Figure 9-8  Contraction patterns in 24 regions of the anterior septum and anterior left ventricular (LV) free wall during pacing at the LV lateral wall. Each graph in the array represents circumferential strain versus time in a specific region. The left ventricle is displayed as if cut down the septum and folding the surface out onto the page. For the sake of clarity, the posterior half of the left ventricle is not shown, but the distribution of strains in this part of the LV wall is a mirror image of that in the anterior wall. Rows represent different levels from base (top) to apex (bottom) of the left ventricle, and columns are circumferential positions from midseptum (far left) via the LV anterior wall to the LV lateral wall (far right). Tics at the vertical axis denote a strain of 0.1 (≈10% shortening); tics on the horizontal axis denote 100 msec. Note the early shortening near the pacing site (*) and the prestretch in regions remote from the pacing site. (Modified from Prinzen FW, Peschar M: Relation between the pacing induced sequence of activation and left ventricular pump function in animals. Pacing Clin Electrophysiol 25:484-498, 2002.)

Fiber length

and Tau, are strongly influenced by pacing.78-80 Parameters of auxotonic relaxation (rate of segment lengthening or LV volume increase) are also lower during ventricular pacing than during sinus rhythm, but the difference is less pronounced than for the isovolumic relaxation parameters.73,78 As a consequence of the slower contraction and relaxation, isovolumic contraction and relaxation phases last longer, thus leaving less time 3

Cardiac output: liters/minute

Both RV pacing and LBBB reduce systolic and diastolic function. These effects are independent of changes in preload and afterload. This conclusion is based on results from studies using preparations in which preload and afterload were controlled,65 as well as those in which preload- and afterload-independent indices of ventricular function were determined68,70-72 (Figs. 9-10 and 9-11). The negative mechanical effect of dyssynchronous activation has been observed under various loading conditions73 and during exercise.74 Ventricular pacing also causes a deterioration in pump function in patients with coronary artery disease,75 as well as those with already-impaired LV function76 (Fig. 9-12). Impaired regional and global cardiac pump function has been observed in patients and animals with LBBB, even if LBBB was not accompanied by other cardiovascular diseases.53,54 Therefore, it appears that under all circumstances, dyssynchrony is an important, independent determinant of cardiac pump function. Although LV dP/dtmax is a preload-dependent parameter, it is also sensitive to changes in ventricular activation sequence, as shown both in animal models and in humans34,76-78 (see Fig. 9-12). Thus, LV dP/ dtmax appears to be an appropriate marker of dyssynchrony-induced changes in LV global contractile function and its correction by any pacing technique. Additionally, the rate of ventricular relaxation is slower during ventricular pacing, with more pronounced detrimental effect on relaxation, in patients with impaired ventricular function76 (see Fig. 9-12). Isovolumic relaxation parameters, such as LV dP/dtmin

Fiber length

Figure 9-9  Upper panels, Maps of local external work during normal (right atrial), left ventricular (LV) free wall, and right ventricular (RV) apex pacing in a dog heart. External work values are presented in color II scale; for values, see scale bar; asterisk denotes pacing site. The LV wall is represented as a circle with the base located at the outer contour and the apex in the middle. Lower panels, Fiber length–fiber stress diagrams from three regions are displayed, indicating the shape of the diagram in normally activated (left), early-activated (middle), and late-activated (right) regions. (Modified from Prinzen FW, Hunter WC, Wyman BT, et al: Mapping of regional myocardial strain and work during ventricular pacing: experimental study using magnetic resonance imaging tagging. J Am Coll Cardiol 33:1735-1742, 1999.)

improved the maximal rate of rise of LV pressure (LV dP/dtmax) without increasing myocardial oxygen consumption, indicating improved efficiency.69 Efficient conversion of metabolic energy to pumping energy is of particular interest in patients with compromised coronary circulation, because higher oxygen consumption over longer periods (as in β-agonist therapy) leads to higher mortality. Effect of Asynchronous Activation on Systolic and Diastolic Pump Function

Fiber stress

0 0.2

Fiber stress

Circumferential strain

S

Base

209

2

1

0

0

5

10

15

20

25

30

LAmean cm H2O Atrial pace

Ventricular pace

Figure 9-10  Relationship between cardiac output and mean left atrial pressure (LAmean) during atrial pacing (open circles) compared with ventricular pacing (solid circles) in an intact dog model. For any given mean LA pressure, the cardiac output is higher with atrial pacing than with ventricular pacing. (From Mitchell JH, Gilmore JP, Sarnoff SJ: The transport function of the atrium: factors influencing the relation between mean left atrial pressure and left ventricular end diastolic pressure. Am J Cardiol 9:237, 1962.)

210

SECTION 2  Clinical Concepts

40

SR

Pace

20

0

30

40

50

60

70

80

90

LV volume (mL) Figure 9-11  Pressure-volume (P-V) diagrams during sinus rhythm (SR) (purple lines) and left ventricular (LV) pacing (Pace) (blue lines) before and during caval vein occlusion. The end-systolic P-V relation (straight lines) shifts rightward during ventricular pacing. These data indicate that the left ventricle operates at a larger volume under all loading conditions, and that at a given LV volume, LV pressure development and stroke volume are reduced. (Modified from Van Oosterhout MFM, Prinzen FW, Arts T, et al: Asynchronous electrical activation induces inhomogeneous hypertrophy of the left ventricular wall. Circulation 98:588-595, 1998.)

for ventricular filling and ejection.53,65,81 Therefore, it is not surprising that cardiac output and systolic arterial and LV pressures are also affected by a dyssynchronous activation. In general, stroke volume is affected more than systolic LV pressure,72,77 presumably because baroreflex regulatory mechanisms partly compensate the decrease in blood pressure. This idea is supported by the finding of higher catecholamine levels82 and greater systemic vascular resistance74 during ventricular pacing. With regard to the changes in stroke volume, it is important to note that ventricular pacing, especially RV apex pacing, can induce mitral regurgitation, as has been demonstrated in animals83,84 and patients.85,86 In addition to reduced stroke volume at unchanged preload, ejection fraction is usually found to be depressed during ventricular pacing74,87 as well as in LBBB.88 Similarly, ventricular pacing can increase pulmonary wedge pressure.74 The negative inotropic effect of ventricular pacing under various loading conditions is clearly illustrated by a rightward shift of the LV function curve, that is, the relationship between cardiac output and mean atrial pressure71 (see Fig. 9-10). Later studies showed a rightward shift of the end-systolic pressure-volume (P-V) relation (see Fig. 9-11), thus suggesting that for each end-systolic pressure, the LV must operate at a larger LV volume.70,72,89 Cause of Reduced Pump Function An obvious cause of reduced pump function during abnormal electrical activation is the asynchronous contraction of the different parts of the ventricular muscle. Using a mathematical simulation, Suga et al.90 divided the LV wall into two elements with similar contractile properties. Contractility of the whole left ventricle decreased considerably when the asynchrony exceeded 100 msec. This concept is supported by experimental observations in dogs with LBBB. In studies using combinations of pacing sites, minimal intra-LV asynchrony, assessed by endocardial electrical activation mapping, consistently led to the highest LV dP/dtmax.36 In addition to intraventricular asynchrony, interventricular asynchrony may affect ventricular pump function. An extensive overview of studies on RV and LV pacing showed that RV pacing consistently reduces LV function more than LV pacing.91 Even at similar levels of intraventricular asynchrony, LV dP/dtmax is lower during RV pacing than during LV pacing.36 This suggests important roles for interventricular asynchrony and associated interventricular

Electrical and Structural Remodeling Asynchronous electrical activation not only has acute mechanical consequences, but also influences cardiac structure and function over time. A particular adaptation of the heart to ventricular pacing occurs when the abnormal activation created by pacing is stopped but repolarization remains abnormal, a phenomenon called cardiac memory.98 Several investigators found evidence that, as in the nervous system, short-term cardiac memory (<1 hour) involves changes in ion channels and phosphorylation of target proteins,99 and that long-term cardiac memory (≈3 weeks of pacing) involves altered gene programming and protein expression.98,100 Interestingly, in humans, cardiac memory appears to reach a steady state within 1 week,101 which appears to be quicker than in dogs.98 Costard-Jäckle and Franz102 have demonstrated opposite changes in repolarization in regions remote from and close to a pacing site starting within 1 hour of ventricular pacing. The repolarization abnormalities related to cardiac memory also have a mechanical counterpart. During sinus rhythm immediately after a period of ventricular pacing, relaxation is disturbed.103 Moreover, systolic function deteriorates between 2 hours and 1 week after ventricular pacing has been stopped (Fig. 9-13). After ventricular pacing is stopped for 1 week, it takes 1 1 2 days for ejection fraction to 2000 *p  .0005 1500

*

*

*

*

1000 500 0

90 80 70 60 50 40 30 20 10 0

T (msec)

60

dP/dt dP/dt (mm Hg/sec)

LV pressure (mm Hg)

80

mechanical interaction. This ventricular interaction is also expressed by abnormal septal motion in asynchronous ventricles.52 Also, the pathway of activation may be a determinant of ventricular function. The sequence of activation from apex to base appears to be important. As already mentioned, the ventricular myocardium is activated from apex to base during normal sinus rhythm. Among all LV pacing sites, the LV apex appears to keep LV pump function closest to that seen during atrial pacing.26,59,92-94 The observation that the addition of pacing sites during LV apex pacing does not improve and sometimes even reduces LV function emphasizes the importance of a ventricular activation directed from apex to base.94 An additional cause of the reduction in pump function during asynchronous activation is mitral valve insufficiency.83,85 This diminishes LV pump function directly by reducing the volume ejected into the aorta and indirectly by reducing LV cavity volume. The immediate cause of mitral valve regurgitation during RV apex pacing is papillary muscle desynchronization.95 In addition, papillary muscle tethering forces may be increased,96 and the transmitral pressure gradient may be decreased.97

Normal Impaired Normal Impaired Normal Impaired LV LV LV LV LV LV dP/dt Atrial pacing

dP/dt

Isovolumic relaxation time (T)

AV sequential pacing

Figure 9-12  Hemodynamic effects of atrial and atrioventricular (AV) sequential pacing in patients with normal and impaired left ventricular systolic function. Note that +dP/dt (maximal rate of rise in LV pressure) is significantly lower in both groups of patients with AV sequential pacing than in patients with atrial pacing. However, only in patients with impaired LV systolic function −dP/dt (maximal rate of decline in LV pressure) is significantly lower, and the isovolumic relaxation time significantly longer, during AV sequential pacing than during atrial pacing. (Modified from Bedotto JB, Grayburn PA, Black WH, et al: Alterations in left ventricular relaxation during atrioventricular pacing in humans. J Am Coll Cardiol 15:658, 1990.)



9  Basic Physiology and Hemodynamics of Cardiac Pacing

Ejection fraction (%)

75 70 65 60 55 50 45 40 1 wk atrial pacing

1 wk ventricular pacing 50

100

150

Atrial pacing 200

250

Time (hours)

Figure 9-13  Left ventricular (LV) ejection fraction before pacing, during short-term and midterm (1 week) ventricular pacing, and during the first days after restoration of normal sinus rhythm. During the entire experimental period, the heart rate was kept at 80 bpm with the use of fixed-rate pacing at the right atrium and, during the week of ventricular pacing, dual-chamber pacing with an AV interval of 100 msec. (Data derived from Nahlawi M, Waligora M, Spies SM, et al: Left ventricular function during and after right ventricular pacing. J Am Coll Cardiol 44:1883-1888, 2004.)

return to its prepacing value.104 It seems likely that changes in the function of an ion channel, such as the L-type calcium channel,99 underlie the reduction in ejection fraction during the first week of ventricular pacing. Therefore, electrical and contractile remodeling appears to occur soon after the onset of asynchronous activation. In addition to the long-term effect on cardiac memory, longerlasting (>1 month) ventricular pacing and LBBB lead to major structural changes, such as ventricular dilatation and asymmetrical hypertrophy.54,70,105 The ventricular dilatation appears related to the LV operating at a larger volume. Global hypertrophy may be induced by this global dilatation, as well as by the greater sympathetic stimulation82 (Fig. 9-14). The asymmetry of hypertrophy appears to be more prominent during LV pacing than during RV pacing and LBBB.54,70,106 However, regardless of pacing site and underlying conduction disturbance, the asymmetry of hypertrophy is characterized by a more pronounced growth of the late-activated myocardium, that is, the same region that shows enhanced contractile performance because of early systolic prestretch. It has been observed in juvenile canine hearts that ventricular pacing leads to fiber disarray,107 which may also be the result of the regionally different and abnormal distribution of mechanical work.31,48,58,59,108 In ventricles undergoing long-term pacing,70 myocyte diameter is increased in late-activated regions, but the number of capillaries per myocyte is unchanged (van Oosterhout and Prinzen, unpublished data). Therefore, the diffusion distance for oxygen increases, potentially leading to compromised oxygenation, which is known to render hypertrophic myocardium susceptible to ischemia. Resynchronization has been shown to result in improved coronary reserve, suggesting that the structural remodeling at the microvascular level may also be reversible.109 An alternative explanation is that CRT could be simply decreasing the oxygen demand in the late-activated regions. In conditions such as infarction and hypertension, ventricular remodeling appears, at least initially, to be meant to compensate for the loss of function and the increased load, respectively. However, dilatation and hypertrophy do not reduce the asynchronous activation induced by pacing or LBBB; if any, they increase it.54,70 The reasons are the longer path length of impulse conduction (dilatation) and the larger muscle mass to be activated (hypertrophy). Moreover, chronic asynchronous activation also reduces expression of gap-junction channels110 and induces lateralization of these channels.111 Akar et al.112 have related the reduced expression in nonfailing and failing asynchronous hearts to slower impulse conduction, especially in the late-activated regions. Moreover, they showed that action potential duration increases

211

in the late-activated myocardium. These molecular changes may render the myocardium not only more asynchronous, but also more susceptible to arrhythmias. Even more complicated molecular changes occur in animal models with combined asynchronous activation and heart failure, such as that caused by rapid ventricular pacing. Under these conditions, betaadrenergic receptors are downregulated,113 as are many ion channels, such as most potassium and calcium channels, whereas other channels display a regional difference in expression, such as the L-type calcium channel.114 Moreover, in both patients with LBBB115 and animal models of dyssynchrony,113 several proteins related to or facilitating the apoptotic process show unique regional distribution, further confirming the hypothesis that abnormal electrical activitation sequence may contribute to remodeling in heart failure patients. All these processes may lead to a vicious circle in which dilatation and hypertrophy further reduce LV pump function, either directly or by increasing asynchrony of activation (see Fig. 9-14). As is the case after myocardial infarction, initial compensatory hypertrophy may result in heart failure many years later. Clinical evidence for such an important role of asynchronous activation in the development of heart failure is discussed later.

Physiologic Effects of Heart Rate and Atrioventricular Synchrony CORRECTION OF BRADYCARDIA BY PACING Cardiac pacing is the only effective treatment for symptomatic bradycardia. Permanent pacemakers have been implanted since the early 1960s to prevent death or syncope caused by ventricular asystole. Although asystole is clearly to be prevented, bradycardia does not always immediately lead to pump failure. This fact can be derived from studies on young patients with congenital AV block. As a compensation for the bradycardia, these patients have enlarged hearts and an increased ejection fraction.116,117 In a 10-year follow-up study, 9 of 149 patients demonstrated dilated cardiomyopathy.117 At the onset of acquired complete heart block, cardiac output is smaller than with normal heart rate.118 As a result, compensatory increases in sympathetic tone and end-diastolic ventricular volume occur, leading to higher atrial rates, enhanced ventricular contractility, and augmented stroke volume. Over the short term, the canine heart adapts to bradycardia by increasing diastolic volume and neurohumoral activation, which results in BiV hypertrophy118-120 (Table 9-1). Asynchronous electrical activation Dyscoordinate contraction Reduced pump function Longer Increased conduction Rightward shift Neurowall stress path length P-V relation humoral Molecular/ activation cellular Ventricular dilatation Regional changes mechanical differences Hypertrophy Asymmetrical hypertrophy Reduced efficiency

Reduced coronary reserve

Susceptibility ischemia

Figure 9-14  Possible relationships among the various consequences of asynchronous activation of the ventricles, resulting from ventricular pacing or conduction disturbances, and the deterioration of pump function over time. P-V, Pressure-volume. For details, see text.

212

SECTION 2  Clinical Concepts

9

Hemodynamic Effects of Ventricular Pacing in Complete Heart Block (CHB)

9-1 

Hemodynamic Parameter Ventricular rate Cardiac index Stroke volume Sympathetic tone Ventricular contractility Systemic vascular resistance Atrial rate

CHB with Slow Ventricular Response

8

Ventricular Pacing in CHB

↓ ↓ ↑ ↑ ↑ ↑ ↑

↑ ↑ ↓ ↓ ↓ ↓ ↓

Relative change (%)

LV cavity volume

50 25 QTc-time

0

2

4

6

8

10

12

14

5 4 3 1

75

0

6

2

Although these adaptations initially manage to return cardiac output to the level before the onset of AV block, this compensatory effect usually fades after 2 months of heart block. After 4 months, contractility is back to baseline, but LV cavity volume and corrected QT interval (QTc) continue to increase118 (Fig. 9-15); the latter increase is associated with an enhanced susceptibility to arrhythmias.119 If complete AV block persists for more than 4 months, congestive heart failure (CHF) ensues.120 Development of CHF may originate from the ventricular dilatation and hypertrophy initiated early after the start of heart block. The importance of the duration of overload is confirmed by the observation that the frequency of CHF symptoms in patients with complete heart block correlates with the duration of heart block. CHF has been described during chronic complete heart block even in patients with normal ventricular function. Brockman and Stoney121 reported that in two thirds of their patients with complete heart block and CHF, symptoms were relieved by ventricular pacing alone, and no additional medical therapy was required. Breur et al.122 showed that, in patients with congenital heart block, implantation of a pacemaker resulted in less LV dilatation than in the untreated control group, in which the LV cavity gradually dilated. Similarly, in dogs with AV block, LV dilatation gradually progresses over time (see Fig. 9-15), whereas ventricular pacing readily reverses the LV dilatation and hypertrophy caused by AV block.93 However, longer-lasting ventricular pacing in hearts with AV block may lead to LV dilatation and hypertrophy secondary to the effect of asynchronous activation. Two studies on young adults who received pacing for approximately 10 years showed a high preponderance of CHF and LV dilatation.123,124 These data strongly suggest that in the patient with AV block, the short-term and long-term effects of RV pacing are different. Therefore, the decision whether or not to start pacemaker treatment using RV pacing or the choice of the pacing site should be made carefully (see later).

LV contractility

7 No. of patients

TABLE

16

Time (weeks) Figure 9-15  Effect of atrioventricular (AV) block on left ventricular (LV) cavity volume (structural remodeling), LV contractility (contractile remodeling), and corrected QT interval (QTc-time) (electrical remodeling) in canine hearts, relative to their values before onset of AV block. Note the gradual and continuing increase in LV cavity volume over time, the early-onset and maintained electrical remodeling, and the biphasic contractile remodeling. (Modified from Peschar M, Vernooy K, Cornelussen RN, et al: Structural, electrical and mechanical remodeling of the canine heart in AV-block and LBBB. Eur Heart J 25 Suppl D:6165, 2004.)

0 51–55 56–60 61–65 66–70 71–75 76–80 81–85 86–90 Optimal ventricular rates at rest Figure 9-16  Histogram of optimal ventricular pacing rates at rest as defined by cardiac outputs determined by dye dilution in patients with complete heart block. (From Sowton E. Hemodynamic studies in patients with artificial pacemakers. Br Heart J 26:737, 1964.)

Early studies suggested that the maximal increase in cardiac output during ventricular pacing at rest occurs at rates between 70 and 90 bpm125-127 (Fig. 9-16). Further rises in rates result in either no additional increase or a decrease in cardiac output accompanied by greater peripheral vascular resistance. In contrast, there is growing evidence that lower heart rates may be particularly beneficial. Indeed, in animal models it appears that bradycardia promotes angiogenesis.128,129 Moreover, Lei et al.130 demonstrated that reducing normal heart rates by 20% to 30% through the continuous infusion of alinidine in a post– myocardial infarction (post-MI) rat model increased basic fibroblast growth factor, its receptor, and expression of related proteins. Bradycardia in this model also increased capillary length density in the border zone of the infarct by 40% and in the remote zone by 14%. Bradycardia also increased arteriolar density in the septum by 62%. These changes translated to a 23% greater coronary reserve, a smaller increase in post-MI left LV volume, and a greater preservation of ejection fraction in the bradycardic animals. Although the relationship between optimal resting ventricular pacing rate and ventricular dysfunction has not been systematically evaluated, it is likely that different resting ventricular rates are required in patients with systolic and diastolic dysfunction. For example, Ishibashi et al.131 reported that in elderly hypertrophic patients, even 25% increases in heart rate, as induced by atrial pacing, significantly reduced cardiac mechanical efficiency. The investigators went as far as to suggest that coronary perfusion itself in elderly hypertensive hypertrophic patients is negatively affected by an increase in heart rate (Fig. 9-17). They recommended that in the treatment of elderly hypertensive patients, the control of the heart rate in addition to the control of blood pressure might be helpful in minimizing the occurrence of myocardial ischemia and might slow down the subsequent progression to heart failure. Since then, their recommended treatment has become common practice with the use of β-blocker therapy to prevent MI and heart failure. In addition to maintaining resting heart rate in the physiologic range, pacemaker therapy preferably allows the heart rate to rise during exercise. A rise in heart rate with greater workload improves exercise capacity. Over the years a number of sensors have been incorporated into VVIR pacemakers. Many investigators have compared the exercise hemodynamics between sensor-driven VVIR pacing and VVI pacing. The percentage improvements of heart rate and exercise duration during symptoms-limited exercise in the VVIR and VVI pacing modes from a number of larger series for currently available rate-adaptive systems showed variable but consistent increases of 69% in rate and 32% in exercise tolerance when patients received VVIR pacing. Notably, most of the early comparative series of patients given VVI and VVIR pacing are rather small and have mixed patient populations in terms of symptoms of heart failure and degree of ventricular dysfunction.



9  Basic Physiology and Hemodynamics of Cardiac Pacing

External mechanical efficiency (%)

50

40

30

20

10

(r  0.56, p  0.02) 0 40

60

80

100

120

Beats/min Figure 9-17  The relationship between heart rate and external mechanical efficiency of the heart (measured as left ventricular work divided by myocardial oxygen consumption expressed as energy using a conversion factor of 2.06), for basal condition (closed circles) and different atrial pacing rates (open circles, 25% above basal heart rate; triangles, 50% above basal heart rate) in elderly hypertensive patients.

The use of an atrial sensing electrode to synchronize ventricular stimulation with intrinsic sinus node discharge (atrial synchronous ventricular pacing–VAT) allowed the implementation of a more physiologic form of pacing. The addition of ventricular sensing capability to this pacing mode resulted in the development of the VDD mode as a combination of VAT plus VVI. Both VVI and VDD modes of pacing prevent syncope due to bradycardia, but conventional VVI units are unable to respond to exertion by raising the pacing rate to meet the demand for a greater cardiac output. The importance of increasing heart rate to augment cardiac output with exercise has been clearly documented in several studies.132-134 Because the improvement in exercise performance is caused predominantly by an increase in heart rate, it is expected that favorable hemodynamic results will be obtained by rate-adaptive systems. Consequently, the modern approach to the management of complete symptomatic heart block involves the use of pacemaker units that provide not only basic ventricular pacing support, but also adaptation of the pacing rate to physiologic needs. A major concern, especially in compromised hearts, is that heart rate should not increase too much, because shortened diastolic filling time, reduced LV compliance at higher rates of ventricular pacing, and increased systemic vascular resistance may limit cardiac output.116,126 Rowe et al.135 found other reasons not to increase pacing rate too greatly. They examined cardiac output and coronary blood flow at low (47 bpm), intermediate (77 bpm), and high (117 bpm) ventricular pacing rates in patients with complete heart block. Cardiac output rose with an increase from low to intermediate pacing but diminished at the high rate. However, coronary blood flow and cardiac oxygen consumption rose progressively with increasing heart rates, thus indicating lower pumping efficiency at the high rate. OPTIMAL ATRIOVENTRICULAR INTERVAL Importance of Atrioventricular Synchronization Basically, the role of the atrial contraction is to help maintain a laminar flow of blood from the venous system to the ventricles; this role is

213

tightly integrated with the movement of the ventricles to create a smooth motion of blood across the system. To perform its role, the atrium changes its function from that of a conduit during early filling to that of a booster pump during atrial systole, to that of a reservoir during ventricular systole. These changes are closely related to longitudinal displacement of the anulus of the mitral and tricuspid valves along the long axis of the heart. Indeed, the anulus is displaced toward the apex of the ventricles during systole and toward the atria during diastole, thus significantly contributing to the smooth movement of blood from the atria to the ventricles. This displacement keeps the overall volume of the heart almost constant during systole and diastole.136 An optimally timed atrial contraction (before isovolumic contraction phase) maximizes LV filling and thereby its output by virtue of the Frank-Starling relation. Too early atrial contraction (as in long PQ times and dual-chamber pacing with long AV intervals) causes a loss of the booster pump function of the atrium. Moreover, early atrial contraction may initiate early mitral valve closure, thereby limiting ventricular diastolic filling time. This observation can be explained by the importance of three factors collaborating to achieve optimal closure of the AV valves. First, the ending of transvalvular flow at the end of the atrial contraction makes the valvular leaflets approach one another. Second, at the beginning of ventricular contraction, the anulus of the AV valves contracts, as do the papillary muscles that hold the leaflets. Simultaneously, at the start of ventricular contraction, ventricular pressure rises above atrial pressure, and the valves close. When these factors are misaligned, an opportunity for diastolic and systolic mitral regurgitation is created. In the early years of pacing, the electrical stimulus was selectively applied to the ventricle (ventricular single-chamber pacing, or V pacing). With V pacing, the contraction of atria and ventricles is uncoupled, leading to an atrial contribution to LV filling that varies from beat to beat. This also results in large beat-to-beat variations in stroke volume, systolic pressure, and other hemodynamic variables.137 The introduction of sequential AV pacing resulted in more regular heart beats and improved hemodynamics in both animals78,138,139 and patients.140 With advances in pacemaker technology, appreciation of the importance of maintaining AV synchrony has improved, and programming of optimal AV intervals in patients with dual-chamber pacemakers is of great importance.141 Figure 9-18 summarizes the main consequences of an optimal AV timing. Because filling time is limited, especially at high heart rates, the atrial contribution to stroke volume is more prominent at high than at low heart rates.142 Patients with Normal Left Ventricular Function The AV interval that maximizes resting cardiac output during dualchamber pacing varies widely among patients, usually reported as 125 to 200 msec. In patients with pacemakers, cardiac output increased by 4% to 20% when the AV interval was lengthened from 0 to between 100 and 130 msec.140,143-145 Optimal AV interval can be determined in most patients during dual-chamber pacing by means of various invasive and noninvasive measurements. Most often, optimal AV interval is assessed by LV outflow recording with Doppler echocardiography (Fig. 9-19). Doppler-derived cardiac outputs at varying AV intervals can be used to fine-tune the AV interval. In patients with preserved ventricular function, this method is highly reliable, is sensitive to small output changes, and has a low interobserver variability. The optimal AV interval determined with Doppler echocardiography correlates well with the optimal AV interval determined with radionuclide ventriculography. Factors that may influence the optimal AV interval in different patients are summarized in Table 9-2. In addition to interpatient variability, other factors may influence determination of optimal AV interval in the same patient, such as heart rate, paced or sensed atrial event, posture, and drugs.146-150 Moreover, in a specific patient’s lifetime, there may be situations in which the optimal AV interval is different from the AV interval considered to be physiologic. For example, AV

