Cardiovascular Magnetic Resonance, Second Edition

Cardiovascular Magnetic Resonance Cardiovascular Magnetic Resonance SECOND EDITION Warren J. Manning, MD Professor o...

Author: Warren J. Manning MD | Dudley J. Pennell MD FRCP FACC

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Cardiovascular Magnetic Resonance

Cardiovascular Magnetic Resonance SECOND EDITION

Warren J. Manning, MD Professor of Medicine and Radiology Harvard Medical School Section Chief, Non-Invasive Cardiac Imaging Beth Israel Deaconess Medical Center Boston, Massachusetts

Dudley J. Pennell, MD, FRCP, FACC, FESC Professor of Cardiology National Heart and Lung Institute, Imperial College Director, Cardiovascular MR Unit Royal Brompton Hospital London, United Kingdom

Saunders / Elsevier Philadelphia, PA

An Imprint of Elsevier Inc. 1600 John F. Kennedy Blvd. Ste 1800 Philadelphia, PA 19103-2899 CARDIOVASCULAR MAGNETIC RESONANCE, Second Edition

ISBN 978-0-443-06686-3

Copyright # 2010, 2002 by Saunders, an imprint of Elsevier Inc. 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 Cardiovascular magnetic resonance / [edited by] Warren J. Manning, Dudley J. Pennell. – 2nd ed. p. ; cm. Rev. ed. of: Cardiovascular magnetic resonance / Warren J. Manning, Dudley J. Pennell. – 1st ed. 2002. Includes bibliographical references and index. ISBN 978-0-443-06686-3 1. Heart–Magnetic resonance imaging. I. Manning, Warren J. II. Pennell, Dudley J., 1958- III. Manning, Warren J. Cardiovascular magnetic resonance. [DNLM: 1. Cardiovascular Diseases–Diagnosis. 2. Magnetic Resonance Imaging–methods. 3. Diagnostic Techniques, Cardiovascular. WG 141.5.M2 C2664 2010] RC683.5.M35M364 2010 616.10 207548–dc22 2010003032

Acquisitions Editor: Rebecca Schmidt Gaertner Editorial Assistant: David Mack Design Direction: Ellen Zanolle

Printed in the United States of America. Last digit is the print number: 9

8 7 6 5 4

3 2 1

To the joys and inspirations of my life— Susan Gail, Anya, Sara, Isaac, and Elie ——WJM To my parents Terence and Joan for ever being full of pride despite the vicissitudes of age, Elisabeth for always exceeding the singular prowess proffered at our wedding—despite denial, and Indigo Lucy Li-Ling for inspiring love and joy in boundless measure (how many times today Indi?). ——DJP

CONTRIBUTORS

Contributors Silvia Aguiar, MD Translational and Molecular Institute, Departments of Radiology and Medicine (Cardiology), Mount Sinai School of Medicine, New York, New York Chapter 25: Atherosclerotic Plaque Imaging: Aorta and Carotid Timothy S. E. Albert, MD Medical Director, Cardiovascular Diagnostic Center, Salinas Valley Memorial Healthcare System, Monterey, California; Assistant Consulting Professor of Medicine, Duke University Medical Center, Durham, North Carolina Chapter 2: Clinical Cardiovascular Magnetic Resonance Imaging Techniques Francisco Alpendurada, MD Consultant, CMR Unit, Royal Brompton Hospital, London, United Kingdom Chapter 39: Cardiac and Paracardiac Masses Andrew E. Arai, MD Senior Investigator, National Heart, Lung, and Blood Institute, National Institutes of Health, Bethesda, Maryland Chapter 18: Acute Myocardial Infarction: Cardiovascular Magnetic Resonance Detection and Characterization Robert S. Balaban, PhD Scientific Director, Division of Intramural Research, National Heart, Lung, and Blood Institute, National Institutes of Health, Bethesda, Maryland Chapter 1: Basic Principles of Cardiovascular Magnetic Resonance Jeroen J. Bax, MD Professor of Cardiology, Department of Cardiology, Leiden University Medical Center, Leiden, The Netherlands Chapter 8: Special Considerations for Cardiovascular Magnetic Resonance Safety, Electrocardiographic Setup, Monitoring, and Contraindications Nicholas G. Bellenger, MD Consultant Cardiologist, CMR Unit, Royal Brompton Hospital, London, United Kingdom Chapter 14: Assessment of Cardiac Function

David A. Bluemke, MD Professor, Radiology and Medicine, Johns Hopkins University School of Medicine, Baltimore; Director, Radiology and Imaging Sciences, National Institutes of Health Clinical Center; Senior Investigator, National Institute of Biomedical Imaging and Bioengineering, Bethesda, Maryland Chapter 35: Pulmonary Artery Cardiovascular Magnetic Resonance Rene´ M. Botnar, PhD Professor of Cardiovascular Imaging, Chair of Cardiovascular Imaging, Imaging Sciences Division, King’s College London, London, United Kingdom Chapter 21: Coronary Artery and Vein Imaging: Methods Chapter 26: Atherosclerotic Plaque Imaging: Coronaries Jens Bremerich, MD Head, Cardiothoracic Radiology, University Hospital Basel, Basel, Switzerland Chapter 30: Cardiovascular Magnetic Resonance in Complex Congenital Heart Disease Peter T. Buser, MD Professor of Cardiology, Chairman, Department of Cardiology; Head, Cardiac Imaging, University Hospital Basel, Basel, Switzerland Chapter 30: Cardiovascular Magnetic Resonance in Complex Congenital Heart Disease Peter Caravan, PhD Assistant Professor, Radiology, Harvard Medical School; Assistant in Chemistry, Athinoula A. Martinos Center for Biomedical Imaging, Massachusetts General Hospital, Boston, Massachusetts Chapter 6: Cardiovascular Magnetic Resonance Contrast Agents Jonathan Chan, MD Advanced Cardiovascular Imaging Fellow, Department of Medicine, Cardiovascular Division, Beth Israel Deaconess Medical Center, Boston, Massachusetts Chapter 22: Coronary Artery Imaging: Clinical Results Michael L. Chuang, MD Advanced Cardiovascular Imaging Fellow, Department of Medicine, Cardiovascular Division, Beth Israel Deaconess Medical Center, Boston, Massachusetts Chapter 11: Normal Cardiac Anatomy, Orientation, and Function

Cardiovascular Magnetic Resonance vii

CONTRIBUTORS

Albert de Roos, MD Professor of Radiology, Department of Radiology, Leiden University Medical Centre, Leiden, The Netherlands Chapter 29: Cardiovascular Magnetic Resonance of Simple Congenital Cardiovascular Defects Rohan Dharmakumar, MD Department of Radiology, Northwestern University, Chicago, Illinois Chapter 42: Magnetic Resonance Assessment of Myocardial Oxygenation Adam L. Dorfman, MD Assistant Professor, Department of Pediatrics and Communicable Diseases; Assistant Professor, Department of Radiology, University of Michigan Medical School, Ann Arbor, Michigan Chapter 31: Complex Congenital Heart Disease: Infant and Pediatric Patients Christopher K. Dyke, MD Consulting Associate, Duke University Medical Center, Durham, North Carolina; Adjunct Faculty, University of Washington School of Medicine, Seattle, Washington; Director, Alaska Heart Cardiovascular MRI and CT Centers, Alaska Heart Institute, Anchorage, Alaska Chapter 2: Clinical Cardiovascular Magnetic Resonance Imaging Techniques Robert R. Edelman, MD William B. Graham Chair, Department of Radiology, North Shore University Health System; Professor of Radiology, Feinberg School of Medicine, Northwestern University, Evanston, Illinois Chapter 34: Cardiovascular Magnetic Resonance Angiography: Carotids, Aorta, and Peripheral Vessels Michael D. Elliott, MD Director, Cardiovascular MR/CT, St. Vincent Heart Center, Indianapolis, Indiana Chapter 2: Clinical Cardiovascular Magnetic Resonance Imaging Techniques Zahi A. Fayad, PhD Professor, Department of Radiology and Medicine (Cardiology); Director, Translational Molecular Imaging Institute, Mount Sinai School of Medicine, New York, New York Chapter 25: Atherosclerotic Plaque Imaging: Aorta and Carotid David Firmin, MD Professor of Physics, Royal Brompton Hospital and Imperial College, London, United Kingdom Chapter 7: Blood Flow Velocity Assessment Chapter 10: Use of Navigator Echoes in Cardiovascular Magnetic Resonance and Factors Affecting Their Implementation

viii Cardiovascular Magnetic Resonance

Mark A. Fogel, MD Associate Professor of Pediatrics and Radiology, The University of Pennsylvania School of Medicine; Director of Cardiac Magnetic Resonance, The Children’s Hospital of Philadelphia, Philadelphia, Pennsylvania Chapter 9: Special Considerations: Cardiovascular Magnetic Resonance in Infants and Children Herbert Frank, MD Professor in Internal Medicine; Director, Department of Internal Medicine, Landesklinikum Tulln and Medical University of Vienna, Vienna, Austria Chapter 39: Cardiac and Paracardiac Masses Matthias G. Friedrich, MD Associate Professor, Libin Cardiovascular Institute of Alberta, University of Calgary, Calgary, Alberta; Director, Stephenson CMR Centre, Foothills Medical Centre, Libin Cardiovascular Institute of Alberta, University of Calgary, Calgary, Alberta, Canada Chapter 38: Cardiomyopathies Tal Geva, MD Professor of Pediatrics, Harvard Medical School; Chief, Division of Noninvasive Imaging, Department of Cardiology, Children’s Hospital Boston, Boston, Massachusetts Chapter 31: Complex Congenital Heart Disease: Infant and Pediatric Patients James W. Goldfarb, PhD Assistant Professor of Biomedical Engineering, Stony Brook University, The Heart Center, Saint Francis Hospital, Roslyn, New York Chapter 34: Cardiovascular Magnetic Resonance Angiography: Carotids, Aorta, and Peripheral Vessels David Grand, MD Warren Alpert Medical School of Brown University, Department of Diagnostic Imaging, Rhode Island Hospital, Providence, Rhode Island Chapter 35: Pulmonary Artery Cardiovascular Magnetic Resonance John D. Grizzard, MD Associate Professor, Section Chief, Non-invasive Cardiovascular Imaging, Department of Radiology, Virginia Commonwealth University, Richmond, Virginia Chapter 2: Clinical Cardiovascular Magnetic Resonance Imaging Techniques Craig A. Hamilton, PhD Associate Professor, Department of Biomedical Engineering, Wake-Forest University School of Medicine, WinstonSalem, North Carolina Chapter 15: Wall Motion Stress Cardiovascular Magnetic Resonance: Ischemia, Viability, and Prognosis

Robert M. Judd, PhD Associate Professor of Medicine and Radiology; Co-Director, Duke Cardiovascular Magnetic Resonance Center, Duke University, Durham, North Carolina Chapter 2: Clinical Cardiovascular Magnetic Resonance Imaging Techniques

Sanjeet R. Hegde, MD, MRCPCH, MBBS Clinical Research Fellow, Evelina Children’s Hospital, Guy’s, St. Thomas NHS Foundation and Trust Division of Imaging Sciences, King’s College London, London, United Kingdom Chapter 44: Pediatric Interventional Cardiovascular Magnetic Resonance

Jennifer Keegan, PhD, MSc Principal Physicist and Honorary Senior Lecturer, Royal Brompton and Imperial College, London, United Kingdom Chapter 10: Use of Navigator Echoes in Cardiovascular Magnetic Resonance and Factors Affecting Their Implementation Chapter 23: Coronary Artery and Sinus Velocity and Flow

Charles B. Higgins, MD Professor of Radiology, University of California, San Francisco Medical School; UCSF Medical School, San Francisco, California Chapter 30: Cardiovascular Magnetic Resonance in Complex Congenital Heart Disease

Philip J. Kilner, MD Consultant CMR Unit Royal Brompton Hospital, London, United Kingdom Chapter 37: Valvular Heart Disease

Agnes E. Holland, MD Washington Radiology Associates, Washington, DC Chapter 34: Cardiovascular Magnetic Resonance Angiography: Carotids, Aorta, and Peripheral Vessels Peng Hu, PhD Cardiac MR Center, Beth Israel Deaconess Medical Center, Boston, Massachusetts Chapter 21: Coronary Artery and Vein Imaging: Methods W. Gregory Hundley, MD Professor, Departments of Biomedical Engineering, Internal Medicine, Molecular Medicine, and Radiology, WakeForest University School of Medicine, Winston-Salem, North Carolina Chapter 15: Wall Motion Stress Cardiovascular Magnetic Resonance: Ischemia, Viability, and Prognosis Fabien Hyafil, MD, PhD Assistant Professor, Departments of Nuclear Medicine and Cardiology, Inserm 698 “Cardiovascular Remodeling,” Bichat University Hospital, Assistance PubliqueHopitaux de Paris, University Paris, Paris, France Chapter 25: Atherosclerotic Plaque Imaging: Aorta and Carotid Hu¨seyin Ince, MD Department of Internal Medicine I, Divisions of Cardiology, Pneumology, and Intensive Care, University Hospital, University of Rostock, Rostock, Germany Chapter 33: Thoracic Aortic Disease Michael Jerosch-Herold, PhD Associate Professor of Radiology, Harvard Medical School; Director of Cardiac Imaging Physics, Department of Radiology, Brigham and Women’s Hospital, Boston, Massachusetts Chapter 4: Myocardial Perfusion Imaging Theory

Hee-Won Kim, PhD Assistant Professor, Department of Radiology, Keck School of Medicine, University of Southern California, Los Angeles, California Chapter 40: Cardiac Transplantation Raymond J. Kim, MD Associate Professor of Medicine and Radiology, Co-Director, Duke Cardiovascular Magnetic Resonance Center, Duke University, Durham, North Carolina Chapter 2: Clinical Cardiovascular Magnetic Resonance Imaging Techniques Won Yong Kim, MD, PhD Associate Professor, Department of Cardiology and MR Center, Aarhus University Hospital, Aarhus, Denmark Chapter 26: Atherosclerotic Plaque Imaging: Coronaries Stephan Kische, MD Department of Internal Medicine I, Divisions of Cardiology, Rheumatology, and Intensive Care, University Hospital, University of Rostock, Rostock, Germany Chapter 33: Thoracic Aortic Disease Kraig V. Kissinger, RT(R)(MR) Senior Technologist, Cardiac MR Center, Beth Israel Deaconess Medical Center, Boston, Massachusetts Chapter 21: Coronary Artery and Vein Imaging: Methods Grigorios Korosoglou, MD University of Heidelberg, Department of Cardiology, Heidelberg, Germany Chapter 13: High Field Cardiovascular Magnetic Resonance Christopher M. Kramer, MD Professor of Medicine and Radiology; Director, Cardiovascular Imaging Center, University of Virginia Health System, Charlottesville, Virginia Chapter 19: Acute Myocardial Infarction: Ventricular Remodeling

Cardiovascular Magnetic Resonance ix

CONTRIBUTORS

Thomas H. Hauser, MD, MMSc, MPH Assistant Professor of Medicine, Harvard Medical School; Director of Nuclear Cardiology, Beth Israel Deaconess Medical Center, Boston, Massachusetts Chapter 22: Coronary Artery Imaging: Clinical Results Chapter 32: Pulmonary Vein Imaging

CONTRIBUTORS

Lucia J. M. Kroft, MD, PhD Academic Radiologist, Department of Radiology, Leiden University Medical Center, Leiden, The Netherlands Chapter 29: Cardiovascular Magnetic Resonance of Simple Congenital Cardiovascular Defects Robert J. Lederman, MD Senior Investigator, Translational Medicine Branch, Division of Intramural Research, National Heart, Lung, and Blood Institute, National Institutes of Health, Bethesda, Maryland Chapter 43: Interventional Cardiovascular Magnetic Resonance Debaio Li, PhD Professor of Radiology, Northwestern School of Medicine; Professor of Biomedical Engineering, Northwestern School of Engineering, Chicago, Illinois Chapter 42: Magnetic Resonance Assessment of Myocardial Oxygenation Alicia M. Maceira, MD ERESA Hospital General de Castellon, Castellon, Spain Chapter 14: Assessment of Cardiac Function Chapter 28: Cardiovascular Magnetic Resonance Assessment of Right Ventricular Anatomy and Function Heiko Mahrholdt, MD Division of Cardiology, Robert-Bosch-Krankenhaus, Stuttgart, Germany Chapter 20: Myocardial Viability David Maintz, MD Professor of Radiology, Department of Clinical Radiology, University of Munster, Munster, Germany Chapter 26: Atherosclerotic Plaque Imaging: Coronaries Warren J. Manning, MD Professor of Medicine and Radiology, Harvard Medical School; Section Chief, Non-invasive Cardiac Imaging, Beth Israel Deaconess Medical Center, Boston, Massachusetts Chapter 11: Normal Cardiac Anatomy, Orientation, and Function Chapter 21: Coronary Artery and Vein Imaging: Methods Chapter 22: Coronary Artery Imaging: Clinical Results Chapter 26: Atherosclerotic Plaque Imaging: Coronaries Constantin B. Marcu, MD Cardiologist, Director of Advanced Cardiac Imaging, Hospital of Saint Raphael, Yale University, New Haven, Connecticut Chapter 24: Coronary Artery Bypass Graft Imaging and Assessment of Flow Raad H. Mohiaddin, MD Consultant CMR Unit Royal Brompton Hospital, London, United Kingdom Chapter 27: Assessment of the Biophysical Properties of the Arterial Wall Chapter 37: Valvular Heart Disease x Cardiovascular Magnetic Resonance

Eike Nagel, MD Professor of Clinical Cardiovascular Imaging; Chair of Cardiovascular Imaging, King’s College London, Division of Imaging Sciences, BHF Centre of Excellence, NIHR Biomedical Research Centre and Wellcome Trust, EPSRC Medical Engineering Centre, King’s College London, London, United Kingdom Chapter 17: Comparison of Perfusion and Wall Motion Cardiovascular Magnetic Resonance Imaging Stefan Neubauer, MD Professor of Cardiovascular Medicine, University of Oxford; Honorary Consultant Cardiologist, John Radcliffe Hospital, Oxford, United Kingdom Chapter 41: Cardiovascular Magnetic Resonance Spectroscopy Reza Nezafat, PhD Assistant Professor of Medicine, Harvard Medical School, Boston; Director, Translational Research, Cardiac MR Center, Beth Israel Deaconess Medical Center, Boston, Massachusetts Chapter 21: Coronary Artery and Vein Imaging: Methods Christoph A. Nienaber, MD Director, Department of Internal Medicine I, Divisions of Cardiology, Rheumatology, and Intensive Care, University Hospital, University of Rostock, Rostock, Germany Chapter 33: Thoracic Aortic Disease Thoralf Niendorf, PhD Professor for Experimental Ultra-High-Field MRI; Charite´—University Medicine; Director, Berlin Ultrahigh-Field MR Faculty, Max-Delbru¨ck Center for Molecular Medicine, Berlin, Germany Chapter 3: Advanced Cardiovascular Magnetic Resonance Imaging Techniques: Spiral, Radial, and Parallel Imaging Noriko Oyama, MD, PhD Assistant Professor, Department of Radiology, Hokkaido University Hospital, Sapporo, Japan Chapter 36: The Pericardium: Normal Anatomy and Spectrum of Disease Ingo Paetsch, MD Assistant Professor of Internal Medicine, Lecturer of Internal Medicine/Cardiology, Charite´ Medical School, Berlin; Director of Cardiovascular Magnetic Resonance Unit, German Heart Institute, Berlin, Germany Chapter 17: Comparison of Perfusion and Wall Motion Cardiovascular Magnetic Resonance Imaging Rajan A. G. Patel, MD Assistant Clinical Staff, Ochsner Clinic Foundation, New Orleans, Louisiana Chapter 19: Acute Myocardial Infarction: Ventricular Remodeling

Ronald M. Peshock, MD Professor of Radiology and Internal Medicine; Assistant Dean for Informatics, University of Texas Southwestern Medical Center at Dallas, Dallas, Texas Chapter 11: Normal Cardiac Anatomy, Orientation, and Function Dana C. Peters, PhD Assistant Professor, Department of Medicine, Harvard Medical School; Scientific Director of the Cardiac MR Center, Beth Israel Deaconess Medical Center, Boston, Massachusetts Chapter 1: Basic Principles of Cardiovascular Magnetic Resonance Chapter 32: Pulmonary Vein Imaging Sven Plein, MD, PhD Consultant Cardiologist, Wellcome Intermediate Clinical Fellow, Academic Unit of Cardiovascular Medicine, University of Leeds, Leeds General Infirmary, Leeds, United Kingdom Chapter 12: Comprehensive Cardiovascular Magnetic Resonance in Coronary Artery Disease Gerald M. Pohost, MD Professor, Keck School of Medicine; Professor, Viterbi School of Engineering, University of Southern California; Professor, School of Medicine, Loma Linda University, Loma Linda; Director, Cardiovascular Imaging, Westside Medical Imaging, Los Angeles, California Chapter 40: Cardiac Transplantation Andrew J. Powell, MD Associate Professor of Pediatrics, Harvard Medical School; Senior Associate in Cardiology, Department of Cardiology, Children’s Hospital Boston, Boston, Massachusetts Chapter 31: Complex Congenital Heart Disease: Infant and Pediatric Patients Chirapa Puntawangkoon, MD Research Fellow, Section on Cardiology, Department of Internal Medicine, Wake-Forest University School of Medicine, Winston-Salem, North Carolina Chapter 15: Wall Motion Stress Cardiovascular Magnetic Resonance: Ischemia, Viability, and Prognosis Amish N. Raval, MD Assistant Professor of Medicine, Division of Cardiovascular Medicine, Department of Medicine, University of Wisconsin School of Medicine and Public Health, Madison, Wisconsin Chapter 43: Interventional Cardiovascular Magnetic Resonance

Reza S. Razavi, MD, MBBs Professor of Paediatric Cardiovascular Science, Division of Imaging Sciences, King’s College London; Consultant Cardiologist, Guy’s and St. Thomas NHS Foundation Trust, London, United Kingdom Chapter 44: Pediatric Interventional Cardiovascular Magnetic Resonance Wolfgang G. Rehwald, PhD Senior Scientist, Cardiovascular MR Research and Development, Siemens Healthcare, Siemens Medical Solutions USA, Inc., Customer Solutions Group, Chicago, Illinois Chapter 2: Clinical Cardiovascular Magnetic Resonance Imaging Techniques Arno A. W. Roest, MD, PhD Fellow, Pediatric Cardiology, Department of Pediatric Cardiology, Leiden University Medical Center, Leiden, The Netherlands Chapter 29: Cardiovascular Magnetic Resonance of Simple Congenital Cardiovascular Defects James H. F. Rudd, PhD, MRCP HEFCE Senior Lecturer, Division of Cardiovascular Medicine, University of Cambridge; Honorary Consultant Cardiologist, Addenbrooke’s Hospital, Cambridge, United Kingdom Chapter 25: Atherosclerotic Plaque Imaging: Aorta and Carotid Juerg Schwitter, MD Associate Professor, University Hospital Lausanne; Director, CMR Center of the University Hospital Lausanne— CHUV, Cardiology, University Hospital Lausanne, Lausanne, Switzerland Chapter 16: Stress Cardiovascular Magnetic Resonance: Myocardial Perfusion Udo P. Sechtem, MD Chairman, Department of Cardiology, Robert-BoschKrankenhaus, Stuttgart; Associate Professor of Medicine and Cardiology, University of Tu¨bingen, Tu¨bingen, Germany Chapter 20: Myocardial Viability Burkhard Sievers, MD Department of Medicine/Cardiology, Heart Center Dresden, University Hospital, Dresden, Germany Chapter 2: Clinical Cardiovascular Magnetic Resonance Imaging Techniques Daniel K. Sodickson, MD, PhD Vice-Chair for Research, Department of Radiology; Director, Center for Biomedical Imaging; Associate Professor of Radiology, Physiology, and Neuroscience, New York University Langone Medical Center, New York, New York Chapter 3: Advanced Cardiovascular Magnetic Resonance Imaging Techniques: Spiral, Radial, and Parallel Imaging Cardiovascular Magnetic Resonance xi

CONTRIBUTORS

Dudley J. Pennell, MD Professor of Cardiology, National Heart and Lung Institute, Imperial College; Director, Cardiovascular MR Unit, Royal Brompton Hospital, London, United Kingdom Chapter 14: Assessment of Cardiac Function Chapter 23: Coronary Artery and Sinus Velocity and Flow Chapter 28: Cardiovascular Magnetic Resonance Assessment of Right Ventricular Anatomy and Function

CONTRIBUTORS

Elmar Spuentrup, MD Department of Radiology, University Hospital, University of Cologne, Cologne, Germany Chapter 26: Atherosclerotic Plaque Imaging: Coronaries Matthias Stuber, PhD Professor, University of Lausanne, Center for Biomedical Imaging (CIBM); Director, CIBM, Centre Hospitalier Universitaire Vaudois, Lausanne, Switzerland Chapter 5: Cardiovascular Magnetic Resonance Tagging Assessment of Left Ventricular Diastolic Function Chapter 13: High Field Cardiovascular Magnetic Resonance Nicholas R. Teman, MD House Officer, Department of General Surgery, University of Michigan Health System, Ann Arbor, Michigan Chapter 40: Cardiac Transplantation Ernst E. van der Wall, MD Professor of Cardiology; Head, Department of Cardiology, Leiden University Medical Center, Leiden, The Netherlands Chapter 8: Special Considerations for Cardiovascular Magnetic Resonance: Safety, Electrocardiographic Setup, Monitoring, and Contraindications Albert C. van Rossum, MD, PhD Professor of Cardiology; Chairman of the Department of Cardiology, VU University Medical Center, Amsterdam, The Netherlands Chapter 24: Coronary Artery Bypass Graft Imaging and Assessment of Flow

xii Cardiovascular Magnetic Resonance

Anja Wagner, MD Department of Cardiology, Hahnemann University Hospital, Drexel University College of Medicine, Philadelphia, Pennsylvania Chapter 2: Clinical Cardiovascular Magnetic Resonance Imaging Techniques Thomas F. Walsh, MD Senior Assistant Resident, Department of Internal Medicine, Wake-Forest University School of Medicine, Winston-Salem, North Carolina Chapter 15: Wall Motion Stress Cardiovascular Magnetic Resonance: Ischemia, Viability, and Prognosis Norbert Wilke, MD Clinical Associate Professor of Medicine (Cardiology); Clinical Associate Professor of Radiology, Cardiovascular MR and CT Services, University of Florida, Gainesville, Florida Chapter 4: Myocardial Perfusion Imaging Theory Rolf Wyttenbach, MD, PhD Chairman, Department of Diagnostic and Interventional Radiology, Ospedale Regionale di Bellinzone e Valli (EOC), Bellinzona, Switzerland Chapter 30: Cardiovascular Magnetic Resonance in Complex Congenital Heart Disease Susan B. Yeon, MD, JD Assistant Professor of Medicine, Harvard Medical School, Cardiovascular Division, Beth Israel Deaconess, Boston, Massachusetts Chapter 26: Atherosclerotic Plaque Imaging: Coronaries Chapter 36: The Pericardium: Normal Anatomy and Spectrum of Disease

Since William Harvey’s discovery of the circulation, the assessment of cardiac structure and function has been a major goal of physicians responsible for the management of patients with heart disease. The first techniques were quite primitive—palpation of the arterial pulse, observation and direct auscultation of the precordium—followed by Rene´ Lae¨nnec’s development of the stethoscope, the first tool to amplify the senses. A paradigm shift occurred with the discovery of the x-ray by Wilhelm Roentgen at the end of the 19th century, allowing the heart to be imaged for the first time. The 20th century can be considered the century of cardiac imaging, beginning with angiography and arteriography, both requiring the injection of contrast material through a catheter. The development and progressive refinement of noninvasive approaches—echocardiography, radionuclide imaging, and computed tomography— advanced cardiology enormously. Cardiovascular magnetic resonance imaging, however, may finally fulfill the dreams of cardiologists. The technique offers a veritable treasure trove of information about the heart and circulation. It not only provides superior displays of chamber structure but allows the identification of myocardial necrosis, infiltration, and fibrosis. Both valvular performance and ventricular function can be measured. Magnetic resonance angiography allows assessment of vascular function and tissue perfusion. Recently, both the walls and luminae of the major coronary arteries have been visualized. When

combined with the ability to determine myocardial viability and to measure cardiac high-energy phosphate stores noninvasively by magnetic resonance spectroscopy, these techniques seem certain to improve enormously the care of patients with ischemic heart disease. It is no longer fanciful to think that coronary arteriography, now performed on more than 1.5 million patients in the United States each year, will become obsolete. Just as it became necessary for physicians entrusted with the care of cardiac patients to become familiar with cardiac roentgenography at the beginning of the 20th century, it now has become vital for their successors to do the same with cardiac magnetic resonance imaging. Warren J. Manning and Dudley J. Pennell and their talented contributors to Cardiovascular Magnetic Resonance have taken a significant step toward enabling clinicians to accomplish this. They deserve the appreciation of cardiovascular specialists of all types for producing this detailed, authoritative, yet eminently readable book. A century ago, great strides were being made in cardiac roentgenography and the future promise was enormous. The situation is similar with cardiac magnetic resonance imaging today. This field is advancing at such a breathtaking speed that I can’t wait for the second edition. Eugene Braunwald, MD Boston, Massachusetts

Cardiovascular Magnetic Resonance xiii

FOREWORD TO THE FIRST EDITION

Foreword to the First Edition

PREFACE

Preface Cardiovascular magnetic resonance (CMR) is a medical imaging field that excites great interest because it combines superb image quality with new techniques for probing the cardiovascular system in novel ways. What surprises is the versatility of the technology: blood flow, angiography, assessment of atherosclerosis, myocardial perfusion, focal necrosis, oxygen saturation, technology, and chemical composition are among the measurements that are being refined for clinical use, in addition to the well-known “gold standard” capabilities of CMR in defining anatomy and ventricular function. Such potential comes at a price, however, as this technology is not quickly learned, and highquality clinical practice needs experience. Professional didactic and clinical training is required for all newcomers to the field and to maintain cutting-edge competency.

The aim of this book is to provide instruction in the current clinical practice of CMR, while also highlighting areas of clinical potential, which are presently in varying stages of development. If we succeed in drawing new investigators and clinicians to enter the field or illuminating new areas for those already involved, then we will have achieved our objective—the healthy growth of competent and motivated practitioners in CMR for the benefit of clinical science, patient care, and the advancement of the field. The reader should be forewarned that CMR is constantly developing, and no text can include all of the most recent developments; however, the foundations provided will serve the reader for many years to come. The future of CMR is bright! Join us in the adventure! WJM, DJP

Cardiovascular Magnetic Resonance xv

ACKNOWLEDGMENTS

Acknowledgments It takes a town to create a book. Like any large endeavor, this text is a success because of the effort of many individuals to whom we owe thanks. These include the outstanding contributions of our primary authors and their collaborators; Rebecca Schmidt Gaertner, our editor at Elsevier Health Sciences; Linda R. Garber, our production editor; our office assistants, Iris Wasserman at the Beth Israel Deaconess Medical Center and Fei Wang at the Royal

Brompton Hospital; the multitude of CMR mentors, trainees, and colleagues who have stimulated and educated us over the past two decades; and most of all to our families, for allowing us the time to pursue this endeavor amongst the myriad of other activities that occupy our daily lives. To all, we express our thanks and high level of appreciation and gratitude. WJM, DJP

Cardiovascular Magnetic Resonance xvii

2D 3D

two-dimensional three-dimensional

A2C A4C ACE-I ACS ADP AMI ARVC ASD AT 1R AT 2R ATP a.u. AUC AV AVM

apical two chamber apical four chamber angiotensin-converting enzyme inhibitor acute coronary syndrome adenosine diphosphate acute myocardial infarction arrhythmogic right ventricular cardiomyopathy atrial septal defect angiotensin 1 receptor angiotensin 2 receptor adenosine triphosphate arbitrary units area under the curve atrioventricular arteriovenous malformation

BOLD

blood oxygen level dependent

C CABG CAD CE CE-CTA

carbon coronary artery bypass graft coronary artery disease contrast enhanced contrast enhanced computed tomography angiography CE-MRA contrast enhanced magnetic resonance angiography CFR coronary flow reserve CHD congenital heart disease CMP cardiomyopathy CMR cardiovascular magnetic resonance CMRS cardiovascular magnetic resonance spectroscopy CO cardiac output CNR contrast-to-noise ratio CP creatine phosphate CS circumferential shortening CSPAMM Complementary SPAtial Modulation of Magnetization CSI chemical shift imaging CT computed tomography DC DCM DCMR DORV DPG DSE DTPA

direct current dilated cardiomyopathy dobutamine stress cardiovascular magnetic resonance double outlet right ventricle diphosphoglycerate dobutamine stress echocardiography diethylenetriamine pentaacetic acid

Dy

dysprosium

ECF ECG ECM EDV EF EPI ESV

extracellular fluid electrocardiogram extracellular matrix end-diastolic volume ejection fraction echo planar imaging end-systolic volume

FFR FID FSE

fractional flow reserve free induction decay fast spin echo

Gd gadolinium Gd-DTPA gadolinium diethylenetriamine pentaacetic acid GRE gradient recalled echo H HCM HDL He HLA Hz

hydrogen/proton hypertrophic cardiomyopathy high density lipoprotein helium horizontal long axis Herz

ICD IMA IMH iNOS IR IV

implantable cardiac defibrillator internal mammary artery intramural hematoma inducible nitric oxide synthetase inversion recovery intravenous

LA LAD LCX LDL LGE LM LPA LV LVEF LVOT

left atrium/atrial left anterior descending coronary artery left circumflex coronary artery low density lipoprotein late gadolinium enhancement left main coronary artery left pulmonary artery left ventricle/ventricular left ventricular ejection fraction left ventricular outflow tract

MI MIP Mn MO MPA MPHRR MR MRA MRI MRS

myocardial infarction maximum intensity projection manganese microvascular obstruction main pulmonary artery maximum predicted heart rate response magnetic resonance magnetic resonance angiography magnetic resonance imaging magnetic resonance spectroscopy Cardiovascular Magnetic Resonance xxiii

COMMON ABBREVIATIONS USED IN THE TEXT

Common Abbreviations Used in the Text

COMMON ABBREVIATIONS USED IN THE TEXT

MTT MVO2

mean transit time myocardial oxygen consumption

NOS NSF NYHA

nitric oxide synthetase nephrogenic systemic fibrosis New York Heart Association

OCT

optical coherence tomography

P PA PCI PCr PDA PDE PDW PET PLA PSA PTCA PWV

phosphorus pulmonary artery percutaneous coronary intervention phosphocreatine patent ductus arteriosus phosphodiesters proton density weighted positron emission tomography parasternal long axis parasternal short axis percutaneous transluminal coronary intervention pulse wave velocity

QCA

quantitative cororonary angiography

RA RCA RCM RF ROC ROI RPA RPP RV RVEF RVOT

right atrium/atrial right coronary artery restrictive cardiomyopathy radiofrequency receiver operator characteristic region of interest right pulmonary artery rate pressure product right ventricle/ventricular right ventricular ejection fraction right ventricular outflow tract

xxiv Cardiovascular Magnetic Resonance

SENSE SMASH SNR SPAMM SPECT SPIO SSFP

SENSitivity Encoding SiMultaneous Acquisition of Spatial Harmonics signal-to-noise ratio SPAtial Modulation of Magnetization single photon emission tomography small particle iron oxide steady state free precession

T T1 T1w T2 T2w TD TE TEE TGA TI TMR TOF TR TTC TTE

Tesla longitudinal relaxation rate T1 weighted transverse relaxation rate T2 weighted delay time echo time transesophageal echocardiography transposition of the great arteries inversion time transmyocardial laser revascularization tetraology of Fallot repetition time 2,3,5-triphyltetrazolium chloride transthoracic echocardiography

USPIO

ultrasmall particle iron oxide

VA VENC VLA VNC VSD VSMC

ventriculoarterial velocity encoding range vertical long axis ventricular noncompaction ventricular septal defect vascular smooth muscle cells

Xe XFR

xenon X-ray fluoroscopy

Basic Principles of Cardiovascular Magnetic Resonance Robert S. Balaban and Dana C. Peters

This introduction to the basic principles of cardiovascular magnetic resonance (CMR) describes the concepts of T1, T2, and T2* and image formation, and describes some common CMR pulse sequences and parameters. These basic ideas often require some time and rereading to fully appreciate, but are necessary to understand the fundamental properties that determine image properties. The human body is composed mostly of water and also a lot of fat. The water (H20) and fat contain many hydrogen atoms. Hydrogen atoms in turn are made up of a proton (1H, the hydrogen nucleus) and an electron. Hydrogen protons are in very high concentration in the body, roughly 100 molar. Because of this abundance, the nuclear magnetic resonance signal can be used to create a distribution map, or image, through magnetic resonance imaging (MRI). MRI depends on the detection of the intrinsic angular momentum, or spin, of protons that is a basic property of matter. Hydrogen protons have spin, and all nuclei with spin interact with magnetic fields. In the absence of a magnetic field, the hydrogen spins are randomly oriented (Fig. 1-1A). However, if placed in a large magnetic field (called B0), the water spins (i.e., the 1H nuclear magnets) partly align with this applied magnetic field, much like iron filings (see Fig. 1-1B), with larger magnetic fields causing greater alignment of the spins. In contrast to the iron filings analogy, however, the interaction of the angular momentum (spin) of the hydrogen protons with the B0 field results in a rotation of the spin angular momentum around the axis of the magnetic field (see Fig. 1-1C). This is why the hydrogen nuclei in MRI are also referred to as spins: they spin, or precess, around the B0 field. The frequency of precession (v) is an important fundamental property of the spins in a magnetic field and is defined by the Larmor equation: n ¼ g B=ð2pÞ

(1)

where v is the precessional frequency in cycles/sec or Hertz (Hz), g is called the gyromagnetic ratio and is related to the mass and charge of the water proton, g/2p ¼ 42.58 106 Hz/Tesla (T; 1 Tesla ¼ 10,000 gauss) for 1H, and B is the applied magnetic field. The precessional frequency for water protons at 1.5 T (a common CMR field strength) is v ¼ 63.87 106 Hz, or roughly 64 MHz (about the frequency of an FM radio station and one of the reasons why the CMR environment must be shielded from FM radio waves).

Molarity of 1H can be estimated as approximately (2 moles hydrogen/mole H20) (1mole H20/18 g tissue). 1000 g/L (density of the body) 100 mole/L.

The Larmor equation (Equation 1) is the basis of CMR imaging. To determine the location of different magnetic spins (or 1 H), the magnetic field, B, is made to vary linearly with position using specialized coils of wire (called gradient coils) inside the magnet. This results in a precessional frequency (v) that is also a linear function of position in the scanner. By measuring the number of spins precessing at each frequency, a magnetic resonance (MR) image is created,1 as discussed later.

DETECTION OF THE MAGNETIC RESONANCE SIGNAL Alignment with the Main Magnetic Field How can the MRI signal from 1H spins be detected to image the sample? Remember, when a sample (i.e., a patient!) is placed in the main magnetic field (e.g., the bore of the 1.5 T magnet), the hydrogen spins align with that field (see Fig. 1-1B) and begin to precess at the Larmor frequency (see Fig. 1-1C). The direction of the main magnetic field B0 defines the Z-axis direction, also called the longitudinal direction. Because the spins partially align with the B0 field, they create a net magnetic field along the Z-axis (oriented parallel with B0). However, the spins precess, or rotate, around the Z-axis in a random, incoherent fashion, resulting in no net magnetization in the X-Y plane (also called the transverse plane; see Fig. 1-1C). An MR signal is detected by placing a coil (loops of wire) near the sample to detect any spin precession in the X-Y (transverse) plane. This coil is called a receiver coil because it receives signal. In CMR, the receiver coil is usually an array of four or more (as many as 32 to 128)2 coils placed around the chest. However, a receiver coil placed to detect the precession of spins in the XY plane will not detect any signal because the spins are in different positions, or phase, thereby canceling each other (see Fig. 1-1C). Therefore, radiofrequency excitation is necessary.

Radiofrequency Excitation and Magnet Strength To create an MR signal, the water (1H) spins must be rotating in a coherent manner in the X-Y plane. To accomplish this task, another (less powerful) magnetic field (called B1), Cardiovascular Magnetic Resonance 3

1 BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

CHAPTER 1

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

No net 1H magnetic field

Net 1H magnetic field

Random phase of individual spins with net 1H magnetic field

Z

Z

X X X Y Y Y B0 = 0

A

B0

B

B0

C

Figure 1-1 Orientation of nuclear spins. A, With no magnetic field, the spins are oriented randomly, producing no net magnetization. B, The spins orient with an applied magnetic field (B0), producing a net magnetic field aligned along B0. C, The magnetized spins rotate around the B0 field (Z-axis) in a random way, resulting in no net magnetization in the X-Y plane.

at 3.0 T3 has been demonstrated for coronary artery imaging,4,5 late gadolinium enhancement, perfusion,6 and function.7 The excitation field (B1) has a strength of only a few gauss. In addition, the B1 field is only applied transiently (for milliseconds), long enough to temporarily deflect the spins into a plane perpendicular to B0. This effect on the net magnetic field is illustrated in Figure 1-2A. A B1 pulse that drives the magnetization completely into the transverse plane (X-Y plane) is called a 90 pulse, referring to the angle the net 1H magnetization moves relative to the Z-axis. However, in practice, the B1 pulse can be used to flip the magnetization by any angle from 0 to 180 . The process flips the spin magnetization from the longitudinal direction into the transverse plane. Once the B1 field is turned off,

with its strength oscillating at the Larmor frequency, is applied perpendicular to the main magnetic field (Fig. 1-2A). This process is called radiofrequency (RF) excitation because it uses a B1 field rotating at a high (radio wave) frequency. The frequency of the perpendicular B1 field has to precisely match the Larmor frequency of the water 1H to rotate the spins in the presence of the stronger B0 field. This concept is analogous to exciting a piano string with a tuning fork. Only the string with the same resonant frequency as the tuning fork will efficiently absorb the energy from the fork and resonate. The water (1H) spins will rotate around the B1 field into the X-Y plane. The B0 field has a magnitude in the range of 1.0 to 7.0 T. For CMR, a B0 of 1.5 T is most commonly used; however, cardiac imaging

Now coherent spins rotate around Z (B0)

Z Final magnetization

B0

Z X

B0

Initial magnetization

Receiver coil

Initial magnetization Y

Pulsed B1 field oscillating at Larmor frequency 90° X

B1

Voltage

Final magnetization

A

Y RF excitation coil

Time

B

Free induction decay

Figure 1-2 Effects of a perpendicular B1 magnetic field. A, The B1 field oscillating at the Larmor frequency of the spins is absorbed by the spins and causes a rotation around the B1 field axis (Y-axis). B, Once the B1 field is removed, the spins continue to rotate around the Z-axis at the Larmor frequency, but the signal decays (in the X-Y plane) as equilibrium is reestablished. The result is a coherent oscillating magnetic field that is detected with the coil as the free induction decay. The free induction decay is a decaying sinusoidal voltage signal with a frequency equal to the Larmor frequency of the spins. RF, radiofrequency. 4

Cardiovascular Magnetic Resonance

T2 Relaxation and Spin Phase The other form of relaxation, T2 relaxation, is that of the randomization of the phase of the spins. The phase indicates the direction of the spin on the transverse X-Y plane. The phase (f) of a spin in the transverse plane depends on its initial phase, f0, the precessional frequency, and the time, t, which it has spent in the transverse plane. f ¼ f0 þ 2pnt

Because precessional frequency depends on the magnetic field, which changes slightly with location, time, and even molecular environment, each spin on the transverse plane has a different frequency and thus phase, and this phase difference increases with time (Equation 3). Figure 1-3 defines the phase (f) of the spin magnetic field vector as the direction of the transverse magnetization relative to the X-axis. The phase of a spin depends on its frequency and its history of frequencies. In Figure 1-3B, a phase diagram is used to show the changing phase relationships for three spins in a sample that has different magnetic fields within it. The magnetic field strength increases from light to dark. The magnetic field strength always varies within the sample because of imperfections in the magnetic field and sometimes intentionally through application of magnetic field gradients. A phase diagram is a stroboscopic image because it is arbitrarily referenced to one spin’s vector, in this case, a spin in the center region for Figure 1-3B. This stroboscopic effect is analogous to taking a flash picture every time the rotating middle vector reaches the 0 position. Using this approach, the relative position of the spin vectors, or phase, can be easily seen. Immediately after a 90 pulse, all of the spins have the same phase; however, with time, the different spin frequencies cause their relative phase to

T1 Relaxation T1 relaxation, also known as spin-lattice relaxation, is the release of energy to the environment, or lattice, that results in the reestablishment of the magnetization along the Z-axis. Thus, T1 is the time constant by which the longitudinal (or Z) magnetization relaxes to its equilibrium value, M0. After a 90 RF pulse, the Z magnetization (Mz[t]) is initially 0, but regrows with a relaxation time T1 to its equilibrium value (M0). MzðtÞ ¼ M0 ð1 et=T1 Þ

(3)

(2)

where t is time after the 90 pulse. Each tissue type has a unique T1. While T1 measurements differ slightly, a recent report found that for myocardium, the T1 is 1100 msec

Low Field Center Field High Field

X-Y plane phase diagram Y 90°

L

L C

X

Y φ

φ 180°

X

C

H

0° Immediately after 90⬚ pulse, t = 0.