214

SECTION 2  Clinical Concepts

Acute and chronic effects of an optimal atrioventricular delay

• Maintains a laminar flow • Optimizes pumping action • Helps close auriculoventricular valves • Preempts diastolic regurgitation

• Minimizes atrial pressures • Maximizes efficiency of the pump • Optimizes filling time • Minimizes mean ventricular pressure required for optimum preload Chronic effects

Acute effects

• Minimizes ventricular diastolic pressures • Minimizes use of neurohormonal compensatory mechanisms • Minimizes ventricular preloads • Minimizes risk of compensatory remodeling • May decrease risk of stroke

• Increases preload and ventricular diastolic pressures • Increases stroke volume and cardiac output • Increases venous return • Prevents circulatory reflexes • Increases ventricular contractility through Starling effect

Figure 9-18  Hemodynamic effects of programming optimal atrioventricular (AV) delay.

sequential pacing for complete heart block complicating an acute MI or after cardiac surgery yields optimal hemodynamics at AV intervals in the range of 80 to 120 msec. This may be related to a surge of intrinsic catecholamine levels or to administered inotropic drugs and reduced LV compliance in these acutely stressful situations. The programmed AV interval regulates the timing of right atrial (RA) and RV conduction, which in a normal heart without significant atrial electrical disease usually coincides with that of left atrial (LA) and LV conduction. From a hemodynamic point of view, LA and LV conduction is most important, followed by the left-sided AV interval. The physiologic left-sided AV interval depends on several parameters, including the programmed AV interval, latency in atrial capture and sensing, interatrial conduction time, latency in ventricular capture, and interventricular conduction time.151 Because of the complex AVDI  50 ms FVI  14.9 cm

AVDI  175 ms FVI  19.0 cm

AVDI  300 ms FVI  15.3 cm

relationship among these intervals (see Fig. 9-19), significantly different right-sided and left-sided AV intervals in the same patient and in different patients can be recorded. Among these times, the one that most influences the optimal left-sided AV interval is the interatrial conduction time, measured as the time from the atrial pacing spike to the atrial depolarization on the esophageal electrogram or to the A wave on Doppler echocardiography. Patients with near-normal interatrial conduction time (≤90 msec) derive the greatest hemodynamic benefit from programmable AV intervals of about 150 msec; in contrast, in those patients presenting with prolonged interatrial conduction time, the AV interval should be set at 200 msec or longer to maximize cardiac output. Of particular note, programming a short AV interval in patients with prolonged interatrial conduction delay may result in depolarization of the LA after LV mechanical activation. This may produce hemodynamic values equivalent to or worse than those seen in patients with VVI pacing alone. Pacemaker syndrome may also occur under these circumstances. Latency in atrial capture and sensing may lead to different optimal AV intervals for sensed P waves and paced P waves (Fig. 9-20). The magnitude of atrial capture and sensing latencies varies among patients and may be affected by the lead and pacemaker circuitry characteristics, electrode position, tissue interface, amplitude and rate of stimulation, P-wave morphology, myocardial disease, electrolytes and other metabolic factors, and drugs. The optimal AV interval for a sensed P wave is about 30 to 40 msec shorter than the optimal AV interval for TABLE

9-2 

Possible Factors Influencing Optimal Atrioventricular Intervals

Type Interpatient

Figure 9-19  Left ventricular outflow recording at atrioventricular delay interval (AVDI; also AV interval) of 50, 175, and 300 msec during AV sequential pacing at identical heart rates in the same patient. The largest flow velocity integral (FVI) reflecting ventricular stroke volume occurs at AVDI 175 msec. Significant decreases are noted in stroke volume at AVDIs 50 and 300 msec. (From Janosik DL, Pearson AC, Buckingham TA, et al: The hemodynamic benefit of differential atrioventricular delay intervals for sensed and paced atrial events during physiologic pacing. J Am Coll Cardiol 14:499, 1989. Reprinted with permission of the American College of Cardiology.)

Intrapatient

Variables Atrial capture latency Atrial sensing latency Intra-atrial conduction delay Ventricular capture latency Intraventricular conduction delay Underlying myocardial disease Heart rate Catecholamine levels Heart rate Paced or sensed atrial event Catecholamine levels Drugs



9  Basic Physiology and Hemodynamics of Cardiac Pacing Paced P wave AVDI

Programmed AVDI  200 ms Effective PQ 1 60 ms

PQ

Sensed P wave-fixed AV delay AVDI

Programmed AVDI  200 ms Effective PQ  240 ms

PQ

Figure 9-20  Programmed atrioventricular delay interval (AVDI) and effect of PQ interval. Top, When a paced P wave occurs, the effective PQ interval is shorter than the programmed AV delay because of latency in atrial capture. Bottom, When a sensed P wave occurs, the effective PQ interval is longer than the programmed AVDI because of latency in atrial sensing. (From Janosik DL, Pearson AC, Buckingham TA, et al: The hemodynamic benefit of differential atrioventricular delay intervals for sensed and paced atrial events during physiologic pacing. J Am Coll Cardiol 14:499, 1989. Reprinted with permission of the American College of Cardiology.)

a paced P wave followed by a paced QRS at a similar heart rate (Fig. 9-21). However, differences between optimal AV intervals for a paced P wave and a sensed P wave may be as great as 100 msec in some individuals. New dual-chamber pacemakers incorporate programmable features designed to adjust the AV interval automatically in response to paced versus sensed P waves and in response to the atrial rate. It is important to recognize that the physiologic AV interval is never equivalent to the programmed AV interval for either a paced or a sensed atrial DVI II

A 150 V

event. The physiologic PQ interval begins with the onset of the P wave; however, with a sensed P wave, the programmed AV interval begins when the native P wave is sensed (so with a delay compared with onset of atrial activation), and with a paced P wave, the programmed AV interval begins with the atrial pacing spike, which is followed by further atrial depolarization. The importance of variation in AV interval during exercise in patients given physiologic pacing who have normal ventricular function is still controversial. Because the major determinant of augmentation of cardiac output during exercise is heart rate, the contribution of atrial systole and relative timing of the atrial and ventricular contraction may diminish in significance. Studies analyzing the effects of variation in AV interval during exercise have yielded conflicting results.152-156 There are two major limitations to these studies: the relatively small sample size of most and the use of a fixed AV interval during each exercise test, but a different AV interval between exercise efforts. There is evidence that a rate-adaptive AV interval, which is automatically decremented in response to an increased atrial rate, is beneficial to cardiopulmonary performance during exercise.157 In patients with chronotropic incompetence and high-level AV block, Sulke et al.158 performed randomized double-blind crossover assessments of rate-adaptive and different fixed AV interval settings during 2 weeks of normal activity and performance on an exercise treadmill in patients undergoing DDDR pacing. There was a subjective improvement in quality of life with rate-adaptive AV intervals compared with fixed AV intervals, and patients preferred the rate-adaptive settings (Fig. 9-22). Notably, the longest AV interval (250 msec) was least preferred and was associated with the highest symptom prevalence. Moreover, exercise duration was not significantly different in any setting in DDDR mode but was significantly reduced in DDD mode.158 Patients with Depressed Left Ventricular Function In general, relaxation is slower in failing hearts than in normal hearts. Consequently, less blood enters the ventricles during the rapid filling phase, as can be observed from lower E waves on Doppler echocardiograms. Therefore, failing hearts are more dependent on properly timed atrial contraction than normal hearts. This has been demonstrated in comparisons of patients with decompensated heart failure and compensated nonvalvular heart disease.125 This knowledge clarifies why

LA LA V 35

A pace M

V pace

VDD II

A 150 V LA

Esoph lead

LA V 110

A sense M V pace

100

LA V 35

LA V 110

Left A-V extension in VDD 75 (110-35)

Figure 9-21  Interatrial conduction and change in the timing of left atrial (LA) depolarization with pacing mode change from DDD to VDD. The time delay between right atrial pacing artifact and LA depolarization is 115 msec (top), resulting in an LA-to-ventricular sequence of only 35 msec. With mode change to VDD (bottom), the LA-to-ventricular sequence extends by 75 msec, resulting in an LA-to-ventricular sequence of 110 msec. (From Wish M, Fletcher RD, Gottdiener JS, Cohen AI: Importance of left atrial timing in the programming of dual-chamber pacemakers. Am J Cardiol 60:566, 1987.)

Visual analog scale score (%)

Esoph lead

215

90

*

80

70

*

60

50 125 ms

175 ms

250 ms

RR AVD

AV delay setting in DDDR mode Figure 9-22  Exercise capacity ( ) and patient’s perception of general well-being ( ) during everyday activity at fixed AV delay settings of 125, 175, and 250 msec and rate-responsive AV delay (RR AVD) settings in DDDR mode. , P < .05; m, P < .03; *, P < .01. (Modified from Sulke AN, Chambers JB, Sowton E. The effect of atrioventricular delay programming in patients with DDDR pacemakers. Eur Heart J 13:464, 1992.)

216

SECTION 2  Clinical Concepts

ATRIAL PACE

AV PACE 60

AV PACE 180

.2 m/s 150

CO 5.2 l/m

CO 2.4 l/m

CO 3.0–1/m

0

A

B

C

Figure 9-23  Mitral flow-velocity curve and simultaneous left atrial (LA) and left ventricular (LV) pressure curves in a patient with severe LV dysfunction and long PR interval. A, Atrial pacing with antegrade native conduction and long atrioventricular (AV) delay interval (also AV interval). There is an increase in LV pressure above LA pressure during atrial relaxation and middiastole (arrowhead), resulting in a shortening of diastolic filling time and the onset of diastolic mitral regurgitation (m/s, meter per second). The baseline cardiac output (CO) is 3 L/min (l/m). B, AV pacing at short AV delay interval of 60 msec. Diastolic filling occurs throughout all of diastole. Atrial contraction occurs simultaneously with LV contraction, resulting in a reduction in cardiac output and an increase in mean LA pressure compared with A. C, Atrioventrucular (AV) pacing at the optimal AV delay interval of 180 msec. There is now an optimal relationship between the mechanical LA and LV contractions, so that diastolic filling period is maximized and the mean LA pressure is maintained at a low level, resulting in a rise in LV diastolic pressure and improvement in CO to 5.2 L/min. (From Nishimura RA, Hayes DL, Holmes DR: Mechanism of hemodynamic improvement by dual chamber pacing for severe left ventricular dysfunction: An acute Doppler and catheterization hemodynamic study. J Am Coll Cardiol 25:281, 1995.)

optimal timing of atrial and ventricular stimulation in patients with pacemakers is of special interest in patients with heart failure. Some investigators have shown that pacemakers can actually improve the coupling between atria and ventricles over that without pacing. This issue applies to patients with excessively long PQ times. In two studies of subjects with these findings, hemodynamic benefit was achieved with AV sequential pacing at a physiologic AV interval159,160 (Fig. 9-23). Reducing the prolonged AV interval increased diastolic filling time and reduced diastolic mitral regurgitation (Fig. 9-24). These beneficial effects are striking, because these studies were performed before the era of resynchronization, and the RV apex was used as a ventricular pacing site, presumably the least preferable site from a hemodynamic point of view (see later). The actual role of the atrial contraction in patients with dilated ventricles and low ejection fraction varies among individuals. The booster pump action of the LA is noticeable as a “shoulder” in the LV pressure tracing, and the LV does not start contracting immediately after atrial contraction (Fig. 9-25). If this plateau lasts too long, diastolic mitral regurgitation could occur. About 20% of the patients in

I AEG VEG 20

PCW

PCW 0 AV pace

V pace

Figure 9-24  Surface electrocardiographic lead I, atrial electrogram (AEG), ventricular electrogram (VEG), and pulmonary capillary wedge pressure (PCW) recordings from a single patient during atrioventricular (AV) pacing (AV pace) and ventricular pacing (V pace) at 80 bpm with an AV interval of 150 msec. The cannon A wave is noted (arrow) on the PCW tracings. (From Reynolds DW: Hemodynamics of cardiac pacing. In Ellenbogen KA [ed]. Clinical cardiac pacing, pp 120-161. Cambridge, Blackwell Scientific, 1992. Reprinted by permission of Blackwell Scientific Publications.)

heart failure studies for whom complete hemodynamic data are available show this type of diastolic LV pressure waveform.161-163 Figure 9-25 shows that using an optimized AV interval, in this case with BiV pacing, results in coincidence of the peak of the booster action of the atrial contraction and LV pressure development. Auricchio et al.164 demonstrated that the maximum improvement in aortic pulse pressure, independent of the site being paced, occurs when this coincidence is achieved. Examination and calculation of cardiac output with Doppler echocardiography, as frequently performed in patients with dual-chamber pacemakers, is more problematic in patients with depressed ejection fraction in whom the value of the effect of pacing may be near the rate of uncertainty in the method.165 Thus, the number of repetitions required to obtain a clinically meaningful and reproducible approximation of the optimal AV interval makes this method, as single assessment, almost impractical for patients with depressed ejection fraction. The most utilized echocardiographic parameters for optimizing AV interval in patients with heart failure include echocardiography-based inflow or preload evaluations (e.g., velocity-time integral of transmitral flow or EA duration)166 and measures of systolic function using time-velocity integral at the LV outflow tract.167,168 Comparison of these different methods with invasive LV dP/dtmax to optimize AV interval in patients with CRT showed that mitral inflow velocity-time integral proved most accurate. The formerly used Ritter method, based on calculating the optimal AV timing by analyzing filling patterns obtained during only two AV intervals, has been found to be accurate to optimize AV interval in CRT patients.169 Indeed, this method was created for patients with high-degree block, and some caution is advised in patients with intact AV conduction.170 In these patients the electromechanical delay is not constant and varies as a function of the extent of collision between the intrinsically conducted depolarization wavefront and the wavefront generated by the artificial pacing spike. Therefore, in patients with an intact conduction system, the extent of collision between these two wavefronts will be different at the long and short AV delay intervals tested in the Ritter method, making the assumption required by the method completely invalid—that is, the electromechanical delay is unchanged at any tested AV delay. Another consideration is important when deciding whether to use an inflow-based method for optimization of the AV interval. In patients with ventricular conduction delay who have CRT devices, the echocardiographic inflow-based method of optimizing AV delay interval



9  Basic Physiology and Hemodynamics of Cardiac Pacing

217

Atrial electrogram Ventricular electrogram

232 ms Figure 9-25  The atrial electrogram, right ventricular electrogram, left ventricular (LV) pressure, and aortic pressure tracings of a patient with New York Heart Association (NYHA) Class III heart failure from the Pacing Therapies for Congestive Heart Failure (PATH-CHF-II) trial.149 Note the effect of programming an AV interval of 128 msec in the biventricular device. With this AV interval, the ventricular pressure starts right at the peak of the pressure increase created by the booster action of the left atrium. This AV interval provides the maximum preload at the minimal mean LV diastolic pressure while maximizing the time available for diastole, thus enabling higher heart rate increases without compromising filling time.

128 ms

Aortic pressure

LV pressure Atrial booster pump effect

neither optimizes the synchrony of the ventricular contraction nor accounts for how much preexcitation of the late-activated region is required to restore intraventricular synchrony (see later). ROLE OF VENTRICULAR SYNCHRONY Abnormal asynchronous activation causes abnormal contraction patterns, inefficient and depressed pump function, and ventricular remodeling. Wiggers30 recognized the importance of normal electrical activation of the ventricle for optimal pump function in 1925 (Fig. 9-26). However, it took about 70 years for the concept to gain broad interest. Two teams of investigators regarded the abnormal activation as being of little importance, because AV sequential pacing allowed for significantly better cardiac pump function than V pacing.137,138 Nevertheless, one group showed better pump function during LV apex pacing than during pacing at RV sites.138 In the 1960s, Kosowsky et al.139 compared RV apex pacing with His bundle pacing, the latter maintaining the normal activation but allowing variation of the AV interval. They concluded that AV synchrony and proper sequence of activation are equally important, as confirmed by later animal and human studies.76,77,144 Clinical Consequences of Asynchronous Activation The combination of acute adverse hemodynamic effects and long-term ventricular remodeling (see Fig. 9-14) may explain why abnormal

Figure 9-26  Original recording made by Wiggers, showing that in a paced beat (the second; note the widened QRS complex), left ventricular and aortic pressures are lower than in the preceding and sequential beats. (Modified from Wiggers CJ: The muscular reactions of the mammalian ventricles to artificial surface stimuli. Am J Physiol 73:346378, 1925.)

Ventricular pressure raise starts at peak of atrial booster pump effect

electrical activation and asynchronous electrical activation have major implications for a patient’s clinical status. Studies have shown that morbidity and mortality are higher in patients with long-term RV apex pacing than with atrial pacing.145,171-173 In patients with sinus node disease and good ventricular function, the risk of heart failure was significantly high after more than 7 years in a comparison of atrial pacing and RV pacing.173 In a similar population the risk of hospitalization for heart failure within 3 years increased with percentage of time patients underwent pacing at the RV apex.174 Interestingly, development of heart failure and atrial fibrillation was more sensitive to percentage of pacing than to pacing mode (single- or dual-chamber pacing). In patients who received an ICD, the incidence of heart failure was higher within a year in those with pacing at VVI 70 bpm rather than in the backup mode.175 Although pacing and, even more, heart failure started at a later age in the previous studies, RV pacing in children can also induce impaired pump function. Tantengco et al.123 observed such adverse effects after approximately 10 years of pacing. The dramatic effects of ventricular pacing in otherwise healthy myocardium may be explained by the pronounced structural changes in juvenile hearts.107,176 The deleterious effect of ventricular pacing is relatively easy to discover when start of pacing is exactly known, although spontaneous evolution in a dilated cardiomyopathy cannot be excluded. Whether adverse effects of asynchronous activation are also applicable to LBBB has been less clear, because LBBB is a silent event that is often accompanied by considerable comorbidity.177-179 Longitudinal studies determining cardiovascular mortality and morbidity show that the presence of LBBB always carries a poor prognosis. In a 29-year follow-up study of 3983 pilots, the morbidity and cardiovascular mortality rate among those showing signs of LBBB was 17.2%, and the most common clinical event observed was sudden death without any previous symptoms (17%). These percentages are 10 times higher than those in subjects without LBBB.180 In the Framingham Study, cumulative cardiovascular mortality over 10 years was approximately five times higher in patients with LBBB than in those without LBBB.181 In a population of 110,000, the risk for development of cardiovascular diseases was 21% in the 112 patients with LBBB, compared with 11% in age- and gender-matched controls. Moreover, cardiac mortality was strongly increased in the patients with LBBB. The first direct evidence for an effect of LBBB on pump function came from studies in patients with intermittent LBBB.182,183 Further proof of the negative effect of LBBB on hemodynamics has come from later studies in an animal model of LBBB, created by ablation of the proximal part of the LBB.54,184,185 After ablation, duration of the QRS complex, asynchrony of RV-LV contraction, and asynchrony of

218

SECTION 2  Clinical Concepts

electrical activation within the LV all increase significantly; also, relative to normal conduction values in the same species, these increases are similar in degree to the increases seen in patients with LBBB.24,186 This finding suggests that, although the LBBB in patients may not be caused exclusively by proximal lesions, this model closely resembles the LBBB in patients. All studies on this animal model of LBBB show that LBBB reduces systolic and diastolic LV function and induces paradoxical septal wall motion.54,184,185 Therefore, both RV apex pacing and LBBB seem to be conditions that increase the risk of heart failure and cardiac death, especially in patients with already-compromised function. For this reason, efforts to prevent RV apex pacing or correct LBBB are increasing. With current knowledge of electrical impulse conduction, the best solution may depend on whether the heart has normal or disturbed intrinsic conduction within the ventricles. In the following discussion, various possible strategies are described, with current and possible future applications. Strategies to maintain synchrony of activation are distinguished from those to restore synchrony. Strategies to Maintain Ventricular Synchrony Approaches to achieve or maintain ventricular synchrony include atrial pacing and pacing from alternative ventricular sites, including the His bundle. More recently, LV pacing and BiV pacing have been specifically indicated in patients with ventricular conduction disturbance. Both methods are commonly referred as “cardiac resynchronization therapy” and are discussed separately later. Atrial Pacing.  Atrial pacing maintains normal ventricular activation, although limited to patients with intact AV nodal conduction. In a comparison of atrial pacing with RV apex pacing in patients with sinus node disease (e.g., VVI vs. AAI), Andersen et al.173 reported less development of all-cause mortality and cardiovascular death, better maintenance of ventricular function and perfusion,187 absence of atrial dilatation,188 and a lower incidence of atrial fibrillation.173,189 Although earlier studies compared atrial pacing with single-chamber pacing, later studies in patients with sick sinus syndrome showed a similar disadvantage of ventricular pacing in the dual-chamber mode compared with atrial pacing. Because the risk of AV block in this patient category (0.6%-1.7%/yr)190,191 is considerably lower than that for development of heart failure (as high as 12%),174 atrial pacing is an immediately applicable, easy, and effective way to avoid abnormal ventricular activation. Alternative-Site Ventricular Pacing.  Patients with disturbed AV conduction are pacemaker dependent and require ventricular pacing. His bundle pacing is the way to maintain physiologic ventricular impulse conduction. Animal studies showed that QRS duration was shorter with pacing in the high ventricular septum than with RV apex pacing.77,176,192 In the 1960s, Scherlag’s group showed that His bundle pacing results in the same QRS duration and pressure development as sinus rhythm and atrial pacing and in better hemodynamics than RV apex pacing.139,193 In 2000, Deshmukh et al.194 first reported successful His bundle pacing in patients on a permanent basis. Using improved implant tools, Zanon et al.195 showed that His bundle pacing preserves normal distribution of myocardial blood flow and mechanics. Nevertheless, because of the relatively difficult positioning of the pacing lead into the His bundle, pacing in the vicinity of His bundle is more often used. The literature refers to this position as “high RV septal” or “RV outflow tract” (RVOT), but in some studies, “RVOT” appears to imply pacing at the high RV free wall. Pacing in this region leads to abnormal QRS configuration compared with atrial pacing. Comparative studies of the acute hemodynamic effects of high septal pacing and RV apex pacing showed a moderate beneficial effect of septal pacing on LV pump function, although several studies found no significant difference between high septal and RV apex pacing.196 Combined RV apex and RVOT pacing in patients resulted in shorter QRS duration than RVOT pacing alone but did not lead to further improvement of cardiac output.197

Longer-term studies appear to show a more consistent beneficial effect of high septal pacing. Tse et al.198 randomly assigned 24 patients to either RV apex or high septal pacing. After 18 months of pacing, perfusion defects and regional wall motion abnormalities were less common and ejection fractions were higher in the high septal pacing group. The reports explicitly mentioned that the “RVOT” lead was positioned in the interventricular septum. Unfortunately, however, the exact best site at the RV septum is not well known. One animal study showed that the RV site for best hemodynamic function differs for individual hearts.94 Also, RV midseptal pacing performed similar to RV apical pacing.59 In contrast, single-site LV pacing often results in better cardiac pump function than RV pacing.91,199 Turner et al.200 showed that LV pacing may also lead to more synchronous activation if pacing is performed at longer AV intervals during maintained AV conduction in patients with a relatively narrow QRS complex. Because electrical asynchrony and mechanical asynchrony within the LV wall are similar in RV pacing and LV pacing201 (Fig. 9-27), the less detrimental hemodynamic effect of LV pacing may be explained by better interventricular coupling. This mechanism may also partially account for the beneficial hemodynamic effects of LV pacing in patients with heart failure and AV node ablation.202 In these patients, no fusion can occur between the impulse wavefronts from endogenous conduction and LV stimulation. However, long-term comparative data in some subgroups of patients are still missing, so no definitive conclusions can be drawn. Animal studies have shown that within the left ventricle, the septum and apex may be ideal pacing sites; pacing in these sites was associated with LV pump function close to that during normal sinus rhythm, even

Late diastole

Early systole

Mid systole

End systole

Midseptum RA

RVa LV

RVa

LV

Figure 9-27  Three-dimensional reconstruction of left ventricular wall with myocardial strain represented in color. Data were obtained using magnetic resonance imaging (MRI) tagging in a normal dog heart. Blue indicates contraction (negative circumferential strain), red indicates the reference state (end diastole), and yellow indicates stretch (positive circumferential strain). Data are shown for late diastole and early systole, midsystole, and late systole for right atrial (RA), biventricular (RVa+LV), RV apex (RVa), and left ventricular (LV) pacing. White arrow points to the midseptum. Note the considerable strain differences during both RVa and LV pacing in early systole and midsystole. RVa + LV pacing reduces these differences, but differences are still more pronounced than during RA pacing. (Modified from Wyman BT, Hunter WC, Prinzen FW, et al: Effects of single- and biventricular pacing on temporal and spatial dynamics of ventricular contraction. Am J Physiol 282:H372H379, 2002.)