Z

A

Spins out of phase

270°

B

Time t later

Still later

Time

Figure 1-3 Phase diagram. A, A phase diagram is a view of the magnetization vectors projected into the X-Y plane. B, The magnetic field of a sample varies as shown. Spins located in the low field (L), central field (C), and high field (H) regions are followed in time after a 90 pulse using the phase diagram. The phase diagram presents the phases of the spins relative to C, which is always shown with 0 phase. Note that the L vector lags behind the C vector, whereas the H vector is ahead because it has a higher phase velocity. This difference is caused by the higher frequency of the spins in the higher magnetic field. This process is called dephasing. Cardiovascular Magnetic Resonance 5

1 BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

and 1200 msec at 1.5 T and 3.0 T, respectively, and for blood, the T1 is 1600 msec at both 1.5 T and 3.0 T.8

the only field causing a rotation is B0. Now the net magnetization of the water proton spins is in the transverse plane and coherent for detection. As the spins rotate around the B0 field axis and return to the B0 orientation, an oscillating X-Y magnetic field can be detected with the receiver coil poised as shown in Figure 1-2B. Figure 1-2B shows the signal detected by the receiver coil after placing the spins in the main magnetic field and then flipping them into the transverse (X-Y) plane. The signal oscillates at the Larmor frequency, decays in amplitude with time, and is called a free induction decay (FID). The decay in net magnitude in the transverse plane is known as relaxation. The decay occurs because the spins, having absorbed energy from the transient B1 field, which is now turned “off,” now return to their original state in equilibrium with the B0 field, with spins partially aligned along the Z-axis (see Fig. 1-1C). This occurs via two extremely important processes, called T1 and T2 relaxation. Nature abhors order and seeks to minimize the energy in the system and return to equilibrium with the main magnetic field, B0, through T1 and T2 relaxation. T1 and T2 relaxation forms the foundation of MRI.

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

change. The higher B0 field spins (dashed arrows), are advancing their phase faster than central B0 field spins (solid arrows), whereas the spins (dotted arrows) at low field are lagging behind both. By watching the time development of this process, one can appreciate that the phase and the frequency are related (Equation 3). High-frequency spins change phase rapidly, whereas slower-frequency spins change phase more slowly. However, the initial starting points of each of these signals are independent of frequency. T2 relaxation can be understood as the time for the spins to dephase. After the initial excitation using a B1 pulse, all of the spins on the transverse plane have identical phase, but then they begin to dephase quickly. This dephasing is described by a time constant, T2. The dephasing causes the bulk transverse magnetization to decay and approach 0 because of incoherent phase. Because the magnetization decays, the transverse signal, S(t), also decays in a simple exponential way, with a time constant of T2:

decrease in the net transverse magnetization. Some of this signal decay can be reversed because some of it is caused by static “off-resonance.” T2* is an important process in the heart because both the lung cavity and the deoxygenated blood in the right side of the heart contribute to B0 inhomogeneity around the heart. The combined effect of all dephasing processes, including the molecular spin-spin interactions (T2), and static off-resonance, is called T2*. The myocardial T2 relaxation rate is approximately 50 msec at 1.5 T,9,10 whereas myocardial T2* is approximately 30 msec at 1.5 T.11 T2* is always less than T2 for all tissues because it includes the effects of T2 relaxation and offresonance effects. Later we will describe how to minimize or use these effects to make in vivo measurements.

SðtÞ ¼ S0 et=T2

In tissue, T2* causes very rapid magnetic relaxation of water protons, compared with T2 relaxation. Thus, the decay rate of the FID is actually a measure of T2*. Because T2* is a more rapid process, it limits the time that we can detect the MR signal. It is possible to circumvent and reverse some of the T2* dephasing because some of it is caused by fixed B0 field inhomogeneity. This reversal is accomplished using a B1 field-generated echo, outlined in Figure 1-4. After the 90 pulse, dephasing occurs. Another B1 field is applied to rotate the magnetization 180 around the X-axis. This 180 flip is achieved by applying the B1 field for twice as long or with twice the power as the 90 pulse. This 180 flip causes the spins to rotate into the opposite sector of the transverse plane. On this side of the plane, the spins now drift together (“rephase”), taking the same amount of time to rephase that they were allowed to dephase. This is analogous to two people standing backto-back in a field and then walking apart (90 pulse) at some fixed rate (but different rates for each person). At some later time, they reverse (180 pulse) their direction and now walk toward each other, each at the same “individual” pace. The time it takes them to reach each other

Spin Echo Imaging

(4)

T2 relaxation is also called spin-spin, or transverse, relaxation. It is called spin-spin relaxation because the mechanism of the relaxation process is through the interaction of spins in the sample with each other at or below the Larmor frequency, making this process dependent on the microscopic motions within the sample. To understand many of the issues in CMR, we need to further understand T2. In general, two distinct processes contribute to the dephasing of spins in the heart. The first mechanism is the true spin-spin interaction (T2) that is unavoidable, irreversible, and dependent on the molecular interactions within the sample. The second process is called T2* (“T2 star”) decay and includes T2 relaxation and also relaxation as a result of the inhomogeneity of the main magnetic field (B0) within the sample. As illustrated in Figure 1-3B, if the B0 field is not homogeneous throughout the sample, then the frequency of the spins in different regions will vary (Equation 1). These spins are called offresonance spins. This results in a randomization of the spin phases as they rotate at different frequencies, and a

Dephasing 90°

X

Rephasing

Spin echo

180°

X Y

t=0

X Y

X Y

t=τ (before 180°)

t=τ (after 180°)

Y

t = 2τ

Time Figure 1-4 Effects of a 180 B1 pulse after a 90 pulse. The spins are first excited by a 90 pulse, as in Figure 1-2. After some time in which dephasing occurs (as in Fig. 1-3B), a 180 pulse is applied. The 180 pulse, by flipping the spins 180 around the X-axis, exchanges the phase position of the slower and faster spins, so that the faster spins are now lagging behind the slower spins. The faster spins begin to catch up, and after a time, t, equal to the time of dephasing, the spins are all aligned in the same direction (have the same phase). This process is called rephasing, or refocusing, to form a spin echo. 6

Cardiovascular Magnetic Resonance

Cardiovascular Magnetic Resonance Imaging To create a CMR image, the MR signal intensity from the sample (i.e., patient) must be determined in three dimensions (X, Y, and Z). Thus, for a single image, it is necessary to collect information on X, Y, and Z positions and signal amplitude. The MR signal shown in Figure 1-2B has frequency, amplitude, and phase. These can be used to determine some spatial or amplitude information. However, there is not enough information to create a two- or three-

MAGNETIC RESONANCE IMAGING BLOCK DIAGRAM Time

Slice select

Z

Phase encode

Y

Refocus gradient/RF

Frequency encode, readout

X

Figure 1-5 General imaging scheme for collecting a simple cardiovascular magnetic resonance image. RF, radiofrequency.

dimensional image from a single FID. This is a major limitation of MRI that results in a relatively slow image acquisition rate compared with many other modalities. The simplest imaging experiment can be divided into four stages, as shown in the block diagram in Figure 1-5: (1) slice selection, (2) phase encoding, (3) refocusing echo (for spin echo sequences), and (4) frequency encoding (readout). Each stage is used to encode the MR data with information on the position and amplitude of the water proton (1H) signal. By convention, Z position information is encoded with the slice selection step, X position information is determined in the frequency encoding/readout step, and Y position information is obtained with the phase encoding step. In practice, the frequency encoding (X), phase encoding (Y), and slice selection (Z) directions are rotated in any appropriate direction. MR is tomographic and can create an image of the body along any plane. Indeed, oblique imaging, slicing through a tissue at 45 or any other angle, is possible and common. This property is almost a unique attribute of MRI. For example, computed tomography can provide images in any plane only after reformatting, but always collects images in the axial plane.

Gradients Spatial encoding in MRI is performed using “gradients”: slice select gradients, phase encoding gradients, and frequency encoding gradients. Gradient coils are special coils within the magnet that modify the main magnetic field (Bz), causing it to change slightly in space: Bz ðx; y; zÞ ¼ B0 þ Gx x þ Gy y þ Gz z

Table 1-1 T1 and T2 Values at 1.5 T Tissue

T1 (msec)

T2 (msec)

Myocardium Arterial blood Fat Skeletal muscle Lung

1100 1600 260 880 820

50 250 110 45 140

(5)

where Gx, Gy, and Gz are called gradient strengths and x, y, and z are the spatial coordinates within the magnet. Using gradients, the main magnetic field, with which the spins align and around which they precess, varies with spatial position inside the scanner. Remember, this magnetic field points in the Z direction, and the gradients do not change this direction, but they do increase or decrease its strength spatially. These spatially varying magnetic fields are called Cardiovascular Magnetic Resonance 7

1 BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

will be the same amount of time as they walked apart. This is the nature of the echo that effectively recaptures all of the coherent transverse magnetization that had been destroyed by the B0 field inhomogeneity in the sample. Note that the echo has both a rephasing and a dephasing period. The spins continue to dephase after reaching the coherent echo as a result of ongoing field inhomogeneity, just as the people in the field would walk past each other and continue to “dephase.” A subsequent 180 B1 pulse could be used to refocus these spins again and again. This method of continually refocusing the magnetization with 180 pulses is called fast spin echo (or turbo spin echo), and is used in cardiac imaging. The number of repeated 180 pulses used to refocus the spins is limited because a 180 pulse will not recover the T2 relaxation processes occurring on the molecular level. However, the magnetization dephases according to T2 relaxation, which is much slower than T2* relaxation. These basic relaxation processes, T1, T2, and T2*, are key in generating image contrast in MRI as well as determining the optimal image sequence for gathering information on cardiovascular anatomy, function, and physiology. Table 1-1 provides values of T1 and T2 for some important tissues. What is important is that T1, T2, and T2* vary for different tissues, thereby providing “contrast” in a CMR image. For example, myocardial relaxation properties change with edema, which is present in different disease states.12,13 Overall, the T1 and T2 values are more prolonged with increasing water content, as with edema. Myocardial remodeling, or myocyte replacement with connective tissue, also changes the water relaxation properties because the nature of the macromolecules in contact with water is critical for these relaxation processes. Finally, most exogenous, intravenously injected CMR contrast agents act by shortening either T114-16 or T2*17,18 of the water spins. By appropriately modifying the imaging sequences, these changes in relaxation properties caused by pathology or exogenous contrast agent can be highlighted in the MRI of the heart. How this is accomplished will be described below.

fðx; y; z; tÞ ¼ g Bz ðx; y; zÞ t ¼ g ðB0 þ Gx x þ Gy y þ Gz zÞ t

(7)

Selective Excitation: Position in Z During the slice select stage, RF excitation is performed. One of the spatial coordinates (Z) is determined by only exciting (with the B1) Z magnetization in a selected slice of the sample. The slice selection process is illustrated in Figure 1-6. A linear magnetic field gradient, Gz, is applied to the sample along the slice direction (the Z-axis). The magnetic field gradient causes the spins along this axis to have slightly different frequencies, as defined by the Larmor equation: vðzÞ ¼ ðB0 þ Gz zÞ ðg=2pÞ

Gradient in Z

Excited slice

si

tio

n

+1000 –1000

po

Frequency or B field

–4 +4 Z position

X

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

gradients because they create a linearly spatially varying magnetic field (magnetic field gradient). Of course, because the magnetic field varies with position, the precessional frequency n(x,y,z), and the phase f(x,y,z,t) of the spins also vary with position (Equations 1 and 3). g g n¼ Bz ðx; y; zÞ ¼ ðB0 þ Gx x þ Gy y þ Gz zÞ (6) 2p 2p

Figure 1-6 Slice selection gradient. A slice selection gradient is applied on the Z-axis, causing the B field to vary from less than to greater than B0. This leads to a linear distribution of frequencies as a function of position along the gradient (Z direction). A frequency selective (e.g., 1000 Hz) B1 pulse is used to excite only spins within a frequency band, resulting in selection of a slice (e.g., 4 mm) that is placed in the transverse (X-Y) plane.

The B1 field is used to excite the spins, but only within a range of precessional frequencies. If only a selected band of frequencies is excited (e.g., -1000 to 1000 Hz), then a slice of spins with precessional frequencies in that range (e.g., an 8-mm-thick slice) is put in the transverse plane, not affecting the rest of the sample. The “slice thickness” (here, 8 mm) can be freely chosen by adjusting the Z gradient. Typical slices range from thin (e.g., 0.5 mm) to thick (e.g., 10.0 cm) slabs. The slice position can also be freely chosen. Furthermore, if no Z gradient is applied during the slice selection, this is equivalent to an infinite slab thickness, so that all of the spins in the body will be tipped into the transverse plane. This is called a nonselective RF pulse, and is commonly used in CMR for preparation pulses (see Figs. 1-17 and 1-18). How can the B1 field be used to excite only spins precessing within a range of frequencies? To excite a bandwidth of frequencies, the Z gradient is turned on, and the B1 strength is modulated in time: B1ðtÞ ¼ B10 sinðp t=TÞ=ðp t=TÞ

This shape, sin(t)/t, is called a “sinc” function, and is plotted in Figure 1-7. This B1(t) excites a well-defined band of frequencies of width, Dn ¼ 1/T. This is shown in Figure 1-7. Here, a Fourier transform is used to show how the spins interpret the sin(t)/t pulse in terms of frequency. Modifications to the sinc pulse are possible to optimize the performance and slice definition of this approach.19,20 The slice selection process is summarized in the imaging “pulse sequence diagram” shown in Figure 1-8. A pulse sequence diagram presents all of the gradients, the B1 field excitation, and the data collection event on a timeline. In Figure 1-8, the slice selection gradient is first increased to a constant level by a ramp. The gradient is maintained while the frequency-selective (i.e., slice-selective) B1 field is applied; note its sinc shape. Subsequently, the slice gradient is ramped down and the rephasing gradient is applied to remove the phase introduced by the slice select gradient. After the slice select excitation, only a slice of the sample has been placed in the transverse plane and all of the spins are in phase. The rephasing gradient shown in Figure 1-8 is performed so that all of the spins are “in phase” after the slice selection process. During the 90 pulse, the spins will dephase, depending on where they are in the slice (similar to Fig. 1-3B), because a magnetic field gradient is applied. Thus, another magnetic field gradient of opposite magnitude and half the

FREQUENCY SELECTIVE B1 PULSE Sinc function

B1 amplitude FFT

Voltage T Time

8

Cardiovascular Magnetic Resonance

(8)

1/T Frequency

Figure 1-7 The slice selective B1 pulse. The shape of the slice selective B1 pulse in time is sin (t)/t. This is called a sinc function. It excites a distinct band of frequencies because the Fourier transform of the sinc B1 field is a square wave in frequency space. The excitation of a band of frequencies by the B1, used in conjunction with the slice selective gradient (Fig. 1-6), causes a slice of spins to be placed on the transverse plane. FFT, fast Fourier transform.

PULSE SEQUENCE SHOWING ONLY SLICE-SELECTION Sinc(t) RF excitation Flat top Ramp

Equal area Z slice select Rephasing gradient

Slice select gradient

Y phase encode

X readout gradient

Signal Time

direction: vðxÞ ¼ ðB0 þ Gx xÞðg=2pÞ. This is illustrated in Figure 1-9 for a sphere. The sphere is placed in a main magnetic field, with a magnetic field gradient in the X position. The dashed lines show how the frequency of spins is varying across the sphere. The frequency encoding data provide a signal from which the number of spins at each frequency (i.e., each location) can be determined, using a Fourier transform, which converts time oscillating data into its frequency (or location) components. Using a frequency encoding gradient, the MR signal provides a measure

1 B field

area of the slice selection gradient must be applied to reverse the dephasing caused by the slice selection process (shaded regions in Fig. 1-8). The idea of gradient “area” is very important. The gradient “area” is the product of the gradient amplitude, G, and the duration, Dt, and is proportional to how much phase the spins have accumulated (see Equation 7) because of the gradient. The gradient area to rephase the spins is half the slice select area, because the spins reach the transverse plane exactly coincident with the peak of the RF sinc pulse. By applying the gradient in the opposite direction, the spins are forced back to the same phase as they had before application of the slice select gradient. This small gradient is called a rephasing gradient. In summary, to selectively excite a slice, a linear magnetic field gradient, Gz, is placed along the axis to be sliced. A sinc B1 field pulse is applied, and only the predefined slice within the sample will be placed onto the transverse plane for further modification to create an image.

0

–1

X-axis

Although the Z position is now known for the signal in the FID through the placing of a slice of the sample into the transverse plane, the X and Y positions remain unknown. Recall that to eliminate the effects of T2* on the signal, a spin echo is created by applying a 180 pulse after the 90 slice select pulse. This results in the formation of a spin echo at a time equal to the time interval between the 90 slice select and the 180 pulse (see Fig. 1-4). If a gradient Gx is applied at this time in the X direction (called the frequency encoding direction in MRI), recall that the frequency of the spins will reflect the locations of the spins in the X

# of spins

Frequency Encoding: Position in X –1

0

1

Relative frequency Figure 1-9 Frequency encoding of position. A frequency encoding (or readout) gradient is applied on the X-axis, causing the B field to vary from less than to greater than the main magnetic field. The signal from the spins during the application of this gradient is acquired. The signal provides a measurement of the number of spins precessing at each frequency, thereby indicating the number of spins at each X location. Cardiovascular Magnetic Resonance 9

1 BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

Figure 1-8 Pulse sequence diagram. In a pulse sequence diagram, a timeline is shown for the B1 pulse, the gradients on each axis, and the signal received. This convention is used in all pulse sequence diagrams. Here, the slice selection process is shown, in which a Z gradient is used with a sinc B1 pulse. RF, radiofrequency.

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

of the number of spins precessing at each frequency. Figure 1-9 plots the precessional frequency versus the number of spins at that frequency. This is equivalent to a plot of X position versus the number of spins at that position. The position of these spins along the X-axis is simply detected by determining the frequencies of the spins while a gradient is applied along this axis. Thus, the process is called frequency encoding, and the gradient played out during this period is often termed the readout gradient, or the frequency encoding gradient. As a refinement, another X gradient, called a dephasing gradient, precedes the readout gradient. This dephaser’s area is half the readout gradient area and opposite in sign. The spins are dephased by the dephasing gradient, and then rephased by the readout gradient, causing the spins to be in phase at the center of the readout gradient (see Fig. 1-11). Without a dephasing gradient, the peak signal would occur at the time when the readout gradient was ramping up to the desired gradient amplitude; therefore, the peak signal would not be measured. When a dephaser precedes the readout gradient, the spins refocus to form an echo at the center of the readout gradient. This is what is called a gradient recalled echo (GRE) because the readout gradient is used to refocus the dephasing gradient applied before it. A very thoughtful reader may wonder why the gradients are always “ramped” up and down. The ramps reflect the reality that the magnetic field cannot be changed instantaneously; the change is constrained by the scanner hardware. With current hardware, the gradient coils can be “slewed” to reach their maximum value (4 to 6 gauss/ cm) within approximately 100 to 200 msec, but magnet hardware is always improving. The frequency-encoding component is added to the pulse diagram (Fig. 1-10). First the dephasing gradient is applied with an area equal to half that of the readout gradient. The readout gradient is then ramped up in the opposite direction and held constant for a time, and then the gradient is ramped back down to zero. This is the time during which the signal is acquired (see Fig. 1-10). During this period, the signal is continuously sampled, providing Nx

Object with 3 spins

samples, where the spatial resolution (Dx) is related to Nx, by field of view (FOV)/Nx ¼ Dx. The FOV is the size of the sample (the patient) in the imaging field. Remember that this signal is localized in Z using slice selective RF excitation and localized in X using frequency encoding. The signal measured contains this X and Z information. However, it does not yet contain Y position information.

Phase Encoding: Position in Y Y spatial information is determined through a process called phase encoding. While the spins are still in the transverse plane after the slice selection process, the spins can be phase encoded by transiently applying a magnetic field gradient along the axis chosen for phase encoding (Y-axis by convention). This transient gradient imparts a phase proportional to its area, which also depends on the Y position. However, many phase encoding gradients of 180° pulse Excitation Z slice select Y phase encode gradient X readout gradient

Equal area

Dephasing gradient

Echo Signal

Figure 1-10 Pulse sequence diagram, as in Figure 1-8, with the addition of the readout portion of the sequence.

No gradient

Gy Gy

Y Time

T

Time

T

C

C

Time

C x

B A

C Phase of spins

B

B A

B

A A

Figure 1-11 Phase encoding. A Y gradient is transiently applied for a time, T, using larger and larger gradient strengths (Gy). Three spins, A, B, and C, located at three different Y locations, respond differently to the phase encoding gradient. The gradient does not affect the spin at y ¼ 0 (spin B) because it still precesses at the Larmor frequency. However, the gradient causes spin A to precess more slowly and spin C to precess more quickly, resulting in phase differences proportional to the gradient strength and the time, t. These phase differences allow the locations of the spins to be determined. 10 Cardiovascular Magnetic Resonance

All of the steps are combined into a single acquisition scheme (Fig. 1-12). A separate echo is collected for each phase encoding step. The phase encoding step is most commonly applied immediately after the slice selection process, frequently at the same time that the dephasing gradient for readout occurs (as shown in Fig. 1-12). All of the frequency and phase encoded raw data are combined to create k-space (Fig. 1-13). The k-space of Figure 1-13 represents the raw data collected at each phase encoding step. The phase and frequency encoding directions are labeled. MR imaging can be understood as

TR and TE in a Spin Echo Sequence TR 90º

180º

180º (NS)

90º

Excitation Slice Select Phase encoding

phase encoding strength changes

Readout Signal

TE

TE

Figure 1-12 Pulse sequence diagram with phase encoding. The phase encoding gradient is often applied just before frequency encoding. A cardiovascular magnetic resonance image is acquired by repetition of the pulse sequence multiple times, as shown here, where the 90 gradient, the 180 gradient, and all gradients are shown repeated. The repetition time (TR) and echo time (TE) are defined. During each repeated pulse sequence, the phase encoding strength changes as shown. NS, nonselective.

Raw data collected: k-space

Image

1

Phase encodings (Ky)

Figure 1-13 The frequency encoding data that are collected during each phase encoding step are known as the k-space signal. The k-space signal does not look like an image, but is transformed into an image with Fourier transform (FT).

Raw k-Space Data and the Fast Fourier Transform

FT

Y

0

–1 Frequency encodings (Kx)

X

Cardiovascular Magnetic Resonance 11

1 BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

progressively increasing areas are required to localize a spin in the Y direction, depending on the object size (called an FOV) and the desired spatial resolution (Dy). The number of phase encodings (Ny) is given by the relationship FOV/ Ny ¼ Dy. Recall that to determine the position of a spin within the sample, a gradient is applied and the resulting frequency reports its position. Frequency is simply a measure of how fast the phase is changing in time (Equation 3). Phase encoding is a method for measuring the locations of the spins by measuring their response to different strengths of Y gradient and deducing their frequency (and therefore location) from these measurements. Unlike frequency encoding, which provides the positions in X with a single measurement, phase encoding is performed slowly, over Ny measurements. In successive measurements, the amplitude of the phase encoding gradient is progressively increased and the signal is measured. Each measurement is called a repetition, and the time to acquire a single measurement is called the repetition time (TR). Figure 1-11 describes the phase encoding technique. Consider three spins (A, B, and C), located along Y at three positions. The application of a gradient, Gy, for a short time, t, results in a location-dependent phase (see Equation 3) for the spins. The first phase encoding has zero amplitude and imparts no phase. The signal measured is the bulk signal. In the next phase encoding step, a small gradient is used and the phase difference among the three spins is small. The net signal measured is less than before because the spins are slightly dephased. A larger phase encoding gradient causes even more phase dispersion and less signal. After the signal is collected after multiple phase encodings, the signals can be used to determine Y locations. It is important to realize that for each phase encoding, an entire frequency encoding must be collected. When the readout signal is finally collected, each of the spins’ phases is influenced by its position in Y. In this way, X and Y spatial information is encoded together. Thus, this slow phase encoding process provides the last piece of information, Y-axis position, needed to create the simplest MR image.

Pulse Sequences and Contrast The sequence depicted in Figure 1-12 is a spin echo sequence. Even in this simple sequence, several factors of its design will change the contrast of the MR image, based on T1 and T2 relaxation. Because a 180 refocusing pulse is used in this sequence, the total time that the spins stay in the transverse plane determines the amount of T2 relaxation that will occur (Equation 3). This time is called the time to echo, or echo time (TE) and is measured from the center of the slice select “sinc” pulse to the center of the refocused echo during readout or data acquisition (see Fig. 1-12). Generally, the longer the TE, the more T2 contrast or T2 weighting is generated in the image (Equation 4).

Another sequence timing parameter that influences the signal amplitude is based on T1 relaxation. For a 90 excitation pulse used in spin echo imaging, one must wait approximately five times the T1 value of the tissue to permit the spins to completely regrow to their original Z magnetization (longitudinal magnetization) before applying another pulse to collect the entire MR signal available. If a shorter time is used between slice selective pulses, the spins’ Z magnetization will not fully regrow between pulses, and this leads to a reduction in the MR signal. This reduction in signal is dependent on the T1 of the sample. The longer the T1, the greater the reduction in signal, because less of the magnetization can recover. The time between each slice excitation pulse is the critical factor in this sequence and is called the time to repetition, or repetition time (TR; see Fig. 1-12). As an example of the importance of TE, the MR water signal from normal (T2 80 msec) and acutely infarcted (T2 100 msec) regions of the myocardium is shown in Figure 1-14A. Note that the signal of both tissues decreases with increasing TE, but the infarcted tissue signal decays more slowly because of its longer T2. This results in a contrast, or increased difference, in signal between the normal and the infarcted tissue at prolonged TE. On inspection of the difference between the signals of the two tissue curves, a TE of approximately 50 msec could be selected to optimize the contrast between these two tissue types. Thus, simply by adjusting the imaging parameters, fundamental information on the heart structure can be obtained. Conversely, pathology may be obscured if the imaging parameters are not ideal. The effect of TR on signal amplitude is shown in Figure 1-14B for a spin echo sequence. The effect of TR is illustrated for normal myocardium with a T1 of 0.8 seconds and for a chronic infarct with a T1 of 1.5 seconds. Note that the shorter the T1, the more rapidly the pulses

TE AND TR EFFECTS ON MYOCARDIAL SIGNAL AMPLITUDE 100.00

100.00 Normal

80.00

80.00 Infarct Relative signal

Relative signal

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

“traveling” through k-space. Each frequency encoding acquires a full line of k-space (kx). Each phase-encoding step moves up along the phase-encoding axis (ky) to acquire another frequency encoding line. Once fully sampled, this k-space is converted into an image using a twodimensional fast Fourier transform. Ideally, the raw kspace data contain enough information so that after the two-dimensional fast Fourier transform an image of infinite spatial resolution results. In reality, there are limits on the amount of data that can be collected. The resolution of the MRI image is directly related to the number of phase-encoded echoes collected (Dy ¼ FOV/Ny); therefore, infinite spatial resolution would require infinite time to collect. In CMR, time is limited because of respiratory and cardiac motions, so the trade-off between the spatial resolution required to observe the structure of interest and scan time (i.e., Ny) is critically important. Schemes to increase the efficiency of collecting the phase-encoded data are discussed later.

60.00 Normal 40.00 Difference

20.00

60.00

40.00

Difference

20.00

0.00

0.00 0

A

Infarct

0.02

0.04

0.06

0.08

TE (sec)

0.10

0.12

0.14

0

B

1

2

3

4

5

6

7

TR (sec)

Figure 1-14 A, Effect of echo time (TE) and (B) effect of repetition time (TR) on cardiovascular magnetic resonance signal amplitude for normal and infarcted myocardium. Because the T1 and T2 of infarcted myocardium differ from those of normal myocardium, contrast between these tissues can be created. 12 Cardiovascular Magnetic Resonance

Fast Spin Echo Imaging The relatively long TR required for full T1 relaxation is the major reason why spin echo methods are slow to collect the phase encoded information required to create an MR image. To circumvent this problem, one highly useful approach is to use multiple 180 refocusing pulses and to collect many echoes during each slice selective excitation. This method, called fast spin echo (FSE), or turbo spin echo (TSE), is an important technique in CMR.21 A phase encoding gradient is applied between each 180 pulse, thereby providing multiple phase encoded steps from a single slice selective pulse. For CMR, typically, 16 to 64 echoes are collected for each slice selective pulse, thereby reducing the time to collect a spin echo image by 16 to 64 times, respectively. Because large blocks of time are required to collect all of these echoes, this method is usually restricted to relatively motion-free phases of the cardiac cycle, such as diastole. The inherently high signalto-noise ratio (SNR) of these FSE approaches supports very high spatial resolution images of the heart. In addition, true T2 contrast can be generated by acquiring the central phase encoding data (which largely controls the image contrast) at a specific time within the echo train. This time is called the effective TE. The longer the time interval between the initial 90 and the acquisition of central k-space (i.e., the longer the effective TE), the more T2 weighting will occur. The FSE technique is very commonly used in CMR to visualize anatomy and measure

the sizes of cardiac chambers (see Fig. 1-13, which shows an FSE image).

Double Inversion Recovery (“Black-Blood”) Fast Spin Echo Fast spin echo imaging of the heart is usually performed with a preparation sequence that nulls flowing blood, so that only the ventricular and atrial myocardium are visible. This preparation is performed before FSE acquisition (Fig. 1-15), using a double inversion recovery sequence, also known as a black-blood preparation. This preparation consists of two inversion (180 ) pulses that immediately follow each other. The first inversion is a nonselective inversion pulse that inverts all magnetization. The second inversion is a slice selective inversion pulse that reinverts a slice of interest, which is centered on the imaging slice, but is slightly larger. This preparation is timed so that at the time of acquisition, all blood magnetization has regrown from fully inverted to 0. Any blood that was reinverted because it was within the slice of interest has now flowed out of the imaging slice. Any blood that has flowed into the slice will contribute no signal because it’s nulled. T2-weighted black-blood FSE has recently been introduced for identifying edematous tissue associated with acute infarct22 and for identifying the region at risk.23

Gradient Echo Imaging For imaging the beating heart and other highly dynamic applications, an imaging sequence with a very short TR is useful. For a short TR, lower flip angles must be used because a long TR is necessary for full relaxation from a 90 slice selective pulse. Furthermore, for a very short TR, the 180 refocusing pulse is eliminated. The signal on the transverse plane will then decay with T2* instead of T2, but this is acceptable if the TE is very short (TE < 3 msec). As the TR is shortened, the flip angle must be reduced to provide the optimal SNR per unit time. The flip

Figure 1-15 Pulse sequence diagram for fast spin echo imaging. TE, echo time.

Fast Spin Echo Sequence 908

1808

1808

Repeated n times

1808 Excitation Slice Select

Phase encode

1808

Readout

1808 Signal

TE

TE

TE

Black-blood preparation

Cardiovascular Magnetic Resonance 13

1 BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

can be applied and still maintain the MR signal. Also apparent in Figure 1-14B is that by varying the TR, the contrast or difference between tissues can be altered. Thus, the TR can be used to vary the image contrast based on T1. Note that a TR of 1 second would be optimal in generating the largest difference between normal and infarcted tissue. This is convenient because a typical TR for spin echo and fast spin echo (discussed later) imaging is 1 R-R interval—approximately 1 second—applying the B1 pulse every heartbeat.

a ¼ cos1 ðeTR=T1 Þ

fast gradient recalled echo (GRE, turbo FLASH, or turbo TFE).24 GRE refers to the refocusing of the frequency encoding gradient, as described earlier. The fast GRE pulse sequence is schematically illustrated in Figure 1-16A. Note that to keep the phase information intact, rewinding gradients are applied after each acquisition equal to and opposite the phase encoding gradient. Furthermore, the readout gradient is also rewound so that on all axes, the spins are in phase after each TR. Usually, an extra gradient is applied at the end of the TR on the X or Z axis, and is called a crusher, killer, or spoiler gradient (see Fig. 1-16), which dephases any remaining transverse magnetization so that it does not contribute signal during the next phase encode step. Although fast GRE images have T2* dephasing, this does not usually reduce image quality because a very short TE is achievable (1 to 3 msec), unless a serious source of magnetic field inhomogeneity (e.g., a metallic implant) is present. Furthermore, the GRE sequence can be used to quantify myocardial T2* (by imaging at multiple TE times), which may help detect diseases with myocardial iron deposition, such as thalassemia,25,26 with good reproducibility.27

(9)

This low flip angle RF pulse is also called an alpha pulse and is suitable for short TRs. A short TR is needed for dynamic imaging. For example, collecting 32 phase encoded lines in k-space, each with a TR of 5 msec, requires 160 msec (5 msec 32 phase encoded lines). This time is sufficiently short that 32 TRs can be acquired in the quiescent diastolic period. Thus, data can be collected rather rapidly using a small excitation pulse rather than the 90 pulse. This imaging method, with a short TR, low flip angle (alpha pulse), and no 180 refocusing pulse, is known as

Gradient Recalled Echo TR

a

a

a

a

a

Excitation Slice Select

Equal area

Phase

Crusher

Readout

Signal

Repeated n times

TE

A

Segmented Inversion Recovery

TI ECG 1808 (NS) 1.0 0.0

Mz

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

angle providing maximum signal at steady state (i.e., after many slice selective pulses) is determined by a trade-off. A larger flip angle provides more spins on the transverse plane (more signal), but less Z magnetization is preserved. Therefore, the following echoes have less signal. A smaller flip angle provides less signal on the transverse plane, but more Z magnetization is preserved for subsequent echoes. The maximum signal at steady state is achieved using the Ernst angle for a given TR and T1:

1808 (NS)

GRE

GRE

Fibrosis Myo

Blood

−0.0 1808 (NS)

B 14 Cardiovascular Magnetic Resonance

GRE Repeat n times to obtain single image

Figure 1-16 A, Pulse sequence diagrams for gradient recalled echo (GRE) imaging showing multiple repetition times (TR). B, Inversion recovery sequence, which uses the GRE sequence, acquiring multiple phase encoding lines in each R-R interval. Before each acquisition, a nonselective 180 B1 pulse is used to generate contrast between tissues with different T1 values. The inversion time (TI) is the time between the 180 B1 pulse and data acquisition. The Z magnetizations for blood, myocardium (Myo), and infarct (labeled) are plotted versus time within the sequence. Because the T1 of infarcted myocardium in the equilibrium phase of a contrast injection is shorter than that of normal myocardium, contrast can be generated between these tissues by judicious choice of TI. ECG, electrocardiogram; NS, nonselective; TE, echo time.

MzðtÞ ¼ M0 ð1 et=T1 Þ MzSS et=T1

The fast GRE pulse sequence is also used for contrastenhanced MR angiography, which is the imaging of arteries and veins during the first pass of an exogenous contrast agent.28,29 During the first pass of a gadolinium contrast agent, blood T1 is reduced to approximately 30 msec. The Ernst angle equation (Equation 9) identifies an optimal angle of approximately 30 for a TR of 5 msec. The acquisition of three-dimensional fast GRE provides a high-quality image of vessels (see Chapters 32–35). However, without contrast agent, the SNR of fast GRE imaging can be low because the time between TRs, during which the longitudinal magnetization can regrow, is short. This limitation is partly overcome using balanced steady-state free precession (SSFP).

Inversion Recovery Fast Gradient Recalled Echo: Late Gadolinium Enhancement One important application for fast GRE is detection of myocardial scar or infarction30 through injection of a gadolinium contrast agent, which slowly accumulates in regions of scar. It is detected as a hyperenhanced (“bright”) signal when imaged 10 and 20 minutes after injection.31 The T1 of normal myocardium, blood, and scarred (infarcted) myocardium is roughly 380 msec, 300 msec, and 270 msec, respectively, at 20 minutes postinjection of 0.2 mmol/kg gadolinium (a typical dose and delay time).14,15,32 A T1-weighted sequence, called an inversion recovery sequence, is used to create contrast between these tissues, which have slightly different T1 values. Inversion recovery uses a non-slice-selective 180 pulse, and a delay before imaging is called the time to inversion, or inversion time (TI). Then the raw data are collected using a GRE segmented k-space approach (see Fig. 1-16B).33 Many

Figure 1-17 Balanced steadystate free precession imaging (SSFP). The sequence is identical to gradient recalled echo except that the flip angle for SSFP is high and its sign is alternated in each repetition time (TR). Also, the gradients in the SSFP sequence are fully rephased. TE, echo time.

(10)

SS

where Mz is the steady-state Z magnetization. Figure 1-16B shows the Z magnetization for the inversion recovery sequence and the T1 typical of blood, myocardium, and infarct. An image acquired at a time when the magnetization from normal myocardium has regrown to zero (the optimum TI) provides contrast in which the scar appears bright and the myocardium dark (see Fig. 1-16B).

Balanced Steady-State Free Precession Directly related to the GRE sequence is the balanced SSFP method. Today, almost all CMR imaging of ventricular function at 1.5 T employs this method, and it also is used for anatomic localization, valve visualization, and coronary artery imaging.34 This very old technique35 was more recently revived36,37 because better scanner hardware allows for the short TR (e.g., 2 to 4 msec at 1.5 T) required for balanced SSFP. Balanced SSFP uses a pulse sequence identical to that of GRE (Fig. 1-17), except for two distinctions. First, after each TR, the spins are rephased on each gradient axis (X, Y, and Z) to zero phase. No crushers, killers, or spoilers are used. This keeps the spins completely in phase (except for dephasing as a result of magnetic field inhomogeneities) so that the transverse magnetization can be reused in the next TR. For this reason, the sequence is called balanced. Second, a large flip angle (e.g., 60 ) is used, flipping spins around the positive and negative Y-axis, in alternate TRs. This flip angle alternation scheme reuses the remaining transverse magnetization and mixes it with longitudinal magnetization, for increased signal in each TR. Balanced SSFP provides high SNR of blood and a unique T2/T1 weighting.38 As stated earlier, balanced SSFP is highly sensitive to off-resonance, requires a short TR and a large flip angle, and is challenging at 3 T, where off-resonance creates artifacts7 and RF heating concerns (specific absorption rate) limit the flip angle.38a

BALANCED SSFP +60°

–60°

+60°

–60°

+60°

Excitation Slice select Phase encode Readout TE Signal TR

Repeated n times

Cardiovascular Magnetic Resonance 15

1 BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

heartbeats are required to collect data for a full image. After a 180 pulse (called an inversion pulse), Mz(t) recovers with a relaxation time, T1:

Three-Dimensional Fast Gradient Echo: Magnetic Resonance Angiography

Analogous to reducing scan time by replacing spin echo acquisition by fast spin echo acquisition, the GRE sequence can be accelerated by acquiring a train of gradient-recalled echoes after a single RF pulse. This approach is called multi-echo, or echo planar, imaging (EPI), and was among the first CMR imaging methods described.39 For anatomy with tissues with long T2* (e.g., the head), EPI can result in a complete image collection with a single slice select pulse and a train of gradient echoes (Fig. 1-18). This is the basis for the functional MRI (fMRI) technique that is widely used in brain mapping. To reduce time between echoes, the readout gradient is rapidly reversed with every phase encode step, so that MR signals are collected during both the positive and negative lobes of the readout gradient. This approach is challenging in the heart, where T2* can be short11 compared with the brain. Because of the acquisition of many phase encode steps after one RF excitation, the EPI method results in poor SNR, image trajectory errors that can be corrected,40-42 and image distortions because of off-resonance. However, these limitations can be overcome by collecting only a few

(e.g., 3 to 9) phase encode steps per slice select RF excitation (instead of all phase encode steps) and repeating this process until all of the data have been collected. With state-of-the-art gradient hardware, nine echoes can be collected per slice excitation with a TR of 10 msec. Thus, imaging times on the order of 80 msec can be achieved in low-resolution images (64 128 image resolution over a 15 32 cm FOV) without the magnetization ever spending more than 10 msec in the transverse plane.43 This method is sometimes used for myocardial perfusion imaging, which requires complete acquisition of multiple slices in every heartbeat. A schematic of an electrocardiogram-gated segmented EPI sequence for myocardial perfusion is shown in Figure 1-18. Then data are rapidly acquired using a multi-shot EPI method. An image from several slices is acquired every heartbeat during the first passage of gadolinium contrast agent through the heart. In this sequence, a nonselective 90 pulse is used to saturate all of the spins to provide T1 weighting (Equation 2) before data acquisition. Spiral imaging is similar to EPI, with more image data obtained after a single RF pulse, compared with conventional imaging, by continuously sampling k-space in a spiral pattern (one interleaf is seen in Fig. 1-19A). One or more interleaves (spiral arms) are collected to fully sample kspace. Spiral image quality is affected by off-resonance artifacts and trajectory errors, but it can provide high SNR as

Multi-shot EPI

ECG 90⬚ (NS)

Tdelay ~ 100–200 msec

20⬚

20⬚

Excitation

20⬚

Slice Select Phase encode

Repeated n times for N dynamics

Readout Signal

SPIRAL TRAJECTORY

Figure 1-19 A, Spiral imaging acquires k-space by sampling along a spiraling trajectory. One or more interleaves, rotated with respect to each other, are required to sample k-space fully. This schematic shows one interleaf of an Archimedean spiral. B, Radial imaging acquires multiple projections at equally spaced angles to sample k-space, analogous to computed tomography. The sampling density of the radial trajectory is greater at the center of k-space, so the lower spatial frequencies are always well sampled.