9  Basic Physiology and Hemodynamics of Cardiac Pacing

TABLE

9-3 

Hemodynamic Parameters during VDD Pacing at Various Ventricular Sites in Dogs* Ventricular Site

Parameter Heart rate Stroke volume Stroke work LV dP/dtpos LV dP/dtneg Tau ESLVP EDLVP QRS duration

RV Apex 1.048† 0.930†‡ 0.789†‡ 0.931†‡ 0.799†‡ 1.1† 0.952† 1.137† 1.862†‡

LV Lat 1.03 0.991‡ 0.980‡ 0.926†‡ 0.840† 1.189†‡ 0.978 1.231† 1.997†

LV Post 1.02 0.873†‡ 0.869‡ 0.927†‡ 0.849† 1.199†‡ 0.983 1.215 1.820†‡

LV Apex 1.03 1.05 1.08 0.99 0.94 1.06 0.96 1.3 1.940‡

LV Septum 1.01 1.03 1.00 1.00 0.93 1.04 0.99 1.22 1.670†

Data from Peschar M, de Swart H, Michels KJ, et al: Left ventricular septal and apex pacing for optimal pump function in canine hearts. J Am Coll Cardiol 41:1218-1226, 2003. *All values are expressed relative to the corresponding value during sinus rhythm. For all parameters, values during left ventricular (LV) apex and LV septum pacing were not significantly different from those in sinus rhythm, except QRS duration. Right ventricular (RV) apex pacing leads to the most dramatic reduction in LV systolic and diastolic function. †P < .5 compared with SR apex pacing. ‡P < .05 compared with LV apex pacing. ESLVP/EDLVP, End-systolic/end-diastolic left ventricular pressure; LV Lat, lateral left ventricle; LV Post, posterior left ventricle.

with pacing at short AV intervals94,203 (Table 9-3). Studies in children now support the superior performance of LV apex pacing over RV pacing with respect to hemodynamic function93 and echocardiographic indices of dyssynchrony204 and reverse remodeling.205 LV septal pacing is not yet clinically feasible, because positioning a lead in the LV cavity carries a risk of embolization and stroke.206 In an animal study, Mills et al.59 advanced a lead from the right ventricle through the interventricular septum to reach the LV septal endocardium. Even after 4 months of pacing, LV septal pacing provided cardiac pump function and efficiency similar to that during natural conduction. The degree of asynchrony induced by single-site ventricular pacing can be reduced by pacing at two or more sites simultaneously, preferably opposite to each other. This “multisite” pacing, however, reduces QRS duration by no more than 20% compared with single-site pacing.50 Animal studies showed that BiV pacing does lead to better LV pump function than RV pacing.50,94,207 Also, in patients with atrial fibrillation and AV node ablation, hemodynamic performance was better during BiV pacing than during single-site LV or RV pacing.208 It is important to note that in patients with normal ventricular conduction, such novel pacing sites (LV apex, RVOT, His bundle, multisite) at best prevented a reduction of pump function compared with atrial pacing or sinus rhythm, but never improved pump function. Also, left intraventricular mechanical asynchrony is only partially normalized, even with BiV pacing.209 Multisite pacing cannot outmatch the quick impulse conduction and its extensive spread through the network of Purkinje fibers. Therefore, avoiding any ventricular pacing would be preferable in hearts with normal ventricular conduction. Minimizing Ventricular Pacing One approach to avoiding any type of ventricular pacing has been to program fixed, long AV intervals; however, this approach resulted in occasional retrograde AV conduction and higher risk of arrhythmias in one third of a select population.210 For many years, some pacemakers have incorporated an algorithm based on the programming of two AV delays: a short, physiologic AV delay and a longer AV delay. These devices start pacing with the longer AV delay and continue at this setting if they find intrinsically conducted ventricular beats. If the intrinsic conduction is not present or fails to occur, the devices automatically switch to the shorter and more physiologic AV delay. When functioning at the shorter AV delay, these devices periodically extend the AV delay to test for intrinsic conduction at programmable intervals. This type of algorithm is called AV search hysteresis.211

219

Rather than prolonging the AV delay, an algorithm called managed ventricular pacing (MVP) has been implemented in two studies.212,213 With this algorithm, when intrinsic activity is present, only the atrium is paced; the ventricle is monitored on a beat-to-beat basis to verify intact AV conduction. Only in cases of transient and persistent loss of AV conduction does the MVP algorithm pacemaker switch to the DDDR mode. The mode intermittently tests for return of normal AV conduction. During atrial fibrillation, the device operates in the DDIR mode. With this approach, the cumulative percentage of ventricular pacing was reduced from a mean of 74% to 4%, with 80% of the patients being paced on the ventricle for less than 1% of the time. In a study with 1065 patients, MVP reduced the risk of developing persistent atrial fibrillation from 12.7% to 7.9% over a mean follow-up of 1.7 years.214 Because many patients undergo pacing for sinus node disease, algorithms such as AV search hysteresis and MVP are promising for maintenance of ventricular synchrony in patients with predominantly intact AV and intraventricular conduction (Fig. 9-28). However, a recent study indicates that conservative programming of lower rate and rate response is advisable to minimize risk of AV decoupling and uncoupling as well as ventriculoatrial (VA) coupling.215

Strategies to Restore Synchrony: Cardiac Resynchronization Therapy Biventricular pacing devices, first implanted in March 1993,216 were approved by the U.S. Food and Drug Administration (FDA) in 2001. The beneficial hemodynamic effects of BiV pacing during that smallscale study have been reproduced many times.161,217-220 The consistency of the results of small but prospective randomized trials, such as Pacing Therapies for Congestive Heart Failure (PATH-CHF),161,162 Multisite Stimulation in Cardiomyopathy (MUSTIC),221 Cardiac Resynchronisation in Heart Failure (CARE-HF),222 and Multicenter InSync Randomized Chronic Evaluation (MIRACLE),223 have removed any doubt

Ventricular conduction

Normal

Abnormal

AV conduction

Yes

No

Atrial MVP

His bundle LV (apex) RV septal BiV (especially if RVA pacing)

Discontinue RVA pacing if present

BiV LV

(Check mechanical asynchrony) Figure 9-28  Flow diagram showing most physiologic mode and/or site of pacing. Potential adverse effects of ventricular pacing and conduction disturbance, such as left bundle branch block (LBBB), on cardiac pump function. Ventricular pacing can induce nonphysiologic atrioventricular (AV) as well as interventricular (biventricular, BiV) and intraventricular asynchrony. Assuming normal PQ times, LBBB causes only abnormal interventricular and intraventricular asynchrony. Effects of AV asynchrony are explained in text; LV, left ventricular; MVP, managed ventricular pacing; RVA, right ventricular apex.

220

SECTION 2  Clinical Concepts

LAO 60°

Y

111 msec

Y

137 msec

52 msec

46 msec

2

X

Total Ventricular Activation Time: 157 msec RV Activation Time: 57 msec LV Activation Time: 105 msec

Y

2

X

Total Ventricular Activation Time: 189 msec RV Activation Time: 85 msec LV Activation Time: 137 msec

95 msec

110 msec

2

Total Ventricular Activation Time: 205 msec RV Activation Time: 130 msec LV Activation Time: 145 msec

Figure 9-29  Color-coded electroanatomic isochronal maps of right ventricular (RV) and left ventricular (LV) activation in three patients with heart failure and left bundle branch block. The earliest ventricular activation site is recorded at the RV anterolateral region (red spot) in all three cases. After almost 40 to 60 msec, a single LV septal breakthrough site is noted (not shown). The latest-activated region is the LV posterolateral wall (blue to purple isochronal lines). The total activation endocardial time and endocardial RV and LV activation times are shown for each patient.

about the feasibility and clinical efficacy of CRT in highly symptomatic patients with heart failure. Large, multicenter, randomized controlled clinical trials (RCTs) have consistently demonstrated significant reductions in mortality and hospitalization rates with CRT in patients who had moderate to advanced heart failure despite optimal pharmacologic therapy and who presented with ventricular conduction delay, sinus rhythm, and depressed LV ejection fraction.222 PHYSIOLOGY OF CARDIAC RESYNCHRONIZATION THERAPY With the introduction of CRT, the use of cardiac pacemakers has been transformed from “just” artificial maintenance of the heart rhythm to restoration of pump function. Although the primary objective of BiV pacing may be to reduce intra-LV asynchrony, it can also be used to establish optimal interventricular coupling as well as AV coupling. The latter especially applies to left-sided AV coupling, because in LBBB the atrio-LV delay is usually prolonged. Parsai et al.224 showed that patients respond to CRT if at least one of the mechanisms of CRT is active.224 The predominant intraventricular conduction abnormality in about one third of patients with heart failure is LBBB, where activation spreads from the RBB to the RV wall and, after transseptal conduction, within the LV from the septum to the LV lateral wall (Fig. 9-29). Thus, pacing at the LV lateral wall is used to create an activation wavefront, which moves in the opposite direction from a spontaneously occurring activation wavefront.16,26,201 Consequently, during BiV pacing, two activation wavefronts are generated and merge approximately in the middle.36,209 A similar effect is obtained by single-site LV pacing using an AV interval that allows merging of the intrinsic activation originating from the RBB with the wavefront derived from the LV pacing lead. This merging wavefront leads to less electrical asynchrony in comparison with LBBB.36 Because of the tight E-C coupling in the heart, CRT also improves coordination of contractions between the cardiac chambers and within the LV wall, as shown with different imaging techniques55,225 (see Fig. 9-27). More specifically, CRT uniformly prolongs the time to maximum contraction55 and thus restores a more coordi-

nated contraction pattern, resulting in a more homogeneous distribution of regional loading conditions.56 The improved synchrony leads to improvement of cardiac pump function, as determined by LV dP/dtmax, pulse pressure, cardiac output, and ejection fraction.* Such improved systolic pump function is achieved at unchanged or even lower filling pressures, denoting a true improvement of ventricular contractility through better coordination of contraction. Moreover, Nelson et al.69 showed that better coordination of contraction improves mechanical pump function while slightly decreasing myocardial energy consumption, suggesting that CRT improves the efficiency of the cardiac pump. Further improvement in pump function may be mediated by a reduction in mitral regurgitation225 and prolongation of diastolic filling time. These beneficial effects occur almost immediately after the start of resynchronization.36,161 Another possible mechanism for the improvement in LV pump function by BiV and LV pacing is that it improves ventricular interaction and relieves external constraint,228,229 a mechanism that could work even in patients with a narrow QRS complex.229 The idea is that early LV contraction also allows early LV filling. In hearts with high central venous pressures, where presumably the RV occupies much of the pericardial space, the early LV activation may allow the LV to fill first, and thereby better, thus recruiting more of its Frank-Starling mechanism.228 A variety of cardiac and extracardiac processes triggered by CRT are responsible for its long-term beneficial effect. First, the improved pump function reduces neurohumoral activation, an effect evidenced by an increase in heart rate variability and a reduction in plasma brain natriuretic peptide (BNP) levels.230 Furthermore, the improved contractility and pump efficiency at a smaller end-diastolic volume reduce mechanical ventricular stretch. This latter reduction and the probably associated reduction in neurohumoral activation may well explain the beneficial reverse-remodeling effect of CRT.225,231 As elegantly demonstrated by Yu et al.,225 such reverse remodeling points to structural *References 36, 161, 219, 220, 226, 227.



9  Basic Physiology and Hemodynamics of Cardiac Pacing

Resynchronization

patients with only mild heart failure or impairment of LV function.233,234 These data indicate that resynchronization cures “dyssynchronopathy” independent of the severity of heart failure. Most CRT studies focus on the left ventricle, but a few indicate that RV function may also be improved by better synchronization of the electrical activation. This would especially apply to right bundle branch block (RBBB). Three-dimensional electroanatomic mapping data have provided the rationale for implementing CRT in these patients.235 However, rather than focusing on LV-based pacing, resynchronization of RBBB requires RV pacing. This also applies to children with repaired congenital defects such as tetralogy of Fallot, who present with rightsided heart failure.236,237

Desynchronization

1000

dP/dt (mm Hg/s)

900

*

800 † 700

†

DETERMINATION OF MECHANICAL ASYNCHRONY

600

†

500

Reverse remodeling

400 225 † Left ventricular volume (mL)

221

200

*

* *

175

*

150 †

*

125

CRT

100 Baseline 1wk

1mo

*

3mo

* * off- off-1wk off-4wk immed

Figure 9-30  Acute hemodynamic and long-term reverse-remodeling effects of cardiac resynchronization therapy (CRT). (Modified from Yu CM, Chau E, Sanderson JE, et al: Tissue Doppler echocardiographic evidence of reverse remodeling and improved synchronicity by simultaneously delaying regional contraction after biventricular pacing therapy in heart failure. Circulation 105:438-445, 2002.)

improvement in the myocardium. These investigators showed the increasing reverse-remodeling effect of resynchronization over time by measuring the time course of LV cavity volume and LV dP/dtmax during the first 3 months after the start of CRT and 1 month after CRT was temporarily stopped, when some beneficial effect of CRT was still noticeable, indicating tissue adaptation (Fig. 9-30). Study in canine LBBB hearts showed that isolated LBBB induces LV dilatation as well as hypertrophy and that these derangements are almost completely reversed by BiV pacing.54,57 Interestingly, BiV pacing reversed the hypertrophy in the LV wall, especially in the late-activated LV lateral wall.57 This reduction in hypertrophy is important because hypertrophy gives rise to various molecular changes,232 resulting in greater risk for contractile failure and arrhythmias. Indeed, in dogs with LBBB, where heart failure was induced by 6 weeks of rapid atrial pacing, calcium homeostasis was affected most in the LV lateral wall.114 Notably, both hemodynamic improvement36,185 and reverse remodeling57 have been observed to a similar extent in dogs with LBBB (with and without heart failure) and in patients with primary and secondary dilated cardiomyopathies. Although CRT was originally indicated for patients with severe heart failure (NYHA Class III and IV), several studies demonstrated that CRT can also induce reverse remodeling in

The primary goal of CRT is to restore atrioventricular, interventricular, and intraventricular synchrony. Because QRS duration is a poor predictor of the response to CRT,225,238,239 mechanical asynchrony was thought to be a better index for selection of CRT patients. Although studies have reported that measurement of mechanical asynchrony is a better predictor of the response to CRT than duration of the QRS complex, most were small and observational. Interventricular asynchrony can be determined from the time difference between pulmonary and aortic valve openings with conventional echocardiography186 or from the time to peak velocity in the RV and LV walls on tissue Doppler imaging (TDI). Other approaches are determination of the phase difference of RV and LV wall motion using nuclear imaging53 and use of the time or phase difference with increased RV and LV pressures.184,240 For practical reasons, the echocardiographic approach is most often used. Intraventricular asynchrony usually requires more sophisticated equipment. The most detailed mechanical activation maps have been obtained through mapping of myocardial shortening (strain) patterns with magnetic resonance imaging (MRI) tagging48,201,209,241 (see Fig. 9-8). These studies have shown that the pattern of deformation gradually changes with increasing distance from the site of earliest activation (asterisk in Fig. 9-8). With the less expensive noninvasive method of M-mode echocardiography, the timing difference of maximal inward motion at the septum and posterior wall, or the septal-posterior wall motion delay (SPWMD),238 can be calculated. The disadvantages of this simple technique are that peaks of septal wall motion may not be precisely defined, and M-mode sections perpendicular to the septum may be difficult to acquire. With color TDI and with pulsed-wave TDI, the time to peak systolic velocity can be measured in different regions of the myocardium.225,242,243 Local contraction can be derived only through subtraction of velocities in nearby regions. Such analysis on digitally stored color-coded tissue Doppler images yields myocardial strain and strain rate.97 Such information is equivalent to that obtained with MRI tagging, but with the great advantage that TDI can also be employed in the presence of a pacemaker. Strain measurements differentiate better than TDI between active systolic contraction and passive displacement. Likewise, Breithardt et al.97 demonstrated more consistent behavior of strain than of velocity signals. However, image acquisition and analysis are timeconsuming and operator dependent, and subtraction of two strain rate signals can create inaccuracies. Therefore, the clinical applicability of strain-rate imaging is limited. A more recent development is the measurement of two-dimensional strain using speckle tracking. This technique uses the same principle as MRI tagging: measurement of the deformation of the tissue by displacement of markers in the tissue. Rather than markers artificially applied, as in MRI tagging, speckle tracking uses the differences in gray level naturally occurring in cardiac tissue as markers. Depending on the echocardiographic section used, longitudinal, circumferential, or radial strains can be measured. Although this technique shows promise and may have less technical limitations than TDI, the various echocardiographic systems appear to handle the strain signals differently in terms of tracking and spatial and temporal filtering, and the

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Strain

0

?

?

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+

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equipment allows only minimal interference by the user. Also, only a few studies showed a true validation of the technique, and this was not in conditions comparable to LBBB or CRT. Nevertheless, speckle-tracking measurements have provided promising results. Mechanical dyssynchrony determined by this technique appeared to predict the response to CRT better than TDI,244 at least partly because strain measurement by speckle tracking is independent of the angle between ultrasound beam and the ventricular wall. TDI measures the velocity of displacement of the myocardial segments relative to the ultrasound transducer and in the direction of the ultrasound beam. Likewise, the longitudinal velocity, as usually determined in basal regions, is the sum of all velocities developed in all regions between the transducer and a particular area. Analysis of strain signals may also be less difficult to interpret than the frequently multiphasic velocity signals.245 TDI mechanical asynchrony is also observed in a considerable number of healthy volunteers. Mechanical Asynchrony or Discoordination? With mechanical asynchrony the focus over the last decade, the results of two randomized multicenter studies were unexpected. The PROSPECT and RethinQ studies showed poor predictive value for mechanical asynchrony as an index of CRT response in patients with wide and narrow QRS complex.246,247 These disappointing results can be explained only partly by technical limitations of the TDI technique. Even when using the “gold standard” technique in cardiac deformation measurements, MRI tagging, determination of mechanical asynchrony (i.e., timing difference in onset or peak shortening) did not predict CRT response.248,249 A better approach for achieving a mechanical predictor of CRT response may be to analyze discoordination of contraction. Whereas measurement of asynchrony requires determining the onset or peak of shortening in a certain region, discoordination is assessed by analyzing the strain curves over a large period of the cardiac cycle (e.g., systole or diastole). Reported indices include the CURE index,245 internal stretch fraction (ISF),249 and septal rebound stretch (SRS)55 (Fig. 9-31). All three parameters provided a good prediction of CRT response, and better than using mechanical asynchrony measures in two studies.55,249 THE SITE OF PACING From a theoretical point of view, the “optimal” sites of pacing are those that establish the greatest reduction in total activation time. This requires that the slowly propagating electrical wavefront originating from the RBB is replaced by or fused with one or more activation wavefronts originating elsewhere. In current practice, CRT is most often achieved by pacing the RV and the LV simultaneously or sequentially. The RV pacing site used is usually the RV apex, and recent trials failed to show improved CRT efficacy during RV septal versus RV apical pacing.250,251 The preferred LV pacing site is often the latest-activated region, which in LBBB is usually the basal part of the posterolateral wall.39,40

Figure 9-31  Schematic representation of the determination of mechanical asynchrony (left panel) and discoordination (right panel). Mechanical asynchrony is often calculated from the time between the time to peak shortening in the septum (green) and lateral wall (red). Discoordination can be assessed as internal stretch fraction (ISF) or septal rebound stretch (SRS). AVC, aortic valve closure; AVO, Aortic valve opening.

By using extensive epicardial mapping and MRI strain analysis, Helm et al.252 demonstrated that in dyssynchronous canine hearts, the area of late electrical and mechanical activation indeed closely matched the area that caused maximal LV dP/dtmax increase during pacing. This region spans about 40% of the LV free wall, especially the lateral wall. In similar studies, Rademakers et al.253 showed that in the presence of an infarction, the optimal site of LV pacing depends on the scar location. Gasparini et al.254 showed that placing the LV lead close to the ventricle’s basal portion conferred greater improvements in LV ejection fraction, quality of life, and distance walked in 6 minutes, similar to those reported for lateral lead placement. Their findings confirm previous observations by Butter et al.,218 who showed that both anterior and lateral wall pacing sites improve LV contractility and pump function. However, pacing from the lateral wall resulted in a significantly larger increase in LV contractility; moreover, in about one third of patients, pacing from the anterior wall significantly depressed LV function. It is important to note that Butter paced the LV from a “middle-apical” position. High-resolution mapping data help explain these apparently discrepant findings (see Fig. 9-29). In patients with LBBB, the basal region of the heart is always the last region to be activated, with minor differences between the superior part of the left ventricle (around aortic root) and posterior and posterobasal regions;40 the latter regions were the pacing sites selected by Gasparini.254 In contrast, the anterior middle-apical part of the left ventricle is usually early-activated, representing the LV breakthrough site, and activation of the middle-apical part of the lateral wall, one of the pacing sites tested by Butter et al.,218 is significantly delayed. Accounting for the differences in activation time in individual patients and using TDI to locate the region with the latest mechanical activation, Ansalone et al.255 investigated whether concordance of LV lead position and latest-activated region influences the outcome of CRT. Indeed, the greatest effect on functional parameters was seen when the LV lead was in the latest-activated region. Although optimization of the pacing site is often difficult when the transcoronary venous access is used, positioning the LV lead through limited thoracotomy may be helpful. During this procedure, LV pressure-volume analysis showed the importance of choosing the best position of the LV lead.226 During a coronary sinus approach, the left ventricle is paced at the epicardial surface, whereas under physiologic conditions, excitation of the LV initiates at the endocardium.11 In a canine LBBB model, endocardial LV pacing during CRT consistently reduced electrical asynchrony compared with conventional epicardial CRT (Fig. 9-32). In parallel with these electrophysiologic improvements, endocardial CRT also improved systolic LV pump function and decreased dispersion of repolarization compared with epicardial LV pacing at the same site.33 Additionally, the hemodynamic effects for endocardial sites were less dependent on location and AV delay than epicardial sites. LV endocardial pacing in patients can be established through a transatrial septal approach, shown to be technically feasible and effective during acute and midterm follow-up.206,256-259 However, its long-term effects



9  Basic Physiology and Hemodynamics of Cardiac Pacing LBBB

EPI-BiV pacing

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Figure 9-32  Three-dimensional reconstruction of electrical activation times of the left and right ventricles during baseline LBBB (left) and biventricular (BiV) pacing with conventional epicardial (EPI) and novel endocardial (ENDO) left ventricular electrode position. Data from epicardial contact mapping and endocardial non-contact mapping in a canine heart with LBBB induced by radiofrequency ablation.59

compared to epicardial pacing and the development of systemic thromboembolic complications need to be investigated further. Given individual variations in etiology, severity, patterns of delayed ventricular activation, location of scar regions, and extent of mitral regurgitation in heart failure, it seems unlikely that one pacing site will “fit all.” This is further exemplified by studies in canine hearts with LBBB and with chronic myocardial infavetion, where the optimal site of LV pacing depended on the infarction site.253 Not surprisingly, pacing inside scar tissue, as in patients with a posterolateral scar, provides little benefit.260 Although stimulating the latest-activated regions makes sense, increasing evidence supports that pacing other areas may be at least as effective. In dogs, acute and chronic LV apical and LV septal pacing maintained regional cardiac mechanics, contractility, relaxation, and efficiency near native levels.59 Derval et al.259 investigated 11 LV pacing sites (10 endocardial, 1 CS) in 35 nonischemic, dilated cardiomyopathy patients but failed to reveal a single optimal pacing site. However, selecting the best location individually doubled LV dP/dtmax compared with conventional coronary sinus pacing. Given the electrical similarity of the spontaneously occurring LBBB and the LBBB induced by RV pacing, several centers have explored the feasibility of “upgrading” already-implanted RV pacemakers to BiV units or, in the case of a new implant, to use BiV pacing from the beginning. The results appear to depend on the clinical setting. In patients with already-implanted RV pacing leads, the upgrade leads to improvement in cardiac function and clinical outcome.261,262 However, the situation is less clear when RV and BiV pacing are compared in patients with atrial fibrillation immediately after AV nodal ablation.263,264 In these crossover studies, parameters such as quality of life, New York Heart Association (NYHA) diagnostic class, and 6-minute walking distance showed modest differences between RV pacing and BiV pacing. However, echocardiographic measures demonstrated moderate improvements with BiV over RV pacing.263 These findings are supported by experimental data showing that both RV pacing and BiV DDD pacing are able to completely reverse LV hypertrophy and dilatation induced by chronic AV nodal ablation in dogs.265 Together, these data suggest that the effects of ventricular pacing should be considered in light of the disease being treated. Another important, yet unresolved issue is whether LV pacing alone is sufficient for adequate CRT. In dogs and patients with LBBB, LV pacing improves LV pump function at least as much as BiV pacing.36,161,219,220 LV pacing also leads to clinical improvement at up to 1 year of follow-up.266,267 The promising effect of pacing at this single site can be explained by the fusion between the LV pacing–triggered wavefront and the wavefront created by intrinsic conduction through the RBB.36 The Device Evaluation of CONTAK RENEWAL 2 and

EASYTRAK 2: Assessment of Safety and Effectiveness in Heart Failure (DECREASE HF) study was a large, prospective randomized trial comparing single-site LV pacing to simultaneous and sequential BiV pacing.268 The three groups (each >100 patients) showed similar echocardiographic reverse remodeling after 3 and 6 months. Compared to LV pacing, simultaneous BiV pacing trended toward greater improvement in LV size. Importantly, however, the same AV delays were programmed for all groups, at approximately half the intrinsic PQ time. Therefore, during single-site LV pacing, the heart was predominantly depolarized by the activation wavefront originating from the LV electrode. Without electrical fusion with an activation wavefront from the RBB, the optimal effect of single-site LV pacing may be missed.36 Even if LV pacing were sufficient for proper resynchronization, other factors would continue to warrant RV lead placement in many patients. In particular, current implantable defibrillator systems require an RV lead for tachyarrhythmia sensing and high-voltage therapies. Moreover, if a patient undergoing CRT also requires optimization of the AV interval because of prolonged PQ time or backup ventricular pacing, an RV lead would be important to allow pacing in case of dislodgement of the LV lead, which occurs in 5% to 10% of patients. ATRIAL PACING AND SENSING Given the rationale of CRT, continuous ventricular pacing should be provided to maximize pump function and cardiac efficiency. Many people equivocally equate a pacemaker with a resynchronizer, but even though they are remarkably similar in function and characteristics, the two devices fundamentally differ in one respect. The primary goal of a pacemaker is to restore the rhythm when absent; a secondary, less important, and even questionable goal (because of extreme difficulty in maintaining correct sequence of activation in ventricles during ventricular pacing) is to restore the proper AV timing in patients with first-degree AV block. In contrast, the primary goal of a resynchronizer is to resynchronize all the time, whether intrinsic activation is present or not. This difference creates opposite goals for the two devices in many situations. A minimum heart rate is required to maintain adequate heart function without triggering heart failure. A minimum heart rate is also required to enable the heart to generate enough cardiac output to sustain the body’s metabolic needs. Nevertheless, we now know that cardiac efficiency is higher at lower heart rate (see Fig. 9-17) and that lower heart rates favor angiogenesis. A first attempt to define an adequate heart rate could be taken by stating that given these two opposing driving forces, the optimal heart rate for a given patient would be the lowest rate that creates sufficient cardiac output to satisfy basic metabolic needs in that patient at a given point. For a patient with normal

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LA RV PR150ms

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AV delay for optimum ventricular resynchronization

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Tricuspid and mitral closure Figure 9-33  Left atrial (LA) contraction (contr) after mitral valve closure. Top panel, Schematic illustration of a patient undergoing atrial pacing, in whom sensing in both ventricles has an interatrial conduction delay of 100 msec (ms) and, for the sake of simplicity, an equal mechanical delay. The patient’s intrinsic atrioventricular (AV) interval (marked as PR in the figure) is 150 msec. The figure illustrates how after a small electromechanical delay, the right and left atrial (RA and LA) contractions follow their electrical activations. This patient also has an intraventricular conduction defect, which makes the septum (marked as RV) activate 110 msec earlier than the lateral wall of the left ventricle (marked as LV). Similarly, both activations are followed after a small electromechanical delay by the septal and left lateral wall contractions. The mechanical asynchrony of the LV prevents the LA from contracting (contr) against a closed mitral valve. Bottom panel, Septum and left lateral wall are resynchronized with the optimum AV interval. Note clinically significant increase in filling time created by resynchronization. BV, Biventricular.