Ky-axis

0.5

0

0

–0.5

–0.5 –0.5

A

Figure 1-18 Segmented echo planar imaging (also called multi-echo imaging; EPI), in which a portion of the frequency encodings are collected after each B1 pulse. Here, a typical pulse sequence for myocardial perfusion is shown, in which eight frequency encodings are collected after each small B1 pulse (here with a 20 flip angle). The acquisition is repeated until all of the data are acquired for an image. Before the multi-shot EPI sequence, a nonselective 90 B1 pulse is used to generate contrast between tissues with different T1 values. The delay time (Tdelay) is the time between the 90 B1 pulse and the data acquisition. ECG, electrocardiogram; NS, nonselective.

RADIAL TRAJECTORY

0.5

Ky-axis

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

Echo Planar Imaging, Spiral, and Radial

0 Kx-axis

16 Cardiovascular Magnetic Resonance

0.5

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B

0 Kx-axis

0.5

imaging was the first imaging method,1 and it has regained some popularity because of trajectory improvements and the recent developments in reconstruction of undersampled MR data51,52 using sparsity constraints. In conclusion, these basic principles of CMR are the foundation of the many techniques of CMR imaging of normal and pathologic cardiac anatomy and function.

References 1. Lauterbur P. Image formation by induced local interactions: examples employing nuclear magnetic resonance. Nature. 1973;242:190–191. 2. Niendorf T, Hardy CJ, Giaquinto RO, Gross P, Cline HE, Zhu Y, Kenwood G, Cohen S, Grant AK, Joshi S, Rofsky NM, Sodickson DK. Toward single breath-hold whole-heart coverage coronary MRA using highly accelerated parallel imaging with a 32-channel MR system. Magn Reson Med. 2006;56(1):167–176. 3. Gharib AM, Elagha A, Pettigrew RI. Cardiac magnetic resonance at high field: promises and problems. Curr Probl Diagn Radiol. 2008;37(2):49–56. 4. Stuber M, Botnar RM, Fischer SE, Lamerichs R, Smink J, Harvey P, Manning WJ. Preliminary report on in vivo coronary MRA at 3 Tesla in humans. Magn Reson Med. 2002;48(3):425–429. 5. Nezafat R, Stuber M, Ouwerkerk R, Gharib AM, Desai MY, Pettigrew RI. B1-insensitive T2 preparation for improved coronary magnetic resonance angiography at 3 T. Magn Reson Med. 2006;55 (4):858–864. 6. Gebker R, Jahnke C, Paetsch I, Kelle S, Schnackenburg B, Fleck E, Nagel E. Diagnostic performance of myocardial perfusion MR at 3 T in patients with coronary artery disease. Radiology. 2008;247(1):57–63. 7. Schar M, Kozerke S, Fischer SE, Boesiger P. Cardiac SSFP imaging at 3 Tesla. Magn Reson Med. 2004;51(4):799–806. 8. Sharma P, Socolow J, Patel S, Pettigrew RI, Oshinski JN. Effect of GdDTPA-BMA on blood and myocardial T1 at 1.5T and 3T in humans. J Magn Reson Imaging. 2006;23(3):323–330. 9. Botnar RM, Stuber M, Danias PG, Kissinger KV, Manning WJ. Improved coronary artery definition with T2-weighted, free-breathing, three-dimensional coronary MRA. Circulation. 1999;99 (24):3139–3148. 10. Foltz WD, Yang Y, Graham JJ, Detsky JS, Dick AJ, Wright GA. T2 fluctuations in ischemic and post-ischemic viable porcine myocardium in vivo. J Cardiovasc Magn Reson. 2006;8(3):469–474. 11. Reeder SB, Faranesh AZ, Boxerman JL, McVeigh ER. In vivo measurement of T2* and field inhomogeneity maps in the human heart at 1.5 T. Magn Reson Med. 1998;39(6):988–998. 12. Croisille P, Revel D, Saeed M. Contrast agents and cardiac MR imaging of myocardial ischemia: from bench to bedside. Eur Radiol. 2006;16 (9):1951–1963. 13. Scholz TD, Martins JB, Skorton DJ. NMR relaxation times in acute myocardial infarction: relative influence of changes in tissue water and fat content. Magn Reson Med. 1992;23(1):89–95. 14. Goldfarb JW, Mathew ST, Reichek N. Quantitative breath-hold monitoring of myocardial gadolinium enhancement using inversion recovery TrueFISP. Magn Reson Med. 2005;53(2):367–371. 15. Klein C, Nekolla SG, Balbach T, Schnackenburg B, Nagel E, Fleck E, Schwaiger M. The influence of myocardial blood flow and volume of distribution on late Gd-DTPA kinetics in ischemic heart failure. J Magn Reson Imaging. 2004;20(4):588–593. 16. Parker DL, Goodrich KC, Alexander AL, Buswell HR, Blatter DD, Tsuruda JS. Optimized visualization of vessels in contrast enhanced intracranial MR angiography. Magn Reson Med. 1998;40(6):873–882. 17. Mathur-De Vre R, Lemort M. Invited review: biophysical properties and clinical applications of magnetic resonance imaging contrast agents. Br J Radiol. 1995;68(807):225–247. 18. Weissleder R, Elizondo G, Wittenberg J, Rabito CA, Bengele HH, Josephson L. Ultrasmall superparamagnetic iron oxide: characterization of a new class of contrast agents for MR imaging. Radiology. 1990;175(2):489–493. 19. Conolly S, Glover G, Nishimura D, Macovski A. A reduced power selective adiabatic spin-echo pulse sequence. Magn Reson Med. 1991;18(1):28–38. 20. Meyer CH, Pauly JM, Macovski A, Nishimura DG. Simultaneous spatial and spectral selective excitation. Magn Reson Med. 1990;15 (2):287–304.

21. Hennig J, Nauerth A, Friedburg H. RARE imaging: a fast imaging method for clinical MR. Magn Reson Med. 1986;3(6):823–833. 22. Abdel-Aty H, Zagrosek A, Schulz-Menger J, Taylor AJ, Messroghli D, Kumar A, Gross M, Dietz R, Friedrich MG. Delayed enhancement and T2-weighted cardiovascular magnetic resonance imaging differentiate acute from chronic myocardial infarction. Circulation. 2004;109 (20):2411–2416. 23. Aletras AH, Tilak GS, Natanzon A, Hsu LY, Gonzalez FM, Hoyt RF, Jr., Arai AE. Retrospective determination of the area at risk for reperfused acute myocardial infarction with T2-weighted cardiac magnetic resonance imaging: histopathological and displacement encoding with stimulated echoes (DENSE) functional validations. Circulation. 2006;113 (15):1865–1870. 24. Haase A, Matthaei D, Hanicke W, Frahm J. Dynamic digital subtraction imaging using fast low-angle shot MR movie sequence. Radiology. 1986;160(2):537–541. 25. Anderson LJ, Holden S, Davis B, Prescott E, Charrier CC, Bunce NH, Firmin DN, Wonke B, Porter J, Walker JM, Pennell DJ. Cardiovascular T2-star (T2*) magnetic resonance for the early diagnosis of myocardial iron overload. Eur Heart J. 2001;22(23):2171–2179. 26. Westwood M, Anderson LJ, Firmin DN, Gatehouse PD, Charrier CC, Wonke B, Pennell DJ. A single breath-hold multiecho T2* cardiovascular magnetic resonance technique for diagnosis of myocardial iron overload. J Magn Reson Imaging. 2003;18(1):33–39. 27. He T, Kirk P, Firmin DN, Lam WM, Chu WC, Au WY, Chan GC, Tan RS, Ng I, Biceroglu S, Aydinok Y, Fogel MA, Cohen AR, Pennell DJ. Multi-center transferability of a breath-hold T2 technique for myocardial iron assessment. J Cardiovasc Magn Reson. 2008;10(1):11. 28. Prince MR, Yucel EK, Kaufman JA, Harrison DC, Geller SC. Dynamic gadolinium-enhanced three-dimensional abdominal MR arteriography. J Magn Reson Imaging. 1993;3(6):877–881. 29. Zhang H, Maki JH, Prince MR. 3D contrast-enhanced MR angiography. J Magn Reson Imaging. 2007;25(1):13–25. 30. Kim RJ, Wu E, Rafael A, Chen EL, Parker MA, Simonetti O, Klocke FJ, Bonow RO, Judd RM. The use of contrast-enhanced magnetic resonance imaging to identify reversible myocardial dysfunction. N Engl J Med. 2000;343(20):1445–1453. 31. Simonetti OP, Kim RJ, Fieno DS, Hillenbrand HB, Wu E, Bundy JM, Finn JP, Judd RM. An improved MR imaging technique for the visualization of myocardial infarction. Radiology. 2001;218 (1):215–223. 32. Weinmann H, Laniado M, Mutzel W. Pharmacokinetics of GdDTPA/ Dimeglumine after intravenous injection into healthy volunteers. Physiol Chem Phys Med NMR. 1984;16:167–172. 33. Edelman RR, Wallner B, Singer A, et al. Segmented turboFLASH: method for breath-hold MR imaging of the liver with flexible contrast. Radiology. 1990;177(2):515–521. 34. Weber OM, Martin AJ, Higgins CB. Whole-heart steady-state free precession coronary artery magnetic resonance angiography. Magn Reson Med. 2003;50(6):1223–1228. 35. Oppelt A, Graumann R, Barfuss H, Fischer H, Hart W, Shajor W. FISP–a new fast MRI sequence. Electromedica. 1986;54:15–18. 36. Duerk JL, Lewin JS, Wendt M, Petersilge C. Remember true FISP? A high SNR, near 1-second imaging method for T2-like contrast in interventional MRI at .2 T. J Magn Reson Imaging. 1998;8(1):203–208. 37. Heid O. True FISP cardiac fluoroscopy. In: Fifth Annual Proceedings of International Society of Magnetic Resonance in Medicine. Vancouver, BC, Canada; 1997. 38. Scheffler K, Lehnhardt S. Principles and applications of balanced SSFP techniques. Eur Radiol. 2003;13(11):2409–2418. 38a. Scha¨r M, Kozerke S, Fischer SE, Boesiger P. Cardiac SSFP imaging at 3 Tesla. Magn Reson Med. 2004;51:799–806. Cardiovascular Magnetic Resonance 17

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a result of high SNR efficiency. Spiral imaging has been shown to have advantages in coronary artery imaging and real-time imaging.44-46 Radial imaging is another trajectory that acquires k-space data as radial spokes (Fig. 1-19B shows 16 radial spokes, or projections), analogous to computed tomography. It has been applied to the heart, especially using undersampling of k-space for fast imaging.47–50 Radial

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39. Mansfield P, Pykett I. Biological and medical imaging by NMR. J Magn Res. 1978;29:355–373. 40. Peters DC, Derbyshire JA, McVeigh ER. Centering the projection reconstruction trajectory: reducing gradient delay errors. Magn Reson Med. 2003;50(1):1–6. 41. Reeder SB, Atalar EA, Faranesh AZ, McVeigh ER. Referenceless interleaved echo-planar imaging. Magn Reson Med. 1999;41(1):87–94. 42. Duyn JH, Yang Y, Frank JA, van der Veen JW. Simple correction method for k-space trajectory deviations in MRI. J Magn Reson. 1998;132(1):150–153. 43. Epstein FH, Wolff SD, Arai AE. Segmented k-space fast imaging using an echo-train readout. Magn Reson Med. 1999;41(3): 609–613. 44. Nayak KS, Pauly JM, Yang PC, Hu BS, Meyer CH, Nishimura DG. Real-time interactive coronary MRA. Magn Reson Med. 2001;46 (3):430–435. 45. Yang PC, Meyer CH, Terashima M, Kaji S, McConnell MV, Macovski A, Pauly JM, Nishimura DG, Hu BS. Spiral magnetic resonance coronary angiography with rapid real-time localization. J Am Coll Cardiol. 2003;41(7):1134–1141.

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46. Hardy CJ, Zhao L, Zong X, Saranathan M, Yucel EK. Coronary MR angiography: respiratory motion correction with BACSPIN. J Magn Reson Imaging. 2003;17(2):170–176. 47. Larson AC, White RD, Laub G, McVeigh ER, Li D, Simonetti OP. Selfgated cardiac cine MRI. Magn Reson Med. 2004;51(1):93–102. 48. Peters DC, Ennis DB, McVeigh ER. High-resolution MRI of cardiac function with projection reconstruction and steady-state free precession. Magn Reson Med. 2002;48(1):82–88. 49. Bi X, Park J, Larson AC, Zhang Q, Simonetti O, Li D. Contrastenhanced 4D radial coronary artery imaging at 3.0 T within a single breath-hold. Magn Reson Med. 2005;54(2):470–475. 50. Barger AV, Grist TM, Block WF, Mistretta CA. Single breath-hold 3D contrast-enhanced method for assessment of cardiac function. Magn Reson Med. 2000;44(6):821–824. 51. Lustig M, Donoho D, Pauly JM. Sparse MRI: the application of compressed sensing for rapid MR imaging. Magn Reson Med. 2007;58(6): 1182–1195. 52. Block KT, Uecker M, Frahm J. Undersampled radial MRI with multiple coils: iterative image reconstruction using a total variation constraint. Magn Reson Med. 2007;57(6):1086–1098.

Clinical Cardiovascular Magnetic Resonance Imaging Techniques Wolfgang G. Rehwald, Anja Wagner, Timothy S. E. Albert, Burkhard Sievers, Christopher K. Dyke, Michael D. Elliott, John D. Grizzard, Raymond J. Kim, and Robert M. Judd

It has now been more than 20 years since the first magnetic resonance (MR) images of the human heart were described. Although few would deny that the quality of cardiovascular MR (CMR) images has dramatically improved since that time, it has nevertheless been only in the last 5 to 10 years that large numbers of dedicated CMR clinical services have opened in the United States. Although organizations such as the Society for Cardiovascular Magnetic Resonance (SCMR; www.scmr .org) are working toward scan standardization, there is often variability among CMR centers as to what actually comprises a “routine clinical CMR.” In this setting, a chapter that attempts to describe all of the possible permutations of pulse sequences, scan protocols, and interpretation methods can quickly become unwieldy. This chapter focuses on the specific clinical CMR imaging techniques that are used routinely at the Duke Cardiovascular Magnetic Resonance Center (DCMRC). More detailed sequences and details are found in the respective chapters addressing each setting or pathology. The dedicated CMR clinical service at the DCMRC first opened in July 2002. Since that time, the authors have experimented with and evaluated many approaches to CMR and have directly observed their effects on overall clinical volume. Clinical volume at the DCMRC has grown steadily since the service first opened and currently includes more than 3000 CMR studies per year (Fig. 2-1). For the last several years, the distribution of CMR tests performed has remained relatively constant (Fig. 2-2), but may vary greatly depending on the expertise and patient population at other CMR centers.1 Nearly 50% of our clinical volume includes pharmacologic stress testing.2 The DCMRC definition of a CMR stress test can be described briefly as a group of three individual scans: 1. Left ventricular (LV) cines (to examine systolic function) 2. Adenosine stress/rest perfusion (to examine coronary flow reserve) 3. Late gadolinium enhancement (LGE; to examine viability/infarction) An additional 25% of the clinical volume is referred for viability testing. The DCMRC definition of a CMR viability test can be described briefly as a group of two individual scans: 1. LV cines (to examine systolic function) 2. LGE (to examine viability/infarction) It is immediately apparent that the only difference between a CMR viability test and a CMR stress test is that the former does not include perfusion imaging. Together, these two tests account for nearly three fourths of annual procedures.

Through the years, the DCMRC philosophy about how to perform CMR has evolved into the concept of a CMR “exam menu” that not only has allowed streamlining of the clinical service but also has made it considerably easier to teach CMR to cardiology fellows and level 2 trainees. The underlying idea of the exam menu (Fig. 2-3) is that most, if not all, routine clinical CMR procedures can be performed simply by combining one or more items from a predetermined menu and performing these scan protocols/pulse sequences in a sequential and consistent manner, analogous to the performance of a transthoracic echocardiogram or physical examination. Thus, both the CMR stress test and the CMR viability test are simply combinations of items from the exam menu. Virtually every study includes two or more of the items listed in Figure 2-3. Perhaps more importantly, only rarely is it necessary to perform a scan not listed on this menu. Because of issues related to the development of nephrogenic systemic fibrosis, all patients should have a renal function assessment before the administration of gadolinium contrast (see Chapter 6). For those with moderate or severely depressed renal function (estimated glomerular filtration rate, eGFR < 30 mL/min/1.73 m2), alternatives to gadolinium contrast should be sought. Accordingly, the remainder of this chapter focuses on each of the items on the exam menu, with an emphasis on providing practical information not typically provided in other CMR textbooks. The conclusion provides typical examples of how the items on the exam menu can be combined to answer common diagnostic questions in cardiology, such as the detection of coronary artery disease and the evaluation of aortic disease. The sections within this chapter are presented in order of increasing technical complexity, whereas items in the exam menu (see Fig. 2-3) are ordered from most to least frequently used in clinical practice. For consistency in this chapter, a “test” is defined as one of the components of the pie chart seen in Figure 2-2 and a “scan” is defined as one of the items on the exam menu seen in Figure 2-3.

SCOUTS (“SCAN” ¼ “SCOUT”) The goal of CMR scout scanning is to establish the shortand long-axis views of the heart and to confirm the optimal position of the anterior and posterior elements of the thoracic coil. Because of patient-to-patient anatomic variation, Cardiovascular Magnetic Resonance 19

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CHAPTER 2

700 CMR procedures (quarterly)

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Figure 2-1 Clinical volume of the Duke Cardiovascular Magnetic Resonance Center since its inception in 2002.

800

600 500 400 300 200 100 0 2002- 2002- 2003- 2003- 2003- 2003- 2004- 2004- 2004- 2004- 2005- 2005- 2005Q3 Q4 Q1 Q2 Q3 Q4 Q1 Q2 Q3 Q4 Q1 Q2 Q3

Congenital evaluation (7%)

CMR EXAM MENU (“SCANS”)

CMR angiography (20%) Stress testing (47%)

1 −− Scout 2 −− Left ventricular/right ventricular cine 3 −− Stress–rest myocardial perfusion 4 −− Late gadolinium enhancement 5 −− Morphology 6 −− Angiography 7 −− Flow/velocity Figure 2-3 The Duke Cardiovascular Magnetic Resonance Center “Exam Menu.”

Cardiac viability testing (26%) Figure 2-2 Overall makeup of the clinical volume. CMR, cardiovascular magnetic resonance.

both the short- and long-axis cardiac views lie at arbitrary angles with respect to scanner coordinates and are therefore referred to as “double oblique” planes. The first step in acquiring these double-oblique views is to acquire images along the axes of the scanner (i.e., axial, sagittal, and coronal planes) passing through the thorax. Figure 2-4 shows typical examples of scout images. Once images along the scanner axes have been acquired, they can be used to prescribe a single oblique, for example, perpendicular to the axial image of Figure 2-4 going through the LVapex and midmitral annulus while parallel to the interventricular septum (dashed yellow line). From this view, one can then prescribe a double oblique, for example, oriented on the true short axis of the LV. Thus, from the perspective of the scanner operator, the goal of scouting is simply to acquire a series of images similar to that of Figure 2-4 and including short- and long-axis images of the LV. The pulse sequence used to scout is based on steadystate free precession (SSFP). The underlying concept of SSFP was described in the mid 1980s,3 but only since the late 1990s has CMR scanner hardware been capable of achieving the SSFP magnetization state. The SSFP pulse 20 Cardiovascular Magnetic Resonance

sequence timing diagram is characterized by an elegant simplicity in which there is symmetry around the data acquisition window on all three gradient axes (Fig. 2-5). The axis labeled “slice” selects the slice to be imaged. The waveforms on the “read” axis create the MR signal as an echo (see “signal” axis). During the positive portion of the “read” waveform, the MR signal is digitally sampled. The “phase”-encoding axis imposes a different phase on each echo that allows spatial encoding of the second dimension called “y” in Figure 2-6. In practice, the primary advantages of SSFP imaging are rapid acquisition speed and resultant images that show very high signalto-noise ratio (SNR) and contrast-to-noise ratio (CNR) in the blood pool and surrounding myocardium. Accordingly, SSFP is used not only for scouts but also for cine CMR imaging. Acquiring any CMR image involves filling the raw data space, referred to as “k-space.” Figure 2-6 shows the process of filling k-space as a series of “lines” starting at the top and proceeding to the bottom. Each line is actually one echo (see enlarged box) acquired during the read event, for example, of the sequence in Figure 2-5. The echo runs in direction x. Direction y is the phase-encoding direction. For an SSFP pulse sequence running on a typical modern CMR scanner, the time needed to acquire each k-space line is approximately 3 msec. Thus, for 100 k-space lines (typical for scouts), the total image acquisition time

Coronal

Sagittal

Figure 2-4 Scout images. Typical examples of axial, sagittal, and coronal scouts.

is approximately 300 msec (100 3 msec). This is fast enough to effectively freeze heart motion, provided that the image data are acquired during diastole when the ventricles are relatively quiescent. Acquiring the image data in diastole requires the scanner hardware to be “triggered” to the cardiac cycle. This is typically achieved by accurately detecting the R-wave of the electrocardiogram (ECG), a process that was previously difficult because of the ECG distortion caused by the magnetohydrodynamic effect of pulsatile blood in the aorta. Vector ECG and other solutions have largely minimized this issue.4 Assuming a heart rate of 60 bpm (R-R duration of 1000 msec), the first “event” the scanner hardware must play out is a delay time of 500 msec (to get to diastole), followed by 300 msec of image data acquisition. Figure 2-7 shows cardiac gating for two successive scout images. Typically, 27 scout images are acquired to define the thoracic contents, including 9 parallel images in each of the axial, coronal, and sagittal imaging planes.

MORPHOLOGY (“SCAN” ¼ “MORPHOLOGY”) Morphology scanning is used less often than cine CMR, but the technical aspects of morphology, particularly as it relates to cardiac gating, are only modestly more complex than for scouting, so they are discussed next. The “goal” of morphology scanning is essentially to create a series of parallel slices that “bread loaf” the thoracic cavity to examine vascular anatomy. Any orientation can be acquired, but frequently only axial planes (or axial in addition to coronal and sagittal planes). It is often desirable to use both “black-blood” imaging (Fig. 2-8) and “brightblood” imaging (Fig. 2-9). The CMR pulse sequences used for these are half-Fourier single-shot fast spin echo (HASTE) and SSFP, respectively. Figure 2-10A shows the pulse sequence timing diagram of the fast spin echo sequence, and Figure 2-10B shows its predecessor, the spin Cardiovascular Magnetic Resonance 21

2 CLINICAL CARDIOVASCULAR MAGNETIC RESONANCE IMAGING TECHNIQUES

Axial

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

Symmetry +α

–α TR

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Time

Figure 2-5 Steady-state free precession (SSFP) pulse sequence timing diagram. The SSFP consists of rapid gradient echoes whose magnetization is preserved across multiple k-space lines, resulting in images with a high signal-to-noise ratio despite a short scan time. RF, radiofrequency; TR, repetition time.

One k-space line Acquiring each line takes ~3 msec (SSFP). Thus, 100 lines requires ~300 msec

y

x

Figure 2-6 Filling of raw data space in Cartesian fashion. The raw data space for CMR is often referred to as “k-space” and must be filled line by line to generate an image. Each line contains one echo. Each echo is acquired after application of a different phaseencoding step (direction y), and thus, each one contains different information. K-space is filled using, for example, the pulse sequence of Figure 2-5. Then the raw data undergo a twodimensional Fourier transformation, resulting in an image of, for example, Figure 2-4. SSFP, steady-state free precession.

echo sequence. HASTE is a special variant of the fast spin echo sequence in which only part of k-space is sampled to reduce acquisition time. The timing diagram for SSFP was described earlier (see Fig. 2-5). Typically, three stacks of HASTE and SSFP images are acquired, one for each orthogonal scanner plane (axial, coronal, and sagittal). Once the images are acquired, the scanner operator can then step through each series of image slices and 22 Cardiovascular Magnetic Resonance

understand the patient’s vascular anatomy and plan further scan planes. Cardiac gating for SSFP morphology images is essentially the same as that used for scout imaging: acquire the entire image in a single heartbeat during diastole (Fig. 2-11). For example, a set of 40 SSFP images for morphology would require 40 successive heartbeats. Cardiac gating for HASTE is essentially the same as for SSFP except that the images are typically acquired every other heartbeat (e.g., 80 total heartbeats are required to complete 40 images). The reason for this relates to the mechanism responsible for making the blood appear black in HASTE images. Morphology imaging typically takes 3 to 5 minutes and is acquired without any breath holding. For example, the acquisition of 40 slices with both SSFP and HASTE would take 40 þ 80 ¼ 120 heartbeats, or approximately 2 minutes. Black-blood imaging is often advantageous for the examination of morphology in that it allows one to clearly distinguish the inner portion of the vessel wall from the blood. In essence, black-blood HASTE can be achieved by carefully combining CMR physics with the physiology of rapidly moving blood. The underlying reason why the blood appears black in HASTE images is not reflected in the timing diagram of Figure 2-10. It is related to the preparatory radiofrequency (RF) pulses. The basic concept is seen in Figures 2-12 and 2-13. Soon after the R-wave, a brief (10 msec) nonselective 180 RF pulse is applied that causes all of the magnetization in the body to flip upside down (“inversion” pulse). Immediately afterward, a second sliceselective RF pulse is applied that causes only the magnetization within the to-be-imaged slice to flip back to where it started (“re-inversion” pulse). The net effect of these two RF pulses is that all of the protons outside of the slice have an inverted magnetization, whereas all protons within the to-be-imaged slice do not. At this point, the scanner simply waits several hundred milliseconds (see Fig. 2-13), and during this period, the heart contracts and pushes blood out of the slice (orange in Fig. 2-12). This blood is then replaced with fresh blood (green in Fig. 2-12) previously located outside the slice. The net effect is that, when the actual image acquisition stage begins in diastole, the fresh blood now in the slice “remembers” that it was inverted shortly after the R-wave. Its T1 recovery curve crosses zero (see Fig. 2-13), consequently produces no signal, and thus appears black. In summary, black-blood HASTE is made possible by a clever combination of a “memory” effect related to MR imaging physics (T1 recovery) as well as the physiology of moving blood. Such images are helpful in detecting abnormal large vessel anatomy, especially when combined with bright-blood SSFP images at the same location.

CONTRACTILE FUNCTION (“SCAN” ¼ “CINE”) Contractile function is a fundamental part of the CMR examination. Cine CMR is used for global and regional LV and right ventricular (RV) wall motion assessment. It is highly accurate and reproducible for ventricular volume, ejection fraction, and mass measurements.5–7 With the use of SSFP sequences, cine CMR has become widely

R

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Figure 2-8 The “goal” of “black-blood” half-Fourier single-shot fast spin echo imaging. Shown is a series of parallel images across the entire thoracic cavity that can then be inspected to assess the vascular anatomy. Usually, sagittal, axial, and coronal stacks are acquired.

accepted as the noninvasive gold standard for contractile function assessment. It has been used as an end point for the evaluation of LV remodeling8–10 and as a reference method for other imaging techniques.11–18 The “goal” of cine CMR is to capture a movie of the beating heart to visualize its contractile function. Figure 2-14

shows a representative example of a midventricular shortaxis slice during eight different time points within the cardiac cycle. Typically, 20 to 30 cine frames are acquired with 30- to 50-msec temporal resolution. Images are acquired with SSFP (see Fig. 2-5). The advantage of SSFP over other pulse sequences, such as the gradient-recalled echo (GRE) Cardiovascular Magnetic Resonance 23

2 CLINICAL CARDIOVASCULAR MAGNETIC RESONANCE IMAGING TECHNIQUES

Figure 2-7 Cardiac triggering for scout imaging consists of a delay after the electrocardiographic (ECG) R-wave and then rapid acquisition of the entire image (all k-space lines) in diastole. Imaging within a single heartbeat is made possible by the intrinsic speed of the steady-state free precession pulse sequence.

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

Left

Head

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Right

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Anterior

Sagittal

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Figure 2-9 The “goal” of “bright-blood” steady-state free precession imaging. The same image planes used for half-Fourier single-shot fast spin echo are acquired so that the two forms of images can be compared.

sequence, is its high SNR, fast speed, and excellent blood/ myocardium CNR that greatly facilitates the identification of the endocardial border. Typically, a stack of multiple closely spaced (contiguous or with 1- to 2-mm gaps) short-axis slices (6 to 8 mm thick) is acquired to provide full coverage of the left and right ventricles. Short-axis views, perpendicular to the long-axis views, can be planned on long-axis scout images, as described earlier. In addition, cine images can be obtained in multiple long-axis views, such as the two-chamber, three-chamber, or four-chamber orientations. A key consideration for cine imaging is that current CMR scanner hardware is not adequate to acquire cine images with high spatial resolution and high temporal resolution during a single cardiac cycle in real time. For example, for a cine with 30 frames, each with 96 lines, a total of 30 96 ¼ 2718 lines must be acquired. Considering that the acquisition time for one line is approximately 3 msec, the acquisition time to create a complete dataset of cine images would be 2718 3 msec ¼ 8154 msec, considerably longer than the typical R-R interval. Imaging speed could be reduced by acquiring fewer phase-encoding lines, but lower spatial resolution or lower temporal resolution would result. For cine imaging, each heartbeat is assumed 24 Cardiovascular Magnetic Resonance

to be “identical” to a subsequent heartbeat such that image quality can be improved by the use of k-space segmentation. Segmented k-space data acquisition allows one to collect only part of each movie frame over 8 to 10 consecutive cardiac cycles and then to combine the data to form a cine loop. The underlying concept is similar to that used in gated single photon emission computed tomography. Figure 2-15 shows the scheme used for CMR in which only a fraction of k-space for any given movie frame is acquired during any single heartbeat. This fraction is referred to as a “segment” (e.g., 10 k-space lines per segment in Fig. 2-15) and typically is adjusted by the scanner operator such that an adequate number of movie frames (generally 20 to 30) fit within the patient’s R-R interval. Accordingly, the full k-space data for any one movie frame actually consist of multiple segments acquired during successive cardiac cycles. The data are then automatically reassembled during image reconstruction to form a movie showing a single cardiac cycle. For example, the scheme in Figure 2-15 acquires 30 segments of 10 lines each in every cardiac cycle. Assuming that the scanner operator has chosen to obtain 90 k-space lines for each movie frame, the scan will acquire data across 90/10 ¼ 9 heartbeats. For example, the

180°

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3rd echo

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Figure 2-10 Pulse sequence timing diagram for (A) a fast spin echo sequence and (B) conventional spin echo. In essence, the fast spin echo consists of a series of spin echoes. Additional postprocessing steps are required to obtain the image. The spin echo sequence (B) is the predecessor of the fast spin echo sequence (A). The blood appears black in the resulting images because of the preparatory pulse described in the text and in Figures 2-12 and 2-13. RF, radiofrequency; TE, echo time; TR, repetition time.

k-space slice 1

MYOCARDIAL PERFUSION AT STRESS AND REST (“SCAN” ¼ “PERFUSION”)

k-space slice 2 Heartbeat 1 Heartbeat 2

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Trigger

Figure 2-11 Electrocardiogram (ECG) triggering with each image acquired in a single diastole, similar to scout imaging. To skip systole, a delay time is inserted after the R-wave is detected.

green segment is always acquired immediately after the R-wave and is used to reconstruct the first movie frame, the orange segment is acquired next and becomes the second movie frame, and so on. In practice, cine imaging is performed during breath holding and takes 10 to 12 seconds for each cardiac slice acquisition (e.g., one short-axis cine). Cine imaging can be performed at rest and with pharmacologic stress with low-dose dobutamine (viability) or graded high-dose dobutamine (stress).19,20 For these higher heart rate scans, high temporal resolution is especially important.

Encouraged by a number of clinical studies,2,21–23 adenosine stress/rest myocardial perfusion CMR has gained clinical attention. Perfusion CMR accurately diagnoses coronary artery disease with high sensitivity and specificity. Rest perfusion CMR, in combination with LGE, is important for distinguishing true perfusion defects on stress images from artifacts. The subject of artifacts is discussed elsewhere.2 In the current protocol, the rest myocardial perfusion scan is routinely performed in all patients who receive gadolinium contrast. The “goal” of myocardial perfusion scanning is to create a movie showing the wash-in of contrast media (typically gadolinium-based) with the blood during its initial pass through the myocardium (“first-pass perfusion”). The CMR pulse sequence that is most commonly used is a GRE sequence. Figure 2-16 shows an example of an adenosine stress/rest perfusion scan in a patient with a stress-induced perfusion defect. The blue slice is shown at three representative time points: (1) before contrast arrival; (2) at the time of contrast arrival in the right ventricle; and (3) shortly after contrast arrival in the LV. During adenosine stress, perfusion defects appear as dark regions (i.e., hypoenhancement) surrounded by bright contrastenhanced normally perfused myocardium. In the corresponding rest perfusion images, perfusion to the hypoenhanced areas may be relatively preserved and thus appear Cardiovascular Magnetic Resonance 25

2 CLINICAL CARDIOVASCULAR MAGNETIC RESONANCE IMAGING TECHNIQUES

90°

Systole

R-wave (end-diastole)

Diastole

Nonselective inversion Slice-selective “re-inversion” Imaged slice Figure 2-12 Physiology of “black-blood” fast spin echo imaging. Immediately after electrocardiogram trigger, a nonselective inversion (green) is applied, followed instantly by a slice-selective re-inversion (orange). The prepared slice (orange) changes shape and position during cardiac contraction, but returns to its original geometry during diastole, when the blood signal is about zero (leading to black blood in the image) and the heart has little motion. Nonblack re-inverted blood (orange) is expelled from the slice during systole. During diastole, an image is obtained from a slice (blue) that is thinner than and lies inside the prepared slice (orange). ECG, electrocardiogram.

• HASTE = Half-Fourier Single-Shot Turbo Spin Echo • Typically used with a dark blood pulse.

ECG

All lines acquired for one complete image.

Noninverted myocardium Selective re-inversion

Nonselective inversion

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

ECG

90180 180 180 180 180 180 180

Time needed to null blood Time d

te nver

i wing Inflo

d bloo

Figure 2-13 Cardiac triggering for “black-blood” fast spin echo imaging. Image acquisition starts when the magnetic relaxation curve of the inverted blood in the cavity is passing through zero (i.e., when the blood is black). The center lines of k-space are acquired first (“centric reordering”). These lines contain information about image brightness and contrast. Hence, they are acquired when the blood is black, at the beginning of the acquisition train.

26 Cardiovascular Magnetic Resonance

1

2

3

4 Late diastole

5

6

7

8

Figure 2-14 The “goal” of cine CMR. A typical midventricular series shows short-axis cine steady-state free precession images at the level of the papillary muscles. The cine series often contains 20 to 30 images. Eight are shown here because of space constraints.

Figure 2-15 Cardiac gating for cine imaging. Raw data lines are acquired in segments over the course of six heartbeats. The segments are then sorted into the k-spaces of the 30 movie frames. Multiple lines form one segment, multiple segments form one k-space, and Fourier transformation of one k-space yields one movie frame. ECG, electrocardiogram.

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30 segments acquired each heartbeat

Each segment contains 10 lines

10 lines

Ordered after cardiac phase: 1

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normal, indicating normal resting perfusion. The reduced coronary flow reserve is caused by hemodynamically significant coronary artery stenosis. Perfusion imaging data are acquired continuously (with each R-R interval) rather than in the segmented manner previously described for cine CMR. Each single image is acquired within 100 to 150 msec. Because of R-R interval

30

Cardiac phase

time constraints, perfusion imaging is typically performed throughout the entire cardiac cycle. All k-space lines for each of four to five short-axis slice locations are acquired during each cardiac cycle (Fig. 2-17) to characterize the first pass of the contrast agent throughout the entire LV myocardium. Because of time constraints, each anatomic level is acquired at a different phase of the cardiac cycle. Cardiovascular Magnetic Resonance 27

2 CLINICAL CARDIOVASCULAR MAGNETIC RESONANCE IMAGING TECHNIQUES

Early systole

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

Stress perfusion

slice 1

slice 2

slice 3

slice 4 Time slice 1

slice 2

slice 3

Rest perfusion

slice 4

Figure 2-16 The “goal” of perfusion imaging. Stress (top) and rest (bottom) perfusion images are shown. In this example, four slices (red, green, blue, and yellow) are acquired during each R-R interval. An apical slice (blue) is shown at different time points throughout image acquisition (3 of 50 time frames are shown: before contrast arrival, at the time of contrast arrival in the right ventricle, and shortly after contrast arrival in the left ventricle). Two stress perfusion defects appear as dark regions (hypoenhancement, arrows). They are surrounded by bright contrast-enhanced normally perfused myocardium. The defects are reversible because they are not present (no hypoenhancement, arrows) during rest perfusion.

The need to acquire multiple images within a single cardiac cycle is the primary reason why image quality is diminished for perfusion imaging. The total number of slices that can be acquired within one cardiac cycle is limited by the R-R interval of the patient. For example, for a heart rate of 80 bpm, the R-R interval is approximately 750 msec, and five short-axis slices, each acquired within 150 msec, could theoretically be acquired. For higher heart rates, the number of short-axis slice locations would be reduced. In addition, the heart rate may increase somewhat during adenosine infusion, so the scanner operator must allow some extra time within the resting R-R interval to account for this possibility. Altogether, perfusion imaging is typically performed for 40 to 50 seconds. To enhance the differences in image intensity between poorly and normally perfused myocardium, a saturation pulse (90 pulse) is applied as preparation before each image acquisition (Fig. 2-18). This 90 pulse tips the longitudinal magnetization into the transverse plane. The transverse magnetization then is dephased (spoiled) by a spoiler gradient. The contrast agent (e.g., gadolinium-diethyl triamine pentaacetic acid [DTPA]) shortens T1, recovery of the longitudinal magnetization, with normal perfusion and higher contrast concentration (relatively shorter T1) 28 Cardiovascular Magnetic Resonance

than in regions with reduced perfusion and lower contrast concentration (relatively longer T1, but still shorter than T1 without contrast). Figure 2-18 shows the relaxation curves for both regions. The timing between the saturation pulse and image read-out strongly influences the difference in image intensity (contrast) between normal and hypoperfused myocardium. The optimal time for image acquisition is shown at location “b” in Figure 2-18. In practice, however, the time between saturation pulse and the first slice acquisition is less than optimal (e.g., location “a”) to allow acquisition of multiple slices within each cardiac cycle. The basic concept of GRE timing is shown in Figure 2-19. The pulse sequence starts with a slice-selective gradient on the slice-encoded axis and an RF pulse. The RF pulse tips the longitudinal magnetization by the flip angle a into the transverse plane to make it available for reception by CMR coils. The flip angle is typically set to 20 to 30 , which is nearly optimal (“Ernst angle”) in the setting of the rapid train of RF pulses used in GRE imaging. The underlying principle of GRE is that the spins are dephased by a first gradient and then rephased by a second gradient with opposite polarity to form a gradient echo, which is then recorded by the CMR receiver coils. As soon as data collection has finished, spoiler gradients are applied to destroy the remaining transverse magnetization

Acquire multiple entire slices during each heartbeat R ECG Trigger

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2

3

4

5

Beat 2

Beat 3

Beat 4

Beat 5

Slice 1 Slice 2 Slice 3 Slice 4 All lines for each image (one slice and one time point) are acquired in a single shot.

Figure 2-18 Magnetization recovery. The recovery of the longitudinal magnetization is faster in myocardium with normal perfusion because of T1 shortening as a consequence of high-contrast agent concentration. Relatively speaking, T1 is prolonged in myocardium with reduced perfusion. To maximize the difference in signal intensity between myocardium with normal and reduced perfusion, the optimal time for image acquisition is when the difference between both curves is largest (b).

Mz

Myocardium with normal perfusion has shorter T1 (recovers faster, looks brighter)

+M0

c b

a

Myocardium with relatively reduced perfusion has longer T1 (recovers slower, looks darker) Time

90° RF Gradient Spoil transverse magnetization

before the next RF excitation pulse is applied. This is important for avoiding image artifacts that can result from mixing of longitudinal and transverse magnetization. For stress myocardial perfusion imaging, a dose of adenosine 140 mg/kg/min is typically used for at least 2 minutes. The adenosine causes coronary vasodilation mediated by the A2A receptor24 and has a very short half-life of less than 10 seconds. Adenosine acts to maximally dilate the

Readout one entire image (when curves have maximum separation)

distal arteriolar bed. In the absence of epicardial stenoses, adenosine can dilate these blood vessels and increase coronary blood flow four- to fivefold.25 For epicardial coronary arteries with significant stenoses, however, these arterioles are already fully dilated (in the absence of adenosine) because of the autoregulation mechanism that aims to restore normal flow impeded by a stenosis. Thus, further vasodilation with adenosine administration does not occur. Cardiovascular Magnetic Resonance 29

2 CLINICAL CARDIOVASCULAR MAGNETIC RESONANCE IMAGING TECHNIQUES

Figure 2-17 Image acquisition in the perfusion scan. Images are acquired in a continuous stream (single shot) over 50 to 60 heartbeats. Depending on the duration of the R-R interval and cardiovascular magnetic resonance scanner hardware, up to five slices can be acquired during each R-R interval.

RF

Slice

Destroy transverse magnetization to avoid mixing of longitudinal and transverse magnetization

A A A Spoiling

Phase Phase encoder B

Read

B

B

ADC Dephasing Signal Dephasing

Time

Although symptoms such as chest discomfort can occur, serious side effects (e.g., bronchospasm or atrioventricular block) are uncommon. (See Chapters 16 and 17.) Figure 2-20A shows the timeline for a comprehensive CMR stress test. After scout and resting cine imaging, adenosine is infused for at least 2 minutes. This minimum infusion duration is chosen based on physiologic studies in humans, showing that, on average, maximum coronary

0

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The resulting heterogeneity in blood flow between regions supplied by coronary arteries with stenoses (little or no flow increase) and territories supplied by normal epicardial arteries (multifold increased flow) is observed as a difference in myocardial contrast agent concentration and thus image intensity. Regional myocardium supplied by the diseased vessels results in a hypoenhancement pattern (see Fig. 2-16) compared with normally perfused myocardium.

Contrast injection 2

Refocusing

IMAGING

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

Figure 2-19 Pulse sequence timing diagram for gradient-recalled echo imaging. ADC, analog digital converter; RF, radiofrequency.

Rephaser to keep spins aligned

α

25

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LGE imaging 40

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Saline flush Breath holding Image acquisition

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Figure 2-20 Timeline for a comprehensive cardiovascular magnetic resonance stress test. Time points 1 to 6 are referenced in the text. The time scale is not linear. LGE, late gadolinium enhancement. 30 Cardiovascular Magnetic Resonance

contrast agent concentration, and therefore hyperenhancement. Chronic infarcts consist of dense scar with an increased interstitial space between collagen fibers (relative to normal myocardium). This leads to increased distribution volume of the gadolinium-DTPA and consequently increased concentration and resultant hyperenhancement. The “goal” of LGE imaging is to create images with high contrast between the hyperenhanced nonviable tissue and normal myocardium for a clear delineation of the regions (Fig. 2-21). This is currently best achieved by using a segmented inversion recovery GRE sequence.27–30,38,39 Figure 2-22 shows the gating for a segmented inversion recovery sequence. The acquisition of multiple k-space lines in each cardiac cycle allows reductions in imaging times to the point where the entire image can be acquired during a single breath hold of 6 to 10 seconds. The images are acquired in mid-diastole by using a trigger delay to minimize cardiac motion. The magnetization of the heart is prepared by a nonselective 180 inversion pulse to create T1 weighting. The inversion time delay between inversion pulse and data collection (more precisely, the center of kspace) is chosen such that the magnetization of viable (normal) myocardium is near its zero crossing, meaning that these regions appear dark (Fig. 2-23). Nonviable myocardium (acutely infarcted or scar), however, appears bright because of the shorter T1 (faster signal recovery curve in Fig. 2-23) after contrast administration. It is important to adjust the inversion time manually for each image to null normal myocardium to account for washout kinetics of the contrast agent.40 At some centers, free breathing 3D acquisitions have displaced 2D breath hold LGE due to the opportunity for superior spatial resolution.40a

VIABILITY AND INFARCTION (“SCAN” ¼ “LGE”) The clinical implications of viability imaging are steadily growing. Beside the assessment of acute and chronic myocardial infarction, viability imaging27–30 has been used for the prediction of contractile improvement after revascularization,31 for measuring the response to beta-blocker treatment,32 for the differentiation of ischemic versus nonischemic cardiomyopathy,33 and for the diagnosis of various nonischemic cardiomyopathies.34–38 A large body of evidence shows that LGE can differentiate between nonviable and viable tissue.27–30,38,39 Regions of irreversible injury exhibit high signal intensity (hyperenhancement) on T1-weighted images after administration of extracellular CMR contrast agent, such as gadoliniumDTPA. The contrast agent significantly shortens local longitudinal relaxation time, resulting in signal increase. The underlying mechanism of hyperenhancement continues to be the subject of debate. The most likely explanation is that in acutely infarcted regions, the ruptured myocyte membranes allow the extracellular contrast agent to diffuse passively into the intracellular space, resulting in increased tissue-level

Figure 2-21 The “goal” of late gadolinium enhancement viability imaging. The image provides a clear delineation between nonviable (hyperenhanced) and viable (dark) myocardium in this short-axis image. Cardiovascular Magnetic Resonance 31

2 CLINICAL CARDIOVASCULAR MAGNETIC RESONANCE IMAGING TECHNIQUES

blood flow is reached 1 minute after the start of adenosine infusion, and in virtually every patient, it is reached after 2 minutes.26 Then a CMR contrast agent, such as gadolinium-DTPA (0.04-0.075 mmol/kg), is rapidly administered (3.5 mL/sec), followed by a (50 mL) saline flush. The adenosine infusion is stopped as soon as image acquisition is completed. Figure 2-20B shows the specific steps performed during the stress perfusion scan in more detail: (1) Adenosine infusion is started. (2) Two minutes later, image acquisition commences. (3) Contrast agent is injected at a rate of 3.5 mL/sec while adenosine infusion continues (two different intravenous lines are therefore required to avoid administering an adenosine bolus and promoting heart block). The patient still breathes freely. (4) As soon as the contrast medium reaches the RV, the patient is asked to suspend respiration for approximately 15 seconds to characterize the initial wash-in of contrast agent into the LV myocardium. (5) The remaining image acquisition is done during shallow breathing. (6) Image acquisition ends approximately 50 to 60 seconds after it began. This description assumes that images are displayed on the scanner in real time to allow observation of contrast agent arrival in the RV. If the scanner does not provide for real-time image display, breath hold (step 4) should be started five to six heartbeats after initiation of the contrast agent injection. Rest myocardial perfusion is then performed. However, before the rest perfusion scan, a 15-minute delay is required for the contrast agent to clear sufficiently from the blood pool. During this time, additional cine scans or flow imaging for valvular evaluation can be performed. For the rest perfusion scan, an additional dose of gadolinium-DTPA 0.05 to 0.075 mmol/kg is given. LGE imaging may then be performed 10 minutes after the completion of the second perfusion study. The total scan time for a comprehensive CMR stress test, including cines, stress/rest perfusion, and LGE, is approximately 45 minutes.

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

R

Figure 2-22 Electrocardiogram (ECG) triggering of the inversion recovery segmented gradient-recalled echo sequence.

R

R Heartbeat 2

Heartbeat 1 ECG Trigger TI

TI

Segment 1 180° inversion

Segment 2 180° inversion

k-space

Mz

Figure 2-23 Inversion recovery curve for viable and nonviable myocardium. RF, radiofrequency.

Nonviable (infarcted) myocardium has shorter T1 (recovers faster, looks brighter)

TI (normal) +M0 TI (infarct)

Time

180°

Viable (normal) myocardium has longer T1 (recovers slower, looks darker) –M0

TI = inversion time

RF

Gradient Spoil transverse magnetization Readout one segment (when normal myocardium is nulled)

FLOW/VELOCITY IMAGING (“SCAN” ¼ “FLOW/ VELOCITY”) Velocity-encoded (VENC) CMR imaging of blood flow is usually performed to measure velocity in arteries, veins, valves, and shunts. The overall goal is to evaluate the severity of valve regurgitation/stenosis or an intracardiac shunt. Further, LV and RV stroke volume can be calculated by summing up the area under the flow/velocity curve of the ascending aorta and pulmonary artery, respectively. With VENC CMR imaging, a cine series of grayscale images reflecting flow during the different phases of the cardiac cycle is acquired. The gray level is proportional to the velocity of blood into or out of the measured plane. 32 Cardiovascular Magnetic Resonance

Both in-plane and through-plane flow can be assessed. Figure 2-24 shows such a series in the upper row referenced to the ECG. The corresponding anatomic cine frames are seen in the lower row. The VENC image shows bright image intensity where fast velocity is present and gray intensity where minimal flow is present (background and parts of the leaflets that do not open). Flow in the opposite direction would show up in levels darker than the background (black to dark gray). Analogous to cine imaging, each VENC image corresponds to a cardiac phase, and ECG gating is required. The sequence is run in retrogated mode to cover the entire cardiac cycle. This is important for the determination of cardiac output, because skipping the last frame would miss the beginning of systole and hence deliver erroneous volumes.

ECG

Velocity-encoded images

Cine images at corresponding times Figure 2-24 The “goal” of flow (velocity-encoded) imaging. A selection of velocity-encoded images of the aortic valve exhibiting stenosis is shown in the middle. Images are registered with the electrocardiogram (ECG) on top, and corresponding cine frames on the bottom are included as anatomic reference.

The physics of VENC are complex.41,42 Simply put, in the presence of a magnetic field gradient, the spins in the image plane acquire a phase that is proportional to the area under the gradient plotted in a pulse sequence diagram, such as the one shown in Figure 2-25. For nonmoving

Figure 2-25 Pulse sequence timing diagram of the electrocardiogramtriggered velocity-encoded sequence. Gradients on the read axis are flowcompensated. On the slice axis, two gradient waveforms are shown inside the dotted ellipses. One is insensitive to flow (“flow-compensated block”) and one is sensitive toward flow (“flow-sensitized block”). The sequence necessary to acquire one line of data is shown. This is an example for assessment of throughplane flow, but in-plane flow can be imaged as well.

spins, this phase can be rewound (the spins can be rephased) by a gradient of the same magnitude, but the opposite sign. If, however, the spins change position between the first gradient and the rewinding gradient, as is the case for spins in flowing blood, then they experience

ADC Time

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Flow-compensated block

Flow-sensitive block

Cardiovascular Magnetic Resonance 33

2 CLINICAL CARDIOVASCULAR MAGNETIC RESONANCE IMAGING TECHNIQUES

R

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

a gradient during rewinding that is not only opposite in sign, but also different in magnitude. This incomplete or excessive rewinding causes a phase difference that is a linear function of velocity. Displaying this phase as image 1 represents velocity as (grayscale) image. Because the phase of spins is also dependent on many other factors (e.g., the properties of receiver coils), it is necessary to measure a base phase during flow-insensitive gradient waveforms (see Fig. 2-25, left dotted oval). The flow-sensitized waveform is shown in the right dotted oval. By subtracting the base phase from the flow phase and displaying the difference on a grayscale, the images of Figure 2-24, upper row, are obtained. The gray level is proportional to velocity: black corresponds to a maximum backward flow, white corresponds to a maximum forward flow, and gray corresponds to no flow. The range from maximum backward flow to maximum forward flow can be quantified in centimeters per second and must be adjusted by the scanner operator to maximize sensitivity to differing physiologic blood flow velocities. The images allow both a coarse overview and a quantitative assessment of abnormal flow patterns. Typically, flow is acquired in the axial plane at the level of the bifurcation of the pulmonary artery and in the near coronal plane in the proximal pulmonary artery. From these data, LV forward stroke volume, aortic regurgitation, RV forward stroke volume, and pulmonic regurgitation can be readily measured and the pulmonic:systemic (Qp:Qs) ratio can also be determined. Mitral and tricuspid regurgitation can then be calculated as the difference between the respective stroke volume (from the short-axis stack of cine images) and the forward stroke volume (see Chapter 37).

ANGIOGRAPHY (“SCAN” ¼ “ANGIOGRAPHY”) Vascular angiography is often performed independently of CMR stress or viability tests. However, this scan is frequently combined with VENC imaging. Occasionally, it is done in

conjunction with a more detailed cardiovascular examination (e.g., pulmonary vein imaging, aortic root angiography). This section discusses some practical issues with regard to ECGgated contrast-enhanced MR angiography (CE-MRA). This chapter does not discuss different angiography techniques, vessel wall or coronary imaging, or advanced forms of motion correction (“navigator echoes”). The “goal” of CE-MRA is to create a three-dimensional data set of a vessel or vascular bed of interest. Use of a T1shortening paramagnetic contrast agent within the blood pool allows the generation of high signal within the vascular tree relative to surrounding tissue, thus creating a “luminogram.” Figure 2-26 shows two examples of CE-MRA. The images were colored during routine postprocessing. Figure 2-26A shows the left ventricle, the aortic arch, the descending aorta, the truncus brachiocephalicus, and carotid and subclavian arteries. Figure 2-26B shows pulmonary arteries and veins. The timing of the image acquisition was specifically chosen such that the LV and aorta are not yet filled by contrast-rich blood and thus are not visible. This “procedure” would commonly be done as part of a complete evaluation of a patient with vascular disease that would include multiple different scans from the exam menu, in addition to the CE-MRA. Figure 2-27 shows the timeline of the procedure. The noncontrast scan is done for later subtraction from the contrast-enhanced scan to eliminate background. Unlike for lower-extremity MRA, ECG gating is critical for imaging of vascular structures that experience motion as a result of cardiac movement or pulsatile blood flow (e.g., aortic root). In these situations, it is important to acquire imaging data only during diastole (see Fig. 2-7) because motion, even during just a portion of the acquisition, will adversely affect image quality. In the case of highly dynamic structures, such as the ascending aorta, it is important to account for not only the systolic cycle length, but also the period in early diastole when passive relaxation of the aorta occurs because of the elastic recoil of the vessel. This short period of “diastolic vessel relaxation” leads to appreciable vascular motion that can blur the edges of an angiogram and compromise its quality. These considerations were taken into account for obtaining

Ao

LA RV

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B

Figure 2-26 Three-dimensional contrast-enhanced magnetic resonance angiography maximum intensity projection of the aortic root and thoracic aorta. This image was electrocardiogram-triggered for high resolution imaging of the aortic root in a young woman with a history of a bicuspid aortic valve and aortic root dilation. A, anterior view; B, posterior view. Ao, ascending aorta; LA, left atrium; RV, right ventricle. 34 Cardiovascular Magnetic Resonance

Contrast injection

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Figure 2-27 Timeline of a typical contrast-enhanced magnetic resonance angiography (CE-MRA). The time scale is not linear. In addition to scouts and fast spin echo imaging, selected “bright-blood” (steady-state free precession) images are obtained for analysis of extravascular structures and vascular structures that may not be in the angiographic field of view. A noncontrast three-dimensional angiogram is performed before the CE-MRA so that a subtraction image can be generated. A comprehensive study can be completed in 20 to 25 minutes. GRE, gradient-recalled echo; HASTE, half-Fourier single-shot fast spin echo.

the images of Figure 2-26 because they would not be as crisp otherwise. For CE-MRA, the most commonly used pulse sequence is a fast three-dimensional spoiled GRE. This is a variant of the GRE technique that was discussed previously (see Fig. 2-19). The need for three-dimensional data requires application of an additional gradient to encode the additional dimension. This gradient is played in the slice-encoding direction. Compared with two-dimensional GRE, a much larger amount of data is acquired, leading to longer scan times. The use of advanced parallel image acquisition and faster gradient sets has decreased scan times significantly. With the use of modern hardware and sequences, high-resolution ECG-gated CE-MRA can be completed in 10 to 20 seconds. The most common type of angiography used on the clinical service, CE-MRA is easily incorporated into the clinical CMR examination without dramatically prolonging the duration of the study or compromising the quality of the other components.

CONCLUSION This chapter described an approach to routine clinical CMR that is based on selecting one or more scans from an exam menu and assembling these into a test. Figure 2-28 shows how one would assemble a test for a patient referred for the evaluation of coronary artery disease. Specifically, one would perform scouting, cine imaging, perfusion imaging at stress and rest, and finally late gadolinium enhancement imaging. The protocol of Figure 2-28 is identical to the CMR stress test2 discussed earlier. Importantly, and as previously noted, this test now accounts for approximately 50% of the 3000 annual CMR procedures performed at the DCMRC. Similarly,

Scout

LV cine

Perfusion

LGE

Figure 2-28 Combination of menu items for the evaluation of coronary artery disease. LGE, late gadolinium enhancement; LV, left ventricle.

Scout

Morphology

Angiography

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Figure 2-29 Combination of menu items for evaluation of the aorta.

Figure 2-29 summarizes how one would approach CMR scanning of a patient referred for the evaluation of aortic disease. In summary, although the definition of “routine clinical CMR” continues to evolve at the DCMRC and other institutions,1 it is very useful to assemble a short list of predefined individual CMR scan protocols (see Fig. 2-3) from which one can create a “test package” specifically tailored to the diagnostic question. This approach substantially decreases the time needed to scan because the operator does not have to select from among the literally hundreds of buttons on the scanner console while the patient waits idly inside the magnet. Instead, the buttons are predefined for each menu item. This approach improves the throughput of the DCMRC clinical CMR service. Perhaps more importantly, however, this approach dramatically reduces the historically steep learning curve for new physicians interested in learning how to perform CMR. Cardiovascular Magnetic Resonance 35

2 CLINICAL CARDIOVASCULAR MAGNETIC RESONANCE IMAGING TECHNIQUES

≈ 3 min

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References 1. Bruder O, Schneider S, Nothnagel, et al. Euro CMR (European Cardiovascular Magnetic Resonance) registry: results of the pilot phase. J Am Coll Cardiol. 2009;54:1457–1566. 2. Klem I, Heitner JF, Shah DJ, et al. Improved detection of coronary artery disease by stress perfusion cardiovascular magnetic resonance with the use of delayed enhancement infarction imaging. J Am Coll Cardiol. 2006;47:1630–1638. 3. Oppelt A, Graumann R, Barfuss H, et al. Eine neue schnelle pulssequenz fuer die kernspintomographie. Electromedica. 1986;54:15–18. 4. Fischer SE, Wickline SA, Lorenz CH. Novel real-time R-wave detection algorithm based on the vector cardiogram for accurate gated magnetic resonance acquisitions. Magn Reson Med. 1999;361–370. 5. Bellenger NG, Burgess MI, Ray SG, et al. Comparison of left ventricular ejection fraction and volumes in heart failure by echocardiography, radionuclide ventriculography and cardiovascular magnetic resonance: are they interchangeable? Eur Heart J. 2000;21:1387–1396. 6. Grothues F, Moon JC, Bellenger NG, et al. Interstudy reproducibility of right ventricular volumes, function, and mass with cardiovascular magnetic resonance. Am Heart J. 2004;147:218–223. 7. Grothues F, Smith GC, Moon JC, et al. Comparison of interstudy reproducibility of cardiovascular magnetic resonance with two-dimensional echocardiography in normal subjects and in patients with heart failure or left ventricular hypertrophy. Am J Cardiol. 2002;90:29–34. 8. Bellenger NG, Rajappan K, Rahman SL, et al. Effects of carvedilol on left ventricular remodelling in chronic stable heart failure: a cardiovascular magnetic resonance study. Heart. 2004;90:760–764. 9. Friedrich MG, Dahlof B, Sechtem U, et al. Reduction (TELMAR) as assessed by magnetic resonance imaging in patients with mild-to-moderate hypertension: a prospective, randomised, double-blind comparison of telmisartan with metoprolol over a period of six months rationale and study design. J Renin Angiotensin Aldosterone Syst. 2003;4:234–243. 10. Lamb HJ, Beyerbacht HP, de Roos A, et al. Left ventricular remodeling early after aortic valve replacement: differential effects on diastolic function in aortic valve stenosis and aortic regurgitation. J Am Coll Cardiol. 2002;40:2182–2188. 11. Ioannidis JP, Trikalinos TA, Danias PG. Electrocardiogram-gated single-photon emission computed tomography versus cardiac magnetic resonance imaging for the assessment of left ventricular volumes and ejection fraction: a meta-analysis. J Am Coll Cardiol. 2002; 39:2059–2068. 12. Kjaergaard J, Petersen CL, Kjaer A, et al. Evaluation of right ventricular volume and function by 2D and 3D echocardiography compared to MRI. Eur J Echocardiogr. 2006;7:430–438. 13. Persson E, Carlsson M, Palmer J, et al. Evaluation of left ventricular volumes and ejection fraction by automated gated myocardial SPECT versus cardiovascular magnetic resonance. Clin Physiol Funct Imaging. 2005;25:135–141. 14. Schaefer WM, Lipke CS, Standke D, et al. Quantification of left ventricular volumes and ejection fraction from gated 99mTc-MIBI SPECT: MRI validation and comparison of the Emory Cardiac Tool Box with QGS and 4D-MSPECT. J Nucl Med. 2005;46:1256–1263. 15. Thorley PJ, Plein S, Bloomer TN, et al. Comparison of 99mTc tetrofosmin gated SPECT measurements of left ventricular volumes and ejection fraction with MRI over a wide range of values. Nucl Med Commun. 2003;24:763–769. 16. Corsi C, Lang RM, Veronesi F, et al. Volumetric quantification of global and regional left ventricular function from real-time three-dimensional echocardiographic images. Circulation. 2005;112:1161–1170. 17. Dewey M, Muller M, Eddicks S, et al. Evaluation of global and regional left ventricular function with 16-slice computed tomography, biplane cineventriculography, and two-dimensional transthoracic echocardiography: comparison with magnetic resonance imaging. J Am Coll Cardiol. 2006;48:2034–2044. 18. Fischbach R, Juergens KU, Ozgun M, et al. Assessment of regional left ventricular function with multidetector-row computed tomography versus magnetic resonance imaging. Eur Radiol. 2007;17:1009–1017. 19. Potter DD, Araoz PA, McGee KP, et al. Low-dose dobutamine cardiac magnetic resonance imaging with myocardial strain analysis predicts myocardial recoverability after coronary artery bypass grafting. J Thorac Cardiovasc Surg. 2008;135:1342–1347. 20. Jahnke C, Nagel E, Gebker R, et al. Prognostic value of cardiac magnetic resonance stress tests: adenosine stress perfusion and dobutamine stress wall motion imaging. Circulation. 2007;115:1769–1776.

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21. Giang TH, Nanz D, Coulden R, et al. Detection of coronary artery disease by magnetic resonance myocardial perfusion imaging with various contrast medium doses: first European multi-centre experience. Eur Heart J. 2004;25:1657–1665. 22. Paetsch I, Jahnke C, Wahl A, et al. Comparison of dobutamine stress magnetic resonance, adenosine stress magnetic resonance, and adenosine stress magnetic resonance perfusion. Circulation. 2004;110: 835–842. 23. Wolff SD, Schwitter J, Coulden R, et al. Myocardial first-pass perfusion magnetic resonance imaging: a multicenter dose-ranging study. Circulation. 2004;110:732–737. 24. Udelson JE, Heller GV, Wackers FJ, et al. Randomized, controlled dose-ranging study of the selective adenosine A2A receptor agonist binodenoson for pharmacological stress as an adjunct to myocardial perfusion imaging. Circulation. 2004;109:457–464. 25. Wilson RF, Wyche K, Christensen BV, et al. Effects of adenosine on human coronary arterial circulation. Circulation. 1990;82:1595–1606. 26. Rossen JD, Quillen JE, Lopez AG, et al. Comparison of coronary vasodilation with intravenous dipyridamole and adenosine. J Am Coll Cardiol. 1991;18:485–491. 27. Kim RJ, Fieno DS, Parrish TB, et al. Relationship of MRI delayed contrast enhancement to irreversible injury, infarct age, and contractile function. Circulation. 1999;100:1992–2002. 28. Wu E, Judd RM, Vargas JD, et al. Visualisation of presence, location, and transmural extent of healed Q-wave and non-Q-wave myocardial infarction. Lancet. 2001;357:21–28. 29. Fieno DS, Kim RJ, Chen EL, et al. Contrast-enhanced magnetic resonance imaging of myocardium at risk: distinction between reversible and irreversible injury throughout infarct healing. J Am Coll Cardiol. 2000;36:1985–1991. 30. Wagner A, Mahrholdt H, Holly TA, et al. Contrast-enhanced MRI and routine single photon emission computed tomography (SPECT) perfusion imaging for detection of subendocardial myocardial infarcts: an imaging study. Lancet. 2003;361:374–379. 31. Kim RJ, Wu E, Rafael A, et al. The use of contrast-enhanced magnetic resonance imaging to identify reversible myocardial dysfunction. N Engl J Med. 2000;343:1445–1453. 32. Bello D, Shah DJ, Farah GM, et al. Gadolinium cardiovascular magnetic resonance predicts reversible myocardial dysfunction and remodeling in patients with heart failure undergoing beta-blocker therapy. Circulation. 2003;108:1945–1953. 33. McCrohon JA, Moon JC, Prasad SK, et al. Differentiation of heart failure related to dilated cardiomyopathy and coronary artery disease using gadolinium-enhanced cardiovascular magnetic resonance. Circulation. 2003;108:54–59. 34. Mahrholdt H, Wagner A, Judd RM, et al. Delayed enhancement cardiovascular magnetic resonance assessment of non-ischaemic cardiomyopathies. Eur Heart J. 2005;26:1461–1474. 35. Mahrholdt H, Goedecke C, Wagner A, et al. Cardiovascular magnetic resonance assessment of human myocarditis: a comparison to histology and molecular pathology. Circulation. 2004;109:1250–1258. 36. Choudhury L, Mahrholdt H, Wagner A, et al. Myocardial scarring in asymptomatic or mildly symptomatic patients with hypertrophic cardiomyopathy. J Am Coll Cardiol. 2002;40:2156–2164. 37. Simonetti OP, Kim RJ, Fieno DS, et al. An improved MR imaging technique for the visualization of myocardial infarction. Radiology. 2001; 218:215–223. 38. Mahrholdt H, Klem I, Sechtem U. Cardiovascular MRI for detection of myocardial viability and ischaemia. Heart. 2007;93:122–129. 39. Weinsaft JW, Klem I, Judd RM. MRI for the assessment of myocardial viability. Magn Reson Imaging Clin N Am. 2007;15:505–525, v–vi. 40. Wagner A, Mahrholdt H, Thomson L, et al. Effects of time, dose, and inversion time for acute myocardial infarct size measurements based on magnetic resonance imaging-delayed contrast enhancement. J Am Coll Cardiol. 2006;47:2027–2033. 40a. Peters DC, Appelbaum E, Nezafat R, et al. Left ventricular infarct size, peri-infarct zone and papillary scar measurements: a comparison of high resolution 3D and conventional 2D late gadolinium enhancement cardiac MR. J Magn Reson Imaging. 2009;30:794–800. 41. Debatin JF, Ting RH, Wegmu¨ller H. Renal artery blood flow: quantitation with phase-contrast MR imaging with and without breath holding. Radiology. 1994;190:371–378. 42. Pelc LR, Pelc NJ, Rayhill SC, et al. Arterial and venous blood flow: noninvasive quantitation with MR imaging. Radiology. 1992;185: 809–812.

Advanced Cardiovascular Magnetic Resonance Imaging Techniques: Spiral, Radial, and Parallel Imaging Daniel K. Sodickson and Thoralf Niendorf

This chapter is concerned with the basic principles and cardiovascular applications of spiral imaging, radial imaging, and parallel imaging. Arguably, all three approaches earn the moniker of “advanced cardiovascular magnetic resonance (CMR) techniques” because of the nontraditional paths they take through the magnetic resonance (MR) data space. Traditionally, most routine clinical day-to-day MR data acquisitions have been aimed at traversing k-space in a so-called Cartesian pattern, with data points acquired at regular intervals on a two-dimensional (2D) or threedimensional (3D) grid, and with one point and one line of data acquired at a time. Conceptually, at least, such an approach is straightforward to implement and to understand. It mimics the regular gridded structure of pixels or voxels in the images that are ultimately produced, and the criteria for completion of image acquisition are clear: image acquisition stops when all of the points on the target grid are populated by data. However, the organs and the bodies that are imaged are continuous entities, and the data space representing them is at root continuous as well. Thus, nothing prevents us from acquiring data along non-Cartesian paths, so long as those data may be manipulated to yield reliable representations of image contents. Spiral imaging and radial imaging are two examples of non-Cartesian acquisition trajectories that have been studied extensively and used increasingly, particularly for rapid and real-time CMR. Parallel magnetic resonance imaging (parallel MRI), meanwhile, challenges the notion that data acquisition must be a monolithic and sequential path through k-space. Unlike traditional approaches that rely entirely on magnetic field gradients to move from one data point to another in MR acquisition, parallel MRI techniques use arrays of radiofrequency (RF) detector coils to supplement the spatial encoding provided by gradients and to generate multiple data points at once rather than one after the other. The result is an acceleration of MR image acquisition beyond previous limits. For imaging of the cardiovascular system, imaging speed is of course at a premium, and parallel MRI techniques have seen extensive use in cardiovascular applications. Another theme connecting spiral, radial, and parallel MRI is rapid imaging. Cardiovascular applications have been a significant motivating force for the development of

ever more rapid MR imaging techniques over the years, and nowhere are the challenges and the benefits of rapid MRI more apparent than in the field of CMR, where cardiac motion, respiratory motion, and blood flow all complicate imaging. Spiral and radial imaging sequences have proven to be powerful tools for rapid imaging, albeit for rather different reasons (the former because of efficient use of field gradients and the latter because of favorable undersampling behavior, as discussed in more detail later). Parallel MRI, on the other hand, is a general strategy that may be used to accelerate most existing imaging sequences, including spiral and radial sequences. In the sections that follow, the basic principles and typical cardiovascular applications of spiral imaging, radial imaging, and parallel imaging will be surveyed. A concluding section explores future directions.

SPIRAL IMAGING Principles Spiral trajectories1,2 have been used extensively for realtime imaging. A sample spiral trajectory is shown in Figure 3-1B, compared with a Cartesian trajectory shown in Figure 3-1A. Figure 3-2 summarizes the process of data acquisition and image reconstruction for spiral imaging sequences. The spiral path through k-space is accomplished using gradients that oscillate in a coordinated fashion along two in-plane directions (see Fig. 3-2, left). The lack of sharp transitions in the oscillating gradient patterns allows efficient use even of gradients with limited switching rates, and high switching rates may be employed for still greater speed. Unfortunately, the irregular sampling pattern of a typical spiral trajectory precludes the use of a simple fast Fourier transform for image reconstruction. Various specialized reconstruction procedures have been proposed. Regridding algorithms3,4 that map the data to a Cartesian grid are most frequently used. After regridding, a fast Fourier transform may be performed to yield the reconstructed image (see Fig. 3-2, right). Although the regridding step adds somewhat to reconstruction time, with modern Cardiovascular Magnetic Resonance 37

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Figure 3-1 Cartesian (A), spiral (B), and radial (C) data acquisition strategies. Paths through k-space are shown in the top row, with arrows indicating the readout direction. Resulting k-space data matrices after digitization of the cardiovascular magnetic resonance signal are seen in the bottom row.

hardware and software, it may be performed efficiently, such that low-latency display is possible.5 The traditional distinction between phase encoding and frequency encoding directions is lost in spiral sequences, and this has some important consequences for the resulting images. First, errors can accumulate throughout the long spiral readouts, as also occurs in Cartesian echo planar imaging readouts. However, in spiral imaging the effects of these errors tend to be distributed across the image, rather than being concentrated at well-defined aliasing positions or visible as coherent spatial distortions or shifts, as is the case for echo planar imaging. For example, off-resonance effects induced by field inhomogeneities, susceptibility variations, or chemical shift, coupled with a long readout time, may result in blurring of images obtained from spiral acquisitions. An example of spiral blurring as a result of off-resonance effects is shown on the left of Figure 3-3A (with a crisper Cartesian image shown on the right). Such blurring can be reduced with: (1) a priori B0 field maps acquired in extra scans6,7; (2) estimates of the spatially varying offAcquisition

resonance frequency obtained from the spiral data itself8; or (3) frequency selective water excitation/fat suppression.9 A number of variants on spiral imaging sequences are used for various purposes. For real-time implementations, data are typically acquired continuously during a single long spiral readout, which promotes speed but also allows phase errors to accumulate. Multi-shot spiral acquisitions, with interleaved spiral trajectories that follow separate RF pulses, may be used to control these errors, (see real-time spiral images in Fig. 3-3B). A typical spiral trajectory begins in the center of k-space and works its way symmetrically outward, but reverse spiral trajectories with different relaxation weighting have also been explored.10,11 Three-dimensional “stack of spiral” sequences, which apply traditional phase encoding along the direction perpendicular to the spiral, have also been described.12,13

Applications Spiral k-space sampling saw its initial cardiovascular application in the area of breath hold 2D coronary artery CMR, where the speed gain enabled (1 1 5) mm3 spatial resolution for a comparatively small number of slices.2 For this first application, interleaved spiral scanning was used without considering the motion of the coronaries. To follow the motion of the coronary artery in spiral coronary artery CMR, vessel tracking with prospective adjustment of the slice location as a function of the coronary position can be applied.14 To further improve the efficiency of spiral coronary artery CMR techniques with submillimeter in-plane spatial resolution as well as a high image signal-to-noise ratio (SNR), adaptive, subject-tailored, automated tracking of the vessel motion over the cardiac cycle has been used.15 Alternatively, variable density spiral k-space acquisitions were proposed to acquire coronary artery CMR at submillimeter spatial resolution with information for motion compensation obtained directly from the coronary anatomy itself.16 This approach eliminates the need for cardiac Reconstruction

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Figure 3-2 Data acquisition and image reconstruction strategies for spiral imaging. A sample pulse sequence for a spiral acquisition is shown on the left, with various lines showing the timing of radiofrequency (RF) pulses, in-plane gradients Gx and Gy, slice direction gradient Gz, and data reception (labeled ADC for analog-to-digital conversion). Note the smooth oscillation of in-plane gradients. Reconstruction strategies for spiral datasets are shown on the right. The most common approach is to regrid the spiral data onto a Cartesian grid (using well-established interpolation algorithms) and then to perform a fast Fourier transform (FFT) to generate a final image. It is also possible in principle to transform the spiral data directly into an image (dashed line), but this would increase computational burden significantly.

3 ADVANCED CARDIOVASCULAR MAGNETIC RESONANCE IMAGING TECHNIQUES: SPIRAL, RADIAL, AND PARALLEL IMAGING

Figure 3-3 Montage of cardiovascular magnetic resonance (CMR) applications of spiral imaging and comparisons with other trajectories. A, Short axis two-dimensional (2D) steady-state free precession (SSFP) cine images from spiral (left) and Cartesian (right) acquisitions. The spiral image on the left exhibits characteristic blurring as a result of B0 inhomogeneities as well as physiologic motion. For comparison, the Cartesian image on the right displays reduced image distortion and improved delineation of myocardial borders because of shortening of the echo train length. B, Long axis twodimensional cine images (left, systole; right, diastole) derived from CMR fluoroscopy using spiral k-space trajectories (four interleaves) and a frame rate of 15 frames/sec (using sliding window reconstruction). A nominal in-plane spatial resolution of (2.5 2.5) mm2 was used for data acquisition. C, Magnetic resonance angiogram (MRA) of the left coronary artery system obtained with a free breathing, navigated, threedimensional steady-state free precession (SSFP) technique using spiral (left) and Cartesian (right) kspace trajectories. The speed advantage of spiral imaging enabled an in-plane spatial resolution of 0.77 0.77 mm2 (left), whereas the traditional Cartesian approach (right) yielded an in-plane spatial resolution of 1.0 1.0 mm2. D, Visualization of the right coronary artery vessel wall using spiral (left) and radial (right) k-space sampling schemes. Motion artifacts are reduced in radial k-space sampling so that the vessel wall is better delineated in the case of radial k-space sampling. (B, Images courtesy of Gabriele Krombach, RWTH Aachen University, Germany; C, images courtesy of Rene Botnar, PhD, Guy’s and St. Thomas’ Hospital, London, United Kingdom; D, images courtesy of Marcus Katoh, PhD, RWTH Aachen University, Germany.)

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triggering, breath holding, and navigator echoes, although extended acquisition times are needed to overcome the reduced SNR efficiency associated with the method. Free breathing 3D spiral coronary artery CMR affords submillimeter in-plane spatial resolution (see Fig. 3-3C). Recent pulse sequence and hardware developments have served to further reduce the acquisition window length, enabling rapid 2D spiral acquisitions covering a large volume of the heart at a spatial resolution of 1 1 2 mm3.17 Double inversion prepared black-blood spiral imaging in tandem with free breathing, navigator-gated, cardiac triggered CMR facilitates coronary vessel wall imaging18 (see Fig. 3-3D). Assessment of global and regional cardiac function has also been performed with spiral imaging techniques. For example, electrocardiographic (ECG) gating of 2D steadystate free precession (SSFP) spiral acquisitions have been used to achieve full R-R coverage, high temporal resolution, and short scan times.19 However, overall image quality in these spiral studies was inferior to that derived from Cartesian sampling (see Fig. 3-3A). Variable density spiral trajectories can be employed to further reduce motion artifacts in 2D cine CMR.20 Interactive spiral real-time imaging methods, which eliminate the need for ECG gating, have gained importance for dynamic studies, such as assessment of cardiac function, guidance of interventional procedures, or flow measurements. For example, the speed advantage of spiral imaging has been exploited for rapid left ventricular function assessment using free breathing cine CMR without cardiac triggering, supported by real-time reconstruction/ display and interactive section positioning.21 Other studies reported a temporal resolution of 120 msec (24 frames/sec using sliding window reconstruction and display) while accomplishing almost 1 mm in-plane spatial resolution, together with high blood-myocardium contrast, affording excellent visualization of blood pool, myocardial wall, and valve leaflet motion at 1.5 T and 3.0 T.22,23 Real-time high temporal resolution spiral imaging reduces underestimation of the peak velocity in flow velocity imaging as a result of averaging over a shorter period around the peak

Acquisition

amplitude.24 Alternatively, the speed of accelerated spiral imaging can be translated into an improved spatial resolution that decreases partial volume effects in flow velocity imaging that arise from partial occupancy of the voxels with static spins.24

RADIAL IMAGING Principles A variant of the radial trajectory was used to generate the very first MR images.25,26 Radial trajectories were subsequently reintroduced and further explored for their comparative insensitivity to motion.27 A sample radial acquisition trajectory is shown in Figure 3-1C, and Figure 3-4 shows a sample pulse sequence and image reconstruction. As is the case for spiral imaging, the traditional distinction between phase encoding and frequency encoding directions is abandoned in a radial acquisition. For radial trajectories, however, coordinated gradient switching along the two in-plane directions occurs not during but between readouts, and each readout occurs along a straight line, with some angular shift from the last readout. As is the case for spiral trajectories, radial image reconstructions are more complex than for the Cartesian case, typically involving either regridding procedures or projection-reconstruction algorithms (e.g., filtered back projection [FBP]) analogous to those used in X-ray computed tomography (CT). The connections between MR data acquired in a radial trajectory and CT data acquired with a moving X-ray gantry are so close, in fact, that radial MRI is sometimes referred to as projection-reconstruction imaging. The motion insensitivity of radial trajectories results in part from their intrinsic oversampling of the center of k-space, which is traversed by each angular readout, or projection. Because each copy of the k-space center constitutes a low-resolution representation of the imaged field of view, some degree of motion can be averaged out in an image

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Figure 3-4 Data acquisition and reconstruction strategies for radial imaging. A sample pulse sequence for a radial acquisition is shown on the left, labeled as in Figure 3-2. In-plane gradients are switched in a coordinated fashion to accomplish readouts along progressively rotated directions. Reconstruction strategies for radial datasets are shown on the right. Regridding and fast Fourier transform may be performed as for spiral data. Alternatively, projectionreconstruction (PR) approaches similar to those used in X-ray computed tomography (CT) may be used to generate an image. ADC, analog-todigital conversion; RF, radiofrequency.

Figure 3-5 Undersampling behavior for Cartesian (A and C) and radial (B and D) imaging. A, A twofold undersampled Cartesian trajectory. Acquired k-space lines are shown in solid black and omitted lines are shown in dashed gray. B, A twofold undersampled radial trajectory. C, The aliased image resulting from Fourier transformation of the undersampled Cartesian trajectory in (A). Note the coherent replication of structures along the undersampled direction, which clearly obscures important cardiac anatomy. D, The image resulting from regridding reconstruction of the undersampled radial trajectory in (B). Undersampling artifact in the form of radial streaking is visible, and cardiac structures are fully visible.

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positions, producing characteristic radial streaking artifacts (see Fig. 3-5D). For undersampled radial trajectories in particular, it has been recognized that significant degrees of undersampling may be tolerated without obscuring important cardiac or vascular anatomy, and without compromising effective spatial resolution.28,29 (The undersampled radial image at the right in Fig. 3-5D, for example, is preferable to the aliased Cartesian image in Fig. 3-5C.) As a result, smaller datasets may be acquired, restoring the speed and efficiency of radial trajectories. Undersampled radial (or undersampled projection reconstruction) techniques have been combined with real-time SSFP CMR, for example, to yield high contrast between blood and myocardium.30–32 Further accelerations have been achieved using combinations of radial undersampling with multi-echo readouts.32 The undersampled projection reconstruction principle has also been extended to three dimensions in at least two ways: either using a “stack of stars” approach with traditional phase encoding in the slice direction,28 or using a true 3D projection reconstruction approach with radial projections extending along all three directions in what has been called colloquially a koosh ball trajectory. The latter approach, termed vastly undersampled isotropic projection reconstruction (VIPR),33 has been used to achieve dramatic accelerations for angiographic applications in which the field of view is

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Cardiovascular Magnetic Resonance 41

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reconstruction that effectively merges these representations. Somewhat paradoxically for a sequence that is often identified with rapid imaging, however, this same redundancy also represents an underlying inefficiency in data acquisition. In principle, the number of acquired data points needed to gather the outer portions of k-space with a requisite separation is generally larger than for Cartesian acquisitions. Nevertheless, radial trajectories have had a recent resurgence for rapid and real-time imaging applications, in part because of their undersampling behavior.28 When Cartesian acquisitions are undersampled (i.e., when intermediate lines of k-space data are omitted such that the spacing of lines exceeds the Nyquist limit [Fig. 3-5A]), the resulting images show well-defined aliasing artifacts (see Fig. 3-5C). This aliasing results from the inability of the Fourier transform operation to separate certain regularly spaced points in the image plane based only on the undersampled spatial frequency data. The regular pattern of Cartesian undersampling results in coherent overlap of spatially separated image regions that can obscure important structures (see Fig. 3-5C). The appearance of undersampling artifacts is notably different in many non-Cartesian trajectories, however. For example, the lack of a regular Cartesian grid structure to the acquired and omitted data in undersampled radial trajectories results in a spreading of aliasing artifact among multiple spatial

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

dominated by narrow vascular structures. In such applications, radial undersampling artifacts may be reduced to a low level by spreading throughout the imaged volume.