50 40

r20.63694

y12.175266.56701x

30 Change in mortality (%)

ventricular function and no heart failure, a more aggressive stance could be taken with respect to programming higher parameters for lower rate limit (LRL) and rate response in the presence of sinus bradycardia or chronotropic incompetence. In the Dual Chamber and VVI Implantable Defibrillator (DAVID) trial, dual-chamber ICD pacing was programmed to an LRL of 70 bpm and short AV intervals; its results strongly suggest that these settings were worse than a simple setting of the ICD to VVI with an LRL of 40 bpm. Many now believe that most of the detrimental effects observed in the DDD arm of this trial resulted from RV pacing creating an asynchronous contraction.175 In the DAVID-II study, VVI backup pacing at 40 bpm was compared with atrial pacing at 70 bpm. The data showed no significant difference between these two settings with respect to clinical outcome and survival as backup pacing.269 Therefore, a heart rate of 70 bpm does not seem to have a detrimental effect in the failing heart, as long as the sequence of conduction is normal. To summarize, the data available strongly suggest caution in programming the pacemaker’s LRL to be higher than the sinus rate in patients with heart failure, whether or not they have sick sinus syndrome. This is especially important because RA pacing may cause interatrial conduction delays, leading to LA contraction against a closed mitral valve and thus lower cardiac output and elevated pulmonary pressure (Fig. 9-33). Moreover, higher-than-necessary heart rates worsen mechanical efficiency, decrease angiogenesis, exacerbate heart failure, and may even increase mortality (Fig. 9-34). In patients who receive an implant (CRT or CRT-D device) primarily for heart failure indications, a general rule is to program the LRL below the sinus rate, unless a clear clinical need to pace the atrium can be documented.

PROFILE

PROMISE FIRST

20 10 0

VHeFT (prazosin)

10 20 30

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40

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50 60

Vesnarinone 70 6 5 4 3 2 1 0 1 2

3

4

5

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9

Change in heart rate (bpm) Figure 9-34  Relationship between the chronic changes in heart rate obtained with several drug therapies and their effect on mortality rates in patients with heart failure. Curiously, there is a strong correlation between mortality benefit and the ability of a drug to achieve a reduction in the heart rate of the patients in the active therapy arm of the trial. HDZ/ISDN, Hydralazine/isosorbate dinitrate; VHeFT, veteran Affairs Vasodilator Heart Failure Trial.



9  Basic Physiology and Hemodynamics of Cardiac Pacing

ATRIOVENTRICULAR OPTIMIZATION As discussed earlier, the AV interval strongly influences ventricular preload and thereby stroke volume in patients with reduced ejection fraction. During CRT, however, the AV interval also affects ventricular pump function by modifying ventricular asynchrony. At long AV intervals, the impulse conducted through the AV node and RBB is allowed to spread into the right ventricle and from there toward the left ventricle, maintaining (part of) the LBBB activation sequence. At shorter AV intervals, however, complete capture occurs from the pacing site(s), allowing resynchronization. The largest improvements in dP/dtmax occur at this point of transition from noncapture to complete capture. The best LV pump function is obtained with a relatively short AV interval, which allows complete activation of the ventricles from the two pacing sites.161 This AV interval may be different from the one resulting in optimized preload164 (Fig. 9-35). In one half of patients, the AV interval for optimal preload is usually longer than that for optimal synchrony; thus it is no surprise that the mean difference between the two AV intervals is close to zero.161 To date, prospective controlled data are still scanty to help determine the relative importance of optimizing synchrony and optimizing preload. Considering that asynchrony is a major cause of depressed pump function in patients with intraventricular conduction abnormalities, and that resynchronization has both short-term and longterm beneficial effects, we postulate that the primary target should be resynchronizing the ventricles, and that optimizing the ventricular inflow is of secondary importance in the majority of patients. The importance of restoration of ventricular asynchrony is supported by studies on the beneficial effect of CRT in patients with permanent atrial fibrillation (AF). In these patients, no preload optimization can be performed, and any possible beneficial effect on preload through CRT is excluded. Nevertheless, CRT has been shown to be effective in patients with AF and wide QRS complexes, especially once the AV node is ablated, abolishing any interference of underlying irregular, higherrate AF spontaneous rhythm. The extent of improvement appears to be comparable to that seen in patients in sinus rhythm, in functional terms as well as long-term prognosis.270-272 To optimize AV delay, invasive and noninvasive determinations of cardiac contractility and stroke volume may be performed. Invasive

225

determination (e.g., using LV dP/dtmax)273 is probably more accurate but is cumbersome and expensive and may carry significant risk for the patient. Noninvasive measurements comprise diastolic inflow measurements (mitral valve E and A wave) and systolic outflow measurements (aortic velocity time integrals, finger arterial pulse pressure). The advantage of outflow measurements is accounting for possible influences of CRT on mitral regurgitation, but mitral inflow is usually easier to measure and more accurate.169 Although validated to optimize AV interval in these patients, these echocardiographic methods remain time-consuming, with unproven clinical effect on large patient cohorts in prospective RCTs. Recent data suggest that the finger plethysmography for assessing arterial pressure and stroke volume is accurate274 and correlates well with echocardiographic methods for AV optimization in these patients.275 Integrated device–based algorithms to optimize AV interval have distinct clinical advantages, and different algorithms have been developed. One algorithm is based on measuring peak endocardial acceleration (PEA), a surrogate of LV dP/dtmax, using a lead with a microaccelerometer at the tip. The tip sensor detects vibration amplitude of the first heart sound, and this value should correlate with LV dP/dt.276 Other AV optimization algorithms are based on a combination of automatic measurements by intracardiac electrograms (EGMs) and surface ECGs268 or through indirect measurement of mechanical timing using surface ECGs or intracardiac signals.277,278 Some studies have found excellent correlation between EGM-based algorithm and invasively derived dP/dtmax268,279 or conventional echocardiography,277 whereas a more recent report demonstrated the reduced accuracy of a device-based interval algorithm optimization compared to conventional echocardiographic methods.278 Particularly relevant in this regard are the results of the large multicenter randomized trial SMARTAV that prospectively compared different modalities of AV interval optimization in CRT recipients, namely empirical AV delay of 120 msec, device electrogram-based algorithm, and echocardiographically guided AV optimization. Neither device electrogram nor echocardiographically based AV optimization conferred greater reduction of left ventricular end-systolic volume during follow-up compared to the empiric approach.279a Other preliminary data were presented about the CLEAR and FREEDOM trials. The CLEAR trial investigated the effect of periodic AV and interventricular (VV) optimization using the PEA

AVL  0 p  0.001 Responders 100 (a) 75 % maximum PP change

(b)

AVL APLS AP RA

10 mm Hg LS

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(c) 100 ms (d)

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AVL (ms) Figure 9-35  Maximum preload is associated with maximum increase in pulse pressure. Left, Upstroke of intraventricular pressure waveform in a patient with NYHA Class III heart failure (PATH-CHF-I trial). (a), During pacing at an intermediate AV interval; (b), during intrinsic propagation through the AV node of the normal atrial depolarization; (c), the ventricular electrogram during intrinsic activation; (d), the atrial electrogram. AVL, Time from peak of intraventricular pressure increase created by booster pump action of left atrium (AP) and the start of the ventricular contraction (LS); RA, right atrial. Right, Relationship between percent increase in aortic pulse pressure (PP) versus baseline (no pacing) obtained during cardiac resynchronization therapy (CRT) at different atrioventricular (AV) intervals (y-axis) versus the time between AP and LS for all the patients in whom hemodynamics improved with CRT in the PATH-CHF-I trial.

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sensor. The clinical composite score had increased in 86% of periodically optimized patients versus 62% in the control group (P < .0007).280 Moreover, the proportion of patients who died from any cause or were hospitalized with heart failure was significantly lower (9%) in the sensor-based optimization group than the control group (25%; P < .001). In contrast, the FREEDOM study, using regular optimization with the Quickopt algorithm, did not improve CRT response.281 Until now the evidence suggests that device-based algorithms for AV optimization may not be equivalent because each may lead to a different outcome. EFFECT OF INTERVENTRICULAR DELAY Most CRT devices now offer independent programming of AV and VV intervals. Studies on the effect of changing VV intervals during BiV pacing have reported conflicting results.208,227,282 Because of variability in the conduction disturbances, as well as ischemic or scarred regions, different timing of stimulation of the two ventricles may be optimal.283,284 In fact, more recent data from larger patient populations indicate the need for optimizing VV interval, with most patients requiring early LV offset. The largest prospective multicenter study evaluating the effectiveness of sequential pacing showed that sequential BiV pacing increased stroke volume and exercise capacity during follow-up to a greater extent than simultaneous BiV pacing for CRT.282 Also, as with Vanderheyden et al.,285 this study confirms that about 60% of patients require early LV offset to optimize CRT. In the Rhythm-II ICD trial, patients with optimal programming of AV and VV interval showed better acute hemodynamic improvements; however, this did not translate into a difference in echocardiographic response after 6 months.286 An optimal setting of the VV interval may have a positive influence. According to the current model of how CRT works, the best resynchronization would be achieved by maximizing the synchrony of the activation through manipulation of the three activation wavefronts originating from the RBB, RV, and LV pacing leads. However, this same reasoning also indicates that an optimal VV interval may depend on the AV interval. The previously cited studies have not tested all the possible variations. Therefore, differences in optimal VV intervals may be caused, for example, by the use of different AV intervals. In clinical practice, performing such elaborate optimization protocol may be impractical and time-consuming, unless a simple technique such as finger blood pressure measurement is used,275,284 with integrated optimization software that simultaneously elaborates and computes the noninvasive hemodynamic data. Optimal use of the extensive possibilities to vary timing of stimulation may be fully applicable only if medical equipment manufacturers provide adequate tools specifically designed and validated for the fine-tuning of AV and VV intervals to provide optimized CRT. It should be emphasized, however, that several prospectively designed studies evaluating the effect of VV timing optimization on symptoms and clinical outcome have failed to show a real

ECG lead II

Pacer rate 67

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Figure 9-36  Radial artery pressure from a patient with complete atrioventricular (AV) block during fixed-rate ventricular pacing. Peak systolic pressure is seen when the PR interval is 270 msec (arrow).

clinical benefit. This might indicate that in the CRT population, VV timing plays a secondary role in modulating response to CRT. It is also possible, however, that a visible individual benefit may be diluted in the general population, or that the biologic variability offsets the minor (yet relevant in the individual patient) effect.

Pacemaker Syndrome When first described, pacemaker syndrome was attributed to ventricular pacing and recognized primarily as a problem of single-chamber ventricular pacemakers. However, with the greater use of dual-chamber pacing modes, it has become apparent that pacemaker syndrome is not unique to single-chamber pacing and may occur in different pacing modalities.287-290 Pacemaker syndrome is an array of cardiovascular and neurologic signs and symptoms resulting from disruption of appropriate AV synchrony (AV dyssynchrony) caused by suboptimal pacing, inappropriate programming of pacing parameters, or upper-limit behavior of AV synchronous pacing systems. The pathogenesis of pacemaker syndrome is complex, involving atrial and vascular reflexes and the neurohumoral system as well as the direct hemodynamic consequences of the loss of atrial systole. Patients most prone to the development of pacemaker syndrome are those with 1 : 1 retrograde ventriculoatrial (VA) conduction (Figs. 9-36 and 9-37) and those with lower stroke volume during ventricular pacing than with sinus rhythm or dualchamber pacing. This syndrome may occur, however, with any mode of pacing that results in permanent or temporary disruption of atrial and ventricular synchronous contraction. Patients treated with CRT devices may also theoretically demonstrate a pacemaker syndrome caused by inappropriate LA and LV timing. Indeed, interatrial conduction time can be significantly prolonged in patients with dilated cardiomyopathy. During CRT, the contraction of the late-activated region (most often located in LV lateral

1.0 m/sec

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Figure 9-37  Doppler aortic flow-velocity integrals during ventricular pacing in a patient with complete heart block. The effect of fortuitously timed P waves (arrows) on aortic flow is demonstrated.



9  Basic Physiology and Hemodynamics of Cardiac Pacing

227

“Shortness of breath”

Figure 9-38  Continuous rhythm strips recorded with a transtelephonic cardiac monitor in a 65-year-old farmer with pacemaker syndrome. This patient’s underlying rhythm disorder was sick sinus syndrome with sinus bradycardia, and he had experienced multiple episodes of atrial fibrillation 1 year earlier. His sinus bradycardia was exacerbated by treatment with quinidine, verapamil, and digoxin. One month after implantation of a VVIR device, the patient returned to his physician complaining of fatigue, weakness, dizziness, malaise, and shortness of breath. After documentation of this rhythm by a cardiac event monitor, he underwent further testing. During VVI pacing, ventriculoatrial (VA) dissociation occurred. During VVI pacing at 90 bpm, the patient’s blood pressure dropped to 90 systolic/68 diastolic mm Hg; during AAI pacing, it was 160/90 mm Hg (at 90 bpm). During sinus rhythm at 58 to 64 bpm, the blood pressure measured 136/70 mm Hg. A DDDR pacemaker was implanted, and within 48 hours, the patient’s symptoms were abolished. He returned to his active, vigorous lifestyle and experienced no subsequent episodes of atrial fibrillation for 1 year. (From Ellenbogen KA, Wood MA, Strambler B: Pacemaker syndrome: Clinical, hemodynamic, and neurohumoral features. In Barold SS, Mujica J [eds]: New perspectives in cardiac pacing. Mt Kisco, NY, Futura, 1992.)

wall) is advanced to coincide with the start of the contraction of the early-activated region (most often in interventricular septum and RV free wall) to maximize mechanical synchrony and thus efficiency.55,69 In a heart where the LA mechanical contraction has been delayed(e.g., because of combined interatrial block and atrial pacing), LA contraction might occur against a closed mitral valve, especially during BiV pacing, in which LV contraction has been advanced with respect to atrial activation (Fig. 9-38). This phenomenon has been described for patients with pacemakers, in whom the issue is less likely, because RV pacing normally delays the mechanical contraction of the left ventricle, increasing the time available for the left atrium to contract before the mitral valve is finally closed by LV contraction and subsequent pressure increase.151 In NYHA functional classes III and IV heart failure, occurrence of the LA contraction after closure of the mitral valve may trigger sudden increases in pulmonary pressures, similar to cannon A waves seen in pacemaker syndrome (usually observable in jugular veins). These increases could lead to acute decompensation and pulmonary edema in a patient with heart failure. It is also important to realize that this situation could be triggered by atrial pacing because of the increased AV delay that this mode of pacing creates. Pacemaker syndrome can be treated with an upgrade of the implanted device or with reprogramming of the parameters to achieve the optimal synchrony of atrial and ventricular contraction. Symptoms associated with pacemaker syndrome are variable in severity and onset (Table 9-4). Although the exact incidence is unknown, moderate to severe symptoms of pacemaker syndrome occur in an estimated 5% to 7% of patients in whom the ventricle is mostly paced (Fig. 9-38). Up to 10% of patients undergoing VVI pacing present with mild symptoms. The pathophysiology of pacemaker syndrome is rather complex, involving hemodynamic, neurohumoral, and baroreceptor changes. In addition to direct hemodynamic effect of loss of AV synchrony or 1 : 1 retrograde VA conduction, atrial and vascular reflexes initiated by atrial distention or elevated atrial pressures may also play a role. Atrial receptors and cardiopulmonary reflexes have been the subject of several in-depth reviews.291,292 Ellenbogen et al.293,294 extensively investigated the role of baroreflex activity during pacemaker syndrome provoked by VA pacing, providing evidence that inadequate systemic sympathetic response plays a key role in the onset of pacemaker syndrome (Fig. 9-39). Indeed, when patients assume an upright position, blood

is pooled in the lower extremities, and the arterial baroreceptors are activated to compensate for the decrease in cardiac output and systolic blood pressure. In some patients, pacemaker syndrome results from the inability to compensate further for the upright posture and augmentation of autonomic tone. In other patients, pacemaker syndrome may result from modification of these vascular responses through the effects of drugs, such as vasodilators and diuretics. In still other patients, pacemaker syndrome may result from activation of inhibitory atrial and cardiopulmonary reflexes that counteract the protective vasoconstrictor reflex (see Fig. 9-39). These responses may be further modified by the production of catecholamines and ANF. Circulating levels of ANF are frequently elevated in patients with complete AV block and bradycardia, being reduced within minutes of programming.295,296 Pacemaker syndrome during AAI or AAIR pacing has been described and emphasizes the role of AV dyssynchrony in the pathogenesis. Patients originally undergoing pacing for sick sinus syndrome may later demonstrate AV conduction abnormalities, or AV nodal conduction may be impaired by drugs with AV node–blocking properties (e.g., digoxin, β-blockers, calcium channel blockers, antiarrhythmics), leading to prolonged and hemodynamically unfavorable atrial R-wave (AR) intervals. AV dyssynchrony may also occur in patients with

TABLE

9-4 

Degree Mild

Moderate

Severe

Symptoms of Pacemaker Syndrome Symptoms Pulsations in neck, abdomen Palpitations Fatigue, malaise, weakness Cough Apprehension Chest fullness or pain Headache, jaw pain Shortness of breath on exertion Dizziness, tiredness, vertigo Orthopnea, paroxysmal nocturnal dyspnea Choking sensation Confusion or alteration of mental state Presyncope Syncope Shortness of breath at rest, pulmonary edema

228

SECTION 2  Clinical Concepts

VVI pacing Sinoaortic baroreceptors

Retrograde VA conduction

↓ C.O. ↓ S.V. ↓ B.P.

↑ L.A.P. ↑ P.A.P.

()

() Vagal cardiopulmonary mechanoreceptors Afferent Input

()

()

()

()

()

()

()

Sympathetic outflow

Effector organs

↑ Epi

Parasympathetic outflow

Vasoconstriction

↑ Tachycardia contractility

Arterial Baroreflex

Bradycardia

Vasodilation

Vagal Cardiopulmonary Reflex

Figure 9-39  Multiple reflex pathways in pacemaker syndrome. Arterial baroreflexes detect a decrease in stroke volume when atrioventricular (AV) asynchrony occurs, leading to sympathetic activation and vasoconstriction. Conversely, AV asynchrony results in higher atrial wall tension and activation of reflex pathways, leading to vagally mediated vasodilatation as well as release of humoral substances (e.g., ANP) that further facilitate these reflexes (e.g., counteracting baroreflex-mediated vasoconstriction). BP, Blood pressure; CO, cardiac output; Epi, epinephrine; LAP, left atrial pressure; PAP, pulmonary artery pressure; SV, stroke volume; VA, ventriculoatrial.

dual–AV nodal pathway physiology during a shift in conduction from the fast pathway to the slow pathway with prolonged AV conduction time. Unfavorable sequencing of atrial and ventricular contractions may occur during AAIR pacing at high sensor-driven rates.288 This situation most often results when the slope determining the sensor-driven rate is programmed too aggressively and out of proportion to the exercise-induced increase in catecholamine levels. As the paced atrial rate progressively increases, the appropriate corresponding decrement in PR interval does not occur, and the interval may even lengthen, Name 25 mm/s 10 mm/mv

eventually resulting in a paced P wave occurring immediately after the preceding R wave (Fig. 9-40). This event most frequently occurs in the earlier stages of exercise and, in some patients, is corrected as the patient continues to exercise. The VDD pacing mode is similar to DDD mode but without the capability of atrial pacing. The VDD mode, previously out of favor, has been revived by the development of CRT.222,297 Atrioventricular dyssynchrony and pacemaker syndrome may also occur during dual-chamber pacing as a result of inappropriate

ID 20-MAY-93 14:38 10-100 Hz Pgm 004B

Name ID: 25 mm/s 10 mm/mv 16-100B

V1 II V5

aVR aVL

Figure 9-40  AAIR pacemaker recording. Patient was receiving β-blockers and experienced pacemaker syndrome while walking. His heart rate increased from 70 to 90 bpm, and his paced atria-sensed QRS (AR) interval increased from 180 msec to more than 500 msec, resulting in shortness of breath. Discontinuing β-blockers and reprogramming the pacemaker to a less “aggressive” slope resulted in a much shorter prolongation of the AR interval with exercise.



9  Basic Physiology and Hemodynamics of Cardiac Pacing

programming, prolonged atrial conduction time, or normal upper rate limit behavior.289,290,298 In AV sequential pacing modes, selection of an inappropriate AV interval may result in an adverse AV sequencing and hemodynamics. Differential AV intervals for paced and sensed atrial events and rate-adaptive AV intervals are useful in some patients for maintaining AV synchrony over a wide range of heart rates at rest and during exercise. The maximum atrial tracking rate is determined by the total atrial refractory period (TARP), which is composed of the AV interval and the postventricular atrial refractory period (PVARP). Long AV intervals or long PVARP intervals are sometimes programmed to avoid pacemaker-mediated tachycardia. When the TARP is too long, the maximal atrial tracking rate is low. Patients may experience symptoms of pacemaker syndrome during exercise when their atrial rate exceeds the maximal atrial tracking rate (MATR), resulting in 2 : 1 AV block and a sudden, marked decrease in cardiac output in patients with complete heart block, or in a lack of ventricular pacing in patients with normal AV conduction (crucial in CRT because therapy would stop). This situation may be improved with reprogramming of shorter AV intervals or PVARP intervals to allow 1 : 1 ventricular tracking at higher atrial rates.

Pacing in Hypertrophic Obstructive Cardiomyopathy In patients with hypertrophic obstructive cardiomyopathy (HOCM), systolic anterior motion of the mitral valve toward the intraventricular septum is associated with LV outflow tract obstruction, often with significant mitral regurgitation. Such patients are extremely sensitive to changes in preload and to conditions that diminish LV filling and result in decreased LV cavity size and increased intraventricular gradient. Thus, these patients may respond quite differently to ventricular pacing than patients with a normal heart. It has long been known that patients with HOCM who require pacing because of bradycardia experience significantly more benefit from pacing modes that preserve AV synchrony than from ventricular pacing.299,300 Moreover, the abnormal sequence of electrical activation following RV pacing may generate a beneficial effect on the septal wall motion in patients with HOCM.160,301,302 RV pacing causes the septum to move away from the free wall during systole, increasing the LV outflow tract diameter and reducing the intraventricular outflow gradient (Fig. 9-41). In addition to widening of the LV outflow gradient, apex pacing may also reduce the Venturi effect and diminish systolic anterior motion of the mitral valve.160 Because of the proposed mechanism of benefit resulting from the altered sequence of a paced ventricular contraction, AV sequential or atrial synchronous pacing must be

229

A

B Figure 9-41  Schematic representation of a monoplane ventriculography. Patient had severe hypertrophic obstructive cardiomyopathy during AAI pacing at a rate of 90 bpm (A) and during DDD pacing at a rate of 90 bpm and an atrioventricular (AV) delay interval of 50 msec  (B). Note that the alteration in ventricular contraction pattern produced by right apical pacing results in less left ventricular outflow tract obstruction during systole. (From Jeanrenaud X, Goy JJ, Kappenberger LK: Effects of dual-chamber pacing in hypertrophic obstructive cardiomyopathy. Lancet 339:1318, 1992.)

performed with a short AV interval to ensure ventricular preexcitation (see Fig. 9-41). Pacing may be especially indicated in patients with HOCM who have significant intraventricular gradients and symptoms despite medical therapy. In these patients, implantation of a sequential AV pacemaker significantly reduces ventricular outflow gradient, improves NYHA functional class and quality of life, and reduces symptoms, including syncope and presyncope. After initial enthusiasm about the role of pacemaker therapy in patients with HOCM, controversy arose later regarding its efficacy and long-term safety.303 Most randomized studies showed flaws in their design.304 Also, at the same time pacemaker therapy was introduced, transcatheter ablation was introduced for reduction of the outflow tract obstruction. The efficacy of the two therapies has not yet been compared. A recent study found promising effects of pacemaker therapy in patients with heart failure and HOCM, including progressive reduction of the outflow tract obstruction and in heart failure class.305 The few existing data suggest that pacing therapy is not beneficial in patients with nonobstructive hypertrophic cardiomyopathy or in those with minimal intraventricular gradients.306 This topic is discussed in more detail in Chapter 15.

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SECTION 2  Clinical Concepts

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Bax JJ, Molhoek SG, van Erven L, et al: Usefulness of myocardial tissue Doppler echocardiography to evaluate left ventricular dyssynchrony before and after biventricular pacing in patients with idiopathic dilated cardiomyopathy. Am J Cardiol 91:94-97, 2003. 243. Garrigue S, Jais P, Espil G, et al: Comparison of chronic biventricular pacing between epicardial and endocardial left ventricular stimulation using Doppler tissue imaging in patients with heart failure. Am J Cardiol 88:858-862, 2001. 244. De Boeck BWL, Meine M, Leenders GE, et al: Practical and conceptual limitations of tissue Doppler imaging to predict reverse remodelling in cardiac resynchronisation therapy. Eur J Heart Failure 10:281-290, 2008. 245. Bilchick KC, Dimaano V, Wu KC, et al: Cardiac magnetic resonance assessment of dyssynchrony and myocardial scar predicts function class improvement following cardiac resynchronization therapy. JACC Cardiovasc Imaging 1:561-568, 2008. 246. Chung ES, Leon AR, Tavazzi L, et al: Results of the predictors of response to CRT (PROSPECT) trial. Circulation 117:2608-2616, 2008. 247. Beshai JF, Grimm RA, Nagueh SF, et al: Cardiac resynchronization therapy in heart failure with narrow QRS complexes. N Engl J Med 357:2461-2471, 2007. 248. Rüssel IK, Zwanenburg JJ, Germans T, et al: Mechanical dyssynchrony or myocardial shortening as MRI predictor of response to biventricular pacing? J Magn Reson Imaging 26:14521460, 2007.