Applications Resilience to undersampling and motion artifacts renders radial trajectories effective candidates for the accurate assessment of global and regional cardiac wall motion. Radial 2D cine SSFP techniques afford image quality comparable to that of Cartesian k-sampling techniques19,21 (Fig. 3-6A). In particular, real-time assessment of cardiac function benefits from the reduced motion sensitivity of radial k-space sampling. This leads to enhanced signal homogeneity in the ventricle while preserving bloodmyocardium contrast, which improves the accuracy of the cardiac functional assessment. The scan efficiency of radial imaging, together with real-time sliding window reconstruction, enables frame rates of up to 20 frames/sec for small matrix sizes (see Fig. 3-6B, left) and approximately 10 frames/sec for high spatial resolution imaging (see Fig. 3-6B, right). Real-time radial cine CMR techniques eliminate the need for ECG gating but offer somewhat limited spatial and temporal resolution compared with conventional segmented radial techniques. Self-gated radial acquisition strategies were introduced to make up for this deficit. Radial self-gated techniques extract the motion synchronization signal directly from the same MR signals used for image reconstruction.34 The self-gating approach can also be extended to derive respiratory gating information directly from the raw imaging data, which enables free breathing segmented cine imaging using radial k-space trajectories.34,35 Free breathing, ungated CMR fluoroscopy is another area that employs the scan efficiency of radial imaging to guide and monitor CMR interventions. Early explorations of real-time imaging for interventions have included: (1) percutaneous intramyocardial application of contrast agents to track and supervise stem cell injections36; (2) safe automatic catheter tracking for real-time CMR-guided catheterization of the aorta, left ventricle, and carotid37; and (3) CMR-guided coronary artery stent placement.38 Radial k-space sampling techniques have increasingly been used for coronary artery CMR. As seen in Figure 3-6C, coronary artery CMR with a radial SSFP technique shows reduced motion artifacts and superior vessel sharpness compared with the Cartesian approach.39 Scan efficiency for free breathing radial acquisitions of the heart can be improved through motion correction techniques40 or through the use of extended acquisition windows during the cardiac cycle (enabled by the motion insensitivity of radial scanning).41 The latter approach, in conjunction with intersegment motion correction using self-guided epicardial fat tracking, holds the promise for rapid free breathing 3D coronary artery CMR with whole heart coverage.42 Four-dimensional coronary artery imaging has been realized using 3D stack of radial acquisitions across the entire cardiac cycle, with a phase sensitive SSFP pulse sequence for persistent fat suppression. The four-dimensional approach allows multiple images at mid-diastole to be averaged, thus enhancing SNR and vessel delineation.43 Inversion recovery preparation modules in conjunction with 3D radial k-space sampling permit blood 42 Cardiovascular Magnetic Resonance

signal suppression in the coronaries and hence are suitable for vessel wall imaging at submillimeter in-plane spatial resolution18 (see Fig. 3-3C). The benefits of radial k-space sampling can be put to use in imaging large and small vessels. The diagnostic value of high isotropic spatial (1.25 1.25 1.25 mm3) and temporal (3 sec/frame) resolution VIPR imaging has been shown in time-resolved contrast-enhanced MR angiography (MRA) of the distal extremity.44 Further VIPR applications include high-spatial-resolution multi-station MRA encompassing the abdomen, thigh, and calf.45 To extend the superior-inferior coverage of peripheral MRA, 3D VIPR acquisitions can be combined with continuous table motion.46 Radial 3D-SSFP imaging combined with a slab-selective inversion prepulse affords flow-targeted MRA without contrast medium application. This approach provides images well suited for vessel geometry assessment and reliable stenosis detection in renal arteries (see Fig. 3-6D).47 A significant improvement in motion artifact suppression, vessel sharpness, and detectable vessel length was found for spin labeling coronary MRA with SSFP and radial k-space sampling.48 Recent work with interleaved and weighted radial imaging has enabled images with multiple contrasts to be obtained from a single dataset. These 2D and 3D methods enable a radial trajectory to be used in conjunction with preparation pulses for detection of myocardial infarction and viability assessment using late gadolinium enhancement (LGE) CMR.49

PARALLEL IMAGING Principles One point to note about all MR data acquisition trajectories discussed so far, Cartesian or non-Cartesian, is that they are inherently sequential. One point and one line of data are acquired at a time. For sequential acquisitions, increased imaging speed is generally accomplished by reducing the time delay between acquired points, or else by acquiring fewer points (undersampling) and tolerating the level of artifact that may result. Unfortunately, there are limits to how quickly sequential data points may be acquired, because field gradients must be switched or RF pulses applied to move from data point to data point. Current technologic and physiologic limits on gradient switching rate and RF power deposition, then, constrain sequential MRI speed. Within these constraints, undersampling is the only remaining prospect for acceleration, but there are also limits to the level of uncorrected undersampling artifact that may be tolerated, even for radial trajectories. Parallel imaging achieves its speed by acquiring more than one data point at a time. How it does this may be understood by analogy to another speed-enhancing imaging approach that employs parallelism, namely, multidetector-row CT (MDCT). (Both parallel MRI and MDCT have served as potent enablers of cardiovascular imaging in recent times, rendering this analogy a particularly apt one for a consideration of CMR techniques.) Figure 3-7 compares MDCT and parallel MRI, with a schematic view

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D Figure 3-6 Montage of cardiovascular magnetic resonance (CMR) applications of radial imaging. A, Short axis electrocardiogram-gated (ECG) two-dimensional (2D) steady-state free precession (SSFP) cine images acquired with radial (left) and traditional Cartesian (right) k-space sampling using a 32-element cardiac coil array and a nominal in-plane spatial resolution of 1.2 1.2 mm2. Radial acquisition provided an enhanced delineation of the endo- and epicardial borders because of its resilience to undersampling, which allowed shorter acquisition windows in each cardiac phase without increasing the breath hold duration. B, Short axis 2D SSFP cine CMR using radial k-space trajectories and a frame rate of 20 frames/sec (left) and 10 frames/sec (right). Data were acquired using a 32-element cardiac coil array and a nominal in-plane spatial resolution of 2.2 2.2 mm2 (left) or 1.4 1.4 mm2 (right). C, CMR of the left coronary artery system obtained with a free breathing, navigated, three-dimensional SSFP technique using radial (left) and Cartesian (right) k-space trajectories. An in-plane spatial resolution of 1.0 1.0 mm2 was used for the acquisition of the radial and Cartesian datasets. The reduced motion sensitivity of radial imaging results in improved vessel delineation. D, Magnetic resonance angiograms of the renal arteries using radial (left) and Cartesian (right) k-space sampling schemes. Signal from the renal parenchyma and veins is completely suppressed by the inversion prepulse used for arterial spin labeling, whereas high-contrast visualization of the renal arteries, including the distal subsegmental branches, is enabled. Motion artifacts are reduced in radial k-space sampling, resulting in an improved vessel geometry assessment and improved detection of stenoses in the renal arteries. (C, Images courtesy of Rene Botnar, PhD, Guy’s and St. Thomas’ Hospital, London, United Kingdom; D, images courtesy of Marcus Katoh, PhD, RWTH Aachen University, Germany.) Cardiovascular Magnetic Resonance 43

Multiple image slices at once

Parallel MRI

Multiple k-space “slices” at once

RF coils

Rotations y ra Xitt em

Gradient steps

er

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

MDCT

Imaged subject Imaged subject Coil 1

Coil 2

Coil 3 k=0 k=2

Detector rows

k=4

Figure 3-7 Comparison of parallel acquisition approaches in multi-detector-row computed tomography (MDCT; left) and parallel magnetic resonance imaging (parallel MRI; right). Components contributing sequential image information are generally labeled in reddish orange, whereas parallel components are labeled in pale yellow. The top row shows a schematic side view of an MDCT scanner and a parallel MRI scanner, respectively. In MDCT, multiple image slices are acquired at once using adjacent rows of distinct X-ray detectors (pale yellow rectangles below the subject). In parallel MRI, multiple radiofrequency (RF) coils (pale yellow rectangles shown in two representative orientations above or below the subject) effectively enable multiple k-space “slices” to be acquired at once. The bottom half shows schematic front views and compares the types of projections of the imaged subject generated using MDCT or parallel MRI. For MDCT, an X-ray emitter is rotated on a gantry, and cone-like projections of the object are recorded in all detector rows from a variety of projection angles. An image of the subject is then reconstructed from this set of projections. In parallel MRI, each measured signal point is an integration or a projection of the imaged volume against the joint spatial modulations produced by field gradients and RF coils. Use of an array of RF coils provides multiple distinct modulations for each gradient step, thereby increasing the number of generalized projections available for image reconstruction and allowing images to be reconstructed from a reduced number of gradient steps. Sample generalized projection functions with complex spatial variation are shown for three coils and three gradient settings or k-space indices. Generalized parallel MRI reconstruction algorithms reconstruct image intensities from the MRI signal data using knowledge of the set of projection and MR functions. The gantry rotations and gradient steps are the sequential components of data acquisition for computed tomography and CMR, respectively, and the detector rows and RF coils represent the corresponding parallel components. (Courtesy of the National Library of Medicine; available at http://www.nlm.nih.gov/copyright.html.)

of each modality shown at the top and a more detailed juxtaposition of spatial encoding mechanisms shown at the bottom. Both modalities use arrays of detectors to accelerate imaging beyond previous sequential limits, and both combine multi-detector acquisition (pale yellow coloring in the figure) with more traditional sequential acquisition strategies (indicated by darker orange shading). In MDCT, multiple sequential projections are gathered by gantry rotation, with additional simultaneous projection information provided by multiple detector rows with coverage of distinct image slices. The ability to acquire multiple image 44 Cardiovascular Magnetic Resonance

slices at once enables coverage of a target imaging volume in a fraction of the time that would be required in the presence of a single detector row. In parallel MRI, gradients generate sequential information in the form of projections against oscillating functions (the usual Fourier projections that make up the acquired k-space matrix). Additional simultaneous information is available from multiple coil array elements, each of whose sensitivity patterns provides a distinct projection function. The extra simultaneous projections from multiple RF coils effectively provide information about multiple k-space “slices” at once. This allows

image contents may be determined, once again with a reduced number of sequential gradient steps. From another point of view, parallel MRI returns to the theme of undersampling introduced earlier in the context of radial imaging, and aims to fill in missing data in undersampled acquisitions with information derived from multiple RF coils. Within this general rubric, a number of particular strategies for parallel MRI have evolved over time. The basic approaches to data acquisition and image reconstruction for several common parallel imaging strategies are seen in Figure 3-8. First, undersampled data from an imaging sequence and k-space trajectory of choice are acquired simultaneously in the multiple elements of a coil array. For Cartesian trajectories, this generally involves the omission of phase-encoding gradient steps in a regular pattern, with only one out of every R lines acquired, where R is the desired acceleration factor. The k-space lines corresponding to omitted gradient steps are of course missing from the data matrix. In Figure 3-8, a case with R ¼ 2 is illustrated, with odd k-space lines (solid red) acquired and even k-space lines (dashed red) omitted. As shown on the right side of Figure 3-8, the missing k-space data may then

Acquisition

Reconstruction

FFT

SMASH, GRAPPA, etc.

SE

NS

E,

GE

M,

SENSE, ASSET, etc.

etc

.

FFT

GRAPPA, PARS, etc.

c.

, et

EM

E, G

NS

SE

Figure 3-8 Data acquisition and image reconstruction strategies for parallel magnetic resonance imaging (MRI). Parallel acquisition may be used with most existing pulse sequences and k-space trajectories. Undersampled data are acquired simultaneously in multiple elements of a radiofrequency (RF) coil array. Each element of a three-element array is shown at the top left, along with corresponding coil sensitivities. Undersampled Cartesian, spiral, and radial trajectories are shown below each array element, with solid black lines indicating acquired data and dashed gray lines indicating omitted data. Various approaches to reconstruction of the multi-coil undersampled data have been described, and a representative subset is indicated with arrows on the right. Simultaneous acquisition of spatial harmonics (SMASH), generalized autocalibrating partially parallel acquisition (GRAPPA), and similar techniques form fully sampled k-space matrices from the multiple undersampled datasets, after which fast Fourier transform (FFT) yields the final image. Cartesian sensitivity encoding (SENSE), array spatial sensitivity encoding technique (ASSET), and so forth perform the FFT first, then operate to unfold aliased component coil images. GRAPPA, parallel imaging with augmented radius in k-space (PARS), and other techniques may be used to regrid arbitrary undersampled non-Cartesian trajectories onto fully sampled grids before FFT. Alternatively, generalized SENSE, generalized encoding matrix (GEM), or kindred techniques transform arbitrary undersampled datasets directly to a final image. The choice of reconstruction technique in practice will be dictated by the particular imaging situation and by the algorithms available on a particular scanner platform. Cardiovascular Magnetic Resonance 45

3 ADVANCED CARDIOVASCULAR MAGNETIC RESONANCE IMAGING TECHNIQUES: SPIRAL, RADIAL, AND PARALLEL IMAGING

the number of time-consuming gradient steps to be reduced, and image acquisition can proceed in a fraction of the time that would have been required otherwise. One notable difference between parallel MRI and MDCT is the comparatively broad sensitivity profiles of RF coils as opposed to the collimated views provided by adjacent rows of X-ray detectors. The RF coil sensitivities fall off smoothly with distance from each coil’s center, such that sensitivities for adjacent coils typically have a significant degree of overlap. (Note that three sample coil sensitivity patterns shown at the right of Fig. 3-7 have been separated for clarity. If they were overlaid on the same field of view, only their peaks would be clearly distinguishable.) This overlap must be accounted for in parallel image reconstructions. One general means of accounting for overlap is to take a cue, once again, from X-ray CT and to formulate the problem as a generalized reconstruction from projections. When the coil-related sensitivity functions and the gradient-related Fourier functions are combined, as shown at the bottom right of Figure 3-7, they form rather odd-looking projection functions. These known projections may be collected into a matrix, and through suitable matrix inversion or other algebraic techniques, internal

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

be filled in using appropriately weighted combinations of acquired data from multiple coils, with component coil weightings chosen based on the known sensitivity patterns of the coils. This is the basis of the simultaneous acquisition of spatial harmonics (SMASH) technique.50 SMASH was the first parallel imaging approach demonstrated in vivo, although various proposals for simultaneous acquisition had been made earlier,51–57 dating back to the time of the first uses of coil arrays for MRI. Various improvements and generalizations of k-space-based parallel image reconstruction strategies have since been described and are somewhat vendor specific. Various improvements and generalizations of k-space–based parallel image reconstruction strategies have since been described. The most commonly used today is the GeneRalized Autocalibrating Partially Parallel Acquisition (GRAPPA) technique.58 Commercial implementations of parallel imaging have kept pace with these developments; for example, Siemens scanners (Siemens Medical Solutions, Erlangen, Germany) use a commercial adaptation of the GRAPPA technique. The far right side of Figure 3-8 shows a set of component coil images obtained by Fourier transformation of regularly undersampled Cartesian datasets. Regular undersampling in this case results in well-defined aliasing in the resulting images (i.e., the left side of the target field of view in each image is folded back around to the right, and vice versa). For any one coil’s image, this aliasing is unavoidable, but when multiple differently aliased images are available, knowledge of each coil’s distinct “view” may be used to undo aliasing artifacts. The Cartesian sensitivity encoding (SENSE) technique59 takes this unfolding approach. Philips scanners (Philips Medical Systems, Best, The Netherlands) use the SENSE algorithm for parallel MRI studies. General Electric systems (General Electric Healthcare, Waukesha, WI) currently provide the array spatial sensitivity encoding technique (ASSET), which uses principles similar to those of SENSE. Other vendors have their own related offerings. For spiral acquisitions, undersampling generally involves the omission of spiral interleaves; for radial sequences, radial projections are omitted (see Fig. 3-8 bottom left). For these non-Cartesian trajectories, undersampling does not result in regular and confined aliasing patterns, and as a result, a Cartesian SENSE unfolding approach is precluded. (This is also true for irregularly undersampled Cartesian trajectories, such as the dense-centered trajectories commonly used for GRAPPA and other self-calibrating parallel imaging techniques.) Techniques such as GRAPPA or parallel imaging with augmented radius in k-space (PARS)60 may be used to combine undersampled non-Cartesian datasets (or irregularly undersampled Cartesian datasets) into fully sampled Cartesian k-space grids that may then be Fourier transformed to yield reconstructed images. Alternatively, generalized approaches with reconstruction from projections may be used to move directly from arbitrary undersampled datasets to final reconstructed images. A suitably generalized SENSE algorithm59,61 may be used for this purpose, as may other techniques, such as sensitivity profiles from an array of coils for encoding and reconstruction in parallel (SPACE RIP),62 the generalized encoding matrix (GEM) technique,63 generalized SMASH,64 and others. These generalized strategies are typically more computationally intensive than their Cartesian counterparts, and they can result in prolonged image reconstruction times. However, efficient algorithms61 46 Cardiovascular Magnetic Resonance

or distributed computing technology65 may be used to achieve prompt reconstruction. Both before and since the advent of parallel MRI, RF coil arrays have seen extensive use for traditional sequential MRI, but for parallel MRI in particular, coil arrays are strictly required. Since the coil sensitivities share the burden of spatial encoding in parallel MRI, the details of array design are also particularly important. For example, to be effective for parallel MRI, array elements must have suitably distinct sensitivity profiles along any direction to be targeted for acceleration, in addition to having a satisfactory SNR over the volume of interest. Much has been written on the subject of array design, and we will not expound further. For an overview, the reader is referred to review articles.66 Various cardiac-specialized arrays have been constructed, and some of these are currently available from vendors. One additional requirement that parallel imaging imposes on MR scanner technology is that multiple receiver channels must be available to capture independent data from different coils. Fortunately, the number of commercially available receiver channels has grown steadily, motivated in part by the fact that the maximum achievable parallel imaging acceleration is equal to the number of independent array elements, and hence the number of independent receiver channels. Two practical considerations for parallel imaging studies are important to consider. First, once a suitable array is selected, calibration of the component coil sensitivity profiles is required. Second, the acceleration achieved using parallel MRI involves particular SNR trade-offs relating both to coil array geometry and to image reconstruction (Fig. 3-9). To perform the sensitivity-encoded parallel image reconstruction, one needs to know the various “views” provided by the array elements. Various calibration strategies have been described, most involving separate acquisition of in vivo reference data in the desired image plane or in an encompassing volume (see Fig. 3-9A). With reference data in hand, varying degrees of image processing may be performed to extract pure sensitivity data59 or to eliminate the contributions of image features shared among all of the coils.63,67 The calibration step does add somewhat to the overall examination time, although rapid low-resolution 3D acquisition may often suffice for an entire series of accelerated acquisitions. However, motion of the patient or the coil array between the time of calibration and the time of an accelerated acquisition can result in calibration errors and image artifacts. This is a particular concern for CMR, for which flexible arrays contoured to the mobile chest wall are commonly used. To address this problem, self-calibrating parallel imaging strategies may be used58,68–70 (see Fig. 3-9B). Self-calibrating approaches incorporate a sensitivity reference directly into the accelerated acquisition (in the form of a small set of additional fully sampled central k-space lines). Therefore, they allow accurate parallel imaging even in the presence of vigorous motion. Among the commonly used techniques currently available from vendors, GRAPPA58 and a modified form of SENSE (mSENSE) are self-calibrating,71 and generalized approaches such as the generalized encoding matrix technique may also be used in a self-calibrating fashion.70 Even with perfectly calibrated sensitivities, the SNR of a parallel imaging study is always reduced compared with an unaccelerated study obtained using the same coil array.59,72 This is a result of both the data acquisition strategies and

SNRaccelerated ¼

SNRunaccelerated pﬃﬃﬃ g R

(1)

Here, R represents the acceleration factor and the square root dependence in Equation 1 reflects the reduced SNR averaging resulting from a reduced number of acquired data points in an accelerated scan. The “geometry factor,” g, on the other hand, is an additional SNR reduction factor characterizing the extent to which the linear combinations used in a parallel image reconstruction amplify noise out of proportion to signal. This noise amplification depends sensitively on the choice of image reconstruction algorithm and on coil array design. The value of g varies spatially across the image plane (see Fig. 3-9C), meaning that the noise background is generally nonuniform in a parallel imaging study, and this should be taken into account in image interpretation.

Coil sensitivity–based parallel imaging strategies may also be supplemented by techniques that operate in the temporal domain. CMR often involves dynamic imaging so that the availability of multiple time frames affords one the opportunity to vary acquisition trajectories as a function of time with otherwise identical acquisition parameters and to take advantage of resulting spatiotemporal correlations. This is the concept behind techniques such as Unaliasing by Fourier encoding the OverLaps using the temporal Dimension (UNFOLD)73 and Broad-use Linear Acquisition Speed-up Technique (k-t BLAST),74 which have been used to achieve accelerations of dynamic imaging without the need for coil arrays and without the g factor–related noise amplification penalty. The k-t BLAST concept allows a more general temporal ordering of data acquisition and uses spatiotemporal correlations measured from sequential or interleaved training data to reassemble image components that are distributed in time and space.74,75 Explicit combinations of spatial information

Low-resolution sensitivity calibration volume g (R = 1)

10 5 0

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g (R = 2)

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5 0

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B acquisition

C

Figure 3-9 Coil sensitivity calibration strategies and noise propagation behavior for parallel imaging. A, External calibration. A lowresolution calibration data volume is acquired separately from the accelerated scan and is processed to yield sensitivity estimates in the target image plane. These sensitivities are then used for parallel image reconstruction. B, Self-calibration. Additional calibration lines (gray) are acquired along with the undersampled accelerated datasets. These calibration lines are generally placed so as to produce a fully sampled region in the center of k-space. The fully sampled region for each component coil’s dataset may then be Fourier transformed or otherwise processed to yield the sensitivity information necessary for parallel image reconstruction. C, Noise amplification with increasing acceleration. A four-element array is shown at the top with its component coil sensitivities. Surface plots of g factor for acceleration factors R ¼ 1 to R ¼ 4 using this array are shown on the left of the simulated images. Peaks in the g factor map correspond to areas of increased noise and hence decreased signal-to-noise ratio in the images. The particular location and shape of noise peaks and valleys are influenced by the choice of coil array geometry, image plane, and acceleration factor. (Reproduced with permission from Sodickson DK. Parallel imaging methods. In: Edelman RR, Hesselink JR, Zlatkin MB, eds. Clinical Magnetic Resonance Imaging. 3rd ed. Philadelphia: Saunders; 2005.) Cardiovascular Magnetic Resonance 47

3 ADVANCED CARDIOVASCULAR MAGNETIC RESONANCE IMAGING TECHNIQUES: SPIRAL, RADIAL, AND PARALLEL IMAGING

the image reconstruction algorithms used in parallel MRI. The scaling of SNR may be expressed as follows:

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

from coil arrays with temporal information include k-t SENSE,76 UNFOLD-SENSE,77 and TSENSE,78 which increase net accelerations for dynamic applications.79–81 Parallel imaging artifacts can result from errors in any of the preparation stages of a parallel imaging study, from

equipment malfunctions, or from “intrinsic” causes, such as motion of the patient or coil arrays. Depending on the acquisition and reconstruction strategy used, these artifacts may be manifested as residual aliasing, increased noise amplification, or diminished temporal fidelity (Fig. 3-10).

A

B

C

D

E

F

Figure 3-10—See legend on next page 48 Cardiovascular Magnetic Resonance

Applications Parallel MRI may be used to improve a wide range of existing imaging studies as well as to enable new ones previously precluded by constraints on scan time. Today, commercial parallel imaging implementations are available for all modern MR scanners, and parallel imaging acquisitions are used routinely in a substantial fraction of all CMR examinations. The increased speed and efficiency associated with parallel MRI may be translated into the following: Shorter examinations Improved spatial resolution and anatomic coverage Improved temporal resolution Enhanced image quality Relaxed physiologic constraints (e.g., RF power deposition, peripheral nerve stimulation, acoustic noise) or physical constraints (e.g., gradient switching rate, dB/dt) Several general areas in which cardiovascular parallel imaging has been applied so far are discussed.

Imaging of Cardiac Anatomy and Structure Imaging of cardiac anatomy and structure using fast spin echo-based techniques benefits from parallel imaging, which helps to limit relaxation-related blurring by allowing reduced echo train lengths. Parallel imaging also enables anatomic imaging and tissue characterization simultaneously via the acquisition of proton density (short echo time) and T2-weighted (long echo time) images in a single

breath hold, thereby reducing the risk of image misregistration compared with the conventional approach.82 Meanwhile, improved image quality is enabled by the synergy between parallel imaging and the improved baseline SNR available at 3 T as opposed to 1.5 T83 (Fig. 3-11A). Perhaps the greatest benefit of parallel imaging for fast spin echo imaging, particularly at higher field strengths, is the capability to reduce the total power deposition by omitting phase encoding steps and corresponding RF refocusing pulses, which can be supplemented by the application of variable flip angles and hyperechoes.84–86 The use of high acceleration factors enabled by many-element coil arrays87,88 promises to allow breath hold 3D black-blood imaging with whole heart coverage, an approach that would eliminate the risk of 2D slice misregistration.

Assessment of Global and Regional Cardiac Function Assessment of global and regional cardiac function requires high muscle-blood contrast, full R-R coverage, high temporal resolution, and short scan times. Parallel imaging strategies serve to improve cardiac functional assessments in various ways: Reduction in the total acquisition time in segmented cine CMR,89 which increases patient comfort and diminishes respiratory artifacts. An acceleration factor of 8 reduces the number of heartbeats required for single-slice imaging from 16 to only 2 (see Fig. 3-11B), which permits single breath hold 2D cine CMR with apex-to-base coverage using multiple slices. Increase in the number of cardiac phases per heartbeat for enhanced temporal resolution to identify more precisely the exact time point of maximal systolic contraction and diastolic filling.89,90 Minimization of motion sensitivity through reduction of the acquisition window duration, which supports the

Figure 3-10 Gallery of potential parallel imaging artifacts. A, Artifacts caused by excessive acceleration. As seen in Figure 3-9C, the overall magnitude and the inhomogeneity of the noise background increase with increasing acceleration, and noise amplification can become severe when acceleration factors approach the number of coils. The dotted white circle on the left highlights noise amplification for an acceleration factor of R ¼ 4 using a four-element array. Noise amplification may be controlled by the use of a more moderate acceleration factor, as shown on the right (R ¼ 2 for the same four-element array). B, Artifacts caused by inappropriate fields of view. Residual aliasing and relatively high noise near the center of the image (dotted white oval) occur if the imaged subject extends beyond the prescribed field of view (FOV) (left, FOV ¼ 24 cm). Image artifacts are not present in the accelerated scans if the target field of view is increased sufficiently (right, FOV ¼ 32 cm). (For both images, the same subregion is shown for clarity of comparison, despite the varying FOV and spatial resolution.) This behavior results from the fact that, in Cartesian sensitivity encoding (SENSE)-based imaging, the overlap of structures in the target field of view leads to ambiguities in the partitioning of intensities among aliased positions, resulting in image artifacts. These artifacts can be removed by using UNaliasing by Fourier encoding the OverLaps using the temporal Dimension (UNFOLD) or other k-t approaches. Some relaxation of the FOV constraint was recently reported to be possible for generalized autocalibrating partially parallel acquisition (GRAPPA) reconstructions.118 C, Artifacts resulting from coverage deficits in an external calibration scan. The accelerated short axis cardiac image on the left used a calibration scan that covers only a limited region marked by the dashed gray rectangle in the sagittal scout image above. This results in blank regions and severe aliasing artifacts highlighted by the dotted white circle. For comparison, the accelerated short axis image on the right used a reference scan that completely covers the prescribed scan plane, resulting in artifact-free images. It is generally recommended to include in the calibration scan the entire region over which coils may have appreciable sensitivity. D, Artifacts caused by coil placement errors. Coil arrays used for cardiovascular parallel imaging should be placed so as to bring the target anatomy within the focus of the array’s sensitivity pattern. Examples of inappropriate (left) and appropriate (right) coil positioning are shown, with target position indicated by solid lines and actual position indicated by dashed lines in the sagittal scout images. For the severe offset shown on the left, reconstructed images are noisy and exhibit considerable uncorrected aliasing, as highlighted by the dotted white circle. E, Artifacts resulting from patient and coil array motion (indicated by different chest wall positions in the axial scout images). Any mismatch between the measured coil sensitivities and the actual sensitivities active during an accelerated scan can result in residual aliasing artifacts (left, dotted circle) that disappear if the measured coil sensitivities match the actual sensitivities (right). Self-calibrating approaches can limit the incidence and severity of these artifacts. F, Artifacts caused by reduced temporal fidelity in accelerated dynamic imaging using spatiotemporal correlations obtained from low-spatial-resolution training data. The temporal fidelity of images reconstructed from accelerated data using k-t approaches decreases with an increasing acceleration factor. Consequently, image blurring is pronounced and tissue border sharpness is reduced in the faster image on the left as opposed to the slower image on the right. The time course over the full R-R interval of a single one-dimensional projection along the dotted line in each image also shows evidence of this blurring (dotted white oval), which has the potential to affect the quantitative assessment of dynamic image data. Cardiovascular Magnetic Resonance 49

3 ADVANCED CARDIOVASCULAR MAGNETIC RESONANCE IMAGING TECHNIQUES: SPIRAL, RADIAL, AND PARALLEL IMAGING

However, by accelerating the acquisition, the use of parallel imaging can reduce the incidence or severity of other common imaging artifacts, such as motional blurring or slice misregistration. This balance should be considered in the planning and interpretation of parallel CMR studies.

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

Anatomic Imaging Triple IR Fast Spin Echo

Double IR Fast Spin Echo Conventional

SENSE (R = 2)

Conventional

SENSE (R = 2)

A Cardiac function using 2D CINE + SENSE Conventional taoq = 16 R-R

SENSE (R = 4) taoq = 4 R-R

SENSE (R = 2) taoq = 8 R-R

SENSE (R = 8) taoq = 2 R-R

B Myocardial perfusion imaging with k-t BLAST Passage right ventricle

Base line

Passage left ventricle

Passage myocardiuim

C Delayed enhancement Conventional

Coronary artery MRA

SENSE (R = 2) + PSIR

D Figure 3-11—See legend on next page

50 Cardiovascular Magnetic Resonance

Breath hold

E

Free breathing

First-Pass Myocardial Perfusion Imaging Using parallel imaging, one- or two-heartbeat temporal resolution has been achieved in saturation-recovery-based techniques used to capture contrast agent kinetics for the assessment of myocardial perfusion.102–105 Moreover, the k-t BLAST and k-t SENSE approach can be used to double the spatial coverage per unit time while preserving

in-plane spatial resolution. Alternatively, k-t BLAST and kt SENSE can be put to use to double the in-plane matrix size without impairing the temporal resolution (see Fig. 3-11C). The combined SNR and speed improvements associated with high-field parallel imaging can be exploited to transition from multi-slice 2D acquisitions to whole heart 3D acquisitions or to enable myocardial perfusion imaging with arterial spin labeling rather than exogenous contrast agent injection.106

Detection of Myocardial Infarction and Assessment of Myocardial Viability The established CMR assessment of ischemic heart disease includes LGE CMR using ECG gated, segmented imaging modules preceded by an inversion recovery preparation to provide consistent high contrast between infarcted and healthy myocardium.107 Accelerated LGE imaging is of clinical importance because unaccelerated approaches have limited spatial coverage of only one to two slices per breath hold, resulting in prolonged examination times of 10 to 15 minutes, with corresponding decay of contrast agent concentration over the course of the examination. Parallel imaging can overcome these difficulties by allowing whole heart coverage in a single breath hold, ensuring uniform suppression of healthy myocardium for all imaged sections. Meanwhile, a phase sensitive reconstruction of inversion recovery technique has been shown to offset the need for perfect evolution time adjustments and to enhance the contrast between healthy and infarcted myocardial tissue.108 This approach doubles the total scan time compared with the conventional one R-R interval approach because it requires two R-R intervals for the acquisition of a T1-weighted inversion recovery dataset and an extra reference image. This trade-off can be compensated by using the time savings inherent in parallel imaging to facilitate short breath hold times (see Fig. 3-11D).

Coronary Artery Cardiovascular Magnetic Resonance Parallel imaging strategies provide several means of improving coronary artery CMR by minimizing the effect of physiologic motion. The use of parallel imaging strategies enables 3D free breathing navigator techniques to be accelerated (see Fig. 3-11E) while preserving image quality

Figure 3-11 Montage of parallel cardiovascular magnetic resonance (CMR) applications. A, Short axis cardiac images obtained with double (left) and triple (right) inversion recovery black-blood fast spin echo imaging at 3.0 T using a conventional unaccelerated approach and a twofold accelerated parallel imaging approach. B, Short axis cine CMR acquired with an unaccelerated (R ¼ 1) conventional two-dimensional (2D) steady-state free precession (SSFP) approach (left) and 2D SSFP with parallel imaging using one-dimensional acceleration factors up to R ¼ 8 (right). The increase in the acceleration factor, together with a constant number of phase encoding steps per cardiac cycle, resulted in a significant breath hold time reduction. The total scan time was 16 heartbeats for the unaccelerated approach and 2 heartbeats for R ¼ 8. Images were acquired using a 32-channel cardiac coil array. C, Selected short axis first-pass perfusion images (spatial resolution of 2.0 2.0 8 mm3) derived from a dataset acquired at three slices per R-R interval and reconstructed using Broad-use Linear Acquisition Speed-up Technique (k-t BLAST) with fivefold acceleration. The one R-R temporal resolution was used to determine the precontrast baseline (far left), to monitor contrast agent arrival in the right ventricle (center left), and to track the passage of the contrast agent through the left ventricle (center right) and the myocardium (far right). D, Short axis views obtained from late gadolinium enhancement (LGE) CMR at 3.0 T. For the conventional approach (left), one R-R interval was used for recovery of the magnetization to achieve a breath hold duration of 12 seconds. The phase sensitive reconstruction of inversion recovery (PSIR) approach (right) required two R-R intervals for full magnetization recovery, which was compensated by using twofold accelerated parallel imaging to keep the breath hold time at 12 seconds. E, Maximum intensity projection (3-mm slice thickness) of an image volume showing the right coronary artery obtained from accelerated (R = 2) ECG-gated, fat-saturated, threedimensional SSFP acquisitions using a breath hold (left) and a free breathing, navigated approach (right). The effective scan time was halved in the twofold accelerated scan while preserving the image quality of the unaccelerated acquisition. SENSE, Cartesian sensitivity encoding. Cardiovascular Magnetic Resonance 51

3 ADVANCED CARDIOVASCULAR MAGNETIC RESONANCE IMAGING TECHNIQUES: SPIRAL, RADIAL, AND PARALLEL IMAGING

visualisation of small, rapidly moving structures, such as cardiac valves.89 Improvement of anatomic coverage without increasing total acquisition time. For example, rapid segmented 3D techniques have been shown to be capable of scanning the entire heart in a single breath hold.91,92 Acceleration of real-time imaging methods for the assessment of cardiac function.93,94 In addition to eliminating the need for ECG gating, accelerated real-time techniques can also be used to track cardiac motion in a free breathing mode. The effect of motion in such studies may be minimized through the use of short acquisition windows and the application of self-calibration for coil sensitivity mapping.95 Improvement of myocardial tagging methods for the assessment of regional myocardial wall motion.96 The potential to suspend the need for breath holding facilitates the monitoring of transient hemodynamics and cardiac mechanics97 and affords the use of true 3D tagging grids for the accurate determination of quantitative 3D motion patterns.98 Acceleration of phase-contrast CMR by reducing the scan time for quantitative flow measurements in the great arteries. This approach supports the detection of congenital heart disease99 and the assessment of flow abnormalities.100 Increase in net accelerations using temporal undersampling strategies. The combination of UNFOLD73 with SENSE has been used to achieve higher acceleration factors for a given number of receiver coils or channels than would otherwise be feasible with SENSE alone.77,81 TSENSE with an acceleration factor of R ¼ 6 (3 2) has facilitated single breath hold 3D cine CMR with whole heart coverage at a temporal resolution of 50 msec.101 The k-t BLAST and k-t SENSE74 approach enables high accelerations, which support single breath hold multislice and whole heart coverage cine acquisitions without prohibitive noise amplification.

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

MRA using HighlY constrained back PRojection (HYPR)

A

From 2D to 3D single breath hold imaging

B Rapid whole heart coverage imaging + reformatting

C

From small volume to moving table whole body imaging

D Figure 3-12—See legend on next page

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FUTURE DIRECTIONS The topic of future directions and the question of what techniques will be considered “advanced” in times to come are tricky topics for a textbook, which can run the risk of

promptly appearing out of date when printed predictions turn out to be off the mark. In a field that evolves as rapidly as MRI has continued to do since its inception, however, it is only prudent to turn our attention briefly to the horizon. One well-established precedent drawn from the severaldecade history of MRI is that imaging speed is likely to continue to increase. Figure 3-12 shows four examples of a current trend toward very high accelerations that may well have a significant effect on CMR imaging in the future. In Figure 3-12A, a simulated 50-fold acceleration is shown for an MRA using the highly constrained back projection (HYPR) technique.114 On the left is the reference image obtained with a 400-projection 2D radial acquisition with a standard FBP reconstruction. FBP reconstruction of a highly undersampled trajectory with only eight projections is shown in the center, and the corresponding HYPR reconstruction is shown on the right. Clearly, the use of eight projections is insufficient for traditional FBP reconstruction, and the image is uninterpretable. However, HYPR incorporates some principles of the k-t techniques discussed earlier, using information from a time series of undersampled radial acquisitions to constrain each time frame. As the fidelity of the HYPR image suggests, extraordinary levels of acceleration may be possible, as long as conditions appropriate for HYPR exist, namely, that temporal information is available and the field of view is occupied sparsely by narrow, high-contrast vessels. Combinations of HYPR principles with 3D VIPR trajectories promise further accelerations for such sparsely populated fields of view. Figure 3-12B and C shows the levels of acceleration that are currently possible and the CMR applications that are thereby enabled using parallel MRI in the absence of these specialized conditions (e.g., for more densely populated fields of view). Both are examples of highly parallel CMR and take advantage of available 32-channel systems and 32-element arrays. Figure 3-12B shows accelerated LGE CMR obtained at 3.0 T. A fourfold acceleration allowed the transition from multiple 2D acquisitions encompassing only one to two slices per breath hold to single breath hold acquisitions with whole heart coverage. Figure 3-12C shows various views from 3D SSFP CMR datasets obtained with eightfold acceleration.87 The high level of acceleration in this case allowed acquisition of a comprehensive axial volume within a single breath hold, enabling visualization

Figure 3-12 Prospects for highly accelerated cardiovascular magnetic resonance (CMR) using undersampled projection reconstruction or parallel MRI. A, Simulated 50-fold acceleration of a magnetic resonance angiogram (MRA) using the highly constrained back projection (HYPR) technique,114 which uses information from a composite image to constrain the reconstruction of each frame in a time series (in this case, containing 16 frames). A reference 400-projection reconstruction image (left) is juxtaposed to an 8-projection projection-reconstruction image reconstructed with a traditional filtered back projection algorithm (center) and an 8-projection HYPR image (right). Unlike filtered back projection, HYPR clearly depicts vascular anatomy despite the high level of undersampling. B, Accelerated parallel MRI for late gadolinium enhancement (LGE) at 3.0 T. Acceleration factors of R ¼ 4 enabled three-dimensional (3D) whole heart coverage in a single breath hold, thereby permitting uniform suppression of the healthy myocardium and preventing slice misregistration. C, Highly accelerated parallel CMR for single-breath-hold whole heart coronary artery CMR.115 A 32-element array (the top half of which is shown schematically on the left) was used to achieve eightfold (4 2) accelerations for an axial 3D steady-state free precession (SSFP) image volume (256 256 60 matrix size, 1.5 1.5 2 mm3 spatial resolution). Several axial images from several sources are shown at center left. Various reformatted views are also shown, including traditional long axis and short axis views (center) as well as depictions of the left and right coronary arteries (center right). The 3D volume rendered view of the right coronary artery (RCA) shown on the right was obtained retrospectively from an automatic segmentation algorithm and attests to the quality of the original individual images as well as the level of contrast achieved with the highly accelerated acquisition. This full complement of information was available after a single 25-second breath hold scan with simple axial planning. D, Schematic depiction of the progression from thin section imaging (left) to rapid volumetric imaging (center) or even to whole body imaging incorporating automatic table motion (right). The general paradigm of highly accelerated imaging with simple axial planning and comprehensive volumetric coverage is highly reminiscent of multi-detector-row computed tomography (MDCT), but incorporates various advantages associated with MRI. 2D, two-dimensional. (A, Images courtesy of Charles Mistretta, University of Wisconsin, Madison.)

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comparable to that of the unaccelerated approach.109 This decreases the total scan time and reduces the susceptibility to patient motion and respiratory drift. Short breath hold 3D coronary artery CMR is challenging because of the competing constraints of spatial resolution, anatomic coverage, and breath hold time. The efficiency advantage of parallel imaging strategies minimizes the effect of cardiac and respiratory motion and facilitates shorter data acquisition windows within the cardiac cycle.110 This is especially beneficial at high heart rates, for which shortening of the mid-diastolic cardiac rest period limits the usable acquisition window. Parallel imaging also allows reductions in the number of cardiac cycles used for in-plane segmentation. This leads to very short breath hold times for each coronary artery acquisition and is especially suited for low heart rates. For either free breathing or breath hold approaches, the time savings associated with parallel imaging can be translated into an improvement of in-plane and through-plane spatial resolution, which results in an improved delineation of proximal and especially distal segments of the coronary arteries.109 Conventional coronary artery CMR studies are generally restricted to targeted thin slabs encompassing a particular segment of the coronary artery tree only. Parallel imaging allows the use of a thicker volume, which supports the visualization of long tortuous segments of the coronary arteries and offers the potential to eliminate localization scans.111,112 The T1 prolongation and SNR gain at high magnetic field strengths, together with the speed benefit of parallel imaging, will prove to be useful for flow-targeted imaging. The SNR improvements promise to be beneficial not only for MR lumography but also for vessel wall imaging.113 As accelerated high spatial–resolution vessel wall imaging is accomplished at 3.0 T, intrinsic contrast mechanisms and specific contrast agents can be used for plaque detection and plaque characterization.

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of the coronary arteries or any other typical cardiac view by subsequent reformatting. The same scanner and array have also been used for accelerated real-time cardiac imaging115 and for rapid volumetric contrast-enhanced MRA at 12- to 16-fold accelerations.88 The trend toward still larger numbers of receiver channels and array elements may only be expected to continue. Acceleration levels may not necessarily increase arbitrarily as channel counts increase because SNR losses are expected to mount rapidly as a function of acceleration caused by fundamental electrodynamic constraints on the shapes of RF coil sensitivities.116,117 However, with sufficient increases in baseline SNR, through the use of high-field MR systems or new hyperpolarized contrast agents, further accelerations may be expected. Of course, one may also combine parallel imaging at more modest levels with highly undersampled radial acquisitions to further push the limits of imaging speed. The examples in Figure 3-12B and C (as well as Fig. 3-12A if HYPR is used with 3D trajectories) represent efforts at rapid volumetric imaging, another significant trend. As imaging speed has increased over time, larger and larger imaging volumes have become accessible at any given spatial and temporal resolution. This is, of course, an appealing prospect for those interested in

comprehensive surveys of, for example, cardiac or vascular anatomy. It is also a trend that is not restricted to CMR. Rapid advances in MDCT are allowing ever more rapid volumetric scans, which have been viewed increasingly as a challenge to CMR. As we have seen, however, CMR is capable of similarly rapid volumetric acquisitions. Indeed, rapid volumetric imaging enables simplified axial planning and a comprehensive volumetric approach (see Fig. 3-12C), which are reminiscent of MDCT. All that remains to complete the analogy is to combine highly parallel MRI with existing moving table approaches to enable true whole body coverage, a progression that is shown in Figure 312D. Of course, comprehensive volumetric CMR boasts a number of advantages over its kindred tomographic modality. By taking advantage of efficient k-space trajectories and highly parallel imaging, MRI can provide some of the speed and simplicity of MDCT, while eliminating concerns about radiation dose and maintaining the tissue specificity and biochemical selectivity of the MR phenomenon. For the optimistically inclined practitioner, it should not be difficult to imagine the comprehensive cardiovascular examination that has long been a holy grail of CMR occupying a scant few minutes. In short, the future of today’s advanced CMR techniques remains promising.

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98. Ryf S, Kozerke S, Boesiger P. 3DCSPAMM tagging accelerated with kt-BLAST. In: Proceedings of the 11th Scientific Meeting of the International Society of Magnetic Resonance in Medicine. Toronto, Ontario, Canada: 2003:1565. 99. Beerbaum P, Korperich H, Gieseke J, Barth P, Peuster M, Meyer H. Rapid left-to-right shunt quantification in children by phase-contrast magnetic resonance imaging combined with sensitivity encoding (SENSE). Circulation. 2003;108:1355–1361. 100. Beerbaum P, Korperich H, Gieseke J, Barth P, Peuster M, Meyer H. Blood flow quantification in adults by phase-contrast MRI combined with SENSE–a validation study. J Cardiovasc Magn Reson. 2005;7:361–369. 101. Kellman P, Larson AC, Zhang Q, Simonetti OP, Arai AE, McVeigh ER. Cardiac CINE 3D true-FISP parallel imaging using auto-calibrating 2D TSENSE. In: Proceedings of the 12th Scientific Meeting of the International Society of Magnetic Resonance in Medicine. Kyoto, Japan: 2004:2120. 102. Atkinson DJ, Burstein D, Edelman RR. First-pass cardiac perfusion: evaluation with ultrafast MR imaging. Radiology. 1990;174: 757–762. 103. Wilke N, Jerosch-Herold M, Wang Y, et al. Myocardial perfusion reserve: assessment with multisection, quantitative, first-pass MR imaging. Radiology. 1997;204:373–384. 104. Ding S, Wolff SD, Epstein FH. Improved coverage in dynamic contrast-enhanced cardiac MRI using interleaved gradient-echo EPI. Magn Reson Med. 1998;39:514–519. 105. Slavin GS, Wolff SD, Gupta SN, Foo TK. First-pass myocardial perfusion MR imaging with interleaved notched saturation: feasibility study. Radiology. 2001;219:258–263. 106. An J, Voorhees A, Chen Q. SSFP arterial spin labeling myocardial perfusion imaging at 3 Tesla. In: Proceedings of the 13th Scientific Meeting, International Society for Magnetic Resonance in Medicine. 2005:253. 107. Kim RJ, Fieno DS, Parrish TB, et al. Relationship of MRI delayed contrast enhancement to irreversible injury, infarct age, and contractile function. Circulation. 1999;100:1992–2002. 108. Kellman P, Arai AE, McVeigh ER, Aletras AH. Phase-sensitive inversion recovery for detecting myocardial infarction using gadoliniumdelayed hyperenhancement. Magn Reson Med. 2002;47:372–383. 109. Huber ME, Kozerke S, Pruessmann KP, Smink J, Boesiger P. Sensitivityencoded coronary MRA at 3T. Magn Reson Med. 2004;52:221–227. 110. Niendorf T, Saranathan M, Lingamneni A, et al. Short breath-hold, volumetric coronary MR angiography employing steady-state free precession in conjunction with parallel imaging. Magn Reson Med. 2005;53:885–894. 111. Weber OM, Martin AJ, Higgins CB. Whole-heart steady-state free precession coronary artery magnetic resonance angiography. Magn Reson Med. 2003;50:1223–1228. 112. Niendorf T, Hardy CJ, Giaquinto RO, et al. Towards single breathhold whole heart coverage coronary MRA using highly accelerated parallel imaging with a 32-channel MR system. Magn Reson Med. 2006;56:167–176. 113. Botnar RM, Bucker A, Kim WY, Viohl I, Gunther RW, Spuentrup E. Initial experiences with in vivo intravascular coronary vessel wall imaging. J Magn Reson Imaging. 2003;17:615–619. 114. Mistretta CA, Wieben O, Velikina J, et al. Highly constrained backprojection for time-resolved MRI. Magn Reson Med. 2006;55: 30–40. 115. Hardy CJ, Darrow RD, Saranathan M, et al. Large field-of-view real-time MRI with a 32-channel system. Magn Reson Med. 2004; 52:878–884. 116. Ohliger MA, Grant AK, Sodickson DK. Ultimate intrinsic signal-tonoise ratio for parallel MRI: electromagnetic field considerations. Magn Reson Med. 2003;50:1018–1030. 117. Wiesinger F, Boesiger P, Pruessmann KP. Electrodynamics and ultimate SNR in parallel MR imaging. Magn Reson Med. 2004;52:376–390. 118. Griswold MA, Kannengiesser S, Heidemann RM, Wang J, Jakob PM. Field-of-view limitations in parallel imaging. Magn Reson Med. 2004;52:1118–1126.