249. Kirn B, Jansen A, Bracke F, et al: Mechanical discoordination rather than dyssynchrony predicts reverse remodeling upon cardiac resynchronization. Am J Physiol 295:H640-H646, 2008. 250. Haghjoo M, Bonakdar HR, Jorat MV, et al: Effect of right ventricular lead location on response to cardiac resynchronization therapy in patients with end-stage heart failure. Europace 11:356-363, 2009. 251. Bulava A, Lukl J: Similar long-term benefits conferred by apical versus mid-septal implantation of the right ventricular lead in recipients of cardiac resynchronization therapy systems. Pacing Clin Electrophysiol 32(Suppl 1):32-37, 2009. 252. Helm RH, Byrne M, Helm PA, et al: Three-dimensional mapping of optimal left ventricular pacing site for cardiac resynchronization. Circulation 115:953-961, 2007. 253. Rademakers L, van Kerckhoven R, van Deursen CJM, et al: Myocardial infarction does not preclude electrical and hemodynamic benefits of CRT in dyssynchronous canine hearts. Circulation Arrhythm Electrophysiol 3:361-368, 2010. 254. Gasparini M, Mantica M, Galimberti P, et al: Is the left ventricular lateral wall the best lead implantation site for cardiac resynchronization therapy? Pacing Clin Electrophysiol 26:162-168, 2003. 255. Ansalone G, Giannantoni P, Ricci R, et al: Doppler myocardial imaging to evaluate the effectiveness of pacing sites in patients receiving biventricular pacing. J Am Coll Cardiol 39:489-499, 2002. 256. Jaïs P, Douard H, Shah DC, et al: Endocardial biventricular pacing. Pacing Clin Electrophysiol 21:2128-2131, 1998. 257. Van Gelder BM, Scheffer MG, Meijer A, Bracke FA: Transseptal endocardial left ventricular pacing: an alternative technique for coronary sinus lead placement in cardiac resynchronization therapy. Heart Rhythm 4:454-460, 2007. 258. Pasquié JL, Massin F, Macia JC, et al: Long-term follow-up of biventricular pacing using a totally endocardial approach in patients with end-stage cardiac failure. Pacing Clin Electrophysiol 30(Suppl 1):31-33, 2007. 259. Derval N, Steendijk P, Gula LJ, et al: Optimizing hemodynamics in heart failure patients by systematic screening of left ventricular pacing sites: the lateral left ventricular wall and the coronary sinus are rarely the best sites. J Am Coll Cardiol 55:566-575, 2010. 260. Bleeker GB, Kaandorp TA, Lamb HJ, et al: Effect of posterolateral scar tissue on clinical and echocardiographic improvement after cardiac resynchronization therapy. Circulation 113:969-976, 2006. 261. Leon AR, Greenberg JM, Kanuru N, et al: Cardiac resynchronization in patients with congestive heart failure and chronic atrial fibrillation: effect of upgrading to biventricular pacing after chronic right ventricular pacing. J Am Coll Cardiol 39:12581263, 2002. 262. Valls-Bertault V, Fatemi M, Gilard M, et al: Assessment of upgrading to biventricular pacing in patients with right ventricular pacing and congestive heart failure after atrioventricular junctional ablation for chronic atrial fibrillation. Europace 6:438-443, 2004. 263. Brignole M, Gammage M, Puggioni E, et al: Comparative assessment of right, left, and biventricular pacing in patients with permanent atrial fibrillation. Eur Heart J 26:712-722, 2005. 264. Leclercq C, Walker S, Linde C, et al: Comparative effects of permanent biventricular and right-univentricular pacing in heart failure patients with chronic atrial fibrillation. Eur Heart J 23:1780-1787, 2002. 265. Peschar M, Vernooy K, Vanagt WYR, et al: Absence of reverse electrical remodeling during regression of volume overload hypertrophy in canine ventricles. Cardiovasc Res 58:510-517, 2003. 266. Touiza A, Etienne Y, Gilard M, et al: Long-term left ventricular pacing: assessment and comparison with biventricular pacing in patients with severe congestive heart failure. J Am Coll Cardiol 38:1966-1970, 2001. 267. Blanc JJ, Bertault-Valls V, Fatemi M, et al: Midterm benefits of left univentricular pacing in patients with congestive heart failure. Circulation 109:1741-1744, 2003. 268. Rao RK, Kumar UN, Schafer J, et al: Reduced ventricular volumes and improved systolic function with cardiac resynchronization therapy: a randomized trial comparing simultaneous biventricular pacing, sequential biventricular pacing, and left ventricular pacing. Circulation 115:2136-2144, 2007. 269. Wilkoff BL, Kudenchuk PJ, Buxton AE, et al: The DAVID (Dual Chamber and VVI Implantable Defibrillator) II trial. J Am Coll Cardiol 53:872-880, 2009. 270. Delnoy PP, Ottervanger JP, Luttikhuis HO, et al: Comparison of usefulness of cardiac resynchronization therapy in patients with atrial fibrillation and heart failure versus patients with sinus rhythm and heart failure. Am J Cardiol 99:1252-1257, 2007. 271. Molhoek SG, Bax JJ, Bleeker GB, et al: Comparison of response to cardiac resynchronization therapy in patients with sinus rhythm versus chronic atrial fibrillation. Am J Cardiol 94:15061509, 2004. 272. Gasparini M, Auricchio A, Metra M, et al: Long-term survival in patients undergoing cardiac resynchronization therapy: the importance of performing atrioventricular junction ablation in patients with permanent atrial fibrillation. Eur Heart J 29:16441652, 2008. 273. Van Gelder BM, Meijer A, Bracke FA: The optimized V-V interval determined by interventricular conduction times versus



invasive measurement by LVdP/dtmax. J Cardiovasc Electrophysiol 19:939-944, 2008. 274. Whinnett ZI, Davies JE, Willson K, et al: Determination of optimal atrioventricular delay for cardiac resynchronization therapy using acute non-invasive blood pressure. Europace 8:358-366, 2006. 275. Van Geldorp I, Delhaas T, Hermans B, et al: Comparison of a non-invasive arterial pulse contour technique and echo Doppler aorta velocity-time integral on stroke volume changes in optimization of cardiac resynchronization therapy. Europace 13:8795, 2011. 276. Tassin A, Kobeissi A, Vitali L, et al: Relationship between amplitude and timing of heart sounds and endocardial acceleration. Pacing Clin Electrophysiol 32(Suppl 1):101-104, 2009. 277. Baker JH, McKenzie Jr, Beau S, et al: Acute evaluation of programmer-guided AV/PV and VV delay optimization comparing an IEGM method and echocardiogram for cardiac resynchronization therapy in heart failure patients and dualchamber ICD implants. J Cardiovasc Electrophysiol 18:185-191, 2007. 278. Kamdar R, Frain E, Warburton F, et al: A prospective comparison of echocardiography and device algorithms for atrioventricular and interventricular interval optimization in cardiac resynchronization therapy. Europace 12:84-91, 2010. 279. Gold MR, Niazi I, Giudici M, et al: A prospective comparison of AV delay programming methods for hemodynamic optimization during cardiac resynchronization therapy. J Cardiovasc Electrophysiol 18:490-496, 2007. 279a.  Ellenbogen KA, Gold MR, Meyer TE, et al: Primary results from the SmartDelay Determined AV Optimization: A Comparison to Other AV Delay Methods Used in Cardiac Resynchronization Therapy (SMART-AV) Trial. Circulation 122:2660-2668, 2010. 280. Delnoy PP, Lunati M, Nagele H, et al: Periodic VV and AV delays optimization in cardiac resynchronization therapy improves patients’ clinical outcome: results from the CLEAR study. Heart Rhythm 5:S55, 2010. 281. Abraham WT, Gras D, Yu CM, et al: Results from the FREDOM trial: assess the safety and efficacy of frequent optimization of cardiac resynchronization therapy. Am Heart J 159:944-948, 2010. 282. León AR, Abraham WT, Brozena S, et al: Cardiac resynchronization with sequential biventricular pacing for the treatment of

9  Basic Physiology and Hemodynamics of Cardiac Pacing moderate-to-severe heart failure. J Am Coll Cardiol 46:22982304, 2005. 283. Sogaard P, Egeblad H, Kim WY, et al: Tissue Doppler imaging predicts improved systolic performance and reversed left ventricular remodeling during long-term cardiac resynchronization therapy. J Am Coll Cardiol 40:723-730, 2002. 284. Whinnett ZI, Davies JE, Willson K, et al: Haemodynamic effects of changes in atrioventricular and interventricular delay in cardiac resynchronisation therapy show a consistent pattern: analysis of shape, magnitude and relative importance of atrioventricular and interventricular delay. Heart 92:1628-1634, 2006. 285. Vanderheyden M, De Backer T, Rivero-Ayerza M, et al: Tailored echocardiographic interventricular delay programming further optimizes left ventricular performance after cardiac resynchronization therapy. Heart Rhythm 2:1066-1072, 2005. 286. Boriani G, Biffi M, Müller CP, et al: A prospective randomized evaluation of VV delay optimization in CRT-D recipients: echocardiographic observations from the RHYTHM II ICD study. Pacing Clin Electrophysiol 32(Suppl 1):120-125, 2009. 287. Levine PA, Seltzer JP, Pirzada FA: The “pacemaker syndrome” in a properly functioning physiologic pacing system. Pacing Clin Electrophysiol 6:279-282, 1983. 288. Den Dulk K, Lindemans FW, Brugada P, et al: Pacemaker syndrome with AAI rate variable pacing: importance of atrioventricular conduction properties, medication and pacemaker programmability. Pacing Clin Electrophysiol 11:1226-1233, 1988. 289. Torresani J, Ebagost A, Alland-Latour G: Pacemaker syndrome with DDD pacing. Pacing Clin Electrophysiol 7:1148-1151, 1984. 290. Pierantozzi A, Bocconcelli P: DDD pacemaker syndrome and atrial conduction time. Pacing Clin Electrophysiol 17:374-376, 1994. 291. Mary DASG: Electrophysiology of atrial receptors. In Hainsworth RMK, Mary DASG, editors: Cardiogenic reflexes, Oxford, 1987, Oxford Scientific, pp 3. 292. Hainsworth R: Reflexes from the heart. Physiol Rev 71:617-658, 1991. 293. Ellenbogen KA, Thames MD, Mohanty PK: Importance of regurgitant fraction for left ventricular end-systolic wall stress to end-systolic volume ratio in patients with chronic mitral regurgitation. Am J Cardiol 65:53-59, 1990.

233

294. Ellenbogen KA, Wood MA, Stambler B: Clinical, hemodynamic and neurohumoral features. In Barold SS, Mujica J, editors: New perspectives in cardiac pacing, NY, Futura, 1992, Mt Kisco. 295. Vardas PE, Travill CM, Williams TDM, et al: Effect of dualchamber pacing on raised plasma atrial natriuretic peptide concentrations in complete atrioventricular block. Br Med J 296:94, 1988. 296. Ellenbogen KA, Kapadia K, Walsh M, et al: Increase in plasma atrial natriuretic factor during ventriculoatrial pacing. Am J Cardiol 64:236-237, 1989. 297. Bristow MR, Saxon LA, Boehmer J, et al: Cardiacresynchronization therapy with or without an implantable defibrillator in advanced chronic heart failure. N Engl J Med 350:2140-2150, 2004. 298. Pieterse MG, den Dulk K, van Gelder BM, et al: Programming a long pace atrioventricular interval may be risky in DDDR pacing. Pacing Clin Electrophysiol 17:252-257, 1994. 299. Shemin RJ, Scott WC, Kastl DG, et al: Hemodynamic effects of various modes of cardiac pacing after operation for idiopathic hypertrophic subaortic stenosis. Ann Thorac Surg 27:137, 1979. 300. Gross JN, Keltz TN, Cooper JA, et al: Profound “pacemaker syndrome” in hypertrophic cardiomyopathy. Am J Cardiol 70:1507-1511, 1992. 301. Jeanrenaud X, Goy JJ, Kappenberger L: Effects of dual-chamber pacing in hypertrophic obstructive cardiomyopathy. Lancet 339:1318-1823, 1992. 302. Nishimura RA, Trusty JM, Hayes DL, et al: Dual-chamber pacing for hypertrophic cardiomyopathy. J Am Coll Cardiol 29:435-441, 1997. 303. Maron BJ: Hypertrophic cardiomyopathy: a systematic review. JAMA 287:1308-1320, 2002. 304. Mohiddin SA, Page SP: Long-term benefits of pacing in obstructive hypertrophic cardiomyopathy. Heart 96:328-330, 2010. 305. Galve E, Sambola A, Saldaña G, et al: Late benefits of dualchamber pacing in obstructive hypertrophic cardiomyopathy: a 10-year follow-up study. Heart 96:352-356, 2010. 306. Cannon RO, Tripodi D, Dilsizian V, et al: Results of permanent dual-chamber pacing in symptomatic nonobstructive hypertrophic cardiomyopathy. Am J Cardiol 73:571-575, 1994.

234

SECTION 2  Clinical Concepts

10  10

Clinical Trials of Atrial and Ventricular Pacing Modes CARSTEN W. ISRAEL

History and Rationale for Different Pacing Modes Cardiac pacing has been developed to treat syncope caused by bradycardia and asystole, initially in patients with permanent third-degree atrioventricular (AV) block. For this purpose, a single lead implanted in the ventricle and pacing at a prespecified rate was sufficient. This simple method of pacing (later coded “VOO 70”) was able to demonstrate a mortality reduction in patients with AV block despite that pacemaker implantation required open-chest surgery at that time.1 Cardiac pacing in permanent third-degree AV block was found not only to prevent syncope, but also to improve symptoms, exercise capacity, and prognosis. It was therefore also used for other patients with bradycardia, such as from intermittent high-degree AV block and sinus node disease (SND). With technical advancement, the limitations of fixed-rate ventricular pacing were overcome by the development of sensors to provide rate-adaptive (R) ventricular pacing. Rate-adaptive pacing was used to improve the hemodynamic benefit of ventricular single-chamber pacing by providing ventricular rates that take into account the metabolic needs of the patient. In patients with SND, however, ventricular single-chamber pacing in VVI and VVI(R) modes did not reduce mortality2 and could even cause symptoms that were later termed pacemaker syndrome.3 To treat SND, atrial single-chamber pacing has been used, particularly in the Scandinavian countries,4 although many physicians remained skeptical about atrial pacing because it could not prevent asystole if AV block occurred.5 However, a number of advantages of atrial versus ventricular single-chamber pacing were observed, including the absence of pacemaker syndrome, better cardiac output and exercise capacity, resolution of mitral regurgitation that could be induced by ventricular pacing, and a lower incidence of atrial fibrillation. Therefore, alternative pacing modes were desired, particularly for patients with SND. The development of dual-chamber pacemakers created hope that DDD pacing would restore and maintain normal AV synchrony and thus provide “physiologic” pacing. It was anticipated that this closer approximation of normal physiology would reduce patient morbidity and mortality and improve quality of life (QOL). After the introduction of dual-chamber pacing, numerous case reports and small studies showed how dual-chamber (physiologic) pacing resolves mitral regurgitation, is associated with lower plasma epinephrine levels, improves exercise tolerance, resolves VVI pacing–associated heart failure, and has other beneficial effects. However, the theoretical advantages of dual-chamber pacing (maintenance of AV synchrony, preservation of sinus node control over heart rate in patients with AV block, maintenance of normal ventricular activation via His-Purkinje system in SND patients) required confirmation in prospective clinical studies given their higher cost, greater complexity, and potential for complications. This was the rationale for trials comparing pacing modes, as outlined in this chapter. Currently, recommendations in international guidelines require confirmation by large, randomized controlled trials (RCTs). However, most RCTs have yielded unexpected results, particularly because the theoretical advantages of dual-chamber pacing have never been clearly demonstrated. All these trials have some results that require further explanation. From these experiences, physicians should consider a

234

number of lessons in the daily practice of implanting and following pacemakers. This chapter presents only a selection of the most important or most representative studies comparing different pacing modes and modalities. COMPARISON STUDIES Ventricular Pacing with Versus Without Rate Response (VVI vs. VVIR) Crossover studies in the 1980s and early 1990s randomized small numbers of patients to fixed-rate ventricular single-chamber pacing (VVI) versus rate-adaptive ventricular pacing (VVIR) with the use of different sensors6-12 (Table 10-1). These studies unequivocally favor rate-adaptive pacing and showed an improvement in submaximal and maximal exercise in terms of objective data (e.g., exercise duration, maximal exercise capacity in watts or duration, VO2 at anaerobic threshold) or subjective data (e.g., pacing mode preference, symptom scores/questionnaires) compared with fixed-rate VVI pacing. This advantage was still present in studies that included patients with SND or retrograde AV conduction. The most striking advantage of VVIR over VVI pacing was observed in patients with AV node ablation for atrial fibrillation. However, these studies were small and not blinded. Therefore, a significant bias in favor of rate-adaptive pacing cannot be excluded. Many patients in these studies were younger than the average age of pacemaker patients. In elderly patients with a sedentary lifestyle and patients with severe heart disease that precludes any exercise above minimal levels (two large groups of pacemaker patients), there is likely no difference between VVI and VVIR pacing. In some patients, VVIR pacing resulted in an increase in angina pectoris. The improvement in exercise capacity in VVIR pacing is based on an increase in heart rate, which increases myocardial oxygen consumption with potentially disadvantageous consequences in ischemic heart disease. In addition, whenever a patient has an intrinsic ventricular rhythm at an adequate rate, this is preferred to avoid deleterious effects of ventricular desynchronization by right ventricular pacing (see later). Lastly, the use of rate-adaptive pacing requires conservative programming of parameters related to the pacing rate, because “overpacing” (i.e., pacing rate faster than needed) is more symptomatic than “underpacing.” Ventricular Pacing with Rate Response Versus Dual-Chamber Pacing (VVIR vs. VDD/DDD/DDDR) With the advent of dual-chamber pacing, a number of trials tested if rate-adaptive ventricular pacing (VVIR) with an artificial sensor driving the heart rate was comparable to dual-chamber pacing, in which the sinus node determines the heart rate. Investigators observed that sensors frequently reacted inappropriately and were therefore inferior to a chronotropically competent sinus node to increase the heart rate linearly according to metabolic needs during exercise.13 Most studies showed a superiority of dual-chamber pacing in terms of exercise capacity, symptoms, and hemodynamic parameters13-20 (Table 10-2). Of note, when patient preference of pacing mode was assessed, there was a clear advantage for AV-synchronous pacing modes; this can be ascertained only in crossover trials. All the randomized large-scale studies comparing single-chamber ventricular and dual-chamber



10  Clinical Trials of Atrial and Ventricular Pacing Modes

TABLE

235

Studies Comparing VVI and VVIR Pacing*

10-1 

Study (year) Lipkin et al.6

Pacing Indication Third-degree AV block (9 of 10 patients with retrograde AV block) Third-degree AV block (6) or SND (5) Third-degree AV block (AV node ablation in 29) AV block (13) or SND (2)

N 10

Sensor Piezocrystal

Control VVI 70 bpm

11

VVI (rate: na)

46

Minute ventilation Different

Result Long-term improvement in breathlessness and increased maximal oxygen consumption with VVIR Similar improvement in oxygen consumption at anaerobic threshold Improved exercise capacity and better quality of life with VVIR

VVI 70 bpm

Improved exercise duration

15

Piezocrystal

VVI 70 bpm

AV block

18

QT interval

VVI 70 bpm

Rossi et al.11

Third-degree AV block

13

VVI 70 bpm

Gammage et al.12

AV block and atrial fibrillation

12

Minute ventilation Piezocrystal

13 of 15 preferred VVIR (less dyspnea, tiredness) Exercise capacity improved significantly on bicycle and on treadmill Exercise tolerance and dyspnea improved with VVIR, patient preference (61%) Angina with VVIR in 2 Improved exercise capacity (maximal exercise time, oxygen uptake at anaerobic threshold) Improved maximum exercise capacity, decreased perceived difficulty of submaximal exercise

Oto et al.7 8

Lau et al.

Smedgard et al.9

VVI 60 bpm

*Small numbers (N) of patients randomized to fixed-rate ventricular single-chamber pacing (VVI) versus rate-adaptive ventricular pacing (VVIR). AV, Atrioventricular; SND, sinus node disease; na, not available.

pacing had a parallel trial design and therefore lack this information (apart from crossover patients, who showed a clear preference for dualchamber pacing mode in PASE and MOST). Ventricular Single-Chamber Pacing Versus Atrial Synchronized Ventricular Pacing (VVI vs. VDD) The question if dual-chamber pacing with ventricular pacing synchronized to atrial activity was superior to single-chamber ventricular pacing was evaluated early. In 16 patients with high-degree AV block, Kruse and Rydén21 found a significant increase in maximal exercise capacity in VDD compared to VVI pacing. Interestingly, the extent of improvement was the same for patients older and younger than 65 years. At comparable exercise levels, the atrial rate was significantly lower in VDD than VVI pacing. In a crossover design, 13 patients underwent invasive hemodynamic measurements after 3 months of VDD and VVI pacing.22 During VDD pacing, cardiac output was significantly higher (>30%), particularly during exercise, because of the heart rate increase in VDD and despite a stroke volume increase in VVI pacing. Arteriovenous oxygen difference was significantly higher during VVI pacing. Arterial blood lactate was significantly higher in VVI than VDD pacing during exercise. Heart size was significantly smaller with VDD pacing. In a questionnaire on symptoms and pacemaker choice, patients preferred the VDD mode. These early observations favoring dual-chamber pacing significantly in all respects over VVI pacing were an important part of initiating the TABLE

10-2 

strong belief in AV synchronous dual-chamber pacing as the “physiologic” and preferred pacing mode. Single-Lead VDD Pacing Versus Dual-Chamber DDD Pacing in Atrioventricular Block The VDD pacing mode offers the advantage of using only one lead, typically tripolar or quadripolar, with two floating rings that sense atrial activity.23 This pacing mode may be similar to DDD pacing with two leads in patients with AV block and normal sinus node function. However, no RCTs have been performed comparing VDD and DDD pacing. A retrospective study analyzed 1214 consecutive patients with pacemaker implantation for AV block (36.5% VVI; 32.9% DDD; 30.6% VDD).24 At implantation, operation and fluoroscopic times were longer in DDD than in VDD and VVI systems. During follow-up of 5 years, complications requiring surgical interventions occurred 2.3 times more frequently for DDD systems than VDD systems, mainly because of a higher incidence of early atrial lead dysfunction. Dual-Chamber Pacing in Chronotropic Incompetence or Paroxysmal Atrial Fibrillation: DDD(R) with Mode Switching There are few studies prospectively comparing different pacing modes within dual-chamber pacing, such as triggered versus inhibited mode. Sulke et al.25 compared DDDR, DDD, DDIR, and VVIR pacing modes in 22 patients with high-degree AV block and chronotropic

Studies Comparing VVIR and DDD(R)/AAI(R)

Study Lau et sl.13 Jutzy et al.14 15

Menozzi et al.

Oldroyd et al.16 Linde-Edelstamet al.17 Lau et al.18 Lau et al.19 Deharo et al.20

Pacing Indication Third-degree AV block

N 10

Study Mode DDD

Control Mode VVIR

SND (12) or third-degree AV block (2) High-degree AV block

14

DDDR

VVIR

14

DDD

VVIR

Third-degree AV block, normal sinus node function High-degree AV block

10

DDD

VVIR

17

DDD

VVIR

SND (18) or third-degree AV block (15) SND

33

DDDR, DDD

VVIR

15

DDDR, AAIR

VVIR

Third-degree AV block, normal sinus function

18

DDD

VVIR

Result Cardiac output (rest, exercise) and systolic blood pressure better with DDD Exercise capacity, cardiac output, and maximal oxygen consumption better with DDDR Double-blind; patients preferred DDD; cardiac output at rest, ANP, and symptom scores better with DDD ANP better with DDD; no difference in symptom scores Less symptoms in DDD; no difference in exercise time, central hemodynamics, norepinephrine, or myocardial oxygen consumption Double blind; quality of life better in DDD and DDDR More arrhythmias (PVCs, AF) and palpitations in VVIR; better blood pressure in AAIR/DDDR; no significant difference in NYHA Trend toward better well-being with DDD; no significant difference in CPx parameters

AF, Atrial fibrillation; ANP, atrial natriuretic peptide; CPx, cardiopulmonary exercise testing; NYHA, New York Heart Association functional class; PVCs, premature ventricular contractions; SND, sinus node disease.

236

SECTION 2  Clinical Concepts

incompetence. In a randomized double-blind crossover design, patients were paced in each of these modes for 4 weeks. Patients were asked their preference, received three questionnaires for subjective assessment, and underwent treadmill testing, mental stress testing, testing of everyday activities (e.g., suitcase lifting, staircase climbing), and echocardiographic evaluation. Scores for subjective grading of well-being were highest for DDDR pacing, similar for DDD and DDIR modes, and lowest for VVIR pacing. DDDR pacing was the preferred mode in 59% of patients, VVIR mode was the least acceptable mode for 73% of patients. Treadmill testing was best in DDDR mode; nontracking modes underresponded to mental stress; and all rate-adaptive modes overreacted to staircase descent. Cardiac output was best in DDDR mode and non-tracking modes showed varying values in patients with AV block. This study suggested that rate-responsive dual-chamber tracking mode is superior to nontracking dual-chamber pacing, dualchamber pacing without rate-response, and single-chamber ventricular pacing in patients with AV block and chronotropic incompetence, based on both objective and subjective measurements. Another study compared different pacing modes in 48 patients with AV block and paroxysmal atrial tachyarrhythmias.26 Patients received dual-chamber systems programmed to DDIR mode for 30 days. In a randomized double-blind crossover manner, devices were then programmed for 4 weeks each to DDDR with mode switching, DDDR with limitation of the upper tracking rate, and VVIR. Patients received transtelephonic self-activated devices to transmit electrocardiograms whenever they perceived symptoms. In addition, at the end of each 4-week study period, devices were interrogated; patients underwent 24-hour Holter recordings, received three symptom questionnaires, and at the end of the study were asked for their preferred pacing period. Patients with intolerable symptoms were allowed to cross over early to the next pacing mode. Different devices with fast- and slowreacting mode-switching algorithms were used. In this study, one third of patients requested early crossover from VVIR to another mode because of intolerable symptoms, most likely pacemaker syndrome. In DDDR pacing without mode switching but limited upper tracking rate, 19% of patients requested early crossover (29% of patients with documented atrial tachyarrhythmias during this period). Only one patient with DDDR and mode switching requested early crossover because of inappropriate tracking of atrial tachyarrhythmia. Patients with fast-reacting mode-switching algorithms in particular preferred DDDR over VVIR mode. Symptom questionnaires favored DDDR with mode switching over DDDR with limited upper racking rates, DDIR, and VVIR pacing. The authors conclude that dual-chamber pacing with tracking and mode switching to a nontracking mode, whenever this automatic function is fast and reliable, is the preferred pacing mode for patients with AV block and paroxysmal atrial tachyarrhythmias (Fig. 10-1). Dual-Chamber Pacing with Versus Without Rate Response (DDD vs. DDDR) The need for a sensor in patients with chronotropic incompetence has received little attention after dual-chamber pacing became available and resolved this problem for patients with AV block and normal sinus function. Most studies subsequently compared different sensors.27,28 Additionally, the definition of “chronotropic incompetence” varied widely. An early study evaluated cardiopulmonary exercise testing with the DDDR versus DDD pacing modes in eight patients.29 Unfortunately, this was a mix of patients, five with AV block and three with sinus node disease, and five with and three without chronotropic incompetence. Chronotropic incompetence was defined as a maximum heart rate at exercise before pacemaker implantation of less than 110 beats per minute (bpm), but not whether it was caused by AV block (resolved with DDD pacing alone) or sinus node dysfunction (requires rateadaptive pacing). In a crossover study, patients were randomized to DDD or DDDR, each for 3 weeks, and performed an exercise test and completed a symptom evaluation form at the end of each study period. Rate-adaptive pacing was associated with a higher oxygen uptake at

19%

14%

67%

A

Dual-chamber mode None

Ventricular single-chamber mode

33%

0%

B

Dual-chamber mode None

67%

Ventricular single-chamber mode

Figure 10-1  Patient preference of pacing mode. A, Overall patient preference. B, Preference in patients without atrial tachyarrhythmias during the study period.

the anaerobic threshold and maximum exercise. However, exercise time and subjective symptoms were not different in the whole patient group, only in patients with chronotropic incompetence, and symptoms of exercise intolerance were better with DDDR pacing. In a study in 63 patients with a DDD pacemaker, a symptom-limited exercise test was performed 1 month after implantation.30 With chronotropic (sinus node) incompetence defined as a maximum heart rate of less than 60% of the age-predicted maximum or less than 100 bpm, 25 patients were then randomized to perform exercise testing in DDD or DDDR mode on two consecutive days. To minimize bias, patients and physicians were blinded with respect to the pacing mode. Paired analysis revealed a statistically significant improvement in the DDDR mode, with higher heart rates at maximum exercise (113 vs. 84 bpm), longer duration of total exercise and time to anaerobic threshold, and higher oxygen uptake at maximum exercise and anaerobic threshold. Only one large RCT performed according to the principles of evidence-based medicine (EBM) specifically assessed the benefit of rate-adaptive pacing in the dual-chamber mode in patients with sinus rhythm.31 Unfortunately, the single-blind Advanced Elements of Pacing Randomized Controlled Trial (ADEPT) included a mix of patients with SND (~66%) and AV block. After an exercise test 1 month after implantation, 872 patients were enrolled. To be eligible, patients needed to complete at least stage 2 on the Chronotropic Assessment Exercise Protocol and to have a heart rate at peak exercise of 80% or greater of the age-predicted maximum. The mean age was 71 years,



however, so patients with a maximum heart rate not exceeding 120 bpm at maximum exercise could be included. Patients were randomized to DDD versus DDDR pacing (dual-sensor system; maximum adaptive rate, 220 bpm minus age) with rate-adaptive AV delay shortening in a parallel study design. After 12 months, general QOL (measured by Short-Form 36 General Health Survey [SF-36], Ferrans and Powers Quality of Life Index) and cardiovascular functional status (Specific Activity Scale and Duke Activity Status Index) were assessed by questionnaires, together with the maximum exercise time at treadmill testing 6 months after randomization. ADEPT found no specific or general differences in scores of the activity and QOL questionnaires. The total exercise time was not different between DDD and DDDR pacing. Independent of discussions about the ideal sensor(s), severe limitations included the following: • The definition of “chronotropic incompetence” was too broad to demonstrate an improvement by rate-adaptive pacing. Previous studies suggest that rate-adaptive pacing can improve exercise capacity only in patients with a maximum sinus rate less than 90 to 100 bpm.32,33 Although patients with a maximum heart rate of 80% or less of the age-predicted maximum have by definition chronotropic incompetence, this is too mild to be improved by rate-adaptive pacing. • Patients with AV block causing limitation of heart rate increase at maximum exercise (one third of patients enrolled in ADEPT) will substantially benefit from DDD pacing alone. If sinus node function is normal, no further improvement can reasonably be expected from the addition of rate-adaptive pacing. • Because of the negative consequences of right ventricular pacing, programming the AV delay to 150 to 200 msec in patients with SND, particularly activating a rate-adaptive AV delay–shortening algorithm, may have severely biased the results and limited exercise capacity, particularly during DDDR pacing. CONCLUSION Insufficient data from large RCTs exist to clarify if and to what extent rate-adaptive pacing in addition to the DDD mode improves exercise capacity in patients with chronotropic sinus node incompetence. For clinical reasons, only advanced chronotropic incompetence (i.e., sinus node unable to achieve rates >90-100 bpm during exercise) should be taken into account; only in this situation will pacing have a significant potential to improve exercise capacity and relieve symptoms. Caution must be used in patients with coronary artery disease in whom (paced) heart rates greater than 120 bpm may induce ischemia. Sensors should be programmed conservatively (i.e., with limited increase in heart rate during exercise) because “overpacing” is more symptomatic than pacing with rates a little too slow for metabolic needs.

Atrial Versus Ventricular Single-Chamber Pacing in Sinus Node Disease In 1994, Andersen et al.34 published the first RCT on pacing mode. A total of 225 patients with SND were randomized to receive a singlechamber pacemaker in the right atrium or ventricle. Of note, the study excluded patients with AV block of any degree (including first-degree AV block with PQ interval >220/260 msec), Wenckebach block at less than 100 bpm, atrial fibrillation (AF) more than 50% of the time (or with pauses >3 seconds or ventricular rates <40 bpm), or bifascicular bundle branch block. In total, 827 of 1052 screened patients were excluded from study participation, predominantly because of AV block (n = 552). Study outcome parameters were mortality, AF, thromboembolism (TE), and heart failure (HF). From retrospective studies, the authors expected AF in 30%, TE in 15%, and mortality in 15% of patients in the ventricular pacing group. The study intent was to demonstrate a 50% reduction in AF and HF and a 66% reduction in TE and mortality with atrial single-chamber pacing, which would have required 100 patients in each group. This rather optimistic prediction

10  Clinical Trials of Atrial and Ventricular Pacing Modes

TABLE

10-3 

237

Clinical Events in Atrial versus Ventricular Pacing for Sick Sinus Syndrome (Danish Study)*

Clinical Event All-cause mortality Cardiovascular mortality Stroke (cerebrovascular accident) NYHA functional class Permanent atrial fibrillation

Ventricular (n = 115) 50% 34% 23% Mean 1.3 8%

Atrial (n = 110) 34%† 17%† 12%† Mean 1.5† 19%†

From Andersen HR, Nielsen JC, Thomsen PE, et al: Long-term follow-up of patients from a randomised trial of atrial versus ventricular pacing for sick-sinus syndrome. Lancet 350:1210-1216, 1997. *Results from long-term follow-up. †P < .05 for lower incidence respectively; better value with AAIR.

of the capabilities of physiologic pacing reflects the widespread belief in its superiority over ventricular pacing at the time of conception of this study. During a follow-up of 3.3 years, all-cause mortality was similar (22% with ventricular vs. 19% with atrial pacing). There was no significant difference in the risk of AF (in the ECG of at least one follow-up visit in 23% of patients with ventricular vs. 14% with atrial pacing). However, TE was significantly more frequent with ventricular pacing (17% vs. 5%; P = .008). At the same time, ventricular pacing was associated with a significant increase in left atrial diameter compared to baseline and to patients randomized to atrial pacing. Andersen et al.35 extended the follow-up until the last patient completed 5 years in the assigned pacing mode. By that time, the first patients enrolled into the trial had completed an 8-year follow-up (mean, 5.5 years). This prolonged follow-up demonstrated a significantly lower mortality in atrial versus ventricular pacing (relative risk reduction [RRR], 34%), mainly from a reduction in cardiovascular mortality (Table 10-3). Similarly, atrial fibrillation was less likely in the atrial pacing group (RRR, 46% vs. ventricular pacing) as were TE events (RRR, 53%). Only 4 of 110 patients developed AV block during the study period (annual rate, 0.6%), two of whom had a right bundle branch block at baseline. The authors detected retrograde conduction in 63% of patients randomized to VVI pacing. CRITICAL APPRAISAL OF DANISH STUDY The Danish study34,35 was the first large-scale RCT on pacing mode selection, at a time when the discussion was whether it was ethically possible to implant VVI(R) pacemaker systems in SND patients. Also, it is as yet the most positive RCT in favor of atrial-based pacing. Compared with RCTs on dual- versus single-chamber ventricular pacing, however, the number of patients in the Danish trial was rather small; the expectations of a 50% AF and HF reduction and even 66% TE and mortality reduction by atrial versus ventricular single-chamber pacing were too optimistic to show a significant difference in such a small patient group. Prolongation of follow-up to cope with this error in sample size calculation was not prespecified, thus arguably not allowed in a prospective trial. Therefore, results resemble a “post hoc” analysis and are only hypothesis generating. However, all results at 3.3 and 5.5 years point to atrial pacing being associated with more favorable clinical results than ventricular pacing. In no study outcome parameter was ventricular better (significantly or nonsignificantly) than atrial pacing. The Danish trial proposed effective criteria to select patients for AAIR pacing with low risk of AV block. The need to measure the Wenckebach block point at study entry in addition to the PQ interval, however, is not clear. During the study, no development of AV block was predictable by a progressive reduction in the heart rate at the Wenckebach block point.36 However, only a few patients who had documented second-degree or third-degree AV block received an upgrade from AAI to DDD. An observational study found an annual incidence of upgrade from AAI to DDD of 1.7% as a result of documented AV block or AF with slow ventricular rate in 399 patients followed for 4.6 years.37 Also of note was a 0.8% additional annual reoperation rate for the addition of a ventricular lead, mainly because

238

TABLE

10-4 

SECTION 2  Clinical Concepts

Evaluation Parameters for Impact of Pacing Modes on Cardiac Performance

Area Clinical

Exercise capacity Echocardiographic

Clinical events Invasive hemodynamics Blood serum Coronary artery perfusion Arrhythmias

Parameters Dyspnea (NYHA class) Symptom scores Edema, dose of diuretics Systolic blood pressure Pacemaker syndrome Quality of life (questionnaires, scores) Patient preference Exercise duration, maximum workload Anaerobic threshold, maximum oxygen uptake Mitral regurgitation Cardiac output LV end-diastolic diameter, LV ejection fraction/fraction shortening Left atrial diameter Mortality (cardiovascular, all-cause) Thromboembolism/stroke Heart failure hospitalization Cardiac output Pulmonary capillary wedge pressure, V wave Myocardial oxygen consumption Atrial natriuretic peptide, brain natriuretic peptide (Nor-)Epinephrine levels Positron emission tomography (PET) Scintigraphy Atrial fibrillation Ventricular arrhythmias

LV, Left ventricular; NYHA, New York Heart Association.

of atrial lead problems. Therefore, 2.5% of AAI systems were surgically changed to VVIR or DDDR per year. There is renewed interest in the Andersen et al.34,35 data because of the deleterious effects of unnecessary right ventricular pacing observed in the DDDR mode38 (see later). In this context, observations from the Danish trial on development of heart failure with VVI pacing in sinus node disease are important.39 The Danish trial observed an increase in New York Heart Association (NYHA) functional class during follow-up in 31% of patients in the ventricular versus 9% in the atrial group (P < .0005). The incidence of crural edema and dyspnea, together with the need for diuretics during the 5.5-year study period, was significantly higher in patients with ventricular pacing, and the left ventricular fractional shortening decreased significantly only in the ventricular group, not in the atrial pacing group. At the same time, the left atrial diameter increased significantly more during ventricular pacing, from 34 to 41 mm. Interestingly, despite the left atrial dilatation in ventricular pacing in SND, a higher incidence of AF was evident only after 5.5 years, not at 3.3 years of follow-up, despite that TE events were observed more frequently in ventricular pacing already at 3.3 years. Of note, 12 TABLE

10-5 

of the 39 patients who developed TE during follow-up never had documentation of AF, neither before randomization nor during follow-up.40 Documentation may have missed AF in a significant number of patients because only 12-lead ECG was used at follow-up. Pacemaker memory functions likely would have documented a much higher incidence of AF.41

Ventricular Versus Dual-Chamber Pacing NONRANDOMIZED EARLY STUDIES Starting from impressive observations of improved hemodynamics with VAT and VDD compared to VVI pacing in patients with AV block,22 many publications documented the benefits of “physiologic” versus ventricular pacing from the earlier 1980s to the late 1990s. Various parameters were used to define the presence or absence of this benefit (Table 10-4). Outcomes of dual-chamber or atrial versus ventricular single-chamber pacing particularly note hemodynamic improvement with dual-chamber pacing and range from case reports to patient series.42 Of these multiple outcome parameters, four were frequently chosen as endpoints in prospective, large-scale RCTs: mortality, stroke, atrial fibrillation, and heart failure. Whereas studies of the early 1980s usually focused on hemodynamic parameters, the concept of clinical studies on pacing shifted toward investigating the impact of different pacing modes on clinical events such as mortality, particularly after the CAST study, in which antiarrhythmic drugs successfully suppressed ventricular ectopy but were associated with an increased mortality. Mortality Retrospective early trials showed a significant reduction of mortality for atrial and dual-chamber pacing compared to ventricular singlechamber pacing, predominantly in patients with sinus node disease43-52 (Tables 10-5 and 10-6). However, mortality rates varied widely for VVI (15%-48%) compared to atrial or dual-chamber pacing (7%-27%).50 In high-degree AV block, Alpert et al.53 found no difference in allcause mortality between 132 patients with ventricular compared to 48 patients with dual-chamber systems. Similarly, in 148 patients with AV block (74 with VVI and 74 with VDD pacemaker systems), LindeEdelstam et al.54 found no difference in mortality in the entire patient group. However, both studies found that in patients with preexisting heart failure, VVI pacing was associated with a significantly higher mortality than VDD pacing. A study of 1245 patients (77% with AV block; ventricular pacing in 97%, dual-chamber pacing in only 3%), suggested that survival was related to the mode of pacing.48 Analyzing 391 patients with AV block or carotid sinus hypersensitivity, Hesselson

Survival in Patients with Sinus Node Disease in Relation to Pacing Mode* Physiologic

Study/Year Rosenqvist43/1988 Sasaki44/1988 Bianconi45/1989 Santini46/1990 Stangl47/1990 Zanini48/1990 Nürnberg49/1991 Hesselson50 1992 Sgarbossa51/1993 Tung et al.52/1994 Weighted means and totals

Patients (N) 89 19 153 214 110 53 37 308 395 36 1414

Follow-up (mo) 44 20 44 61 52 45 41 30 66 44 50

Ventricular Patients (N) 79 25 150 125 112 57 93 193 112 112 1058

*Results from retrospective observational studies. †P ≤ .05. PHYS, Physiologic pacing mode; RRR, relative risk reduction; VENT, ventricular pacing mode.

Follow-up (mo) 47 35 59 47 54 40 41 44 66 44 49

Death (%/yr) PHYS 2.1 7.6 3.8 2.8 4.0 2.5 7.1 8.3 3.9 8.3 4.8

VENT 5.9 8.2 6.3 7.7 6.2 5.3 13.8 12.8 7.5 16.5 9.4

RRR (%) −64† −7 −40 −64† −35 −53 −49† −35† −48† −50† −49



239

10  Clinical Trials of Atrial and Ventricular Pacing Modes

TABLE

10-6 

Survival in Patients with Atrioventricular Block and Other Indications in Relation to Pacing Mode* Physiologic

Study/Year Alpert53/1986 Linde-Edelstam54/1992 Zanini48/1990 Hesselson50/1992

Patients (N) 48 74 53 299

Ventricular

Follow-up (mo) 60 65 45 30

Patients (N) 132 74 57 92

Death (%/yr)

Follow-up (mo) 60 65 40 44

PHYS 6.0 6.6 2.5 10.4

VENT 7.4 5.9 5.3 14.9

RRR (%) −19 +11 −53 −30†

*Results from retrospective observational studies. †P ≤ .05. PHYS, Physiologic mode; RRR, relative risk reduction; VENT, ventricular mode.

et al.50 demonstrated a significant mortality benefit during a 7 year follow-up with DDD (54%) versus VVI pacing (33%). Atrial Fibrillation and Thromboembolism Starting in the 1980s, almost all retrospective studies showed a lower incidence of AF in patients with atrial/dual-chamber pacing compared to ventricular single-chamber pacing43-52,54-59 (Table 10-7). This was particularly evident for SND. Heart Failure Several studies identified ventricular versus “physiologic” pacing mode as an independent predictor of mortality in patients with congestive heart failure (CHF),54 whereas others did not confirm this observation.51 Overall, retrospective studies suggest a large improvement in clinical outcomes with the use of atrial and dual-chamber pacing versus ventricular pacing systems, with a 45% RRR in mortality, 62% reduction in HF, 66% reduction in stroke, and a 72% reduction in AF (see Table 10-7). These data were used to calculate sample size in the Danish trial of AAI versus VVI pacing.34 Medicare Trial As with other investigators in the late 1980s, Lamas60 noted the problem of “retrograde,” or ventriculoatrial (VA) conduction during VVI pacing at Brigham and Women’s Hospital (Boston). This was a common phenomenon in pacemaker patients and a significant cause of adverse symptoms after VVI pacemaker implantation. Also, only a properly timed atrial systole contributed positively to systolic ventricular performance by the Frank-Starling mechanism which states that the extent of systolic myocardial fiber shortening depends on the degree TABLE

10-7 

of diastolic fiber stretch, or preload. Therefore, with the important atrial contribution to cardiac performance, dual-chamber pacing with restoration of normal cardiac physiology (AV synchrony and appropriate rate response) may be advantageous in patients with AV block or SND. Lamas et al.61 sought to confirm the theoretical advantages of dualchamber pacing in a large, retrospective analysis of a 20% random sample of all Medicare beneficiaries 65 or older who received a first pacemaker implantation (1988-1990) and whose vital status information was available for at least 1 year (n = 36,312). Dual-chamber pacemaker selection was an independent predictor of survival at 1 and 2 years, with a relative mortality risk reduction of 18%, even after controlling for potentially confounding factors. However, choice of dualversus single-chamber pacing mode was biased by patient age, gender, indication for pacing, presence of CHF, and AF. Additionally, implantation in the western United States, rural provider, and hospitalization in a facility with 500 beds or more, a private hospital, or a hospital with a catheterization laboratory were also associated with dual- instead of single-chamber pacemaker selection. For this potential bias, the authors performed a prospective trial. RANDOMIZED STUDIES In the early 1980s, double-blind crossover studies prospectively compared VVI to dual-chamber pacing. Perrins et al.62 randomized 13 patients with AV block and a VDD pacemaker to dual-chamber or VVI mode. They found a better exercise capacity and reduction of symptoms during the period with dual-chamber pacing, as documented by patients’ diaries and monthly symptom scores. Similarly, Kristensson

Pacing Mode, Atrial Fibrillation (AF), Stroke/Thromboembolism (CVA/TE), Heart Failure (HF), and Mortality* AF (%/yr)

Study Sutton55 Markewitz56 Ebagosti57 Langerfeld58 Rosenqvist43 Sasaki44 Bianconi45 Feuer59 Santini46 Stangl47 Zanini48 Nürnberg49 Hesselsson50 Linde-Edelstam54 Sgarbossa51 Tung52 Total/mean

P-yr 3239 592 408 1265 636 112 1301 807 1582 981 389 444 2498 799 2789 543 18385

P 1.4 3.1 2.0 0.9 1.8 0 4.8 2.5 1.4 1.5 1.0 4.8 1.4 – – – 1.9

V 6.8 10.0 5.2 5.4 12.1 12.3 7.9 4.5 11.8 4.2 5.3 11.0 5.3 – – – 7.0

CVA/TE (%/yr) R −85† −69 −62† −83 −85† −100† −39† −44† −88† −64 −81† −56 −74 – – – −72

P – – 0 – 3.3 0 1.2 – 0.5 – 0 0 – – – – 0.9

V – – 1.4 – 3.9 6.8 2.1 – 2.6 – 2.1 3.1 – – – – 2.7

HF (%/yr) R – – −100 – −15 −100 −43 – −81 – −100 −100 – – – – −66

P – – – – 3.9 2.6 – – – – 0.5 – – – – – 2.6

*Results from nonrandomized observational studies. †Multivariate analysis identified ventricular pacing as an independent predictor of chronic atrial fibrillation and stroke. P [PHYS], physiologic pacing mode; P-yr, Patient-years; R [RRR], relative risk reduction; V [VENT], ventricular pacing mode.

V – – – – 9.4 9.6 – – – – 1.6 – – – – – 6.8

Death (%/yr) R – – – – −59† −73 – – – – −69 – – – – – −62

P – – 3.5 – 2.1 7.6 3.8 2.2 2.8 4.0 2.5 7.1 9.0 5.8 3.9 8.3 4.8

V – – 4.8 – 5.9 8.2 6.3 3.9 7.7 6.2 5.3 13.8 13.6 7.0 7.5 16.5 8.9

R – – −27 – −64† −7 −40 −44 −64† −35 −53 −49† −34 −17 −48 −50 −45

240

SECTION 2  Clinical Concepts

et al.63 randomized 44 patients to VVI or VDD for 3 weeks for each mode. They observed a statistically significant improvement of maximal exercise tolerance of 14% with AV synchronous pacing. Arterial lactate, respiratory rates, and perceived exertion ratings during submaximal exercise were higher on ventricular inhibited pacing. Symptom scores during the 3-week periods with VVI pacing were worse than with dual-chamber pacing. Most patients had a better functional class during AV synchronous pacing and preferred this pacing mode. Importantly, these two studies attempted to reduce any bias by double-blinding patients and the physicians who performed exercise testing. Also, in-patient comparison was done between ventricular and dual-chamber pacing modes, not available in later trials, which used a parallel trial design to avoid carryover effects. In one of the first prospective randomized, intention-to-treat trials on pacing mode selection, Mattioli et al.64 compared ventricular (VVI or VVIR) with physiologic pacing (AAI, DDD, DDDR, VDD) in 210 patients (SND in 110, AV block in 100) with a median age of 77 years. Only six patients received an AAI pacemaker. Patients with previously known AF were excluded from this study. A total of 56 patients (27%) developed AF during follow-up. The incidence of AF was 10% at 1 year, 23% at 3 years, and 31% at 5 years. The investigators observed an increased incidence of chronic AF in patients receiving ventricular pacing (P < .05). No difference was observed in the incidence of chronic AF between patients paced for AV block and patients paced for SND. In the subgroup of SND patients, the incidence of chronic AF was significantly less in patients receiving a physiologic than in those receiving a ventricular pacemaker (see Table 10-7). Pacemaker Selection in the Elderly Study (PASE) After their observations in the Medicare trial, Lamas et al.65 randomized 407 patients 65 or older with sinus rhythm and an indication for permanent pacing (SND or AV block) to single- or dual-chamber pacemaker therapy in the Pacemaker Selection in the Elderly Study. All patients received a dual-chamber device programmed to VVIR or DDDR according to the randomization result. The primary endpoint was health-related QOL (measured by SF-36); average age was 76. During the 30-month follow-up, QOL improved significantly after pacemaker implantation, but without a significant difference between VVIR and DDDR pacing modes, neither in QOL nor in clinical outcomes (cardiovascular events or death; Table 10-8). However, this refers to the intention-to-treat analysis while “on-treatment”, 26% percent of patients assigned to ventricular pacing were crossed over and paced in the dual-chamber mode because of symptoms of pacemaker syndrome. The PASE subgroup of patients with SND had a significantly better QOL in some subscales of the SF-36 and a better cardiovascular functional status with dual-chamber pacing than with ventricular pacing. A trend toward lower all-cause mortality, lower incidence of AF, and reduced incidence of the combined endpoint of stroke, hospitalization for heart failure, or all-cause mortality favoring dual-chamber pacing was seen in patients with SND, but not in those with AV block (see Table 10-8). The authors concluded that pacemaker therapy improves QOL, but that the benefits associated with dual-chamber pacing versus

TABLE

10-8 

Clinical Events in Pacemaker Selection in the Elderly Study (PASE) All Patients (n = 407)

Event All-cause mortality Atrial fibrillation

VVIR 17% 19%

DDDR 16% 17%

SND (n = 175) VVIR 20% 28%

DDDR 12%* 19%*

AV Block (n = 201) VVIR 15% 11%

DDDR 17% 16%

Data from Lamas GA, Orav J, Stambler BS, et al: Quality of life and clinical outcomes in elderly patients treated with ventricular pacing as compared with dual-chamber pacing. N Engl J Med 338:1097-1104, 1998. *Trend (P < 0.10) toward lower incidence with DDDR. SND, Sinus node disease.

ventricular pacing were moderate and observed only in the subgroup of SND patients. Critical Appraisal.  PASE was the first large multicenter RCT on dualchamber versus ventricular single-chamber pacing performed according to EBM guidelines. The results were primarily interpreted as disappointing, showing only a marginal or even no advantage of dualchamber pacing. However, the most relevant clinical finding may be the crossover rate of 26% from VVIR to DDDR pacing because of symptoms potentially caused or worsened by single-chamber pacing. In PASE, all patients had received dual-chamber devices, and randomization to single- or dual-chamber pacing was performed by device programming (“software randomization”). Patients randomized to VVIR pacing thus were fortunate because crossover did not require an additional surgical intervention, only reprogramming. Therefore, it was argued that crossover may have been performed because of bias on the side of physicians who attributed symptoms during VVIR pacing inappropriately to “pacemaker syndrome,” despite that per protocol, investigators were required to discuss each case individually with the principal investigator. An analysis of causes and results of crossover from VVIR to DDDR pacing, however, showed that the most common reason for crossover was “severe pacemaker syndrome” (77%), defined as a constellation of fatigue, dyspnea, effort intolerance, paroxysmal nocturnal dyspnea, neck fullness, presyncope, cannon A waves, blood pressure (BP) reduction with ventricular pacing, and CHF initiation or exacerbation.66 The most common individual symptoms were fatigue (62%), effort intolerance (64%), dyspnea (55%), and presyncope (34%). The diagnosis “pacemaker syndrome” or “intolerance to VVIR” in these patients with crossover is supported by the observation that, in contrast to other patients, their QOL deteriorated significantly after pacemaker implantation before crossover and improved significantly after reprogramming to dual-chamber pacing. The use of β-blockers, the presence of nonischemic cardiomyopathy, and systolic BP less than 110 mm Hg during ventricular pacing (relative risk of crossing over = 2.6) predicted intolerance to VVIR. However, the sensitivity of a decrease in paced systolic BP to less than 110 mm Hg at implantation for predicting intolerance to VVIR pacing was only 36%, the positive predictive power only 48%. The presence of VA conduction at implant did not predict the development of intolerance to VVIR pacing, although VA conduction or cannon A waves were present in almost 50% of crossover patients. Similar to results of later trials, the incidence of AF was significantly higher in patients with sinus node disease randomized to VVIR compared to DDDR (28% versus 16% at 18 months). There was no difference between VVIR and DDDR with regard to the development of AF in patients with AV block. In conclusion, the PASE trial suggests that symptoms of a pacemaker syndrome during VVIR pacing are frequent in an elderly population with sinus rhythm and sinus node disease or atrioventricular block. If the crossover rate of 26% translates into 26% of patients requiring surgical reintervention if they are implanted with a ventricular singlechamber device, this complication rate is clinically unacceptable and not cost-effective. Mode Selection Trial in Sinus Node Dysfunction (MOST) The group of investigators of PASE followed their observation that differences in outcome parameters between VVIR and DDDR were (if present at all) more pronounced in sinus node disease, performing a pacing mode comparison in SND patients.67 A total of 2010 patients were randomized to dual-chamber (1014) or ventricular pacing (996) and followed for 33 months. The incidence of the primary endpoint (all-cause mortality or nonfatal stroke) did not differ significantly (21.5% in DDDR, 23.0% in VVIR). Adjusted analyses revealed that in patients assigned to DDDR pacing, the risk of AF was 23% lower (21% vs. 27%) and HF scores were better. Heart failure hospitalization was reduced by 27%, the composite endpoint death, stroke, or HF hospitalization was reduced by 15% (Table 10-9). Dual-chamber pacing resulted in significantly better QOL for six of the eight SF-36 subscales