Myocardial Perfusion Imaging Theory Michael Jerosch-Herold and Norbert Wilke

The concept of injecting a tracer into the bloodstream and detecting its transit and distribution in the heart muscle for the assessment of myocardial perfusion is well established in nuclear cardiology and X-ray densitometry. Both exogenous, injected contrast agents and endogenous contrast mechanisms have been used to assess perfusion with cardiovascular magnetic resonance (CMR). The use of a gadolinium-based contrast agent for the assessment of myocardial perfusion with CMR has been extensively validated and successfully applied in patient studies. Recent developments, in particular, the introduction of parallel imaging and high-field (3 Tesla) CMR, have made it possible to combine the requirements for spatial and temporal resolution for myocardial perfusion imaging during the first pass, with multi-slice coverage. The need for quantitative analysis of perfusion studies is also receiving increasing acceptance.1 This chapter reviews the theoretical foundations of myocardial perfusion imaging with CMR to convince the reader that the techniques have matured to a point where they are applicable in clinical studies, despite the additional time required for quantitative analysis. There is already compelling evidence that CMR is superior to nuclear imaging for the assessment of myocardial perfusion.2

THE PHYSIOLOGIC BASIS FOR MEASURING MYOCARDIAL PERFUSION Under normal conditions, the blood flow resistance of the coronary circulation is determined primarily by the myocardial microcirculation, meaning vessels that are smaller than 300 mm in diameter. The adequate supply of oxygen and metabolites to the myocytes is tightly coupled to myocardial blood flow. Adequate and approximately constant blood flow is maintained through autoregulation and can compensate under resting conditions for up to 80% diameter coronary artery stenosis.3,4 With more severe narrowing in an epicardial vessel, and in the absence of significant collateral flow, the distal perfusion bed is fully vasodilated, even under resting conditions, and no further augmentation of blood flow is feasible. In healthy subjects, myocardial blood flow can increase three- to fourfold with maximal vasodilation.5 This means that differences in myocardial blood flow between a region subtended by a stenosed coronary artery and the territory of a normal coronary artery are amplified with maximal vasodilation.

Myocardial perfusion imaging during pharmacologic vasodilation (e.g., with adenosine or dipyridamole) rests on the physiologic observation that the hemodynamic significance of a lesion is most apparent during maximal vasodilation.3,4 A related measure can be obtained in the catheterization laboratory with an intravascular Doppler flow probe by measuring the coronary flow reserve to assess lesion severity6 or by measuring the fractional flow reserve,7 defined as mean distal coronary artery pressure divided by the aortic pressure during maximal vasodilation. These functional tests of coronary and myocardial blood flow overcome the known limitations of diameter vessel lumen measurements by projection of invasive X-ray angiography for determining the hemodynamic significance of epicardial lesions. Furthermore, myocardial perfusion imaging can be used to assess functional impairments in the microcirculation. For example, both myocardial perfusion reserve and coronary flow reserve were abnormally low in women with microvascular dysfunction and without hemodynamically significant epicardial lesions.8,9 There is also evidence from epidemiologic studies that CMR perfusion imaging detects subclinical disease and silent ischemia in subjects without a history or symptoms of atherosclerotic disease. The myocardial perfusion reserve in response to adenosine was found to be associated with coronary risk factors,10 and the perfusion reserve decreases with increasing coronary calcium burden.11 In asymptomatic patients, a lower myocardial perfusion reserve is associated with decreased regional left ventricular function12 and also with lower regional myocardial strain.13 With the development of CMR as a new imaging modality for myocardial perfusion imaging, it has become possible to probe with sufficient spatial resolution for more subtle indicators of regional myocardial ischemia. Blood flow across the myocardial wall is not uniform, but instead favors the subendocardium to accommodate its higher workload and higher rate of oxygen consumption.14,15 Under normal conditions, the ratio of endocardial to epicardial blood flow is approximately 1.15:1. With a coronary artery stenosis, blood flow is first diverted away from the subendocardial layer and the endocardial to epicardial blood flow ratio is often less than 1:1, in particular, under stress.16,17 With myocardial ischemia, the subendocardial layer is accordingly most susceptible to necrosis. This potential advantage of myocardial perfusion imaging can be sufficiently appreciated only if the imaging modality provides spatial resolution on the order of 2 mm or better. With CMR, it has become feasible to detect flow impairments limited to or most accentuated Cardiovascular Magnetic Resonance 57

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in the subendocardial layer.18,19 The spatial resolution of conventional imaging modalities, such as positron emission tomography (PET) and single photon emission computed tomography (SPECT), was insufficient to detect blood flow deficits limited to the subendocardial layer. More specifically, the sensitivity of a myocardial perfusion imaging technique is directly related to its spatial resolution and the ability to discern transmural variations of flow.18,20 The use of CMR offers the unique possibility of quantitatively assessing perfusion, viability, and function with high accuracy to distinguish stunned, hibernating, and infarcted myocardium. Bolli and colleagues showed that even small differences in blood flow during ischemia result in large differences in postischemic function, suggesting that the ability to quantify flow in the low flow range is of importance in predicting the probability of postischemic recovery.21 The extent and incidence of microvascular obstruction observed with CMR was associated with the duration of ischemia before coronary intervention.22

FIRST PASS IMAGING WITH EXOGENOUS TRACERS The success of applying CMR to detect the first pass of an injected contrast agent and to detect perfusion abnormalities starts with an understanding of the contrast mechanisms that allow detection of the contrast agent. The transit of a contrast agent through the vasculature and tissue leads to changes in the longitudinal (T1) and transverse (T2)* relaxation time constants for the detected 1H signal: the contrast reagent is not detected directly, but rather through its effects on the signal from 1H nuclei, mostly in water and lipids. The chelates of paramagnetic ions used as contrast reagents have a relatively high magnetic susceptibility that causes, on a microscopic scale, magnetic field inhomogeneities and also shortens the T2* relaxation rate of water. The contrast reagent also shortens the T1 magnetization recovery after a radiofrequency pulse has disturbed the equilibrium state. The CMR methods for perfusion imaging can be subdivided into T1- and T2*- weighted techniques: T1-weighted techniques produce signal enhancement during the transit of the contrast agent, whereas T2*-weighted techniques cause signal loss. For contrast agents confined to the vascular bed, and in the absence of significant organ motion, T2*-weighted perfusion imaging gives rise to relatively larger signal changes than T1-weighted techniques because the magnetic susceptibility effects of the contrast agents extend beyond the capillaries. However, T2*-weighted imaging techniques have an inherent sensitivity to motion, leading to signal loss in the presence of motion. For cardiac perfusion imaging, T1-weighted techniques such as gradient echo imaging with short echo times and a magnetization preparation for optimal T1-weighting have therefore steadily gained preference. This chapter provides a brief overview of three major sequence techniques for T1weighted perfusion imaging of the heart. They are, in historical order, spoiled gradient recalled echo imaging (GRE), multi-shot T1-weighted echo planar imaging (EPI), and gradient echo imaging with steady state free precession (SSFP). 58 Cardiovascular Magnetic Resonance

Single-shot GRE with magnetization preparation is well suited for T1-weighted, quantitative myocardial perfusion imaging. A magnetization preparation for T1 weighting can take the form of an inversion pulse or a saturation pulse; either of them is generally applied in a non–slice-selective mode so that blood flowing into the image plane during the subsequent image acquisition has been subjected to the same T1 preparation as myocardium. A magnetization preparation with a saturation pulse can be repeated in a sequential multi-slice sequence for each slice, and this will result in the same degree of T1 weighting for each slice. After the magnetization preparation, images for one or more slices are rapidly read out within approximately 200 to 300 msec, or even less with parallel imaging techniques. Typical sequence parameters for such rapid readouts are a repetition time (TR) per phase encoding step of 2.0 msec or less, an echo time (TE) of 1 msec or less, a receiver bandwidth on the order of 800 to 1000 Hz/pixel, and an in-plane spatial resolution of 2 to 3 mm. Because of the short TE, the signal is relatively insensitive to flow and magnetic susceptibility variations. With a linear ordering for the phase encoding steps from high to low spatial frequencies, the image acquisition only needs to be delayed by 10 to 20 msec after a saturation recovery preparation. With a combination of short TR and TE times, the signal initially increases linearly with contrast agent concentration, or as a function of the relaxation rate constant. (The relaxivity of the contrast reagent is not appreciably different between blood and tissue.) Eventually, the signal ceases to increase because of the opposite effect of T2* on the signal at higher contrast concentrations, and also because of the limited dynamic range for T1-related signal changes imposed by the pulse sequence technique. Figure 4-1 shows an example of a CMR perfusion study with a spoiled GRE technique. Because EPI eliminates the need for radiofrequency excitation before each phase encoding step, it is one of the fastest imaging methods for freezing heart motion (50 to 100 msec for single-shot acquisitions). In an EPI pulse sequence, an initial radiofrequency excitation creates a coherent precession of transverse magnetization, and a train of gradient echoes is then generated by applying a rapidly oscillating magnetic field gradient along the readout direction. Each of these gradient echoes is preceded by a short, blipped gradient pulse, applied along a direction perpendicular to the oscillating magnetic field gradient. The blipped phase encoding gradients between echoes advance the trajectory in the frequency acquisition space (k-space) perpendicular to the readout direction. Instead of just a single line, after each radiofrequency excitation, one acquires a set of parallel lines in frequency space. The T2*-related decay of the gradient echo amplitudes in the echo train results in an increasing sensitivity of later echoes to T2* and motion, and causes blurring. The effective echo time can be an order of magnitude longer for EPI sequences compared with fast GRE sequences (e.g., 30 to 40 msec vs. 1.2 to 2.0 msec for fast gradient echo imaging). Unfortunately, a longer effective TE gives rise to magnetic susceptibility and flow artifacts in the ventricular cavity. Furthermore, flow- and susceptibility-related artifacts in the ventricular cavity make it difficult to measure the arterial transit of the contrast agent. To partially capture the speed advantage of EPI and to minimize T2*- and flowrelated artifacts, one reduces the length of the echo trains

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Figure 4-1 The gradient echo cardiovascular magnetic resonance images show the first pass of an extracellular contrast agent (0.03 mmol/kg Omniscan, GE Healthcare Princeton, NJ) in a healthy volunteer, with appearance of the contrast first in the right ventricle (RV), followed by the left ventricle (LV), and finally leading to myocardial contrast enhancement. The contrast agent was injected after acquisition of approximately four pre–contrast images. Images were acquired in the short axis view at the level of the papillary muscle. The arrows show the correspondence between images and the time course of changes in signal intensity for a region of interest in the left ventricle (solid circles) and a myocardial segment in the posterior wall (diamonds). Characteristics of the signal curve for a tissue region, such as its increased dispersion, lower amplitude, and reduced rate of signal enhancement compared with the ventricle, result from transit through the coronary microcirculation, where the volume of distribution is limited to either the extracellular or the intravascular space. a.u., arbitrary units; IV, intravenous.

and the images are read out with several radiofrequency excitations/echo trains.23 This results in a 30% to 40% speed advantage compared with conventional GRE sequences and is particularly useful at field strengths of 1.5 T or lower. (At 3.0 T and higher, any increase in TE beyond the minimum allowed by the gradient system nearly inevitably gives rise to susceptibility artifacts, in particular, during passage of a contrast bolus.) The signal-to-noise ratio and the contrast-to-noise ratio in GRE images are relatively low, when a short TR and wide receiver bandwidths are used. The maximum signal intensity is reached with relatively low flip angles. (The flip angle corresponding to maximum signal intensity is referred to as Ernst angle.) To overcome this limitation, one can use an ingenious scheme, referred to as SSFP, to preserve the coherence of the transverse magnetization between TRs in the pulse sequence and efficiently convert magnetization between transverse and longitudinal orientations. With SSFP, the Ernst angle can approach 90 (i.e., one can reach the maximum theoretical signal amplitude after each radiofrequency excitation while still repeating the excitations with very short TRs). In fact, the TRs should be as short as feasible because the susceptibility of the SSFP technique to any off-resonance shifts increases with TR and also with flow.24 Off-resonance frequency shifts result in dark bands at locations where the frequency shift (in radians per second) times TR (in seconds) equals an odd

multiple of p/2. Herein lies the Achilles heel of the SSFP technique, and not surprisingly, the passage of a contrast bolus can exacerbate frequency shifts and artifacts. As absolute frequency shifts increase with field strength (assuming the same relative field homogeneity at different field strengths), one can best reap the advantages of SSFP techniques for perfusion imaging at field strengths of up to 1.5 T. Because of the relatively good signal-to-noise ratio, the SSFP techniques can produce more appealing image quality,25 but the use of the SSFP technique for myocardial perfusion imaging must be approached with caution because the artifacts on SSFP images can be deceptive and mimic hypoperfusion. Ultrafast T1 measurements would provide the most accurate estimates of contrast agent concentration, and with the advent of parallel imaging, this approach may become practical. Chen and coworkers developed such a T1 fast acquisition relaxation mapping (T1-FARM) method to obtain single-slice T1 maps of the heart with 1-sec resolution.26,27 With direct measurement of T1, the problems associated with the saturation of the signal with increasing contrast agent dosage can be avoided. This should allow for more accurate quantification of blood flow with usage of higher contrast agent dosages. Figure 4-2 shows two signal curves for a region of interest in the left ventricle measured during bolus injection of 0.075 mmol/kg gadolinium diethyl triaminepentaacetic acid (Gd-DTPA) with the Cardiovascular Magnetic Resonance 59

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Figure 4-2 Comparison of signal time curves for a region of interest in the left ventricle (LV) obtained in a canine with quantitative T1 imaging (T1-FARM) and T1-weighted, saturation recovery prepared fast gradient echo (TR/TE ¼ 2.4/1.2 msec; flip angle ¼ 18 ) and a gadolinium diethyl triaminepentaacetic acid (Gd-DTPA) dosage of 0.075 mmol/kg at 1.5 T. The curves were normalized so the TurboFLASH and T1-FARM recirculation peaks were equal. a.u., arbitrary units, (Data courtesy of Z. Chen, C.A. McKenzie and F.S. Prato, St. Joseph’s Hospital, London, Ontario.)

T1-FARM and the saturation recovery prepared GRE techniques, respectively. For better temporal resolution, to achieve multi-slice coverage, and to allow simpler image reconstruction, in the future, parallel imaging methods will be used to accelerate T1-FARM image acquisition. All current CMR techniques offer higher spatial resolution than is available with nuclear imaging techniques. An illustration is shown in Figure 4-3 from studies in an experimental animal model.

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Parallel imaging refers to the use of independent receiver channels for the parallel acquisition of data from multiple elements of a receiver coil array (see Chapter 3). Parallel imaging techniques have led to a leap forward in image acquisition speed. Each element of a receiver coil array has a characteristic spatial sensitivity profile. In the ideal case in which the signal profiles do not overlap, it becomes possible to localize a signal within the dimensions of the coil element, without gradient encoding. By measuring the local coil profiles, one can reduce the number of phase encoding steps and during image reconstruction replace the missing phase encodings with information related to the coil profiles. The algorithms can operate on the sensitivity encoded images calculated for the individual coil elements (SENSE)28,29 or on the spatial frequency data acquired with the coil elements (simultaneous acquisition of spatial harmonics [SMASH]).30 Theoretically, the number of phase encoding steps can be reduced by a factor that equals the number of independent coil elements in the phase encoding direction. In practice, and in particular for imaging techniques with relatively poor signal-to-noise ratio, such as those used for myocardial perfusion imaging, speed-up factors on the order of 2 are more realistic.31 Such a speed-up results in image acquisition times with GRE techniques on the order of 100 to 150 msec/image, with an in-plane spatial resolution on the order of 2 mm. During pharmacologic stress and with heart rates of approximately 100 bpm, one can cover the heart with 4 to 5 slices during stress and up to 10 slices at rest.31

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Figure 4-3 Images from cardiovascular magnetic resonance imaging (CMR), positron emission tomography (PET), and (TTC) staining in a porcine model in which obtuse branches of the left circumflex coronary artery had been ligated 4 weeks before combined MRI and PET studies. From CMR acquired during the first pass of an intravascular iron oxide contrast agent (NC100150 injection, Nycomed), the one shown at left corresponds to the highest peak signal enhancement in tissue and indicates (pointed out by arrow) a subendocardial perfusion defect in the posterior segment, in agreement with 2,3,5-Triphenyltetrazolium chloride (TTC) staining shown at right with sub-endocardial absence of staining in sector highlighted by arrow. 13NH2 positron emission tomography (middle image) was carried out 3 hours before the MRI study and shows a fixed defect in the posterior segment (arrow), suggesting a transmural infarct. Fluorodeoxyglucose-positron emission tomography images (not shown) also indicated irreversible damage in the posterior segment, in disagreement with the findings from MRI and TTC staining. ANT, anterior; RV, right ventricle. 60 Cardiovascular Magnetic Resonance

WATER EXCHANGE AND ITS EFFECTS ON MYOCARDIAL CONTRAST ENHANCEMENT With 1H magnetic resonance imaging, the presence of a contrast agent is detected indirectly through the change in the 1 H relaxation recovery rates, T1 and T2. The effect of a contrast agent on T1 and T2 within a single uniform closed space can be described by its relaxivity, r1,2, using the following equation: T1;2 ¼ r1;2 ½CR. Here, [CR] denotes the concentration of the contrast reagent (e.g., in units of millimoles), and r1 and r2 are the relaxivity in units of 1/ (mM s). We only focus here on T1 effects because of the predominant usage of T1-weighted techniques for myocardial perfusion imaging. In a compartmentalized system such as myocardial tissue, with barriers that are permeable to water, one has to consider not only the local interaction of water with the contrast agent within a compartmental space, but also its exchange between, for example, the plasma space and the interstitial space. For an extracellular agent, the exchange of water between the interstitial and intracellular spaces also comes into play. Water exchange across the capillary barrier and the cytolemmal membrane, if sufficiently fast, allows 1H spins to sample different environments during spin relaxation. As a result, the observed relaxation recovery rate is no longer determined solely by local contrast

agent concentration, but depends also on the rates at which H spins exchange between spaces and by the relative volumes of those spaces. Thus, even though a contrast agent may be confined to the vascular space, as in the case of an intravascular contrast agent, the 1H spins outside the vascular space can enter the plasma space and relax at a faster rate than in the absence of water exchange. (The reverse process, of 1H spins leaving the vascular space and relaxing more slowly in the interstitial space, is equally likely.) Two limiting cases are typically considered: (1) the slow or no exchange limit, where the 1H spins relax at a rate intrinsic to the space they dwell in, and (2) the fast exchange limit, where 1H spins in spaces linked by fast water exchange relax at the same rate, which represents a weighted average of the intrinsic relaxation rates. The limiting cases of no exchange and fast exchange provide convenient simplifications for the analysis of myocardial contrast enhancement. If signal curves for myocardial regions of interest are analyzed with the no exchange assumption, then any water exchange will lead to overestimation of the volume of distribution of the contrast agent.37 Conversely, if the signal curves are analyzed with the fast exchange assumption, then a lower rate of exchange, intermediate between the fast and no exchange limits, will result in an underestimation of the contrast distribution volume and myocardial blood flow.38 Which limiting case is closer to the true setting depends on the experimental setting. Previous experimental studies suggest that for the capillary barrier, the rate of water exchange is on the order of 7 Hz or less,39 and a study with an intravascular agent gave an estimate of the intravascular lifetime of a water molecule on the order of 0.3 seconds.40 It would appear that neither the fast exchange limit nor the slow exchange limit approximates well the effects of water exchange in myocardial tissue, at least when the relaxation recoveries proceed undisturbed. Nevertheless, typical T1 time constants are in the range of 100 to 1000 msec in myocardial tissue, depending on contrast agent concentrations, and these T1 values are much larger than the typical TR used for rapid image readout. The application of radiofrequency pulses during a relaxation recovery can be expected to alter the relaxation recovery. Effectively, the rapid application of radiofrequency pulses during a relaxation recovery can reduce the sensitivity of the observed 1H signal to water exchange.41 With a sufficiently high flip angle (20 ), the no exchange assumption is reasonable for the interpretation of the signal curves measured with an ultrafast (TR ¼ 2.4 msec) saturation recovery prepared GRE sequence.42 In principle, it is also possible to incorporate a description of water exchange into a tracer kinetic model, as shown recently by Li and associates.43 The optimal approach to address the combination of contrast enhancement and water exchange remains an area of active research. 1

ENDOGENOUS CONTRAST FOR THE ASSESSMENT OF MYOCARDIAL PERFUSION Approaches have been developed for an assessment of myocardial perfusion that are not based on the use of an exogenous CMR contrast agent, but instead exploit endogenous contrast mechanisms related to blood flow and blood Cardiovascular Magnetic Resonance 61

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Another remarkable advance for fast CMR perfusion imaging is the development of sparse sampling techniques for dynamic imaging applications. An example is the unaliasing by Fourier encoding the overlaps using the temporal dimension (UNFOLD) technique, which involves a trade-off of temporal sampling to allow sparser spatial sampling.32,33 Without any other adjustments, a halving of the field of view would speed up the image acquisition by a factor of two, but would result in foldover artifacts. To unfold the overlaps (unaliasing), Madore and associates32 proposed that complementary sets of phase encoding steps are interleaved in a dynamic imaging series. The alternation between two sets of phase encoding steps results in a modulation of the aliased image components, whereas un-aliased image components remain constant. The aliased image components can be removed by filtering of the temporal frequency spectrum at each pixel location. The UNFOLD approach has been applied by Di Bella and colleagues in myocardial perfusion studies,34 and was found to work well if patients hold their breath because respiratory motion interferes with UNFOLD. A further generalization of sparse sampling in the k-space and time domains is provided by the broad-use linear acquisition speed-up technique (BLAST).35,36 Briefly, with k-t BLAST, one views the acquisition as a sampling process in a higher dimensional k-t space. Different sets of phase encodings are acquired at successive time points, and the repetition period for each phase encoding set equals the k-t BLAST acceleration factor. Image aliasing is removed by use of filtering or prior information, and as a final result, one obtains a series of images over time, t. Although these sparse sampling techniques have only been applied in initial feasibility studies, it is safe to assume that they will play an increasingly prominent role in myocardial perfusion imaging and will allow unprecedented spatial resolution.

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

oxygenation. These approaches fall under the categories of spin labeling,44 blood oxygen level dependent (BOLD) contrast,45,46 and magnetization transfer contrast.47,48 These techniques have to be considered technically more challenging than measurements with exogenous contrast agents, or they offer only an indirect measure of blood flow. With spin labeling, the spins are either inverted or saturated in a slab that is generally located upstream of the imaged slice.49–51 The flow-dependent change of signal intensity caused by the inflow of saturated or inverted spins into the image slice provides a measure of tissue perfusion. The spin labeling method generally relies on the assumption that the net arterial blood flow in the myocardium follows the direction from base to apex. Cardiac motion complicates the interpretation of signal changes in a spin labeling experiment because the labeled spins can be transported into the imaged slice either through blood flow or by through-slice-plane motion of the heart. The BOLD technique offers a measure of hemoglobin saturation that reflects regional oxygen supply and demand (see Chapter 42). Deoxyhemoglobin is paramagnetic, whereas oxyhemoglobin is only diamagnetic. This means that deoxyhemoglobin causes a considerably larger reduction of T2* than oxyhemoglobin and therefore a larger signal intensity attenuation in GRE images. Deoxyhemoglobin can be used as an endogenous intravascular tracer because of the tight coupling between oxygen demand and blood flow. BOLD contrast changes have been observed in the heart after administration of dipyridamole and dobutamine.45,46 The link between BOLD contrast and blood flow depends on the balance between blood flow and oxygen metabolism (i.e., between oxygen supply and demand).52 Spin labeling, BOLD contrast, and magnetization transfer contrast CMR can be carried out in the steady state, meaning that there is no need to capture with ultrafast imaging the transient signal changes observable in the myocardium after injection of an exogenous tracer. Ultimately, these techniques may allow an assessment of myocardial perfusion with high spatial resolution ( 1 mm or less).

QUANTITATIVE EVALUATION OF MYOCARDIAL PERFUSION The images acquired in a myocardial perfusion study with a bolus injection of contrast agent can be qualitatively evaluated based on the level of differential signal enhancement in the myocardium. With a T1-weighted imaging technique, myocardial segments showing reduced signal intensity enhancement during the first pass, relative to other myocardial segments, are interpreted as hypoperfused. This type of qualitative judgment of signal enhancement differences can suffer from substantial observer bias, and small reductions in perfusion can be missed. Although higher dosages of contrast agent would increase the dynamic range, the images are also more likely to be contaminated by artifacts near the endocardial border. Furthermore, global reductions of blood flow caused by diffuse microvascular ischemia will be missed by this approach. Therefore, it becomes necessary to evaluate the contrast state during stress or vasodilation, relative to a baseline or resting state, to uncover global impairments of perfusion. 62 Cardiovascular Magnetic Resonance

With an appropriate imaging technique, such as a fast T1-weighted gradient echo sequence and low contrast agent dosages (e.g., 0.03 mmol/kg Gd-DTPA for an intravenous bolus injection into an antecubital vein), the signal changes are proportional to the local contrast agent concentration. Under these conditions, the signal curves can be interpreted as or transformed into contrast agent residue curves. Absolute units for the signal intensity (or related equivalent such as contrast concentration or relaxation rate), do not matter here, because enhancement in the tissue is interpreted relative to the enhancement in the blood pool, both measured on the same linear scale. In fact, for the purpose of quantifying perfusion, the tissue is often viewed as a linear, stationary system, where enhancement in the tissue can be modeled as a linear response to arterial input of contrast. Over the years, investigators using PET have presented compelling arguments for a quantitative approach to myocardial perfusion imaging, including absolute quantification of myocardial blood flow.53 With CMR, a quantitative analysis of contrast enhancement is based on signal intensity curves, which can be generated for user-defined regions of interest or at the pixel level. CMR perfusion studies require a careful approach for image segmentation and registration. CMR perfusion studies are not motion averaged, the endocardial borders are relatively well defined during transit of contrast through the ventricular cavity, and the large signal enhancement in the ventricular cavity requires accurate image segmentation to avoid spillover effects. Figure 4-4 shows an example of a CMR perfusion study in a patient that was performed for absolute quantification based on model-independent analysis of signal intensity curves.54 In the vein of Zierler’s55–57 original work on the central volume principle, we consider here a tissue region of interest with a single arterial input and a single venous output. The injection of a tracer is described by the variation of tracer concentration at the (arterial) input cin(t). The amount of tracer that remains in the tissue region at any time t, q(t), represents the difference between the amount supplied to the tissue region up to time t, and the total amount of tracer that has exited the region of interest up to time t, cout(t): ðt qðtÞ ¼ F ½cin ðsÞ cout ðsÞ dt

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Figure 4-4 A 71-year-old man presented with a silent non–Q wave anterior myocardial infarction. A, An image from a series of 60 dynamic images taken at rest shows the initial contrast enhancement in the myocardium, with reduced contrast uptake in the anterior and anterior septal segments (black arrow) because of an eccentric 90% stenosis with a lesion length of 10 mm in the proximal left anterior descending coronary artery. B, Late gadolinium enhancement was observed in the anterior and anterior septal segments (white arrow) from midlevel to apex, but not at a more basal level, including the level of the slice in shown in A. C, The upper panel shows the change of signal intensity (after subtraction of the baseline offset) in two myocardial segments (solid circles for lateral-posterior segment and inverted triangles for anterior-septal sector). The curve in the lower panel shows the changes of signal intensity in a central region of the left-ventricular blood pool and depicts the arterial input of contrast. Myocardial blood flow was quantified by model-independent deconvolution of the myocardial signal intensity curves, using an arterial input measured in the left ventricular blood pool. The solid lines in the upper panel of C show the calculated tissue response, which agree well with the measured data points. Blood flow at rest in an anterior septal segment was 0.7 mL/g/min, compared with 1.2 mL/g/min in the posterior wall. The anterior and anterior septal segments appeared hypokinetic on cine magnetic resonance imaging. After cardiovascular magnetic resonance, the patient was referred for coronary angiography. During percutaneous coronary intervention, a drug-eluting stent was deployed in the proximal left anterior descending coronary artery. a.u., arbitrary units; LAT, lateral; LV, left ventricle; RV, right ventricle; SI, signal intensity.

input is a dirac delta impulse, cin(t) ¼ d(t), and requiring that cout(t ¼ 0) ¼ 0 (i.e., the tracer cannot instantaneously reach the output after injection). This property of the impulse response is independent of the vascular and compartmental structure inside the tissue region of interest or the properties of barriers, such as their permeability. For example, leakage of a contrast agent from the capillaries into the interstitial space produces a redistribution of the contrast agent within the region of interest, but does not contribute to flow into and out of the region of interest, as long as transport by diffusion and convection is much slower than by blood flow. For first pass CMR contrast agents such as Gd-DTPA, these assumptions hold up well, but for freely diffusible tracers, they break down. The central volume principle indicates that blood flow can be determined by deconvolution of the measured tissue residue curve with the arterial input curve. The deconvolution analysis is quite sensitive to noise in the data, and one therefore needs to constrain the deconvolution operation. For example, the impulse response, R(t), should be a monotonically decaying function of time because by definition of the impulse input, no additional tracer enters the tissue region after the initial delta function input. Furthermore, the impulse response should be reasonably smooth. Smoothness of the impulse response follows to the degree that the magnitudes of flow, diffusion, and permeation in the heart impose time scales that exclude abrupt signal intensity changes in myocardial tissue after contrast agent

injection. A Fermi function has been used as an empirical model for the impulse response to fit the first pass portion of the signal curves.42,58 The Fermi model of the impulse response has the following equation: A ; (3) 1 þ exp½ðt wÞ=t where A, w, and t are the model parameters, with no direct physiologic meaning. From an empirical standpoint, the shape of the Fermi function provides a reasonable approximation of the shape of the impulse response of an intravascular tracer. The portion of the signal intensity curve up to the peak has higher sensitivity to flow changes than later phases of contrast enhancement, and the early enhancement is relatively insensitive to capillary leakage of contrast agent. This initial part of the signal intensity curve can be approximated well by the Fermi function model to assess blood flow. The Fermi model for the quantification of myocardial perfusion and perfusion reserve was validated by comparison to blood flow measurements with labeled microspheres59,60 and also by comparison to invasive coronary flow reserve measurements,42 with similar regional perfusion in all segments of approximately 1 mL/kg at rest.61 As an alternative to the deconvolution analysis, one can build a mathematical model of tracer transport through tissue. For an extracellular contrast agent, this involves exchange of tracer between vascular and interstitial spaces. For a compartmental model, it is assumed that the tracer is RðtÞ ¼

Cardiovascular Magnetic Resonance 63

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1.19 mL/g/min (lateral-posterior) LAT

these models tends to be rather large. Sensitivity analysis is called for to determine which (preferably few) model parameters need to be adjusted for a best fit of the measured residue curves. The number of degrees of freedom can be reduced by using an intravascular contrast agent.63,64 The realism of spatially distributed models of blood tissue exchange offer an advantage for the identification and quantification of physiologic changes observed indirectly with

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BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

well mixed within each compartment, meaning that the contrast agent concentration equilibrates instantaneously within each compartment. The rapid injection of a tracer will create a tracer concentration gradient from the arterial to the venous side. Spatially distributed models are described by partial differential equations that account for the variation of tracer concentrations within each tissue region as a function of time and at least one spatial variable.62,63 The number of parameters in

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Figure 4-5 A, Conceptual diagram of the spatially distributed, two-region model that was implemented in the JSIM (JSIM, Java-based Simulation and Modeling environment) simulation environment (http://www.physiome.org/jsim/) for analysis of tracer residue curves and quantification of myocardial blood flow. The following model parameters are optimized for fitting to experimental data: plasma flow (Fp), plasma volume (Vp), and capillary permeability surface area product (PScap). B, A tissue contrast residue curve measured in a 56-year-old female volunteer during vasodilation with adenosine (open circles) was analyzed with the model described in A, and the best fit corresponding to a tissue blood flow of 2.8 mL/g/min is shown as a solid line. Additional model curves for flows of 1, 2, and 4 mL/g/min and for otherwise unchanged model parameters are shown as dashed lines to illustrate the sensitivity of the initial contrast enhancement to blood flow. The rate of the initial contrast enhancement (“up-slope”) changes as the flow is varied, whereas the curves cluster together after the first pass, when sensitivity to flow changes is low. Also, for equal increments of blood flow, the peak contrast enhancement changes less as the absolute magnitude of flow increases. C, The measured arterial input was replaced by an impulse input, with approximately the same area under the curve. Now the amplitude of the simulated tissue response to the impulse input increases in proportion to the blood flow (F impulse amplitude), as predicted by Zierler’s central volume principle. a.u., arbitrary units; LV, left ventricle. 64 Cardiovascular Magnetic Resonance

ARTERIAL INPUT FUNCTION Quantification of myocardial blood flow generally requires the measurement of an arterial input function, which serves as reference for analysis of the myocardial contrast enhancement. The term arterial input function refers to the measurement of contrast enhancement at a location before it enters the perfusion bed under study. Under ideal conditions, the input function would be measured immediately before blood enters the perfusion bed, but with CMR, it is not feasible to measure reliably the arterial input, even in the epicardial vessels. Therefore, one has to settle for measurement at a more upstream location, either near the aortic root or in the left ventricle. With contrast-enhanced CMR, one faces a further constraint in the limited dynamic range of contrast enhancement, which results in signal saturation at higher contrast concentrations, as illustrated in Figure 4-6A. Nevertheless, the dynamic range can be tuned to the concentration range with the sequence parameters, and it is indeed possible to achieve approximate linearity between signal change (above precontrast signal) and tracer concentration by using short inversion times ( 50 msec). However, the same tuning to short TI

values is not optimal for capturing myocardial contrast enhancement because the tracer concentrations are lower than in the blood pool and the short time-after-inversion (T1) values decrease the sensitivity to small relaxation rate constant (R1) changes. Figure 4-6 shows the effects on contrast enhancement of the delay time after a saturation preparation. There are essentially three solutions to the problem of signal saturation: (1) Dual bolus technique: one performs two injections with a low and a relatively higher contrast dosage, each optimized to measure contrast enhancement in the myocardium and the blood pool, respectively, but with otherwise identical injection protocols and under stable hemodynamic conditions.59 The signal curves for a blood pool region of interest are scaled according to the ratio of contrast dosages used in the dual bolus protocol, and then further quantitative analysis can be performed on curves for tissue and blood pool with an optimized contrast-to-noise ratio. (2) As a second alternative, one can include in the pulse sequence a rapid low-resolution image acquisition with very short inversion time to measure the arterial input with minimal signal saturation.65,66 Such pulse sequence techniques with arterial input scout impose a small time penalty, but otherwise show promise for keeping the acquisition protocol simple and not requiring dual bolus injections. (3) Finally, it is often feasible to construct calibration curves of signal versus (R1), or tracer concentration, to correct for signal saturation,67 as long as the calibration curves are monotonically increasing as a function of R1. An example of calibrationbased saturation correction appears in Figure 4-6C.

TR/TE = 2.3/1.0 msec Flip angle = 18o

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Figure 4-6 A, Model-based calculations based on the Bloch equations were used to calculate contrast enhancement (CE) in the left ventricular (LV) blood pool for a range of relaxation rate constants (R1), assuming a precontrast T1 value of 1100 msec at 1.5 T. The delay time (TD) between saturation preparation and gradient echo image readout varied between 10 and 300 msec and had a marked effect on the nonlinearity of the contrast enhancement curves. Short TD values are most appropriate for measuring contrast enhancement for a wide range of R1 values, as encountered during the first pass of a contrast agent in the blood pool. Longer TD values result in better contrast-to-noise ratios. The dark blue line corresponds to the ideal case in which contrast enhancement is linearly proportional to R1 and contrast agent concentration. B, Modelbased calibration curves can be used to convert the observed contrast enhancement to the ideal case in which contrast enhancement is linearly proportional to contrast concentration. Such calibration curves can be generated from the simulations shown in A, using the sequence parameters and precontrast T1 values of an actual study. C, With the model-based calibration curve in B, a measured curve of left ventricular contrast enhancement was corrected for saturation effects. Left ventricular contrast enhancement was measured with a bolus of 0.05 mmol/kg gadolinium diethyl triaminepentaacetic acid (Magnevist, Berlex, Wayne, NJ). With this dosage, the peak signal measured in the left ventricle (LV) shows approximately 30% saturation. The contrast concentration change in the blood pool would be underestimated by 30% if the measured signal were taken at face value and not corrected for saturation. TE, echo time; TR, repetition time. Cardiovascular Magnetic Resonance 65

4 MYOCARDIAL PERFUSION IMAGING THEORY

tomographic imaging techniques such as CMR. An example of the use of a tracer kinetic model for analysis of a CMR perfusion study is shown in Figure 4-5, including an illustration of the central volume principle, based on model simulations.

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

It may be possible in the future to control the arterial input of an injected contrast agent in a manner similar to the approach used by arterial spin labeling or in myocardial contrast echocardiography with echogenic bubbles that can be destroyed by ultrasound pulses. For CMR, hyperpolarized contrast media68 could be one possible approach because it is relatively easy to selectively quench the hyperpolarized signal.

PRACTICAL ASPECTS OF CARDIOVASCULAR MAGNETIC RESONANCE PERFUSION IMAGING This chapter has mostly focused on theoretical aspects of myocardial perfusion imaging, and the presentation was in part motivated by a desire to provide useful insights for optimizing myocardial perfusion studies and for avoiding some pitfalls. We now summarize some practical rules of thumb for myocardial perfusion imaging protocols that have proven to be of relevance in practice. 1. Contrast agent injection. Because of the dependence of myocardial contrast enhancement on the arterial input of contrast, it becomes important to inject the tracer sufficiently fast so that different levels of tissue blood flow can be sensitively discriminated. The use of a power injector with an injection rate of 3 mL/sec, in particular, for studies during coronary hyperemia, is important. 2. Temporal resolution of measurements. The temporal resolution should allow updating of images from every slice at a rate equal to the heart rate. This rule of thumb automatically compensates for the higher temporal resolution requirements during stress compared with rest. 3. Spatial resolution. Mild reductions of myocardial blood flow as a result of an epicardial lesion manifest themselves first in the endocardial layer. A spatial resolution of 2.5 mm or less should be used to resolve transmural variations in blood flow. For rest studies, the temporal resolution can be reduced from one R-R interval to two R-R intervals, to accommodate the requirements for spatial resolution, but for hyperemia, the one R-R interval tends to be closer to the critical limit for temporal resolution. 4. The CMR sequence for contrast residue detection. T1weighted imaging can be rendered relatively insensitive to the effects of water exchange by the use of higher flip angles (above the Ernst angle), and this simplifies the conversion of signal time curves into residue curves. There is some evidence that avoidance of high contrast dosages (e.g., > 0.1 mmol/kg for Gd chelates) and high spatial resolution may be effective measures against artifacts and the appearance of a transitional dark rim at the endocardial border. The rationale for these measures is that: (1) the susceptibility difference between blood loaded with contrast and tissue gives rise to signal loss at the endocardial border, which is less likely to be observed with lower contrast dosages; and (2) Gibbs

66 Cardiovascular Magnetic Resonance

ringing, which may be one source of the dark rim artifact, is reduced by increasing the number of sampling points (i.e., spatial resolution, in this case). 5. Measurement of arterial input. For a semiquantitative or quantitative analysis, contrast dosages need to be kept below approximately 0.05 mmol/kg body weight for Gd chelates. 6. Parallel imaging. For near-complete coverage of the heart, the use of parallel imaging techniques has become a blessing and a necessity. Nevertheless, myocardial perfusion studies are relatively signal-intensity-“starved,” and for this reason, acceleration factors have to be used conservatively to avoid excessive noise and blurring. Acceleration factors on the order of two are recommended, although higher factors may work well when SSFP techniques are used. 7. Analysis and modeling. The rate of contrast enhancement in a myocardial tissue region (“up-slope”) has empirically been proven to provide a relatively good surrogate measure of blood flow, but it should be appropriately normalized by a measure of contrast enhancement at the arterial input.69 The resulting perfusion index can be determined for rest and stress, and the ratio can provide an approximate measure of the perfusion reserve. Several CMR equipment vendors have now included this up-slope method in their cardiac perfusion analysis packages. Other parameters, such as the maximum amplitude of contrast enhancement, are less sensitive than the up-slope parameter to changes in tissue blood flow. Absolute measures of myocardial blood flow can be determined if contrast enhancement is linearly proportional to contrast concentration in blood. Blood flow estimates are obtained by deconvolution of the tissue curves with arterial input, either by using a model-independent approach, or by modeling. Currently, these latter approaches for blood flow quantification are used only in a research setting, and commercial software is not yet available.

CONCLUSION In the context of assessing blood flow in the myocardial microcirculation, the CMR first pass imaging technique can address several clinical issues, if the appropriate conditions for spatial and temporal resolution can be met. Valuable quantitative and largely observer-independent information related to the pathophysiology of ischemic heart disease becomes available with myocardial perfusion imaging. CMR perfusion reserve measurements can be applied to assess the degree of atherosclerosis, to assess the functional significance of coronary artery lesions, and to evaluate myocardial viability. In milder forms of ischemic heart disease and ischemic cardiomyopathies, CMR at rest and during stress is a good test to probe for more subtle perfusion limitations that may be limited to subendocardium or that result in only moderate blood flow reductions. For these applications, quantitative CMR methods have been successfully developed and validated to provide a true measure of myocardial blood flow.

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Cardiovascular Magnetic Resonance Tagging Assessment of Left Ventricular Diastolic Function Matthias Stuber

Left ventricular (LV) diastolic function has been recognized as an important factor in the pathophysiology of many common cardiovascular diseases. Dilated and hypertrophic cardiomyopathies, coronary artery disease, and systemic hypertension are all associated with abnormal LV filling dynamics. Diastolic dysfunction has also been increasingly appreciated as a major cause of heart failure, especially in the elderly. Although invasive hemodynamic measures/assessment of diastole are considered the gold standard, echocardiographic methods, including tissue Doppler imaging, have gained greater use in the clinical assessment of LV diastolic function because of their noninvasive acquisition, which greatly facilitates serial assessments. With the advent of parallel imaging, cardiovascular magnetic resonance (CMR) techniques, three-dimensional (3D) data collection, spiral and steady-state free precession (SSFP) imaging, and higher magnetic field strength, the prolonged tag persistence permits easier access to diastole. Together with rapid state-of-the-art software analysis tools, quantification of LV diastolic wall motion can now easily be performed.

CARDIAC MOTION During systole, the heart performs a complex motion analogous to “wringing” a towel, and as a result, the base and apex rotate in opposite directions. There is counterclockwise rotation at the apex and clockwise rotation at the base.1 In parallel, the atrioventricular valvular plane (basal LV and right ventricle) descends toward the apex. (The apex is relatively stationary.) The lateral free wall of the right ventricle performs a more pronounced long axis contraction than the lateral wall of the LV.2 During isovolumic relaxation, myofibrils return to their resting state from the contracted state. This process is accompanied by a rapid untwisting at the apex, whereas the volume and cavity shape of the heart remain nearly unchanged. This rapid untwisting typically lasts less than 75 msec and precedes the early passive filling phase of the ventricles. During this filling phase, practically no rotational components can be seen at the apex of the healthy heart.3

Assessment of Cardiac Rotation/ Motion: Non-Cardiovascular Magnetic Resonance Methods With echocardiographic imaging, the myocardium has relatively poor internal structure because of the absence of structural landmarks, making quantification of parameters, such as rotation, stress, and strain, limited.4,5 Several invasive methods to examine cardiac motion during diastole have been reported. One approach is the surgical/invasive implantation of tantalum markers into the midwall of the myocardium.6 In combination with X-ray angiography, the motion of these markers can then be recorded with high temporal and spatial resolution. Using such an approach, alterations in diastolic untwisting have been observed in patients who have undergone heart transplantation shortly before rejection.7 Although this method is very powerful, it is invasive, requires ionizing radiation, and is inappropriate for routine clinical use. Alternative angiographic “markers,” such as tracking of the bifurcations of the coronaries,8 suffer from the limited number of landmarks and their irregular geometric distribution. Furthermore, they only provide motion information about the epicardial layers of the myocardium.

Assessment of Cardiac Rotation/ Motion: Cardiovascular Magnetic Resonance Methods Similar to echocardiography, conventional cine CMR images do not provide information about the internal structure of the myocardium. However, a CMR myocardial tagging technique, spatial modulation of magnetization (SPAMM), originally proposed by Zerhouni9 and Axel10 and further developed and refined by others, offers the opportunity to assess strain noninvasively. With these methods, the magnetization of the muscle tissue is spatially modulated, or “tagged,” by the application of a specific time series of radiofrequency (RF) pulses and magnetic field gradients. The tagging is typically applied immediately after the R-wave of the electrocardiogram. Images are then Cardiovascular Magnetic Resonance 69

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acquired during successive heart phases in which the tags may be identified as dark lines or grids. Because these tags are spatially fixed with respect to the muscle tissue at the time of the tag application, local myocardial motion can be derived from the translation, rotation, and distortion of the tags on the myocardium. However, because of the relaxation effects, the tags fade rapidly and cannot be reliably detected after end systole (approximately 300 msec after tag application). This is a serious drawback for the quantification of systolic and diastolic dynamics of the heart wall. Another limitation is that this approach does not compensate for through-plane motion. The problems associated with tag fading as well as through-plane motion2 can be overcome by the application of complementary spatial modulation of magnetization (CSPAMM) CMR tagging approaches.3,11 This technique is based on a subtraction of two acquisitions. For both acquisitions, the magnetization of the tissue is modulated or “tagged” locally, and a thin slice and a thick slice are imaged during the subsequent imaging procedure. If the modulation function preceding the second acquisition is inverted with respect to the first tagging procedure, subtraction of the two acquisitions leads to signal that is derived only from the initially tagged thin slice. This approach both prolongs the persistence of the tags until late diastole and assures that the same tissue plane is imaged in the multiple heart phase images. In this chapter, the technique is described in more detail and initial results associated with diastolic motion components of the human heart are discussed.

METHODS Complementary Spatial Modulation of Magnetization and Slice Following The CSPAMM myocardial tagging technique involves the periodic modulation of the magnetization in a thin slice of the myocardium (dz; Fig. 5-1). A sinusoidal modulation of the magnetization is typically performed immediately after the detection of the R-wave of the electrocardiogram.

dz

ds

Figure 5-1 Slice-following principle. An initially tagged planar slice of the thickness dz translates and distorts during the cardiac cycle. A volume of the thickness ds is imaged multiple times during the cardiac cycle. This volume must encompass the potential extent of the motion of the tagged, thin slice. 70 Cardiovascular Magnetic Resonance

Subsequently, a thick slice (ds; see Fig. 5-1) encompassing the expected full extent of motion of the selected thin slice is imaged periodically (n heart phase images) during the cardiac cycle. The procedure, consisting of labeling of a thin slice and subsequent imaging of a thicker volume, is performed twice with an inverted modulation of the magnetization for the second experiment. Subtraction of the two acquisitions leads to an image derived from the signal coming from the labeled part of the magnetization in the thin slice. Because of the subtraction procedure, the signal coming from the thick volume outside the tagged slice (ds; see Fig. 5-1) is suppressed. Therefore, the problem of through-plane motion is eliminated because only a projection of the same tissue elements is visualized in the images. The signal from the tagged thin slice can be decomposed into two parts, with the first part holding the tagging information and a second part (responsible for the fading of the tags) built up as a function of time. For CSPAMM, this second component is also suppressed by the subtraction procedure. As a result, only signal derived from the tagged component of the magnetization remains after subtraction, and fading of the tags is avoided. For a constant tag contrast and a maximized signal-to-noise ratio in systolic and diastolic images, a series of ramped RF excitation angles must be used.11 Typically, double oblique short axis sections of the myocardium are tagged with a slice thickness of 6 to 8 mm. Subsequently, 16 to 20 sequential heart phase images are acquired with a temporal resolution (Dt) of 35 msec. With this high temporal resolution, rapid motion components, such as diastolic untwisting at the apex, can be identified readily. Because the ratio of wanted to unwanted signal components must be optimized, the thickness of the imaged volume (ds; see Fig. 5-1) must be reduced to a minimum. Therefore, it depends on the level of the tagged slice with respect to its level on the long axis. For basal LV images, where a long axis contraction of more than 20 mm may be expected for the lateral free wall of the right ventricle,2 a slice thickness of 30 mm is typically chosen. For equatorial slices, a slice thickness of 25 mm is appropriate, and at the apex, a slice thickness of 20 mm is used because of reduced through-plane motion. For suppression of breathing-induced motion artifacts, a repetitive breath hold scheme3,11 or single breath hold techniques12 can be applied. Considering the location of the relevant tagging information in k-space, a reduced k-space acquisition scheme can be applied.11,13 Hereby, two sets of orthogonally line tagged images are acquired. Subsequent combination of these acquisitions results in grid-tagged images (Figs. 5-2 and 5-3). With this method, acquisition time is significantly reduced and image resolution perpendicular to the line tags is not affected. Although traditionally segmented k-space gradient echo techniques were used for signal readout, both T1 relaxation and serial RF excitations for imaging are responsible for the fading of the tags. T1 can only be increased by going to a higher magnetic field strength, and the number of RF excitations can be reduced by using alternative imaging sequences with fewer RF excitations. Hereby, echo planar imaging12 (see Fig. 5-2) and more recently also spiral imaging14 proved to be very valuable alternatives (see Fig. 5-3). Especially with spiral imaging, a very high

t=72 msec

t=109 msec

t=146 msec

t=183 msec

t=220 msec

t=257 msec

t=294 msec

t=331 msec

t=368 msec

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t=479 msec

t=516 msec

t=553 msec

t=590 msec

t=627 msec

t=664 msec

t=701 msec

t=738 msec

Figure 5-2 Twenty apical left ventricular short axis images in a healthy adult subject acquired with complementary spatial modulation of the magnetization cardiovascular magnetic resonance tagging. The systolic images include frames 1 to 9 (331 msec after the R-wave of the electrocardiogram), and the diastolic images are shown in frames 10 to 20 (368 to 738 msec). The temporal resolution in these images is 37 msec, and two line-tagged acquisitions were combined to create a grid-tagged image off-line. During systole, a counterclockwise apical rotation is seen, followed by a rapid clockwise untwisting in frames 10 to 15 (368 to 553 msec) during early diastole. This precedes the filling phase (frames 16 to 20; 590 to 738 msec) of the ventricles.

temporal resolution, with up to 70 frames/sec (14 msec temporal resolution), could be obtained, whereas a spatial resolution of 1.5 mm was easily possible14 (see Fig. 5-3). Because spiral imaging is susceptible to off-resonance blurring in the images, spectral spatial RF excitations15 are a stringent requirement.14 Together with echoplanar and spiral readouts, the use of SSFP was exploited by Herzka and colleagues,16 who found that SSFP leads to an improved tagging contrast compared with more conventional segmented k-space gradient echo imaging. By combining CSPAMM with SSFP imaging, Zwanenburg and associates17 showed that CSPAMM-tagged images can easily be acquired in a single breath hold. With the advent of parallel and 3D imaging, 3D assessment of myocardial motion based on 3D CSPAMM lattice tagging was reported by Ryf and colleagues.18 Simultaneously, the same authors proposed an extension to a previously reported analysis procedure19 that enables relatively time-efficient analysis and quantification of both systolic and diastolic motion of the heart.20 However, acquisition times were lengthy and only practical in coached breathing patterns in well-trained subjects.

Evaluation For the extraction of motion data from the tagged time series of images, sophisticated image analysis tools are needed. Although the identification of tags was very labor-intensive and took hours in the early years, sophisticated algorithms that reduce quantification of tagged time series of cardiac images to seconds are now readily available,19 and variants have been successfully implemented.20 Quantification of tagged images involves the identification of the tags in each heart phase image. With these automated or semiautomated algorithms, the tags may be identified with an accuracy that exceeds the image resolution.21 If the grid intersection points are identified for all heart phase images, local trajectories on the myocardium (Fig. 5-4) are defined and motion-specific parameters (e.g., strain, rotation, rotation velocity, radial or circumferential shortening or shear between epi- and endocardial muscle layers of the myocardium) can be derived. Using this approach, several studies of healthy subjects and patients with myocardial infarction,22 Cardiovascular Magnetic Resonance 71

5 CARDIOVASCULAR MAGNETIC RESONANCE TAGGING ASSESSMENT OF LEFT VENTRICULAR DIASTOLIC FUNCTION

t=35 msec

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17 msec

332 msec

542 msec

17 msec

332 msec

542 msec

24 msec

339 msec

549 msec

Figure 5-3 Spiral complementary spatial modulation of the magnetization myocardial tagging using a spectral-spatial excitation for fat suppression in each cine frame. Midventricular short and long axis views of a line-tagged and a grid-tagged myocardium in a healthy subject. The tagline distance is 4 mm and the temporal resolution is 35 msec. The time after the R-wave is indicated on the images. On the grid-tagged, short axis view, 50 tagline intersections can be observed.

hypertrophied cardiomyopathy,23 and aortic stenosis with pathologically hypertrophied hearts, as well as athletes with physiologic hypertrophy, were evaluated for apical untwisting during diastole.24 Untwisting velocity, time to peak untwisting velocity, and early diastolic strain can be determined as an index of diastolic function. Time to peak untwisting velocity (untwisting time; Table 5-1) is defined as the time delay between the point in time of minimum inner cavity area and the maximum untwisting velocity (Fig. 5-5). 72 Cardiovascular Magnetic Resonance

RESULTS Images Figure 5-2 shows 20 heart phase images with a temporal resolution of 37 msec acquired at the apex of a healthy subject.12 The grid structure remains visible, with a high contrast up to the last acquisition in late diastole (>700 msec). No fading of the tags is seen in the images. Therefore, the method is well suited for the quantification of diastolic heart wall motion.

At the apex of the healthy heart, a counterclockwise rotation during systole can be seen (see Fig. 5-2, phases 1 to 9; see Fig. 5-5, phase 1). This systolic rotation is followed by a rapid untwisting during isovolumic relaxation (see Fig. 5-2, phases 10–15 and early diastole; see Fig. 5-5, phase 2). This untwisting phase is typically followed by the filling phase of the ventricles, where little rotational component is seen (see Fig. 5-5B, phase 3).3,25 An identical separation of early diastolic apical untwisting and the filling phase of the ventricles is seen in highly competitive athletes with physiologically hypertrophied hearts. Neither the apical peak rotation angle nor diastolic untwisting time is changed in the hypertrophied hearts of athletes compared with healthy control subjects (see Table 5-1). However, among patients with LV hypertrophy as a result of pressure overload/aortic stenosis, a completely different apical rotation pattern is seen (see Fig. 5-5). The end-systolic peak rotation is significantly increased (P < 0.01) in comparison to the athletes or healthy subjects, and no separation of untwisting and filling can be seen during diastole. Untwisting and filling occur simultaneously, and the point in time of maximum untwisting velocity is significantly delayed in these patients (P < 0.01; see Fig. 5-5).

Figure 5-4 End-diastolic apical cardiovascular magnetic resonance image acquired in a healthy adult subject. The grid-tagged image is overlain with the corresponding local trajectories. The arrows start at the beginning of systole and at end diastole.

Table 5-1 Peak Rotation at the Apex, Maximum Rotation Velocity During Diastolic Untwisting and Untwisting Time for 12 Aortic Stenosis Patients, 12 Athletes and 11 Healthy Volunteers* Patients Aortic Stenosis Athletes 22 Volunteers 22

Peak Rotation (Deg)

Untwisting Velocity (Deg/s)

Untwisting Time (% ES)

125 62 72

8029 568 5517

326 178 168

22

14 12 100

10 8 6 4 2 0 –2

Systole

–4

0

A

20

40

60

Diastole

80

Rotation (x100% ES)

Apical rotation angle (degrees)

*The untwisting time is related to the duration of systole. Data are expressed as mean one standard deviation. Modified and reprinted with permission of the American Heart Association. Stuber et al. Alterations in the local myocardial motion pattern in patients suffering from pressure overload due to aortic stenosis. Circulation. 1999;100:363.

1 80 60

20

0

B Aortic stenosis

3

0

100 120 140 160 180 200

Time (% ES)

2

40

Healthy

10

20

30

40

50

60

70

80

90 100

Apical area reduction (x 100% ES) Rowers

Figure 5-5 A, Cross-sectional apical rotation velocity of athletes, healthy adults, and patients with aortic stenosis. The values are mean 1 standard deviation. The time axis is normalized to the end-systolic point of the cardiac cycle. The time point of maximum diastolic untwisting velocity (arrows) is delayed in the patients with aortic stenosis compared with the athletes or the healthy control subjects. Apical rotation velocity is identical in athletes and control subjects. B, Left ventricular rotation-area loop (apical plane) in healthy control subjects, rowers, and patients with aortic stenosis. The loop is separated in isovolumic contraction (1), ejection (2), isovolumic relaxation (3), and filling of the left ventricle. Both rotation and area change of the inner lumen at the apex are related to their maximum values (100%). ES, end systole. Cardiovascular Magnetic Resonance 73

5 CARDIOVASCULAR MAGNETIC RESONANCE TAGGING ASSESSMENT OF LEFT VENTRICULAR DIASTOLIC FUNCTION

Apical Rotation

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

Pressure overload LV hypertrophy results in the addition of new sarcomeres in parallel.26 Furthermore, an increase in the amount of collagen, with a consequently increased elastic stiffness, has been reported.27 This rearranged fiber architecture, together with the increased stiffness of the muscle tissue, may explain the alterations of the diastolic rotation pattern with a prolonged and delayed apical untwisting in these patients. The prolongation of untwisting results in an overlap with early diastolic filling and presumably impedes normal filling. Thus, a prolongation of early diastolic untwisting may be responsible for the occurrence of diastolic dysfunction in these patients. In patients with myocardial infarction, a prolonged untwisting phase with an overlap of diastolic untwisting and filling has also been observed28 and endsystolic apical peak rotation is usually severely reduced.

and susceptibility to diaphragmatic drift or misregistration in serial breath holds must be considered. However, with the availability of parallel imaging, echo-planar or spiral imaging, the number of RF excitations for imaging can be minimized. As an alternative to CSPAMM, more conventional tagging7,31 in combination with SSFP may support an improved tagging contrast-to-noise ratio, as does 3.0 Tesla (T) imaging because of an increase in T1 (from 850 msec at 1.5 T to 1200 msec at 3.0 T for myocardium and from 1200 msec at 1.5 T to 1650 msec for blood at 3.0 T). For these reasons, access to early diastolic myocardial motion may be feasible, even without the availability of CSPAMM. Nevertheless, the investigation of heart wall motion is still the subject of basic research, and appropriate parameters for the quantification of diastolic dysfunction and threshold values for normal subjects remain to be fully defined.

Strain Measurements In a study by Ennis and colleagues,23 systolic and diastolic strain was computed in both healthy adults and in patients with hypertrophic cardiomyopathy. In this study, CSPAMM was extended with a sophisticated phase reordering algorithm (CAPTOR) that enables more complete coverage of the entire cardiac cycle. In these patients, a significantly reduced strain was observed around the circumference of the heart in early diastole. In mid-diastole, no major difference in strain between the patients and the healthy adults was reported. Similarly, the strain rates were significantly reduced in these patients during early diastole. CMR assessment of diastolic function is beginning to be studied in broader populations. A population-based study of the multi-ethnic subclinical atherosclerosis (MESA) study showed evidence of diastolic dysfunction in asymptomatic individuals with LV hypertrophy.29 Despite similar regional systolic strain and strain rate among those with and without LV hypertrophy, the regional diastolic strain rate was significantly reduced among those with LV hypertrophy. A provocative study in a small group suggested that diastolic untwisting dysfunction could be identified by SPAMM after a sedentary rest in previously active adult subjects.30

Limitations The additional value of CMR tagging applied for the quantification of diastolic function remains to be more fully investigated compared with gold standard techniques. Currently, the CMR technique is not widely accessible to clinical cardiologists and the CSPAMM sequence is not currently available on all vendors’ systems. Furthermore, because CSPAMM is based on a subtraction technique, scanning time is doubled

CONCLUSION The CSPAMM myocardial tagging technique is a noninvasive method for the quantification of local heart wall motion. Because of the suppressed fading of the tags and the accessibility to the diastolic phase of the cardiac cycle, CSPAMM is well suited for the characterization of the diastolic portion of the cardiac cycle. Moreover, by the application of a slice-following procedure, the effects of through-plane motion can be suppressed and the same tissue elements can be tracked. Because of the relatively high temporal resolution of the data, rapid cardiac motion components, such as apical diastolic untwisting, can be recorded. Data suggest that pathologic LV hypertrophy in patients with aortic stenosis can be differentiated from hypertrophy in athletic hearts.24 Alterations in diastolic untwisting are seen in patients with pressure overload as a result of aortic stenosis, whereas none are observed in athletes’ hearts. Even though the athletes’ hearts were significantly larger than control subjects’ hearts, apical diastolic untwisting remains unchanged. Similarly, and using CSPAMM, significant alterations in early diastolic strain and strain rate have been reported in patients with hypertrophic cardiomyopathy.23 The current data derived from CSPAMM myocardial tagging clearly suggest that alterations in the diastolic phase of the cardiac cycle can be recorded readily. With the advent of parallel imaging, higher magnetic field strength, and more advanced imaging sequences, persistence of tags can be prolonged, even without CSPAMM, enabling access to early diastolic motion of the left ventricle. Although CMR myocardial tagging requires off-line computer analysis, quantification is now a matter of seconds and the technique may offer a new tool for the evaluation of diastolic wall relaxation in healthy and diseased states.

References 1. Maier SE, Fischer SE, McKinnon GC, Hess OM, Krayenbuehl HP, Boesiger P. Evaluation of left ventricular segmental wall motion in hypertrophic cardiomyopathy with myocardial tagging. Circulation. 1992;86:1919–1928. 2. Rogers Jr WJ, Shapiro EP, Weiss JL, et al. Quantification of and correction for left ventricular systolic long-axis shortening by magnetic resonance tissue tagging and slice isolation. Circulation. 1991;84:721–731. 74 Cardiovascular Magnetic Resonance

3. Fischer SE, McKinnon GC, Scheidegger MB, Prins W, Meier D, Boesiger P. True myocardial motion tracking. Magn Reson Med. 1994;31:401–413. 4. D’hooge J, Heimdal A, Jamal F, et al. Regional strain and strain rate measurements by cardiac ultrasound: principles, implementation and limitations. Eur J Echocardiogr. 2000;1:154–170. 5. Gilman G, Khandheria BK, Hagen ME, et al. Strain rate and strain: a step-by-step approach to image and data acquisition. J Am Soc Echocardiogr. 2004;17:1011–1020.

20. Ryf S, Tsao J, Schwitter J, Stuessi A, Boesiger P. Peak-combination HARP: a method to correct for phase errors in HARP. J Magn Reson Imaging. 2004;20:874–880. 21. Atalar E, McVeigh ER. Optimization of tag thickness for measuring position with magnetic resonance imaging. IEEE. 1994;13:152–160. 22. Nagel E, Stuber M, Lakatos M, Scheidegger MB, Boesiger P, Hess OM. Cardiac rotation and relaxation after anterolateral myocardial infarction. Coron Artery Dis. 2000;11:261–267. 23. Ennis DB, Epstein FH, Kellman P, Fananapazir L, McVeigh ER, Arai AE. Assessment of regional systolic and diastolic dysfunction in familial hypertrophic cardiomyopathy using MR tagging. Magn Reson Med. 2003;50:638–642. 24. Stuber M, Scheidegger MB, Fischer SE, et al. Alterations in the local myocardial motion pattern in patients suffering from pressure overload due to aortic stenosis. Circulation. 1999;100:361–368. 25. Rademakers FE, Buchalter MB, Rogers WJ, et al. Dissociation between left ventricular untwisting and filling: accentuation by catecholamines. Circulation. 1992;85:1572–1581. 26. Grossman W, Jones D, McLaurin LP. Wall stress and patterns of hypertrophy in the human left ventricle. J Clin Invest. 1975;56:56–64. 27. Hess OM, Lavelle JF, Sasayama S, Kemper WS, Ross J. Diastolic myocardial wall stiffness of the left ventricle in chronic pressure overload. Eur Heart J. 1982;3:315–324. 28. Matter C, Mandinov L, Kaufmann P, Nagel E, Boesiger P, Hess OM. [Function of the residual myocardium after infarct and prognostic significance]. Z Kardiol. 1997;86:684–690. 29. Edvardsen T, Rosen BD, Pan L, et al. Regional diastolic dysfunction in individuals with left ventricular hypertrophy by tagged magnetic resonance imaging: the Multi-Ethnic Study of Atherosclerosis (MESA). Am Heart J. 2006;151:109–114. 30. Dorfman TA, Rosen BD, Perhonen MA, et al. Diastolic suction is impaired by bed rest: MRI tagging studies of diastolic untwisting. J Appl Physiol. 2008;104:1037–1044. 31. Axel L, Dougherty L. Heart wall motion: improved method of spatial modulation of magnetization for MR imaging. Radiology. 1989;172:349–350.

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6. Ingels Jr NB, Daughters GT, Stinson EB, Alderman EL. Measurement of midwall myocardial dynamics in intact man by radiography of surgically implanted markers. Circulation. 1975;52:859–867. 7. Yun KL, Niczyporuk MA, Daughters GT, et al. Alterations in left ventricular diastolic twist mechanics during acute human cardiac allograft rejection. Circulation. 1991;83:962–973. 8. Potel MJ, Rubin JM, MacKay SA, Aisen AM, Al-Sadir J, Sayre RE. Methods for evaluating cardiac wall motion in three dimensions using bifurcation points of the coronary arterial tree. Invest Radiol. 1983;18:47–57. 9. Zerhouni EA, Parish DM, Rogers WJ, Yang A, Shapiro EP. Human heart: tagging with MR imaging: a method for noninvasive assessment of myocardial motion. Radiology. 1988;169:59–63. 10. Axel L, Dougherty L. MR imaging of motion with spatial modulation of magnetization. Radiology. 1989;171:841–845. 11. Fischer SE, McKinnon GC, Maier SE, Boesiger P. Improved myocardial tagging contrast. Magn Reson Med. 1993;30:191–200. 12. Stuber M, Spiegel MA, Fischer SE, et al. Single breath-hold slicefollowing CSPAMM myocardial tagging. MAGMA. 1999;9:85–91. 13. McVeigh ER, Atalar E. Cardiac tagging with breath-hold cine MRI. Magn Reson Med. 1992;28:318–327. 14. Ryf S, Kissinger KV, Spiegel MA, et al. Spiral MR myocardial tagging. Magn Reson Med. 2004;51:237–242. 15. Meyer CH, Pauly JM, Macovski A, Nishimura DG. Simultaneous spatial and spectral selective excitation. Magn Reson Med. 1990;15:287–304. 16. Herzka DA, Guttman MA, McVeigh ER. Myocardial tagging with SSFP. Magn Reson Med. 2003;49:329–340. 17. Zwanenburg JJ, Kuijer JP, Marcus JT, Heethaar RM. Steady-state free precession with myocardial tagging: CSPAMM in a single breathhold. Magn Reson Med. 2003;49:722–730. 18. Ryf S, Spiegel MA, Gerber M, Boesiger P. Myocardial tagging with 3D-CSPAMM. J Magn Reson Imaging. 2002;16:320–325. 19. Osman NF, Kerwin WS, McVeigh ER, Prince JL. Cardiac motion tracking using CINE harmonic phase (HARP) magnetic resonance imaging. Magn Reson Med. 1999;42:1048–1060.

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CHAPTER 6

Cardiovascular Magnetic Resonance Contrast Agents Peter Caravan

Nearly one third of all magnetic resonance imaging (MRI) scans today (2009) use a gadolinium-based contrast agent. The contrast agent typically makes diseased tissue appear brighter (or in some cases darker) than the surrounding tissue. Cardiovascular applications, such as magnetic resonance angiography (MRA) and functional imaging of myocardial perfusion and viability, are growing rapidly, and represent the bulk of cardiovascular magnetic resonance (CMR) scans that use a contrast agent. The first approved contrast agent, gadopentetate (Gd-DTPA), appeared in 1988, and several other compounds followed. These first contrast agents were extracellular fluid (ECF) agents. Although frequently used in CMR and an essential component of CMR perfusion and late gadolinium enhancement (LGE) assessment of fibrosis, none of these agents has been approved for cardiac applications. Thus, the use of these agents in CMR is considered “off-label.” There are currently several compounds in clinical trials and one recently approved that are intravascular agents designed specifically to enhance contrast-enhanced (CE) MRA. At the preclinical stage, there are exciting advancements in molecular imaging agents, including compounds that detect pH changes, enzymatic activity, specific biomolecules such as fibrin or integrins, and magnetically labeled cells. This chapter focuses first on the general underlying chemistry and biophysics of contrast agents in clinical CMR. The mechanism of action of different classes of contrast agents is described, with examples drawn from CMR applications. Finally, there is a brief survey of novel contrast agents potentially useful for cardiovascular indications that are currently in clinical or preclinical development. The vast majority of MRI agents in clinical use are small molecules based on chelated gadolinium. The bulk of this chapter focuses on gadolinium complexes, including their chemistry, biophysics, and applications. Other exogenous compounds have been used to change signal properties in MRI (e.g., iron particles, hyperpolarized nuclei), and these will be discussed as appropriate to CMR. This chapter assumes that the reader has a basic understanding of CMR vocabulary, and the reader is referred to Chapter 1 for further clarification.

76 Cardiovascular Magnetic Resonance

INTRODUCTION TO THE BIOPHYSICS OF MAGNETIC RESONANCE IMAGING All contrast agents shorten both T1 and T2. However, it is useful to classify MRI contrast agents into two broad groups based on whether the substance increases the transverse relaxation rate (1/T2) by roughly the same amount that it increases the longitudinal relaxation rate (1/T1) or whether 1/T2 is altered to a much greater extent. The first category is referred to as “T1 agents” because, on a percentage basis, these agents alter 1/T1 of tissue more than 1/T2 because of endogenous transverse relaxation in tissue. With most pulse sequences, this dominant T1-lowering effect gives rise to increases in signal intensity, and thus these agents are referred to as “positive” contrast agents. The T2 agents largely increase the 1/T2 of tissue selectively and cause a reduction in signal intensity, and thus they are known as “negative” contrast agents. Paramagnetic gadolinium-based contrast agents are examples of T1 agents, whereas ferromagnetic large iron oxide particles are examples of T2 agents. There are many mechanisms by which contrast agents shorten T1 and T2. Considerable chemistry and biophysics can be applied to understand or predict these mechanisms. However, in many cases, the effect of these mechanisms can be reduced to a single constant, called “relaxivity.” Figure 6-1 shows the effect and definition of relaxivity. Figure 6-1A shows the effect of a typical contrast agent on the relaxation time of two hypothetical tissues, one with T1 ¼ 1200 msec (similar to heart muscle at 1.5 T) and one with T1 ¼ 400 msec. At low concentration (left side of the graph), it appears that the contrast agent has a larger effect (change in T1) on the tissue with the longer T1. At higher concentrations of contrast agent (right side of Fig. 6-1A), both tissues approach approximately the same T1. A simple way to quantify this effect is to consider the rate of relaxation, 1/T1 (sometimes denoted “R1”). In most cases in medical imaging, the contrast agent increases the relaxation rate proportional to the amount of contrast agent: 1 1 ¼ þ r1 ½CA T1 T1o

(1)

20 1200 15 1/T1 (s−1)

T1 (msec)

1000 800 600 400

10

5

200 0

0 0

A

2

3

Contrast agent (mM)

where T1 is the observed T1 with contrast agent in the tissue, T1o is the T1 before addition of the contrast agent, [CA] is the concentration of contrast agent, and r1 is the longitudinal relaxivity, often just “relaxivity.” The conventional units for r1 are mM-1s-1 (per millimolar per second, sometimes Lmol-1s-1). Thus, the slope of 1/T1 as a function of contrast agent concentration (Fig. 6-1B) shows the relaxivity, in this case, 4 mM-1s-1. Figure 6-1B shows that the effect of the contrast agent on the relaxation rate is independent of the initial T1 of the tissue. That is, in terms of relaxation rate, the contrast agent has the same effect, regardless of initial T1. Transverse, or T2, relaxivity is defined in an analogous way: 1 1 ¼ þ r2 ½CA T2 T2o

1

(2)

For all medically used contrast agents, r2 is larger than r1. Relaxivity is dependent on magnetic field strength, is dependent on temperature, and in some instances can depend on protein binding, pH, or even the presence of enzymes. Contrast agent behavior in vivo is seldom as simple as the pure linear effect relaxation rate shown in Figure 6-1. Even in the simple case of pure linear relaxation, the effect of the contrast agent on the CMR image is generally nonlinear. In traditional spin echo sequences, nonlinearity can be a result of T1 saturation or T2 signal loss. Once the contrast agent reduces T1 < repetition time (TR)/2, increasing contrast agent concentration will have little effect on increasing the available longitudinal magnetization because the tissue will have nearly fully recovered the magnetization before the next radiofrequency (RF) pulse. Furthermore, because contrast agents affect both T1 and T2 relaxation, at high enough concentration, the contrast agent will reduce T2 to the order of the echo time (TE), and will then decrease magnetic resonance (MR) image intensity. These effects are seen in Figure 6-2, where signal intensity is plotted versus contrast agent concentration for T1- and T2-weighted spin echo sequences. Figure 6-2 was generated assuming a contrast agent relaxivity of 4 mM-1s-1, typical of most commercial ECF gadolinium agents, and tissue relaxation times typical of myocardium at 1.5 T (T1 ¼ 1200 msec, T2 ¼ 50 msec). For the T1-weighted sequence (TE/TR ¼ 15/600), Figure 6-2A, signal intensity begins to level out at a contrast agent concentration between 0.5 and 1.0

0

4

B

1

2

3

4

Contrast agent (mM)

mM. From Figure 6-1A, this is the range as which the T1 of the myocardium decreased to approximately 300 msec, or TR/2. At concentrations greater than 1 mM, the T1 effect is saturated, and the only imaging effect of the contrast agent is to make T2 shorter and cause signal loss, even on this T1-weighted sequence. Signal is lost because even a T1-weighted sequence has a finite TE, and T2 effects can enter when T2 is short enough. The signal intensity plateau on the T2-weighted scan (TE/ TR ¼ 80/3000), Figure 6-2B, occurs at much lower contrast agent concentration. Because TR is so long, the only real effect of the contrast agent is to reduce (rather than increase) signal intensity on this T2-weighted scan. However, this negative contrast can also be medically useful, and certain contrast agents (notably, the iron-oxide particles) create negative contrast exactly by providing enhanced T2 relaxation, and thus darker images on T2-weighted scans. Increasing the relaxivity (r1 or r2) will have the effect of pushing the simulated curves in Figure 6-2A to the left, which means that peak signal and subsequent signal loss will occur at lower contrast agent concentrations. A more linear response of signal to contrast agent can be achieved with a fast three-dimensional spoiled gradient recalled echo (GRE) sequence. This is seen in Figure 6-2C, where signal intensity is plotted versus contrast agent concentration using the same tissue relaxation times and relaxivities as in Figure 6-2A and 6-2B for a typical fast three-dimensional spoiled GRE sequence, TE/TR/flip ¼ 2.2/9.0/40 . The short TR and very short TE ensure that signal intensity increases across the entire concentration range. At high concentration, the effect of the contrast agent is becoming nonlinear, but signal still is increasing with increasing contrast agent concentration.

Commercial Contrast Agents and Those in Clinical Development The addition of paramagnetic materials to reduce relaxation times goes back to the earliest days of MR. In the 1940s, Bloch and colleagues used ferric nitrate to enhance the relaxation rate of water.1 Exogenous contrast was applied to MRI in 1977 when Lauterbur and associates reported Cardiovascular Magnetic Resonance 77

6 CARDIOVASCULAR MAGNETIC RESONANCE CONTRAST AGENTS

Figure 6-1 Change in (A) longitudinal relaxation time (T1) and (B) longitudinal relaxation rate (1/T1) for typical myocardial tissue (solid line, T1 ¼ 1200 msec at 1.5 T) and short T1 tissue (dashed line, T1 ¼ 400 msec).

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

using manganese dichloride to differentiate normal from infarcted myocardium in dogs.2 Carr and colleagues reported the first use of a gadolinium complex, gadopentetate dimeglumine (Gd-DTPA; Magnevist, Bayer Schering, Berlin, Germany), in patients with brain tumors in 1984.3 By 1988, Gd-DTPA was approved for clinical use.

0.7 0.6 0.5 0.4

Extracellular Agents

0.3 0.2 Spin echo TR/TE/flip = 600/15/90o

0.1 0 0

1

A

2

3

4

Contrast agent (mM)

0.7 Spin echo TR/TE/flip = 3000/80/90o

0.6 0.5 0.4 0.3 0.2 0.1 0 0

1

B

2

3

4

Contrast agent (mM)

0.7 3D GRE TR/TE/flip = 9.0/2.2/40o

0.6 0.5 0.4 0.3 0.2 0.1 0 0

C

1

2

3

4

Contrast agent (mM)

Figure 6-2 Effect of contrast agent on myocardial image intensity on T1-weighted and T2-weighted spin echo scans. (A) T1-weighted spin echo (repetition time [TR] ¼ 600 msec) assumes a patient with 100 bpm heart rate and shows a linear increase of signal only for contrast agent concentration < 0.5 mM. (B) T2-weighted spin echo images (TR ¼ 3000 msec) show only T2 signal loss effects as a result of contrast agent with no T1 enhancement because of the long TR. (C) Typical short-TR fast spoiled gradient recalled echo (GRE) sequence. The very short echo time (TE) and short TR give monotonically increased image intensity across the entire range of contrast agent concentrations typically found in clinical scans. 78 Cardiovascular Magnetic Resonance

The most common contrast agents used clinically are ECF agents. Although several are approved for clinical use, none is specifically approved for cardiac applications. These all behave in a very similar manner, and are typically referred to as “gadolinium” or “gado” agents. Figure 6-3 shows the chemical structures of several approved ECF agents. Chemically, these compounds exhibit three similar features: they all contain Gd, they all contain an 8-coordinate ligand binding to Gd, and they all contain a single water molecule coordination site to Gd. Nomenclature for contrast agents can be confusing: there is a generic name (e.g., gadopentetate dimeglumine), a trade name (e.g., Magnevist, Bayer HealthCare Pharmaceuticals), and usually a chemical code name or abbreviation (e.g., Gd-DTPA). Any of these three names may be used in the scientific literature. The multidentate ligand is required for safety.4 The ligand encapsulates the gadolinium, resulting in a high thermodynamic stability and kinetic inertness with respect to metal loss. This enables the contrast agent to be excreted intact, an important property because these contrast agents tend to be much less toxic than their substituents. For example, the DTPA ligand and gadolinium chloride both have an LD50 of 0.5 mmol/kg in rats (LD50 ¼ dose that causes death in 50% of the animals), whereas the Gd-DTPA complex has a safety margin that is higher by nearly a factor of 20, with an LD50 of 8 mmol/kg for the Gd-DTPA complex.5 The gadolinium ion and coordinated water molecule are essential to providing contrast. The gadolinium(III) ion has a high magnetic moment and a relatively slow electronic relaxation rate, properties that make it an excellent relaxer of water protons. The proximity of the coordinated water molecule leads to efficient relaxation. The coordinated water molecule is in rapid chemical exchange (106 exchanges/sec) with solvating water molecules.6 This rapid exchange leads to a catalytic effect whereby the Gd complex effectively shortens the relaxation times of the bulk solution. The ECF agents have very similar properties, and these are summarized in Table 6-1. They are all very hydrophilic complexes with similar relaxivities and excellent safety profiles, and they can be formulated at high concentrations. On injection, ECF agents quickly and freely distribute to the extracellular space. Administration of any of these agents (with rare exceptions)7 yields similar diagnostic information. There are some differences among the physical properties. The diamide complexes gadodiamide and gadoversetamide have considerably lower thermodynamic stability (log K17 vs. log K > 21 for other Gd complexes).8,9 The nonionic (neutral) compounds (gadodiamide, gadoteridol, gadoversetamide, and gadobutrol) were designed to minimize the osmolality of the formulation. This was prompted by the distinct reduction in toxicity and side effects brought on by the development of

N

N

N O O Gd O O O O O O H H

O O O

N

N

NH O O

Gd O O

H

N NH O O O O O H

Gd-DTPA-BMA (Omniscan) Gadodiamide

Gd-DTPA (Magnevist) Gadopentetate O OH

O

OH N OH

N Gd O

OH2

N

N

O

O

O O

O

Gd O

O

H

Gd-DO3A-butrol (Gadovist) Gadobutrol

O

N NH O O O O O

O

H

Gd-DTPA-BMEA (OptiMARK) Gadoversetamide

O

O O

O N

N

O

N OH

N Gd

OH2

N

N

O

O

Gd O

N

N

NH

O O

O

N

OH2

N O

O

O

Gd-DOTA (Dotarem) Gadoterate

Gd-HPDO3A (ProHance) Gadoteridol

Table 6-1 Approved (January 2009) Magnetic Resonance Imaging Contrast Agents—Relaxivity,120 Osmolality, and Viscosity Generic Name Gadopentetate Gadoterate Gadodiamide Gadoteridol Gadobutrol Gadoversetamide

Product Name

Chemical Abbreviation

r1, 0.47 T 40 C

r2, 0.47 T 40 C

Osmolality* (osmol/kg)

Viscosity* (cP)

Magnevist (Bayer HealthCare Pharmaceuticals) Dotarem (Guerbet Group) Omniscan (GE Healthcare) ProHance (Bracco Diagnostics, Inc.) Gadovist (Bayer HealthCare, Inc.) OptiMARK (Mallinckrodt, Inc.)

Gd-DTPA

3.4

4.0

1.96121

2.9121

Gd-DOTA

3.4

4.1

Gd-DTPA-BMA

3.5

3.8

Gd-HPDO3A

3.1

3.7

Gd-DO3A-butrol

3.7

5.1

Gd-DTPA-BMEA

4.2

5.2

1.35122 (4.02)122 0.79121 (1.90)122 0.63122 (1.91)122 0.57123 (1.39)121 1.11124

2.0122 (11.3)122 1.4121 (3.9)122 1.3122 (3.9)122 1.4123 (3.7)121 2.0124

*All concentrations 0.5 M except those in parentheses, which are 1 M.

nonionic X-ray contrast media. However, for CMR, the injection volumes are much smaller than for X-ray, and thus the overall increase in osmolality after injection of a CMR contrast agent is minimal. Unlike with X-ray contrast, there is no documented safety benefit in using nonionic CMR contrast agents. One benefit of the nonionic compounds is the ability to formulate them at high concentration (1 M)10 without drastically

increasing osmolality or viscosity (see Table 6-1). These highconcentration formulations may be useful in fast dynamic studies, such as dynamic MRA or myocardial perfusion. These ECF agents, as with iodinated preparations used for X-ray and computed tomography, are excreted by the kidneys. As a result, clearance is impaired in patients with abnormal or depressed renal function (discussed later). Cardiovascular Magnetic Resonance 79

6 CARDIOVASCULAR MAGNETIC RESONANCE CONTRAST AGENTS

Figure 6-3 Chemical structures of commercial (United States or Europe) extracellular fluid contrast agents.