10  Clinical Trials of Atrial and Ventricular Pacing Modes

TABLE

10-9 

Clinical Events in Mode Selection Trial in Sinus Node Dysfunction (MOST)

Clinical Event Death Stroke Heart failure hospitalization Atrial fibrillation

VVIR (n = 996) 23.0% 4.9% 12.3% 27.1%

DDDR (n = 1014) 21.5% 4.0% 10.2%* 21.4%*

Data from Lamas GA, Lee KL, Sweeney MO, et al: Ventricular pacing or dual-chamber pacing for sinus-node dysfunction. N Engl J Med 346:1854-1862, 2002. *P < .05 for lower incidence with DDDR.

during the 4 years of follow-up. The authors concluded that in SND patients, dual-chamber pacing does not increase stroke-free survival compared to ventricular single-chamber pacing, but DDDR significantly reduces AF risk and HF symptoms and improves QOL. Critical Appraisal.  The MOST trial had the ambitious aim to assess clinical events (mortality, stroke), AF incidence, and symptoms (HF, pacemaker syndrome, QOL) in VVIR versus DDDR pacing for SND. While the latter two endpoints favored dual-chamber pacing, the first and most important endpoint (mortality) did not demonstrate a significant difference. Therefore, the MOST trial was frequently regarded as a negative trial, showing no significant benefit of dual-chamber compared to ventricular single-chamber pacing. In-depth analyses of mortality68 and stroke risk69 showed no advantage of dual-chamber pacing, and QOL was better in dual-chamber mode in several items, particularly related to physical function.70 However, there were two important lessons from MOST: (1) pacemaker syndrome is frequent in SND and VVIR pacing,71 and (2) unnecessary right ventricular pacing in SND is detrimental.38 A strong argument in favor of dual-chamber pacing is the observation of severe pacemaker syndrome in approximately 20% of SND patients treated by VVIR pacing.71 The investigators in MOST attempted to prevent the major weakness of PASE, where the crossover rate from single- to dual-chamber pacing of 26% was so high that it may have prevented any meaningful intention-to-treat analysis of the difference between the two pacing modes. In MOST, therefore, the diagnosis of pacemaker syndrome strictly required documentation of new symptoms (e.g., dyspnea, syncope) together with VA conduction

or systolic BP reduction of 20 mm Hg or more during VVI pacing compared to intrinsic rhythm. Even in the presence of pacemaker syndrome, crossover of pacing mode had to be confirmed by the clinical coordinating center; other attempts to resolve pacemaker syndrome (e.g., reprogram pacing rate, deactivate sensor) were required; and only if these failed, the centers could ask the principal investigator to review the indication for crossover to dual-chamber mode. Despite these rigid requirements, 20% of patients programmed to VVIR met criteria for pacemaker syndrome and required crossover from single- to dual-chamber. Interestingly, pacemaker syndrome could basically develop at any time after implantation: At 6 months, it was present and required reprogramming in 14% of patients with VVIR pacing, in 16% at 1 year, in 18% at 2 years, and in 20% at 4 years. Slow sinus rates but not the presence or absence of VA conduction before pacemaker implantation predicted the development of pacemaker syndrome. QOL in patients with pacemaker syndrome was significantly worse than in patients without this complication of VVIR pacing, and it improved significantly after crossover to dual-chamber mode. Therefore, pacemaker syndrome with 20% incidence can be regarded as the most frequent complication of ventricular singlechamber pacing in SND, with significant impact on patients’ symptoms and QOL (Fig. 10-2). A second post hoc analysis of MOST detected one of the basic problems of studies on mode selection and changed the practice of cardiac pacing. Sweeney et al.38 found that the percentage of right ventricular (RV) pacing was strongly associated with the risk of developing HF and AF. This refers to patients with SND and narrow QRS at baseline. Interestingly, this association was more striking for dualchamber than single-chamber pacing. From these results, hypothetically, a “badly timed” (i.e., too early) ventricular stimulus may be more detrimental in dual-chamber pacing, where the pacing mode forces adverse AV timing to be present in every cycle, versus single-chamber ventricular pacing, where (besides pacemaker syndrome with VA conduction) only some cycles have by chance a detrimental AV relation, whereas in most cardiac cycles, atrium and ventricle are dissociated. The potential adverse effects of unnecessary RV pacing and inappropriate AV timing are problems not only in PASE and MOST, but in all trials comparing ventricular single- and dual-chamber pacing and may partly explain the lack of a positive effect of dual-chamber pacing in these studies.

Criterion 1: New or worsened dyspnea, orthopnea, elevated jugular venous pressure, rales, and edema with ventriculoatrial (VA) conduction during ventricular pacing

Criterion 2: Symptoms of dizziness, weakness, presyncope or syncope, and a ≥20 mm Hg reduction of systolic blood pressure when the patient had VVIR pacing compared with atrial pacing or sinus rhythm

23% Figure 10-2  Pacemaker syndrome in MOST. Percentage of patients meeting criterion 1 or 2 or both criteria in Mode Selection Trial in Sinus Node Dysfunction.

241

42%

35% Criterion 1

Criterion 2

Both criteria

242

TABLE

10-10 

SECTION 2  Clinical Concepts

Clinical Events in Canadian Trial of Physiologic Pacing (CTOPP)

Clinical Event (annual) Stroke Heart failure hospitalization Atrial fibrillation

VVIR (n = 1474) 1.1% 3.5% 6.6%

AAIR/DDDR (n = 1094) 1.0% 3.1% 5.3%*

Data from Connolly SJ, Kerr CR, Gent M, et al: Effects of physiologic pacing versus ventricular pacing on the risk of stroke and death due to cardiovascular causes. N Engl J Med 342:1385-1391, 2000. *P ≤ .05 for lower incidence with DDDR.

Canadian Trial of Physiologic Pacing (CTOPP) After PASE in 1998, the Canadian Trial of Physiologic Pacing was the second large RCT on pacing mode selection.72 It enrolled patients scheduled for pacemaker implantation without permanent atrial fibrillation at 32 Canadian centers. Patients were randomized to receive either a ventricular pacemaker or a “physiologic pacemaker” (i.e., atrial single or dual chamber) and followed for an average of 3 years. The primary outcome was stroke or death from a cardiovascular cause; secondary outcomes were all-cause mortality, AF development, and hospitalization for heart failure. CTOPP randomized 2568 patients, 1094 received a physiologic system (dual-chamber or atrial singlechamber pacemaker), and 1474 a ventricular single-chamber device. The annual rate of stroke and death from cardiovascular causes were similar in both groups (5.5% with ventricular pacing, 4.9% with physiologic pacing; P = 0.33) (Table 10-10). The incidence of AF was lower with physiologic versus ventricular pacing (annual rate, 5.3% vs. 6.6%; P = .05) with the difference emerging only 2 years after implantation. All-cause mortality and HF hospitalization were not significantly different, whereas postoperative complications occurred more frequently with physiologic pacing (9.0%), which consisted of atrial singlechamber pacing in only 5% and ventricular single-chamber pacing in another 5% of patients because of problems to implant the atrial lead. The authors of CTOPP concluded that physiologic pacing provides little or no benefit over ventricular pacing with regard to stroke prevention or cardiovascular death. Critical Appraisal.  Of the many questions about CTOPP, the authors cite the most important: in 56% of 1931 eligible patients who were not enrolled, it was the physician who declined participation. Heart failure of NYHA Class II or higher was present in significantly more eligible patients who were not enrolled (49%) than in those who were enrolled (39%). Because many physicians believed dual-chamber pacing to be superior to ventricular single-chamber pacing at patient recruitment in CTOPP, they likely declined to enroll patients when they were concerned about worse outcomes with ventricular single-chamber pacing, namely, HF patients and those who might require pacing most of the time. Therefore, CTOPP probably represents primarily patients with intermittent bradycardia, such as intermittent sinoatrial and atrioventricular block. A post hoc analysis supported the criticism that the lack of benefit of physiologic pacing may derive from CTOPP including many patients with only intermittent and rare bradycardia. Tang et al.73 compared patients with an unpaced heart rate before pacemaker implantation of ≤60 bpm to patients with an unpaced heart rate above 60 bpm. Risk of cardiovascular death or stroke was significantly higher in patients with ventricular single-chamber pacing than with physiologic pacing, if the unpaced heart rate before pacemaker implantation was ≤60 bpm, whereas patients with an unpaced heart rate greater than 60 bpm derived no benefit in this study endpoint from physiologic pacing. Similarly, only patients with an unpaced heart rate ≤60 bpm showed a better exercise capacity, as measured by 6-minute hall walk test with dual-chamber pacing, whereas pacing mode did not affect exercise capacity in patients with an unpaced heart rate greater than 60 bpm.74 One of the major criticisms of CTOPP refers to the lack of experience with atrial leads and dual-chamber pacing by many centers that enrolled a large number of patients.75 The results that centers were

unable to implant an atrial lead in 5% of patients, that incidence of postoperative complications for dual-chamber pacemaker implantation was 9%, and that more than 17% of dual-chamber devices were reprogrammed to VVI at 5-year follow-up reflect the lack of experience in implanting and following dual-chamber devices in some of the centers that participated in CTOPP. As an argument in favor of dual-chamber pacing, a prespecified subanalysis found “chronic atrial fibrillation” (defined as AF lasting >1 week) occurred significantly more often in ventricular single-chamber pacing.76 This trend was even more pronounced in 1474 patients from CTOPP in whom follow-up was extended to a mean of 6.4 years.77 However, even in this extended study period, no difference could be observed between dual-chamber and ventricular single-chamber pacing with regard to cardiovascular mortality and stroke. Additionally, the effect of pacing mode on AF incidence was regarded as less important after publication of the AFFIRM trial.78 In conclusion, CTOPP showed for the first time that even in a study with a large patient group and a long follow-up period, it is difficult to demonstrate an advantage in “hard endpoints” of dual- over singlechamber ventricular pacing. However, CTOPP was frequently regarded as a study with the “apparent intent of the trial to pattern cost control instead of optimize patient care.”75 Even in Canada, it never had an impact on pacing mode selection, which was characterized by a constant increase of the percentage of dual-chamber systems. United Kingdom Pacing and Cardiovascular Events (UKPACE) Trial Complementary to MOST, which randomized patients with SND to ventricular single- or dual-chamber pacing, the UKPACE trial randomized patients with AV block; 2021 patients age 70 or older with high-degree AV block randomly received a ventricular single-chamber or a dual-chamber pacemaker.79 In the single-chamber group, patients were further randomized to receive either fixed-rate or rate-adaptive pacing. The primary outcome was all-cause mortality; secondary outcomes were AF, heart failure, and stroke/transient ischemic attack (TIA)/systemic embolism. During the follow-up of 4.6 years, annual mortality did not differ (7.2% vs. 7.4% in single- vs. dual-chamber group) (Table 10-11). Surprisingly, there were also no significant differences between both groups in the incidence of AF, HF, or stroke/ TIA/embolism. The authors concluded that in elderly patients with high-degree AV block, pacing mode does not influence all-cause mortality or the incidence of cardiovascular events. Critical Appraisal.  It had been expected that results of the UKPACE trial might favor dual-chamber pacing, because patients with highdegree AV block are frequently pacemaker dependent and require ventricular pacing continuously. Patients with SND in MOST were usually not pacemaker dependent and may therefore have had periods when they did not require pacing. This could dilute the advantageous effects of “physiologic” compared to ventricular single-chamber pacing. Elderly patients with AV block enrolled in UKPACE may have been particularly sensitive to suboptimal hemodynamic conditions produced by asynchronous pacing in the VVI and VVIR modes. In their trial design paper, the authors of UKPACE therefore hypothesized that “it may be that such patients have much to gain from the superior hemodynamics of optimal pacing in terms of functional capacity and

TABLE

10-11 

Clinical Events in United Kingdom Pacing and Cardiovascular Events (UKPACE) Trial

Clinical Event (annual) All-cause mortality Cardiovascular mortality Thromboembolism Heart failure hospitalization Atrial fibrillation

VVI/VVIR (n = 1009) 7.2% 3.9% 2.1% 3.2% 3.0%

*P < .05 for higher incidence with DDDR.

DDDR (n = 1012) 7.4% 4.5%* 1.7% 3.3% 2.8%



10  Clinical Trials of Atrial and Ventricular Pacing Modes

Study

Physiologic

Ventricular

Wt%

Danish

39/110

57/115

4.1

0.66 [0.44,0.99]

CTOPP

390/1094

565/1474

40.8

0.92 [0.81,1.05]

PASE

32/203

34/204

2.9

0.94 [0.58,1.52]

MOST

200/1014

204/996

17.8

0.97 [0.8,1.18]

UKPACE

393/1012

387/1009

34.4

1.01 [0.88,1.16]

Overall

1054/3433

1247/3798

100

0.95 [0.87,1.03]

243

HR [95% CI]

Homogeneity: chi-square = 3.99 df = 4 p = 0.41 Association: chi-square = 1.69 p = 0.19 0.1

0.5

1.0

2.0

Hazard ratio Figure 10-3  Pooled estimate of mortality in randomized trials of pacing mode. Danish, Atrial versus ventricular pacing in sick sinus syndrome (Andersen et al.34,35); CI, Confidence interval; CTOPP, canadian Trial of Physiologic Pacing; HR, hazard ratio; MOST, Mode of Selection Trial in Sinus Node Dysfunction; PASE, Pacemaker Selection in the Elderly; UKPACE, United Kingdom Pacing and Cardiovascular Events trial.

quality of life, in addition to the possibility of reduced cardiovascular morbidity and improved survival.”80 The main finding of UKPACE, that dual-chamber pacing did not convey any advantages in terms of mortality, AF, stroke, or heart failure in elderly patients with AV block compared with single-chamber ventricular pacing, was therefore particularly disappointing. UKPACE was the only large RCT that found no difference in AF between single- and dual-chamber pacing. In fact, AF was detected more frequently in dualchamber pacing during the first 18 months of the study, perhaps from a systematic bias. In dual-chamber devices, AF can easily be diagnosed by the atrial lead and electrogram (EGM). In VVI and VVIR devices, the detection of AF is limited to the surface electrocardiogram (ECG) and can be difficult in patients with complete AV block and VVIR pacing.81 Explanations for the contradictory results between a number of hemodynamic studies demonstrating unequivocally a significant advantage of dual-chamber pacing on one side, and the lack of any clinically meaningful benefit of dual-chamber pacing in UKPACE on the other side, are difficult. Three prespecified UKPACE subanalyses (nuclear assessment of left ventricular function, Holter recording to assess prevalence of atrial arrhythmias, assessment of QOL and cognitive function) to date have not been published. Some studies on hemodynamics in pacing, particularly in the context of cardiac resynchronization therapy, have emphasized the importance of timing the AV delay to achieve acceptable results in dual-chamber pacing.82,83 Hemodynamic optimization of the AV delay had not been performed in UKPACE. On the contrary, programming instructions in UKPACE suggested the use of short AV delays (150 msec, with rate-adaptive shortening to 75 msec)80 to facilitate and ensure continuous ventricular pacing. In view of the deleterious effects of RV pacing, this may have negatively affected patients with dual-chamber devices (see Role of Prevention of Unnecessary Right Ventricular Pacing). It was not reported how many patients had only intermittent versus permanent second-degree or third-degree AV block. Continuous ventricular pacing may have been unnecessary and therefore disadvantageous in these patients. Programming of such short AV delays caused a significantly (P < .001) higher percentage of ventricular pacing in patients with dual-chamber pacing: Counters (available in 65% of patients) demonstrated a median of 99% RV pacing in dualchamber mode versus 94% in VVI and 93% in VVIR at 1 month.79 Inappropriately short AV delays can lead to premature mitral valve closure, causing pulmonary venous regurgitation by nonphysiologic

flow reversal during atrial systole.84 Because of device programming, this may have occurred more frequently in UKPACE than in other RCTs of pacing mode, leading to disadvantageous side effects that counteracted and precluded any potential benefit of dual-chamber pacing. META-ANALYSIS OF VENTRICULAR VERSUS ATRIAL-BASED PACING After UKPACE, Healey et al.85 performed a meta-analysis of the large trials on pacing mode selection (Danish study,34,35 PASE,65 CTOPP,72 MOST,67 UKPACE79). As expected, no advantage of dual-chamber pacing was found for mortality, for the combination of mortality and stroke, or for heart failure hospitalization (Figs. 10-3 to 10-8). The incidence of atrial fibrillation was lower in dual-chamber pacing. However, predominantly because of UKPACE results, a subanalysis showed a lower incidence of AF only for dual-chamber pacing in sinus node disease, but not in AV block. Even though this meta-analysis included the impressive number of 7231 patients, results cannot overcome the general limitations of the individual trials, such as selection bias and biased detection of AF when an atrial lead is present.

Atrial Single-Chamber Versus Dual-Chamber Pacing: Clinical Trials In 2004, Masumoto et al.86 performed a retrospective analysis of 196 patients with sinus node disease who were implanted with dualchamber devices programmed to DDD (n = 101) or AAI (n = 95) mode. Consecutive patients implanted between 1979 and 2000 were included if they had no second- or third-degree AV block, a maximum PR interval of 220 msec, and no left bundle branch block or bifascicular block. During a mean follow-up of 8 years, eight AAI devices had to be switched to DDD because high-degree AV block developed (incidence, 1.1%/yr). At 10 years, there was no need to add a ventricular lead in 89% of patients with AAI pacing. Excluding generator exchanges, 18 patients with DDD versus 17 with AAI system underwent reoperation, predominantly for lead failure. All-cause mortality was similar in both groups (10-year survival: 88% in AAI, 93% in DDD), as was freedom from permanent AF (94% in AAI, 91% in DDD) (Table 10-12). The authors concluded that AAI pacing can achieve similar clinical benefits as DDD pacing in sinus node disease with relevant AV or infrahisian conduction disturbances.

244

SECTION 2  Clinical Concepts

Study

Physiologic

Ventricular

Wt%

HR [95% CI]

Danish

25/110

45/115

4.2

0.54 [0.33,0.87]

CTOPP

307/1094

452/1474

47.6

0.91 [0.78,1.05]

MOST

112/982

124/958

15.3

0.88 [0.68,1.13]

UKPACE

266/1011

863/3551

33

1.09 [0.91,1.29]

Overall

710/3197

242/1004

100

0.94 [0.85,1.04]

Homogeneity: chi-square = 8.31 df = 3 p = 0.04 Association: chi-square = 1.64 p = 0.2 0.25

0.50

1.00

2.00

Hazard ratio Figure 10-4  Pooled estimate of cardiovascular mortality or stroke in randomized trials of pacing mode.

TABLE

10-12 

Clinical Events in AAI versus DDD pacing

Pacing Mode Atrial Dual chamber

All-Cause Mortality 12% 7%

Permanent AF 4% 8%

Lead Failure 15% 23%

Data from Masumoto H, Ueda Y, Kato R, et al: Long-term clinical performance of AAI pacing in patients with sick sinus syndrome: a comparison with dual-chamber pacing. Europace 6:444-450, 2004. AF, Atrial fibrillation.

Nielsen et al.87 compared myocardial blood flow and left ventricular ejection fraction (LVEF) in 30 patients with SND randomized to AAI versus DDD pacing. After 22 months of pacing in the assigned mode, myocardial blood flow was assessed with positron emission tomography (PET) and LVEF measured with echocardiography. Patients with dual-chamber pacemakers were temporarily programmed to AAI to obtain an intraindividual comparison. Myocardial blood flow was

similar in the groups with AAI and DDD pacing. However, blood flow significantly improved if the patient’s DDD pacing was temporarily programmed to AAI mode. LVEF was similar to baseline results after 22 months in AAI but had become significantly lower in patients with DDD pacing, from 61% to 56% (P = .01). However, if DDD pacing was temporarily changed to AAI, measurements of LVEF returned to baseline results. The authors concluded that dual-chamber compared with single-chamber AAI pacing has detrimental effects on myocardial blood flow and LVEF, which are still reversible after 22 months. In a single-center pilot study, consecutive patients with sinus node disease were screened and enrolled if they had no bundle branch block (BBB), and the spontaneous PQ interval did not exceed 220 msec (patients ≤70 years) or 260 msec (patients >70 years).88 Of 952 patients with a first pacemaker implantation between 1994 and 1999, 755 were excluded from the study, mainly for AV block (48%), permanent AF (10%), or BBB (5%), and only 2% of patients did not give consent. The 177 enrolled patients were randomized to receive an AAIR pacing system (n = 54) or a DDDR system programmed to a short AV delay

Study

Physiologic

Ventricular

Wt%

HR [95% CI]

Danish

26/110

40/115

4.5

0.54 [0.33,0.89]

CTOPP

224/1094

367/1474

40.3

0.8 [0.68,0.95]

PASE

35/203

38/204

5.3

0.91 [0.57,1.44]

MOST

217/1014

270/996

34.9

0.79 [0.66,0.94]

UKPACE

98/1012

111/1009

15.1

0.88 [0.67,1.16]

Overall

600/3433

826/3798

100

0.8 [0.72,0.89]

Homogeneity: chi-square = 3.26 df = 4 p = 0.52 Association: chi-square = 17.43 p = 3e–05 0.25

0.50

1.00

2.00

Hazard ratio Figure 10-5  Pooled estimate of atrial fibrillation in randomized trials of pacing mode.



10  Clinical Trials of Atrial and Ventricular Pacing Modes

Study

Physiologic

Ventricular

Wt%

HR [95% CI]

105/1094

156/1474

34.5

0.9 [0.7,1.15]

PASE

14/203

24/204

4.9

0.55 [0.28,1.06]

MOST

104/1014

123/996

31

0.82 [0.63,1.06]

UKPACE

111/1012

105/1009

29.7

1.04 [0.8,1.36]

Overall

334/3323

408/3683

100

0.89 [0.77,1.03]

CTOPP

245

Homogeneity: chi-square = 3.76 df = 3 p = 0.29 Association: chi-square = 2.46 p = 0.12 0.25

0.50

1.00

2.00

Hazard ratio Figure 10-6  Pooled estimate of hospitalization for heart failure in randomized trials of pacing mode.

(150 msec, DDDR-s, n = 60) or a long AV delay (300 msec, DDDR-l, n = 63). Echocardiographic (left atrial diameter, left ventricular function and size) and clinical (AF, stroke or arterial embolism, mortality, heart failure) outcomes were assessed in these three modes of pacing in SND. During the study, patients in the DDDR-s group received pacing for a mean of 90%. Patients in the DDDR-l group still received 17% of (a priori unnecessary) ventricular pacing. After a mean follow-up of 3 years, patients randomized to AAIR did not show a significant change in left atrial diameter, left ventricular diameter or LVEF shortening compared to baseline. In contrast, left atrial diameter was increased and left ventricular shortening fraction decreased significantly in the DDDR-s patients. In the DDDR-l group, the left atrial diameter and the left ventricular end-systolic and end-diastolic diameters increased significantly during the 3-year follow-up. Atrial fibrillation in the 12-lead ECG during at least one follow-up visit was significantly less common in AAIR (7.4%) than in DDDR-s (23.3%) or DDDR-l (17.5%).89 During the study, 16 strokes occurred in 14 patients, but no peripheral embolism was observed. Clinical events, including all-cause mortality, cardiovascular mortality, development of CHF, and increased use of diuretics, did not differ among the three groups. However, progression of HF symptoms, although not significantly different among the groups, was seen, with an increase of at least one NYHA class in 31% of patients paced in AAIR mode versus 30% in DDDR-s and 46% in DDDR-l. The study by Nielsen et al.87 would have required 450 patients for an 80% power to detect clinical or echocardiographic differences between AAIR and DDDR (no differences between DDDR-s and DDDR-l were expected). Inclusion was stopped when a multicenter trial (DANPACE) randomizing SND patients to AAIR versus DDDR pacing was initiated. However, the results of this study from Aarhus suggest that in patients with SND, narrow QRS and normal AV conduction dual-chamber pacing compared to AAIR pacing may result in deleterious effects in terms of left atrial enlargement, left ventricular volumes and ejection fraction, and AF incidence. Surprisingly, even in these select patients, programming the AV delay to very long values was unable to prevent a significant proportion of unnecessary right ventricular pacing. The observed proportion of ventricular pacing of 17% had detrimental effects in terms of left atrial dilatation and incidence of AF similar to DDDR pacing with short AV delay. In a study following these observations, Albertsen et al.90 randomized 50 patients with sinus node disease to AAIR or DDDR (AV delay 220 msec paced and 200 msec sensed). Primary endpoint was dyssynchrony (evaluated by tissue-Doppler imaging with at least 1 segment

with delayed longitudinal contraction) and LVEF (evaluated by 3D echocardiography) after 12 months compared to baseline. Physicians performing echocardiography were blinded with regard to pacing mode. Secondary endpoints referred to 6-minute walking test and NT-proBNP. During follow-up, patients with DDDR pacing received ventricular pacing for 66% of the time. Dyssynchrony slightly increased with DDDR pacing (mean of 1.3 delayed segments at baseline vs. 2.1 after 12 months) but not in AAIR pacing. Although LVEF decreased statistically significant with DDDR pacing (63% at baseline vs. 59% after 12 months; no change observed in AAIR pacing), there was no difference between AAIR and DDDR pacing after 12 months in NYHA class or NT-proBNP. These results confirm other observations of deteriorated left ventricular function with unnecessary right ventricular pacing, and suggest the development of left ventricular dyssynchrony as one cause. At the same time, echocardiographic changes from baseline did not translate into significant clinical changes in this small study. The Danish multicenter randomized trial on single-lead atrial versus dual-chamber pacing in patients with sick sinus syndrome (DANPACE) was started in 1999.91 It compared AAIR and DDDR pacing in patients with SND but with no BBB or AV block (except for PR interval ≤220 msec for patients <70 years and ≤260 msec for those ≥70 years). This trial used hardware randomization. Primary endpoint was all-cause mortality; secondary endpoints included detection of atrial fibrillation during scheduled follow-up, progression to permanent AF, incidence of stroke, heart failure, and the need for reoperation. It planned to include 1900 patients followed for a mean of 5.5 years to detect an absolute 6% difference in all-cause mortality with a power of 80%. The trial was terminated after 1415 patients, implanted in 20 centers in Denmark, United Kingdom, and Canada, were followed for a mean of 5.5 years. No patient was lost to follow-up. Patients in the DDDR group had been programmed to an AV delay of ≤220 msec paced and ≤200 msec sensed with rateadaptive shortening. This caused ventricular pacing for a mean of 65% of the time in DANPACE. All-cause mortality (30% in AAIR, 27% in DDDR) was not significantly different (P = 0.53) nor was there a difference in stroke, NYHA functional class, use of diuretics, or hospitalization for heart failure. To the surprise of the investigators, the incidence of paroxysmal AF was significantly lower in patients paced in the DDDR mode (RRR, 21%; P = .024). This difference appeared already within 6 months after randomization and was still observed in a multivariate analysis. However, permanent AF was similar in both pacing modes. Of clinical

246

SECTION 2  Clinical Concepts

SINUS NODE DISEASE Study

Physiologic

Ventricular

Wt%

HR [95% CI]

Danish

25/110

45/115

10

0.54 [0.33,0.87]

CTOPP

141/512

213/689

53.1

0.91 [0.74,1.13]

MOST

112/982

124/958

36.8

0.88 [0.68,1.13]

Overall

278/1604

382/1762

100

0.85 [0.73,1]

Homogeneity: chi-square = 3.92 df = 2 p = 0.14 Association: chi-square = 4.07 p = 0.044 0.25

A

1.00

0.50

2.00

Hazard ratio NO SINUS NODE DISEASE

Study

Physiologic

Ventricular

Wt%

HR [95% CI]

CTOPP

166/582

239/785

43.6

0.9 [0.74,1.1]

UKPACE

266/1010

242/1002

56.4

1.09 [0.91,1.13]

Overall

432/1592

481/1787

100

1 [0.88,1.14]

Homogeneity: chi-square = 2.01 df = 1 p = 0.16 Association: chi-square = 0 p = 0.98 0.80

B

1.00

1.25

Hazard ratio

Figure 10-7  Cardiovascular mortality or stroke, based on indication for pacing. A, Sinus node disease. B, Pacing indications other than sinus node disease.

significance is the observation of a doubling of the need for a reoperation during follow-up in AAIR pacing (hazard ratio = 2.0 with AAIR; 95% confidence interval = 1.54-2.61; P < .001), mainly for the need of a ventricular lead. Only one patient in the DDDR pacing group had to be upgraded to a biventricular system. The authors concluded that DDDR systems programmed to an AV delay of ≤220 msec instead of AAIR systems should be used in sinus node disease. The incidence of AF related to pacing mode is different in DANPACE compared to results from a 2003 study by Nielsen et al.88 In this pilot study in 177 patients, the incidence of AF was significantly lower with AAIR pacing (7.4%) compared to DDDR with long AV delay (17.5%) and DDDR with short AV delay (23.3%). In the SAVE PACe trial, patients had significantly less persistent AF if randomized to a dualchamber mode allowing single nonconducted P waves not to be tracked, to minimize ventricular pacing in the dual-chamber mode, than patients with conventional DDDR pacing92 (see Role of

Prevention of Unnecessary Right Ventricular Pacing). However, patients with SND and conventional dual-chamber pacing in DANPACE received RV pacing for only 65% of the time versus 99% in SAVE PACe. In conclusion, several earlier studies comparing AAIR and DDDR pacing in patients with sinus node disease and narrow QRS complex suggested an advantage of the single-chamber atrial pacing mode that systematically rules out unnecessary ventricular pacing. This was strongly confirmed by post hoc analyses from MOST, in which unnecessary RV pacing was associated with worse clinical outcome. However, two randomized prospective trials (DANPACE and MVP trial; see later) did not confirm a benefit of AAIR pacing (DANPACE) or DDDR pacing with strict prevention of RV pacing in patients with an implantable cardioverter-defibrillator (ICD). Most likely, the development of first-degree AV block with long PR intervals or higher-degree AV block adversely affects hemodynamics if the ventricle is left unpaced. AV



10  Clinical Trials of Atrial and Ventricular Pacing Modes

247

SINUS NODE DISEASE Study

Physiologic

Ventricular

Wt%

HR [95% CI]

Danish

26/110

40/115

6.4

0.54 [0.33,0.89]

CTOPP

150/512

252/689

39.2

0.78 [0.64,0.96]

PASE

17/91

24/85

4.1

0.57 [0.31,1.07]

MOST

217/1014

270/996

50.2

0.79 [0.66,0.94]

Overall

410/1727

586/1885

100

0.76 [0.67,0.86]

Homogeneity: chi-square = 2.84 df = 3 p = 0.42 Association: chi-square = 18.62 p = 1.6a–05 0.25

A

1.00

0.50

2.0

Hazard ratio NO SINUS NODE DISEASE

Study

Physiologic

Ventricular

Wt%

HR [95% CI]

CTOPP

74/582

115/785

42.9

0.84 [0.63,1.12]

PASE

18/112

14/119

7.5

1.46 [0.73,2.94]

UKPACE

98/1011

111/1007

49.6

0.88 [0.67,1.16]

Overall

190/1705

111/1911

100

0.9 [0.74,1.09]

Homogeneity: chi-square = 2.09 df = 2 p = 0.35 Association: chi-square = 1.24 p = 0.27 0.5

B

1.0

2.0

Hazard ratio

Figure 10-8  Atrial fibrillation based on indication for pacing. A, Sinus node disease. B, Pacing indications other than sinus node disease.

block may also enhance the development of AF, thus questioning the concept that RV pacing should always be systematically avoided for its associated ventricular dyssynchrony. Therefore, at present, there are no firm recommendations on the ideal pacing mode in sinus node disease. AAI pacing does not appear as desirable as before DANPACE. Potentially, implantation of a dualchamber device programmed to maintain AAI pacing as long as sufficient for proper AV synchrony may be the best option, avoiding reoperation whenever first-degree AV block or higher develops. The role of first-degree AV block induced by atrial pacing from sites far from the septum (e.g., right atrial appendage, lateral right atrial wall) needs to be further evaluated. Atrial pacing–induced AV desynchronization with first-degree paced AV block may be one mechanism limiting the potential benefits of atrial and dual-chamber pacing in patients with SND.

Dual-Chamber Versus Ventricular Pacing in Defibrillator Patients The possibility of physiologic pacing in patients with an indication for an ICD and dual-chamber pacing was highly desired before it was available.93 However, the first studies on dual- versus single-chamber ICDs were disappointing, failing to show improved discrimination between ventricular and supraventricular tachycardia.94-96 Some newer studies question if single-chamber ICDs are comparable to dualchamber ICDs in prevention of inappropriate therapy.97-99 However, one of the aims of dual-chamber pacing in ICD patients is the reduction in clinical events, particularly heart failure and HF progression caused by bradycardia, which frequently develops during treatment with β-blockers and amiodarone. Even an increased risk of ventricular tachyarrhythmia (e.g., from bradycardia-dependent

TABLE

10-13 

SECTION 2  Clinical Concepts

Clinical Endpoints in Dual Chamber and VVI Implantable Defibrillator (DAVID) Trial at 1 Year

Pacing DDDR (70 bpm) VVI (40 bpm)

All-Cause Mortality 10.1% 6.5%

HF Hospitalization 22.6% 13.3%

Combined 26.7%* 16.1%*

Data from Wilkoff BL, Cook JR, Epstein AE, et al: Dual-chamber pacing or ventricular backup pacing in patients with an implantable defibrillator: the Dual Chamber and VVI Implantable Defibrillator (DAVID) Trial. JAMA 288:3115-3123, 2002. *P < .05. HF, Heart failure.

arrhythmias) has been observed in ICDs that were programmed to a VVI 40-bpm backup pacing mode.100 The Dual Chamber and VVI Implantable Defibrillator (DAVID) trial therefore evaluated if dualchamber pacing reduced the incidence of HF hospitalization and allcause mortality compared to backup ventricular pacing in patients with a conventional indication for ICD therapy and heart failure (LVEF ≤40%) but no standard indication for antibradycardia pacing.101 In 37 U.S. centers, 506 patients were enrolled and randomized to DDDR pacing at 70 bpm versus VVI backup pacing at 40 bpm (software randomization, DDDR systems implanted in all patients). Programming of other parameters, particularly the AV delay, was left to the individual investigator. The mean sinus rate at baseline was 72 bpm, the mean PR interval of patients in DAVID was 184 msec, and 27% of patients had a BBB. At 6 months, 86% of patients received β-blockers and 27%, amiodarone. The trial was prematurely stopped at a mean follow-up of 8 months because, unexpectedly, patients with dual-chamber ICDs had a significantly higher incidence of HF hospitalization (43 [23%] at 1-year follow-up) and death (23 [10%] at 1 year) than patients with single-chamber ICDs (30 and 15, equivalent to 1-year incidence of 13% and 6.5%). This difference reached statistical significance in favor of single-chamber ICDs (Table 10-13). One major difference in patients with single-chamber ICDs was the significantly lower proportion of ventricular pacing. At the 6-month control, device counters documented that VVI 40-bpm back-up pacing was associated with a mean proportion of ventricular pacing of only 0.6% in these patients without bradycardia compared to 59.6% ventricular pacing in the DDDR mode. A post hoc analysis corroborated the hypothesis that right ventricular pacing was the cause of worse outcome in patients with dual-chamber ICDs.102 In 380 patients who had a 3-month follow-up and had not yet reached a study endpoint, a multivariate analysis demonstrated that the percent of cumulative RV pacing was correlated with death or HF hospitalization during further study. If patients with dual-chamber ICDs had RV pacing more than 40% of the time, they had a more than fivefold higher risk of HF hospitalization or death than patients with a dual-chamber ICD and RV pacing for 40% or less of the time (P = .02). Patients with DDDR pacing at 70 bpm and ≤40% RV pacing trended toward a better outcome than those with VVI 40-bpm backup pacing. Another analysis assessed data in patients with “soft” indications for pacing, such as bradycardia, defined as sinus bradycardia of less than 60 bpm, and without a conventional indication for pacing.103 Results similar to the overall analysis were observed in these patients. However, the two soft indications are fundamentally different; patients with firstdegree AV block may have experienced specifically deleterious effects of dual-chamber pacing because they likely were continuously paced (and unpaced in VVI 40-bpm group), whereas patients with sinus bradycardia might have specifically benefited from dual-chamber pacing as long as a sufficiently long programmed AV delay prohibited unnecessary RV pacing. However, this was not reported in this post hoc analysis. These data were partly replicated in the second Multicenter Automatic Defibrillator Implantation Trial (MADIT-II), although differences between pacing modes in a primary prevention ICD trial did not reach statistical significance.104 In 404 patients who received a singlechamber ICD and 313 patients with a dual-chamber ICD for primary prevention of sudden cardiac death after myocardial infarction (LVEF ≤30%), the trend was toward an increase in hospitalization for heart

failure in patients with dual-chamber devices. Because of the nonrandomized assignment to dual-chamber ICDs, however, patients who had received dual-chamber devices were older and had higher NYHA functional class, more frequent prior coronary artery bypass surgery, wider QRS complex, more frequent left BBB, atrial tachyarrhythmias, and higher nitrate urea levels. The DAVID investigators performed a second study with VVI backup pacing at 40 bpm in patients with heart failure, an indication for ICD therapy, and no conventional indication for bradycardia pacing (DAVID-II).105 This time the dual-chamber group was programmed to AAI 70 bpm and tested for “non-inferiority” compared to backup VVI pacing at 40 bpm. In 29 U.S. centers, 543 patients were enrolled and implanted with a dual-chamber ICD programmable to the VVI and AAI pacing modes. Of note, 57 patients from the first DAVID trial who did not reach a study endpoint were also included. During a mean follow-up of 2.7 years, no significant difference was seen in the primary combined endpoint of time to HF hospitalization or death (incidence 11%, 17%, and 25% at 1, 2, and 3 years, respectively). In a subanalysis, asymptomatic sinus bradycardia at baseline was associated with a lower incidence of adverse outcome (12% vs. 26%) in patients randomized to VVI 40-bpm back-up or AAIR 70-bpm pacing.106 The question if dual-chamber ICDs have any potential to reduce adverse clinical events compared to single-chamber ICDs was assessed in the Dual Chamber and Atrial Tachyarrhythmias and Adverse Events Study (DATAS).107 Patients with Class I indication for an ICD (according to ACC/AHA 1998 guidelines) were included if permanent atrial tachyarrhythmias and an indication for dual-chamber pacing were absent. A total of 334 patients were then randomized to single-chamber ICD, dual-chamber ICD, or “simulated single-chamber ICD” (i.e., dual-chamber ICD implanted but programmed to single-chamber mode for pacing and discrimination between supraventricular and ventricular tachyarrhythmia). With crossover of the latter two groups after 8 months, all patients were followed for 16 months. The primary outcome was a composite of all-cause-mortality, need for cardiovascular hospitalization or intervention, two or more inappropriate shocks, and sustained symptomatic atrial tachyarrhythmia for 48 hours or longer. This composite was expressed as a prespecified score. Any one of these endpoints occurred in 65 patients with dual-chamber ICD (58%), 82 patients with single-chamber ICD (74%), and 84 patients with a dual-chamber ICD programmed to single-chamber parameters (76%). Mortality was 4% in dual-chamber, 9% in singlechamber, and 10% in simulated single-chamber ICD patients. The composite endpoint score was 33% lower for patients with dualchamber ICDs (Fig. 10-9). 50

46

45

42

40 Percent

248

38

35 30 25 20 15 10 5

16

13

9 10

12 5

4

3

6

0 Death Dual-chamber

Hospitalization Single-chamber

HF Inappropriate hospitalizations shocks Simulated single-chamber

Figure 10-9  Clinical events in DATAS. All numbers represent percentages in Dual Chamber and Atrial Tachyarrhythmias Adverse Events Study. HF, heart failure.

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10  Clinical Trials of Atrial and Ventricular Pacing Modes

In ICD patients, study results suggest that prevention of unnecessary right ventricular pacing plays a major role. Comparing single- and dual-chamber devices, trials showed a significant increase particularly in HF hospitalization with dual-chamber ICDs that was not caused by the dual-chamber mode itself but by RV pacing. This can lead to severe deterioration of ventricular function, most likely through desynchronization of right and left ventricular contraction, especially in preexisting HF. Of note, studies randomizing ICD patients to single- or dual-chamber pacing enrolled only patients without an indication for antibradycardia pacing. Results therefore do not apply to ICD patients with SND or AV block. Also, patients without structural heart disease where overdrive pacing may play a role (e.g., patients with long QT syndrome) are not represented in studies randomizing patients to single-chamber versus dual-chamber ICDs. Other, “soft” indications have not been systematically studied. Although patients with firstdegree AV block seem to have a particularly high percentage of unnecessary RV pacing with implications for HF progression (see below), sinus bradycardia, which may contribute and aggravate HF in some patients, may have a different outcome if ventricular pacing is prevented. Newer studies and post hoc analyses avoiding unnecessary RV pacing and using dedicated programming for better discrimination between supraventricular and ventricular tachyarrhythmias suggest some advantage for dual-chamber ICDs over single-chamber ICDs with 40-bpm backup VVI pacing.

difference in the development of AF emerged only after 2 years, and in light of the results of AFFIRM,78 was deemed not clinically relevant. At the same time, in patients who required pacing most of the time (“unpaced heart rate <60 bpm”), physiologic pacing significantly reduced stroke, cardiovascular mortality, and even all-cause mortality, rendering CTOPP the most positive trial on pacing mode selection in favor of physiologic pacing.

Critical Appraisal of Pacing Mode Trials

UKPACE

The large randomized trials on pacing mode dramatically changed the perception of “physiologic pacing.” In studies of the late 1980s, the addition of an atrial lead to a ventricular lead was perceived as a great step forward in cardiac pacing. The dual-chamber pacing mode enabled ventricular pacing at a programmable time after atrial sensing (AV-synchronous pacing), which improved cardiac output significantly compared with “asynchronous” ventricular pacing in the VVI(R) mode, as well as reducing mitral regurgitation, right atrial pressure, and pulmonary capillary wedge pressure and resolving pacemaker syndrome in a significant number of patients. Retrospective studies confirmed these hemodynamic advantages of dual-chamber pacing by clinical data. However, these clinical data from retrospective and nonrandomized studies were questioned108,109 when selection bias became evident.61,110 It is disappointing and difficult to explain why the theoretical advantages and the positive results of early studies did not translate into positive results in large, randomized clinical trials. However, some findings deserve special mention. PASE The Pacemaker Selection in the Elderly Study could be seen as evidence that elderly patients with conventional indications for permanent pacing do not seem to derive any benefit from dual-chamber over single-chamber pacing in terms of QOL, cardiovascular events, or mortality. However, PASE could also be viewed as evidence that approximately 25% of patients with sinus rhythm develop severe symptoms with VVIR pacing, which can be resolved by dual-chamber pacing in most cases. The high crossover rate in PASE precludes any firm interpretation of the results. Therefore, PASE does not answer the question if DDDR pacing provides an advantage over VVIR pacing in elderly patients. CTOPP The Canadian Trial of Physiologic Pacing is usually regarded as negative because three of four endpoints (stroke/cardiovascular events, allcause mortality, heart failure) were not reduced with physiologic compared to ventricular pacing. Even the statistically significant

MOST Although the Mode Selection Trial in Sinus Node Dysfunction was negative with regard to a difference in stroke, all-cause mortality, cardiovascular mortality, and HF hospitalization in patients with SND paced in VVIR versus DDDR mode (confirming lower incidence of AF with dual-chamber pacing and showing some superiority of dualchamber pacing in QOL), MOST demonstrated a surprisingly high incidence of severe pacemaker syndrome and, despite special precautions in the study protocol, an even higher crossover rate from VVIR to DDDR as a result of problems with single-chamber pacing. If these results would translate into daily clinical practice, MOST would suggest that every third patient with SND and VVIR pacing would require a second surgery for an upgrade to dual-chamber pacing, an economic and medical disaster. The most important observation in MOST may be the deleterious effect of unnecessary RV pacing (see later).

From the information published, the United Kingdom Pacing and Cardiovascular Events study is the most negative trial on pacing mode selection. It clearly showed that death can be prevented in AV block as effectively by single-chamber as by dual-chamber pacemakers. Any difference between these pacing modes with regard to exercise capacity, symptoms, and QOL remains unclear. The lack of difference between single- and dual-chamber pacing in AV block could be explained by patient selection bias in UKPACE (only 45% of eligible patients were enrolled) and by inappropriately short AV delays (per protocol 75-150 ms) with adverse hemodynamic consequences. The “forced” proportion of ventricular pacing of 99% in the dual-chamber mode which was significantly more than in VVI and VVIR pacing (93 and 94%) may theoretically have further deteriorated clinical results of dualchamber pacing (see Role of Prevention of Unnecessary Right Ventricular Pacing). OTHER ISSUES In addition to interpretations of the findings in the large RCTs on pacing mode selection, the following issues should be considered when evaluating the clinical implications of these trials: 1. Experience of study centers with the more complex, dual-chamber systems (atrial lead dislodgement, atrial over/undersensing, programming issues). 2. Lack of delineation of patient subgroups, such as patients with permanent bradycardia and pacemaker dependence. 3. Insufficient instruments to quantify pacing-associated symptoms, quality of life, and exercise capacity, particularly in elderly patients with comorbidity. 4. Unequal diagnostic tools to detect atrial fibrillation, an endpoint in all trials. The atrial lead allows easy detection of AF with the atrial EGM, whereas in ventricular devices, detection of AF relies on the surface ECG and may be difficult in complete AV block.81 5. Too few trials with sufficient long-term follow-up. The randomized trials on pacing mode selection demonstrate that mortality is similar with ventricular single- compared to dual-chamber pacing. This mainly refers to patients with atrioventricular block; for those with sinus node disease, it has not been convincingly shown that pacing improves survival compared to the natural course. It also became quite evident that dual-chamber pacing is associated with less AF than ventricular single-chamber pacing. However, the central

250

SECTION 2  Clinical Concepts

35

optimized to the patient (lead position, AV delay, pacing rates), (3) dual-chamber pacing is more complex and therefore associated with more pitfalls and potential for complications (atrial lead dislodgement, pacemaker mediated tachycardia), and (4) dual-chamber pacing is not as “innocent” as previously thought because it may cause more poorly timed right ventricular pacing than VVI(R), with potentially deleterious effects.

31 Intolerance to VVI (R) in %

30 26 25 20 15

Role of Prevention of Unnecessary Right Ventricular Pacing

10 4

5

3

0 PASE

MOST

CTOPP

UKPACE

Figure 10-10  Intolerance to VVI(R) in PASE, MOST, CTOPP, and UKPACE trials. Blue columns, Studies with software randomization requiring only reprogramming to crossover to DDD(R). Red columns, Studies with hardware randomization requiring reoperation to crossover to DDDR.

question is whether dual-chamber pacing is associated with improved hemodynamics that translate into less heart failure, less symptoms, and better QOL and exercise capacity; this remains unanswered even after randomizing more than 7000 patients. Beyond criticism about how such “soft endpoints” should best be evaluated (e.g., SF-36 unable to detect subtle, pacing-related QOL differences; symptom scores not sufficiently validated and not used), there are some critical, unresolved issues. First, randomized trials on pacing mode selection suffer from selection bias, particularly if hardware randomization is applied. From theoretical considerations and earlier studies, investigators have ethical concerns regarding whether pacemaker-dependent patients are at an increased risk for adverse events with ventricular single-chamber pacing. It is a strength of published trials (particularly CTOPP and UKPACE, which were hardware randomized) to keep track of why eligible patients were not enrolled into the trial: In CTOPP, only 16% of eligible but not enrolled patients declined consent, whereas in 56% the physician declined.72 Second, in a blinded trial, patient preference would be a powerful argument in favor of or against a pacing mode.25,26,62,63 Unfortunately, but logically, all large RCTs on pacing mode selection had a parallel trial design preventing intraindividual comparisons. Intraindividual comparisons could only be made in crossover patients (because of suspected pacemaker syndrome) from VVIR to DDDR pacing in PASE and MOST, showing a striking preference for DDDR pacing in these select patients. Third, the choice of software versus hardware randomization has both advantages and problems. If a patient “software randomized” to VVIR expresses symptoms, it is easy to reprogram the device to DDDR and see if symptoms are resolved. This may lead to an overestimation of the incidence of “VVIR intolerance” (26-31% in PASE and MOST). On the other hand, in hardware randomization, crossover to dualchamber pacing requires surgical intervention with potential complications, medically and legally. It is therefore evident that the incidence of “VVIR intolerance” is underestimated in hardware-randomized trials (3.1% in UKPACE, 4.3% in CTOPP) (Fig. 10-10). In summary, despite the RCT results on pacing mode selection, the perception that dual-chamber pacing is superior to single-chamber ventricular pacing in sinus rhythm still governs current guidelines111,112 and international physicians’ choices75,113 of pacing mode selection. However, trials have shown that (1) the beneficial effects of dualchamber pacing may have been overestimated (see rationale and sample size calculation of the Danish trial and UKPACE), (2) dualchamber pacing by itself may not convey hemodynamic benefits if not

Surprising data from the DAVID trial showed that compared with VVI 40-bpm backup pacing, DDDR pacing at 70 bpm produced more, not less, heart failure hospitalization and a slight, statistically not significant increase risk of mortality in patients with an ICD and no bradycardia101 (see earlier). In search of an explanation, it became obvious that right ventricular pacing was present in dual-chamber pacing much more frequently than in ventricular backup pacing (56% vs. 3% of patients with paced ventricular rhythm after 6 months). Similar results have been reported (independent of the presence of a single- or dual-chamber ICD) for patients in MADIT-II, where a cumulative proportion of ventricular pacing of more than 50% was associated with an almost twofold increased risk of new or worsened heart failure.114 A number of studies on histology,115 hemodynamics,116 and echocardiography117 have demonstrated deleterious acute and long-term consequences of RV apical pacing, their full clinical impact emerged only after DAVID. The MOST trial enrolled only patients with sinus node disease; apart from 11% of patients with additional second- or third-degree AV block,67 patients who did not require ventricular pacing. Sweeney et al.38 investigated the effect of unnecessary RV pacing in patients with SND. They selected 1339 patients from MOST who had a narrow QRS (<120 msec) at baseline before pacing was instituted. Taken from pacemaker memory functions, 707 patients randomized to DDDR received RV pacing at an average 90% of the time. In contrast, in patients randomized to VVIR, RV pacing was present only 58% of the time. In a Cox proportional hazard model, RV pacing (again, unnecessary in SND patients) was a strong predictor of HF hospitalization in DDDR mode (Table 10-14). For percentages of RV pacing in the AV-synchronous mode up to 40%, every 10% increase in ventricular pacing linearly increased the relative risk increase for HF hospitalization by 54%. RV pacing in the DDDR mode more than 40% of the time versus less than 40% was associated with a 2.6-fold increased risk of HF hospitalization. This association was much more striking for DDDR than for VVIR pacing, in which RV pacing up to 80% of the time had only a minor impact on HF development, whereas pacing more than 80% of the time resulted in a 2.5-fold increased risk of HF hospitalization. Similar to the increased risk of heart failure hospitalization, unnecessary right ventricular pacing also increased the risk of atrial fibrillation.38 Up to 80% to 85% of RV pacing in the DDDR or VVIR modes, the risk of AF linearly increased by approximately 1% for every percent increase in cumulative ventricular pacing. The authors concluded that

TABLE

10-14 

Association of Percentage of Ventricular Pacing (VP) with Heart Failure Hospitalization (HF Hosp) and Atrial Fibrillation (AF) for Dual-Chamber and Single-Chamber Ventricular Pacing DDDR (n = 707)

VP (%) <10 10-50 50-90 >90 Total

HF Hosp 2% 9% 9% 12% 10%

AF 16% 19% 32% 17% 21%

VVIR (n = 6432) HF Hosp 7% 6% 8% 16% 9%

AF 21% 23% 29% 18% 24%

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10  Clinical Trials of Atrial and Ventricular Pacing Modes

unnecessary RV pacing in SND leads to ventricular desynchronization in asynchronous as well as in AV-synchronous RV pacing, and over a broad range of cumulative percentages of pacing, linearly increases the risk of development of HF and AF in patients without BBB. The higher percentage of RV pacing in dual-chamber than in ventricular single-chamber pacing may be one explanation for the lack of superiority of this pacing mode in large trials. A post hoc analysis confirmed this for the DAVID trial.102 The authors dichotomized the group of patients with DDDR pacing at 70 bpm into those with 40% or less versus those with more than 40% RV pacing. In a Cox regression analysis, ventricular pacing 40% or less of the time was associated with slightly fewer clinical events than with VVI 40-bpm backup (P = .07). Patients with ventricular pacing more than 40% of the time in DDDR mode had a 5.3-fold increased risk of a study endpoint (death or HF hospitalization) than patients with 40% or less ventricular pacing. Prevention of unnecessary RV pacing i