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

Blood Pool Agents Because CMR contrast agents are administered intravenously, they are all potentially capable of imaging the blood vessels. The ECF agents described earlier are used routinely, if off-label, for CE-MRA (see Chapters 23, 24, and 34 to 36). One drawback of using ECF agents for MRA is their pharmacokinetics. ECF agents rapidly leak out of the vascular space into all the interstitial spaces of the body. Angiography with ECF agents is thus typically limited to dynamic arterial studies. There has been considerable effort toward designing specific blood pool agents that would be tailored for vascular imaging. The ideal blood pool agent should remain in the vascular compartment and not leak out into the extracellular space. It should be capable of being given as a bolus such that a dynamic arterial image can be obtained. At the same time, it should have sufficient relaxivity and blood half-life to allow high-resolution steadystate images to be obtained. Recently, MS-325 (gadofosveset, marketed as VasovistW and more recently as AblavarW in the United States by Lantheus Medical Imaging, Billerica, MA; Fig. 6-4) was approved, and there are several other blood pool agents that have reached various stages of clinical development. Three approaches have been taken to design blood pool agents: protein binding, increased size, and ultrasmall iron oxide particles. These are discussed later. Both MS-325 (gadofosveset)11,12 and B22956 (gadocoletic acid; Fig. 6-5)13,14 are gadolinium-based compounds that bind reversibly to serum albumin. Albumin is the most abundant protein in plasma, and its concentration is high enough (600 to 700 mM) to bind reversibly most of the contrast agent after injection. Reversible albumin binding serves four purposes: (1) the albumin slows the leakage of the contrast agent out of the intravascular space; (2) the reversible binding still allows a path for excretion–the unbound fraction can be filtered through the kidneys or taken up by hepatocytes; (3) the bound fraction is “hidden” from the liver and kidneys, leading to an extended plasma half-life; and (4) the relaxivity of the contrast

O N

N O Gd O O O O O O O H H

O O O

Gd-BOPTA (MultiHance) Gadobenate

O O O P O N

N

O O O

N O O Gd O O O O O O H H

MS-325 (Vasovist; Ablavar) Gadofosveset Figure 6-4 Chemical structures of other commercial contrast agents with weak (Gd-BOPTA) or strong (MS-325) serum albumin binding.

agent is increased 4- to 10-fold on binding to albumin (discussed later). Gd-BOPTA (gadobenate; see Fig. 6-4) has weak affinity15,16 for albumin (10% bound), which leads to a modest increase in relaxivity relative to the ECF agents.

GdL

LGd O

HO2C

LGd N

N

GdL =

HO

O

N

O

LGd

O

Lys

LGd Lys

Lys Lys

O O

B22956 (Gadocoletic acid)

Lys

O

HN

LGd

Lys = N H

Lys

N

Lys

LGd

Lys

GdL

Lys

Lys

GdL

Lys Lys GdL

Lys

O Lys Lys

Lys Lys

GdL

GdL

N

O

Lys

Lys

GdL

GdL

Lys

GdL Lys Lys

Lys GdL

Figure 6-5 Chemical structures of some investigational gadolinium-based blood pool agents.

Lys

Lys

GdL

Gadomer (aka Gadomer-17)

80 Cardiovascular Magnetic Resonance

Lys Lys

Lys

N

N O O Gd O O O O O O H H

GdL Lys Lys

Lys

LGd N

O Lys

N

Lys

LGd O

O

Lys

Lys

Lys OH2

N

GdL

Lys

Lys

LGd

O

Gd O

NH

Lys Lys

Lys

O H

N

Lys GdL

GdL

Agent type r1 buffer (mM-1 s-1) r1 plasma (mM-1 s-1) % bound plasma

MS-325* Gadofosveset

B22956† Gadocoletic Acid

Gd-BOPTA Gadobenate

Gadomer-17‡

Strong protein binding 6.612 5012 9112

Strong protein binding 6.5125 27125 95125

Weak protein binding 4.4126 9.7127

Increased size 16.5120 19.0120

*Data at 0.1 mM. {Data at 0.5 mM. {Relaxivity per gadolinium.

The binding and relaxivity features of gadolinium-based blood pool agents are listed in Table 6-2. Because binding affinity is moderate to weak for these compounds, the fraction bound to albumin will depend on the concentrations of albumin and the contrast agent.12,17 Immediately after injection, when the concentration of the contrast agent is high relative to albumin, there will be a greater free fraction. As the concentration of the contrast agent begins to stabilize (at 0.5 mM) the fraction bound will become constant. The observed relaxivity will depend on the fraction bound; unlike ECF agents, T1 change in plasma is not linearly related to contrast agent concentration.17,18 Among the albumin-binding agents, MS-325 (gadofosveset) has a somewhat lower albumin affinity than B22956 (gadocoletic acid), although the majority of both is bound under steadystate conditions. The relaxivity of albumin-bound MS-325 (gadofosveset) is higher than that of B22956 (gadocoletic acid). MS-325 (gadofosveset) is mainly renally excreted, whereas B22956 (gadocoletic acid) has significant biliary clearance as well as renal excretion. Early work on blood pool compounds involved gadolinium covalently linked to macromolecules, such as polylysine, dextran, or modified bovine serum albumin.4 The large size of these compounds restricted diffusion out of the vascular space and led to very good vascular imaging properties. However, these compounds cleared very slowly in preclinical studies and there were concerns about a potential immunologic response. This approach was modified by the synthesis of compounds that were large enough to be kept in the vascular compartment, but small enough to be eliminated by glomerular filtration in the kidneys. Gadomer (gadomer-17) is an example of this type of blood pool agent (see Fig. 6-5).19,20 Gadomer is a large, 17 kDa molecule that contains 24 covalently linked gadolinium complexes. The dendritic (branching) approach to synthesis results in a compound that is globular. The per-gadolinium relaxivity of Gadomer is much higher than that of its monomeric units because of the slow tumbling of the molecule (discussed later). The use of multiple gadolinium chelates to increase the size also increases the molecular relaxivity (24 Gd 18.7 ¼ molecular relaxivity of 450 mM-1s-1), which in turn means that lower doses can be given. Gadomer is a neutral hydrophilic compound that has little affinity for plasma proteins and is excreted renally. All of the gadolinium-based vascular agents described earlier are not “true” blood pool compounds. Although far superior to the ECF agents in terms of extravasation and relaxivity, some fraction of the compound leaks out

into the extracellular space. Iron oxide particles, on the other hand, are true blood pool agents. The small particle iron oxide particles (SPIOs) used for liver imaging are large enough to be recognized by the reticuloendothelial system and rapidly removed from the bloodstream. The smallest size fraction of these particles, ultrasmall particle iron oxide (USPIOs), evaded the reticuloendothelial system and could be used to image the blood pool.21,22 Although smaller, these particles are still too large to passively leak out of the vascular space and make very good blood pool agents. Making ultrasmall particles not only changes the biodistribution of the compound, but also changes the relaxation phenomena. SPIOs have a much greater effect on T2 than on T1 (large r2/r1) and are used as T2 or T2* agents. USPIOs have very good T1 relaxation properties (smaller r2/r1) and can be used for brightblood imaging (T1-weighted). The iron oxide particles are not excreted; the iron is eventually resorbed into the body.

RELAXIVITY Because their effect is indirect, CMR contrast agents differ from other diagnostic imaging agents. Water and fat are imaged, and it is the action of the contrast agent on the relaxation properties of the water hydrogen nuclei that generates contrast; in X-ray contrast media and nuclear imaging agents, the effect observed is more direct. Because water is present at a very high concentration (55,000 mM) and the contrast agent is typically at a much lower concentration (0.1 to 1 mM), the contrast agent must act catalytically to relax the water protons for there to be a measurable effect. Relaxivity, r1 and r2, thus describes this catalytic efficiency. Some compounds are better magnetic catalysts than others (have higher relaxivity). Moreover, relaxivity is dependent on the external magnetic field, B0. This section explains these differences and the molecular basis for them. For discrete ions, such as Gd(III) and manganese (Mn[II]), the factors influencing relaxivity are the same; for iron oxide particles, the relaxation mechanism is different and will be treated separately. Relaxivity can be factored into inner- and outer-sphere terms. “Inner sphere” r1 ¼

Dð1=T1 Þ OS ¼ rIS 1 þ r1 ½Gd

(3)

refers to the relaxation enhancement due to the exchangeable waters in the inner sphere. “Outer sphere” refers to relaxivity resulting from water in the second and outer-coordination spheres. This separation is often used because the inner-sphere component is easier to understand from a theoretical framework and the effect of changing specific molecular parameters on relaxivity can be tested. For ECF agents, the inner-sphere and Cardiovascular Magnetic Resonance 81

6 CARDIOVASCULAR MAGNETIC RESONANCE CONTRAST AGENTS

Table 6-2 Albumin Binding and Observed Relaxivities (20 MHz) of MS-325, B22956, Gd-BOPTA, Gadomer-17, at 37 C

q=½H2 O (4) rIS 1 ¼ T1m þ tm Inner-sphere relaxivity is given by Equation 4, where q represents the number of water molecules bound to the metal ion (typically q ¼ 1), [H2O] is the water concentration in millimolar, T1m is the relaxation time of the water that is bound to the metal ion, and tm is the lifetime of the water molecule in the inner sphere. Equation 4 shows that the more sites available for water to bind (q), the more efficient the process. Likewise, the T1 of the bound water (T1m) should be very short. This is indicative of how effectively the metal ion can relax the coordinated water. The lifetime tm of the bound water is the inverse of the water exchange rate (kex ¼ 1/tm). If turnover is slow (tm is long), then it will not matter how efficiently the bound water is relaxed if this relaxation cannot be transmitted to the bulk water. Paramagnetic relaxation (1/T1m) occurs via a dipolar mechanism. At a magnetic field encountered in medical imaging, 1/T1m is shown in Equation 5, where C is a 1 SðS þ 1Þ 3tc (5) ¼C 1 þ o2H t2c T1m r6MH constant,4 S is the spin quantum number, rMH is the metal ion-toproton distance, tc is the correlation time, and oH is the hydrogen Larmor frequency. The product S(Sþ1) is proportional to the magnetic moment. All other factors being equal, the higher the magnetic moment, the more efficient the relaxation. This is why Gd3þ (S ¼ 7/2) is preferred to copper (Cu2þ, S ¼ 1/2). The dipolar effect depends on the distance between the ion and the hydrogen nucleus, rMH, to the inverse sixth power. The inner-sphere water is critical; it has the shortest metal-to-hydrogen distance of water hydrating the metal complex. The correlation time, tc, is the time constant that governs the interaction between the electron spin of the ion and the nuclear spin of the hydrogen. Depending on the ion and the field strength, the correlation time is either the time constant for rotational diffusion of the molecule, tR, or the electronic T1 of the metal ion, T1e. In the MR experiment, nuclei are excited by applying RF energy at the Larmor frequency. To relax these nuclei, there needs to be a resonant source for energy transfer. A paramagnetic molecule tumbling in solution creates a fluctuating magnetic field. Molecules tumble (undergo rotational diffusion) because of thermal energy, but because they also collide with each other, there is a distribution of rotational diffusion rates, with a mean rate characterized by 1/tR. The closer this tumbling rate is to the Larmor frequency, the more efficient the relaxation by the contrast agent. Small molecules, such as Gd-DTPA, tumble very fast, in the gigahertz range (1 GHz ¼ 1000 MHz), but the Larmor frequency for protons at imaging fields is much slower. For example, at 1.5 T, the Larmor frequency is approximately 65 MHz, so relaxation is not as efficient as it could be. Larger molecules, such as proteins, tumble much more slowly. When contrast agents are made to tumble more slowly, relaxivity is increased. Newer contrast agents take advantage of this phenomenon. Lauffer23 pointed out that if small contrast agents could be made to bind noncovalently to protein targets, then their relaxivity would be increased on binding because the contrast agent would take on the rotational characteristics of the protein (receptor-induced magnetization enhancement). MS-325 is an example of a contrast agent designed to exploit the receptor-induced magnetization enhancement effect.12 MS-325 targets the blood protein serum albumin. In the absence of albumin, the relaxivity of MS-325 is approximately 50% greater than that of Gd-DTPA, but in the presence of albumin, the relaxivity is approximately 600% greater than that of Gd-DTPA at 1.5 82 Cardiovascular Magnetic Resonance

40

r1 (mM−1 s−1)

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

outer-sphere contribution to relaxivity is about the same, whereas for newer contrast agents, the inner-sphere component accounts for the majority of overall relaxivity.

30

20

10

0 0.5

1.0 1.5 2.0 B0 (Tesla)

2.5

3.0

Figure 6-6 Magnetic field dependence on relaxivity for MS-325 (gadofosveset; circles) and gadopentetate (Gd-DTPA; squares) in either serum albumin solution (filled symbols) or buffered saline (open symbols) at 37 C.

T. This is seen in Figure 6-6, where the magnetic field dependence on relaxivity is plotted for MS-325 and Gd-DTPA in either serum albumin solution (filled symbols) or in buffered saline. Increasing the molecular size is another way to increase relaxivity. Gadomer is a dendrimer that contains 24 Gd and has a molecular weight of 17,500. The increased size results in slower motion and higher relaxivity. However, the rotational effect is defined by more than just molecular weight. Linear polymers containing Gd have lower relaxivities24,25 than may first be expected because of fast rotation around one axis of the molecule. Electrons undergoing relaxation in a paramagnetic ion also generate a fluctuating field that can cause nuclear relaxation. It is the faster of the two processes (rotation or electronic relaxation) that defines how efficiently the hydrogen is relaxed. For Gd and Mn complexes, electronic relaxation depends on the applied field and it gets slower as the magnetic field is increased. At low fields (< 0.5 T), electronic relaxation governs hydrogen relaxation, whereas at higher fields, rotational diffusion is the dominant mechanism. Gd3þ and Mn2þ have symmetrical electronic configurations (half-filled f and d shells, respectively), and as a result, electronic relaxation is relatively slow. Other ions, such as dysprosium (Dy3þ), have a higher magnetic moment, but are rather poor relaxors because electronic relaxation is so fast. Equation 5 also indicates that at higher fields the T1 relaxation mechanism becomes less efficient as oH2tc2 > 1. The field at which this condition is met is of course dependent on the magnitude of the correlation time. This is seen in Figure 6-6, where the relaxivity of MS-325 in serum albumin solution begins to decrease after a peak at approximately 0.7 T. Figure 6-7 further illustrates the effect of correlation time and field strength on relaxivity. In Figure 6-7, r1 and r2 are simulated over a range of fields encountered in CMR for correlation times of 0.1 nsec (typical of ECF agents), 1 nsec (intermediate motion), and 10 nsec (typical of albumin-bound agents). Figure 6-7 shows that the benefits of very slow rotation are seen at lower field strengths. Note also that r1 does not go to zero because there is also an outer-sphere component to relaxivity11 and the correlation times that govern outersphere relaxivity are quite short. Transverse relaxivity is defined by equations similar to Equations 4 and 5, with the exception that there are other mechanisms that cause r2 to increase at high fields.4,26 The r2 is always greater than r1, and for very slowtumbling systems, the r2/r1 ratio becomes large at high fields. Electronic and nuclear relaxation is described in greater detail in various reviews and books.4,26

100 90 80 70 60 50 40 30 20 10 0

100 90 80 70 60 50 40 30 20 10 0 0

0.5

A

1

1.5

2

2.5

0

0.5

1

B

Bo (Tesla)

Figure 6-7 suggests that slow-tumbling T1 agents become less effective at high fields, but relaxation times for tissue are longer at high fields and the signal-to-noise ratio increases with increased field strength (B0). These factors and the choice of sequence mean that a contrast agent with a lower relaxivity at 3 T than at 1.5 T may still provide greater contrast at 3 T than at 1.5 T. The iron oxide-based contrast agents are not discrete molecules, but crystals of iron oxide (Fe3O4) surrounded by a coating. The individual spins of each iron cooperatively via quantum mechanical interactions build up to give the crystal a very large total spin, and thus relaxivity will be a function of the number of spins. Such a material is called “ferromagnetic,” and its magnetism persists outside the external magnetic field. A weaker form of this is superparamagnetism, in which small particles of iron oxides with aligned spins occur in a magnetic field. Because the particles are small (submicron), the magnetic susceptibility effect is smaller than for large crystals of ferrites. Superparamagnets are no longer magnetic outside of the external field. The iron oxide particles consist of a core of one or more magnetic crystals of Fe3O4 embedded in a coating. Because these are materials, there is a distribution of sizes. USPIOs have a single crystal core and a submicron diameter (e.g., ferumoxtran [AMI-227, SineremW, Guerbet, Roissy, France aka Combidex, Advanced Magnetics, Cambridge, MA] has a crystal diameter of 4.3 to 4.9 nm and a global particle diameter of 50 nm).27 SPIOs have cores containing more than one crystal of Fe3O4 and are larger than USPIOs, but still submicron in size (e.g., ferumoxide [EndoremW, Guerbet, Roissy, France or FeridexW, Bayer Healthcare, Wayne, NJ] has a crystal diameter of 4.3 to 4.8 nm and a global particle diameter of 200 nm).27 USPIOs and SPIOs are small enough to form a stable suspension and can be administered intravenously. The difference in particle size determines their pharmacokinetic behavior. There are no inner-sphere water molecules in iron particles, and relaxation of water arises from the water molecules diffusing near the particle. However, the mechanism of

3

1.5

2

2.5

3

Bo (Tesla)

outer-sphere relaxation differs from that described earlier. One feature is that the crystals have a net magnetization, and as the external field is increased, this magnetization is increased (as is true for Gd, but the effect is much smaller). The modulation of this net magnetization can cause proton relaxation and have field dependence.28 There are some generalities about relaxivity in these particles. For the USPIOs, longitudinal relaxivity (r1) can be quite high and these can function as effective T1 agents. The r2/r1 ratio for USPIOs is significantly larger than for Gd complexes, and r2 increases with increasing magnetic field. When there is aggregation of crystals, which is the case in SPIOs, r1 tends to decrease and r2 increases. Thus, for the particles themselves as well as for aggregates of particles, the ratio of r2/r1 typically increases as the size of the particles or aggregates increases, although the T2 relaxivity as a function of particle size can be quite complicated. The effect of aggregation of crystals is that the aggregate itself can be considered a large magnetized sphere whose magnetic moment increases with increasing field strength. This gives rise to susceptibility effects and the SPIOs can act as T2* relaxation agents. This has important consequences when considering the effects of contrast agent compartmentalization on imaging (discussed later). Table 6-3 gives r1 and r2 values in plasma at 1.5 T and 3.0 T for a range of contrast media. The ECF agents have slightly lower r1 at 3.0 T and low r2/r1 ratios. The weak albumin binders have higher relaxivities, and the slow-tumbling Gd compounds, such as Gadomer or albumin-bound MS-325, have several-fold higher relaxivities than the ECF agents. As Figure 6-7 suggests, the slow-tumbling compounds also have significantly lower r1 at 3.0 T than at 1.5 T and the r2/r1 ratio is increased at 3.0 T. Three iron particle formulations are listed. The SPIOs FeridexW and ResovistW (Bayer Healthcare, Berlin, Germany) have a large r2 and a very large r2/r1 ratio, and this ratio is increased at 3.0 T. The USPIO SHU555C (Bayer Healthcare, Berlin, Germany) has good T1 relaxivity at 1.5 T, but at 3.0 T, the transverse relaxivity dominates.

Table 6-3 Relaxivities120 of Selected Contrast Media (0.25 mM) in Plasma at 1.5 T and 3.0 T at 37 C COMPOUND

1.5 T -1 -1

3T -1

-1

-1

-1

Chemical/Code

Generic name

r1 (mM s )

r2 (mM s )

r1 (mM s )

r2 (mM-1 s-1)

Gd-DTPA Gd-DTPA-BMA Gd-HPDO3A Gd-BOPTA MS-325* Gadomer

Gadopentetate Gadodiamide Gadoteridol Gadobenate Gadofosveset Gadomer-17

4.1 4.3 4.1 6.3 27.7 16.0

4.6 5.2 5.0 8.7 72.6 19.0

3.7 4.0 3.7 5.2 9.9 13.0

5.2 5.6 5.7 11.0 73 25

*Data from Eldredge HB, Spiller M, Chasse JM, et al. Species dependence on plasma protein binding and relaxivity of the gadolinium-based MRI contrast agent MS-325. Invest Radiol. 2006;41:229–243.

Cardiovascular Magnetic Resonance 83

6 CARDIOVASCULAR MAGNETIC RESONANCE CONTRAST AGENTS

Figure 6-7 Effect of rotational correlation time on longitudinal (r1, 7a) and transverse (r2, 7b) relaxivities as a function of field strength. Long correlation time (tR ¼ 10 nsec, solid line), typical of albumin binding, gives high r1 that decreases with increasing field and high r2; intermediate correlation time (tR ¼ 1 nsec, short dashed line) shows relaxivity maximum pushed out to a higher field; short correlation time (tR ¼ 0.1 nsec, long dashed line), typical of extracellular fluid agents, shows low, roughly field-independent r1, r2. Simulations with other parameters typical of gadolinium-based agents.11

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

CONTRAST-ENHANCED TISSUE RELAXATION The extent to which a metal complex influences tissue relaxation rates depends on three factors: 1. The chemical environment encountered by the metal complex. Binding of the agent to macromolecules can cause significant relaxivity enhancement. An example of this is shown in Figure 6-6, comparing the relaxivity of MS325 in buffer solution and in serum albumin solution. 2. Compartmentalization of the metal complex in tissue. Generally, tissue water is compartmentalized into intravascular, interstitial (fluid space between cells and capillaries), and intracellular space, constituting roughly 5%, 15%, and 80% of total water, respectively. Cellular organelles further subdivide the intracellular component. If water exchange between any of these compartments is slow relative to the relaxation rate in the compartment with the longest T1, multiexponential relaxation may result. This can decrease the effective tissue relaxivity of an agent because not all of the tissue water is encountering the paramagnetic center. 3. The magnetic susceptibility of the contrast agent. The contrast agent causes a microscopic field inhomogeneity on a biologic scale of 10 to 1000 nm rather than on the chemical scale of 0.1 to 1 nm. This results in a reduction in apparent T2. The result of the first two effects is that the simple relaxivity equation (Equation 1) often is not valid in a biologic setting; likewise, describing the effects on a CMR pulse sequence with a single tissue relaxivity can be misleading. Care must be taken when trying to estimate the concentration of a contrast agent from signal intensity changes. The actual relaxivity within a compartment might be affected by the biologic milieu. For some compounds with strong or weak albumin binding (e.g., MS-325 [gadofosveset], B22956 [gadocoletic acid], Gd-BOPTA [gadobenate]), local variation in the albumin concentration will affect the amount of contrast agent bound to albumin and thus affect the overall relaxivity. For example, in the plasma space, albumin concentration is typically high (0.6 to 0.7 mM) compared with the extracellular space in the normal heart (0.2 to 0.3 mM). The ECF agents interact very weakly if at all with plasma proteins or membrane structures. Tweedle and colleagues and Wedeking and associates showed that the relaxivity in blood and soft tissue of Gd-DTPA (gadopentetate) and Gd-DOTA (gadoterate) is the same as the relaxivity in aqueous solution.29,30 In extreme settings, the actual relaxivity could vary.31 Except for pathologic situations, most CMR contrast agents in use today are excluded from intact cells. Thus, the contrast agent will be localized to the extracellular space. As a result, the simple relaxivity equations do not necessarily hold. For a Gd-based ECF agent in a test tube, it takes approximately 3 msec for water to diffuse between Gd molecules32; in the time of a typical imaging TR, a given water molecule may interact with thousands or millions of Gd molecules, and all water molecules will interact with approximately the same number of Gd ions. However, if that same ECF agent is compartmentalized solely within the cardiac microvasculature, it takes between 2 and 20 seconds for most of the water in the tissue (85% of it is 84 Cardiovascular Magnetic Resonance

extravascular) to physically diffuse into the microvasculature; most of the water in the tissue does not have the opportunity to be relaxed by the Gd within TR of an imaging acquisition, resulting in a lower signal enhancement than that predicted by Equation 1 and assuming a uniform distribution of contrast agent throughout the tissue. To deal with compartmentalization, the concepts of “water exchange” and “exchange time,” t, between compartments are often used.33,34 The water exchange rate and the size of the compartments will determine the effect of the contrast agent on CMR signal. To illustrate this phenomenon, the two limiting causes of exchange will be described.32,33 In one extreme, water moves so quickly between the biologic compartments that the net effect is as if the contrast agent were uniformly spread throughout the entire tissue. This situation, called “fast exchange,” occurs whenever the exchange rate, 1/t, between the compartments is much faster than the difference in relaxation rates between the compartments.35 This occurs in blood, where red cells have a short water exchange time, on the order of 5 to 10 msec.36 The intact red cell prevents most CMR contrast agents from entering the cell, but as long as the plasma T1 is longer than 20 msec, the two compartments of the blood (plasma þ red cells) are in fast exchange and blood behaves for CMR purposes as if the contrast agent were spread uniformly through the blood. In this case, in general, the effective relaxation rate will be the weighted average of the relaxation in the two compartments. That is, if for compartment i the volume fraction is fi, the initial T1 is T1i and the concentration of agent is Ci (which could be zero), the entire tissue together will behave as shown: 1 1 1 ¼ f1 þ r1 C 1 þ f 2 þ r2 C 2 T1 T11 T12

(6)

In “slow exchange,” the water exchange rate is much slower than the difference in relaxation rates between the compartments. In this case, a single relaxation time, and thus a single relaxivity, is meaningless, because the two microscopic compartments will relax with their own relaxation times. Very few biologic compartments show true slow exchange, whereas the intermediate case, when exchange is neither slow nor fast (“intermediate exchange”) occurs very commonly. With intermediate exchange, relaxation behavior appears biexponential, although both the apparent compartment size and the effective T1 of the two compartments will vary from their true biologic size and T1. It is possible to model the signal intensity behavior as a function of contrast agent concentration to estimate water exchange times in vivo. Although characterizing human tissue as having only one or two compartments is an oversimplification, these types of models have proved useful for explaining the effects of biologic water mobility on contrast-enhanced scans.37 Biologic compartmentalization also results in susceptibility contrast. The contrast agent causes microscopic field inhomogeneities, sometimes called “mesoscopic” inhomogeneities.38 Water diffusion causes the protons to dephase from one another because of the different magnetic fields that they experience. Even in the absence of water diffusion, the field inhomogeneity causes intravoxel dephasing and thus signal loss on GRE images because of the different

NEWER CONTRAST AGENTS The agent Gd-BOPTA (gadobenate) has an approximately twofold higher relaxivity than the other ECF agents. Although it was designed for liver imaging, Gd-BOPTA distributes to extracellular space in the same manner as GdDTPA. The higher relaxivity can result in greater conspicuity and higher reader confidence for detecting lesions.39 Similarly, Gd-BOPTA has shown efficacy in contrastenhanced dynamic MRA and coronary MR.40,41 Recently, the hepatobiliary contrast agent Gd-EOB-DTPA (gadoxetic acid, marketed as Primovist in Europe and Eovist in the United States, Bayer Healthcare, Berlin, Germany) was approved for the detection and characterization of liver lesions. Like Gd-BOPTA, it has weak binding to plasma proteins and an approximately twofold higher relaxivity compared to ECF agents. In principal it should also be suitable for MRA applications but it is not approved for this purpose. The USPIO ferumoxytol (Feraheme, Advanced Magnetics, Cambridge, MA) was approved in 2009 for the treatment of iron deficiency anemia in adult patients with chronic kidney disease. This compound also produces good T1 and T2 contrast and an MRA indication is being pursued.42,43 MS-325 (gadofosveset, formerly marketed as Vasovist, now marketed as Ablavar by Lantheus Medical Imaging, Billerica, MA) was designed specifically for CE-MRA applications and is approved for use in Europe, Canada, and Australia, and subsequently (December 2008) the first contrast agent approved for CE-MRA in the United States. In the United States, gadofosveset is indicated for use as a contrast agent in MRA to evaluate aortoiliac occulusive disease in adults with known or suspected peripheral vascular disease. This agent may well have many uses beyond dynamic MRA. Its extended plasma half-life and vascular retention enables steady-state CE-MRA, offering the potential benefits of imaging multiple vascular beds with a single injection. It allows acquisition of images with higher spatial resolution than is achievable with dynamic MRA, eliminating the need for bolus timing, and it may be possible to administer the agent several minutes before the patient enters the magnet. It remains to be seen how dependent these benefits will be on the ability to do robust artery and vein separation.

NOVEL CONTRAST AGENTS IN DEVELOPMENT Molecular imaging has been defined as “the in vivo characterization and measurement of biological processes at the cellular and molecular level.”44 Combining the high-resolution anatomic and

functional imaging ability of CMR with the specificity of molecular targeting is very appealing, but somewhat limited by the inherent sensitivity limitations of CMR. Examples of molecular CMR contrast agents used for cardiovascular applications are numerous. The vast majority of these agents have only been studied in animal models. A few are described here. Scientists at EPIX Pharmaceuticals (Lexington, MA) designed gadolinium-based agents that target fibrin,45,46 an abundant component of thrombus. These molecules contain a peptide sequence specific for fibrin that has been linked to several gadolinium chelates. This class of compounds has been shown to positively enhance thrombus visualization in animal models of atherosclerotic plaque rupture47 and carotid artery thrombus,48 and in embolic models of thrombus in the coronary arteries,49,50 left atrium,51 and lung.52 Figure 6-8 shows an example of acute coronary thrombus visualization with the experimental fibrin targeted agent EP2104R in a porcine model. This contrast agent has been shown to visualize thrombus in human subjects, one of the first examples of clinical translation of a molecular MR probe.53 The Barnes Hospital-Washington University group developed a platform technology based on large particles prepared from a perfluorocarbon emulsion.54,55 The particles can be functionalized by adding molecules with lipophilic components that are noncovalently incorporated into the particle. Signal generation arises from thousands of lipophilic gadolinium chelates incorporated on the surface. Targeting vectors can similarly be incorporated on the surface to guide the particle’s distribution. This group used antibodies to target the particles to either fibrin56–58 or tissue factor.59 Alternately, small molecules that bind to the avb3 integrin were incorporated and the particle was used to detect overexpression of this integrin that occurs during angiogenesis.60–63 Use of the avb3-targeted particle to deliver fumagillin to inhibit angiogenesis has also been described while monitoring the process by molecular MRI.64 Imaging and characterization of atherosclerotic plaque has been an active area of investigation (see Chapters 27 and 28). There are some new contrast agents that have shown efficacy in animal models of atherosclerosis. One interesting class of compounds is the “gadofluorines.”65–67 These amphiphilic small molecules consist of a Gd chelate linked to a perfluorocarbon chain and a hydrophilic group, such as a sugar. Perfluorocarbons do not associate with lipids and impart interesting biodistribution properties. Gadofluorine M has been shown to localize in plaques induced in Watanabe heritable hyperlipidemic rabbits, but not in the vessel wall of normal New Zealand white rabbits.67 Figure 6-9 shows an image of a rabbit aorta taken 48 hours after gadofluorine M injection. The lesions are clearly delineated on this inversion recovery image and are confirmed with histology. The Mt. Sinai group has taken many approaches to imaging plaque targeting various plaque components. One approach was to use high-density lipoprotein as a carrier for a lipophilic contrast agent.68 The high-density lipoprotein is believed to facilitate transport of the contrast agent to the plaque. They used mixed micelle platforms to nonspecifically target plaques69 and also incorporated antibodies to macrophage-scavenging receptor to home the mixed micelles to macrophage in the plaque.70,71 In collaboration with the Guerbet group, they reported a Gd-based agent targeting matrix metalloproteases in the plaque.72 A Gd-based contrast agent specific to type I collagen was described,73 and this probe was used to show in vivo molecular imaging of fibrosis in a mouse model of healed postinfarction myocardial scarring.74 This probe may find utility in other pathologies with elevated collagen levels. The Massachusetts General Hospital group developed a gadolinium-based agent (MPO-Gd) sensitive to myeloperoxidase (MPO) activity. Ischemic injury of the myocardium causes timed recruitment of neutrophils and monocytes/macrophages, which produce substantial amounts of local MPO. MPO forms reactive Cardiovascular Magnetic Resonance 85

6 CARDIOVASCULAR MAGNETIC RESONANCE CONTRAST AGENTS

microscopic magnetic fields within the voxel. The strength of the perturbing magnetic field is directly proportional to contrast agent concentration and its molar magnetic susceptibility (w). The actual magnitude and even the direction of the magnetic field shifts depend strongly on the size and the shape of the biologic compartment in which the contrast agent resides38; the size of the susceptibility contrast effect depends on how the water diffuses through the tissue.

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

A

B

C

D

Figure 6-9 Gadofluorine-enhanced M (0.1 mmol/kg, 48 hours postinjection) image of atherosclerotic lesions in an 18-month-old Watanabe hyperlipidemic rabbit. Segmented inversion recovery gradient recalled echo cardiovascular magnetic resonance. (Courtesy of Dr. B. Misselwitz, Bayer Schering, Berlin, Germany.) 86 Cardiovascular Magnetic Resonance

Figure 6-8 Acute coronary thrombus visualization with thrombin-specific gadoliniumbased agent, EP-2104R. (A) Double oblique slice from brightblood coronary magnetic resonance angiography sequence (multiplanar reconstruction) demonstrating lumen of left anterior descending (LAD) artery with bright signal (arrows). (C) X-ray angiography demonstrating thrombus in LAD (arrow). (B and D) Double oblique slices from an inversion recovery sequence (multiplanar reconstruction, same orientation as in A, before (B) and 2 hours after (D) administration of the agent. The left anterior descending coronary artery thrombus is seen as a “bright” spot on postcontrast images (arrow in B and D). (Courtesy of Dr. A. Wiethof, EPIX Pharmaceuticals, Cambridge, MA.)

chlorinating species capable of inflicting oxidative stress and altering protein function by covalent modification. MPO-Gd is first radicalized by MPO and then either spontaneously oligomerizes or binds to matrix proteins, all leading to enhanced spin-lattice relaxivity and delayed washout kinetics. MPO-Gd was used to measure inflammatory responses after myocardial ischemia locally and noninvasively in a murine model.75 There is a rich literature describing the use of iron oxide particles as targeted contrast agents. Commercial USPIOs passively accumulate in macrophages. This property has been exploited to provide negative contrast in macrophage-rich atherosclerotic plaques. There appears to be a positive correlation between macrophage uptake and plasma half-life.76,77 Iron oxides have also been used to label cells, and then the cells are tracked in vivo using CMR.78,79 Visualization of mesenchymal stem cells engrafted into the myocardium of a pig has been demonstrated using a 1.5 T clinical scanner.80,81 Injection sites containing greater than 105 cells could be detected in vivo.80 In a mouse model of myocardial infarction, a therapeutic intervention of mouse embryonic stem cells could be followed by visualizing the stem cells and following their effect on LV function over a period of 4 weeks.82 The Massachusetts General Hospital group has also shown that the iron oxide coating material can be chemically modified to introduce new targeting vectors.83–85 They termed this contrast agent platform “cross-linked iron oxide.” For example, annexin V can be grafted onto a cross-linked iron oxide86 to detect apoptotic cardiomyocytes in a mouse model of transient left anterior descending artery occlusion.87 Another CLIO-based probe is targeted

SAFETY The safety of CMR contrast agents, which themselves offer no direct therapeutic benefit, is always a question of appropriate medical concern. The Gd-based chelates had initially been considered among the safest injectable compounds in medical use, with a specific reputation for superior safety in patients with renal dysfunction (compared with an iodinated preparation). This view has changed with the apparent link between Gd contrast and nephrogenic systemic fibrosis (NSF). A very rare, but potentially devastating condition affecting patients with chronic kidney disease, NSF has an estimated prevalence of more than 10% among patients receiving Gd who are on hemodialysis.93 Although presenting primarily with skin thickening, tethering, hyperpigmentation, and disabling joint flexion contractures, patients with NSF can also have multiorgan/systemic fibrosis, leading to organ dysfunction and even death.94,95 The universal feature of NSF is Gd administration and chronic kidney disease, most commonly with an estimated glomerular filtration rate of less than 30 mL/min/1.73m2.95–98 Gadolinium has been shown to be present in skin biopsy specimens of patients with NSF,99–104 and also in internal organs.105 The prevailing hypothesis is that chronic kidney disease results in prolonged exposure to the contrast agent, providing opportunity for the Gd to dissociate from its chelator, and this dissociated Gd is believed to be responsible for the toxic response.106 The least stable contrast agent, Gd-DTPA-BMA (gadodiamide), is the agent that has been most frequently associated with NSF,107–109 but it is now established that other Gd compounds also pose a risk.93,110

Although ECF agents are similar in terms of their imaging efficacy, they differ in terms of how well they bind Gd. This may prove to be an important risk factor for NSF because both the metal (Gd) and the chelate have potential toxic effects5 and have shown acute toxicity in animal studies at high enough doses. These various animal studies have been reviewed,111 and a recent high-dose study in rats showed NSF-like lesions.109 Metal complexes, such as Gd contrast agents, are characterized in terms of their thermodynamic stability and kinetic inertness. Stability is a measure of the affinity of the chelator for the metal ion and is expressed as an equilibrium constant for the association of the chelator (ligand) with the metal. Stability constants for approved agents have the order Gd-DOTA (gadoterate) > GdHPDO3A (gadoteridol) Gd-DO3A-butrol (gadobutrol) MS-325 (gadofosveset) Gd-BOPTA (gadobenate) GdDTPA (gadopentetate) > Gd-DTPA-BMA (gadoversetamide) Gd-DTPA-BMEA (gadodiamide).4 The compounds with lowest thermodynamic stability are the neutral complexes with acyclic chelators: Gd-DTPA-BMA (gadodiamide) and Gd-DTPA-BMEA (gadoversetamide). Stability constants describe equilibrium values, but do not indicate how quickly equilibrium is reached. In the biologic milieu, protons and other metal ions compete to bind the chelator ligand, whereas there are ions such as phosphate and carbonate that have a high affinity for Gd. Under conditions that favor dissociation of the Gd (e.g., low pH), two complexes with comparable stability may have very different rates of dissociation. For example, the macrocyclic complex Gd-HPDO3A (gadoteridol) is an example of a compound that is more kinetically inert to Gd loss than Gd-DTPA (gadopentetate), which has a similar stability constant. The rate of acid-assisted Gd dissociation is 20 times faster for Gd-DTPA (gadopentetate) than for Gd-HPDO3A (gadoteridol).112 Laurent and colleagues measured the rate of Gd loss in the presence of zinc and phosphate under a standard set of conditions for various approved agents.113,114 The macrocyclic complexes GdDOTA (gadoterate), Gd-HPDO3A (gadoteridol), and GdDO3A-butrol (gadobutrol) were the most inert with respect to Gd loss. The compound that released Gd the most rapidly was Gd-DTPA-BMA (gadoversetamide) under the conditions specified by Laurent and colleagues. These differences in kinetics and thermodynamic stabilities may well translate into safety with regard to patients with chronic kidney disease. For instance, there appears to be a much lower incidence of NSF among patients who received Gd-HPDO3A (gadoteridol),115 and this may be a result of its combination of stability and inertness. Although care should be taken when considering the use of Gd-based contrast in renally compromised patients (estimated glomerular filtration rate < 60 mL/min/1.73m2), it is important to keep this risk in perspective. Gd-based contrast agents have been used safely in millions of patients, and to date there have been no reports of Gd-associated NSF in patients with normal renal function. Regarding other side effects, Kirchin and Runge116 reviewed the safety record of clinically approved contrast agents as of 2003. The seven approved Gd agents (Gd-DTPA [gadopentetate], Gd-DOTA [gadoterate], GdBOPTA [gadobenate], Gd-DTPA-BMA [gadoversetamide], Gd-DTPA-BMEA [gadodiamide), Gd-DO3A-butrol [gadobutrol], and Gd-BOPTA [gadobenate]) appeared indistinguishable with respect to their safety profile. The most common Cardiovascular Magnetic Resonance 87

6 CARDIOVASCULAR MAGNETIC RESONANCE CONTRAST AGENTS

to vascular cell adhesion molecule 1 using a peptide specific to this molecule.88 This probe was used to identify inflammatory activation of cells in a mouse model of atherosclerosis. The lack of sensitivity in CMR stems from the very small degree of polarization among the nuclear spins. In a magnetic field there is a net magnetization, but this is small, and approximately 0.0006% of the spins are polarized. A technique called “spin exchange” uses a high-powered laser (also called “optical pumping”) to increase the polarization by four to five orders of magnitude (hyperpolarization).89 Isotopes with long T1 values can be hyperpolarized and used as contrast agents. The long T1 is necessary to maintain the contrast medium in the hyperpolarized state long enough to obtain an image before the spins relax back to the equilibrium value. Gases often have long T1 values, and isotopes of the noble gases helium (He-3) and xenon (Xe-129) have been used for imaging. Contrast agents with hyperpolarized C-13 have been reported using a C-13-enriched water-soluble compound90 with long relaxation times (in vitro: T1 ¼ approximately 82 seconds, T2 ¼ approximately 18 seconds; in vivo: T1 ¼ approximately 38 seconds, T2 ¼ approximately 1.3 seconds). This could be formulated at a C-13 concentration of 200 mM and hyperpolarized to 15%. The authors used this material for CE-MRA in rats and swine.91 Hyperpolarized C-13 also offers the potential for investigating metabolic pathways. For instance, Merritt and colleagues studied the metabolism of [1-13C]-pyruvate in a perfused rat heart.92 This field is likely to expand in the coming years. A major benefit of hyperpolarized contrast media is the excellent sensitivity with no background (high signal-to-noise ratio). Challenges include the distribution and availability of the hyperpolarization equipment and imaging hardware compatibility for imaging nonhydrogen nuclei (not available on all clinical scanners). However, there is now a commercial polarizer available.

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

adverse events were headache, nausea, taste perversion, and urticaria (hives). Nearly all adverse events with these agents were transient, mild, and self-limiting. Nevertheless, there are reports of serious adverse reactions, including life-threatening anaphylactoid reactions and death. The best estimate puts the rate of these events at between 1 in 200,000 and 1 in 400,000 patient administrations.117 The Gd-based ECF contrast agents have been studied, and most are approved for pediatric use in patients older than 2 years, though there are differences in their approval wording. One distinguishing clinical indicator among the Gd-based agents is that patients receiving Gd-DTPA-BMA (gadoversetamide) or Gd-DTPABMEA (gadodiamide) often show spuriously low serum calcium levels.118 These two contrast agents (but not the other Gd-based agents) appear to interfere with the reagent used in the clinical chemistry test for calcium.119 The safety profile for current CMR contrast agents of course will not necessarily be the same for future compounds. For all contrast agents, the package insert should always be consulted for the latest safety information.

CONCLUSION A large number of contrast agents have been approved for MRI in the last 18 years. In December 2008, the first blood pool agent (gadofosveset) was U.S. Food and Drug Administration–approved for a CMR application. A significant number of new agents, including blood pool and tissue-specific agents that are potentially relevant for CMR, are currently in development. The behavior of these agents, which often is summarized by a single effectiveness parameter, r1, is actually quite complex, in terms of both the underlying chemistry and the biophysics in vivo. Although many elements of that complexity remain active areas of research, both for understanding existing agents as well as for creating new agents and new uses for those agents, the relative safety and ease of use of these agents has brought them into routine use in many medical applications. As clinical CMR expands, no doubt the use of these contrast agents will expand as well.

References 1. Bloch F, Hansen WW, Packard M. The nuclear induction experiment. Phys Rev. 1948;70:474–485. 2. Lauterbur PC, Mendonca Dias MH, Rudin AM. Augmentation of tissue water proton spin-lattice relaxation rates by in vivo addition of paramagnetic ions. In: Dutton PO, Leigh J, Scarpa A, eds. Frontiers of Biological Energetics. New York: Academic Press; 1978:752–759. 3. Carr DH, Brown J, Bydder GM, et al. Intravenous chelated gadolinium as a contrast agent in NMR imaging of cerebral tumors. Lancet. 1984;1:484–486. 4. Caravan P, Ellison JJ, McMurry TJ, et al. Gadolinium(III) chelates as MRI contrast agents: structure, dynamics, and applications. Chem Rev. 1999;99:2293–2352. 5. Gries H. Extracellular MRI contrast agents based on gadolinium. Top Curr Chem. 2002;221:1–24. 6. Powell DH, Ni Dhubhghaill OM, Pubanz D, et al. High-pressure NMR kinetics. Part 74. Structural and dynamic parameters obtained from 17O NMR, EPR, and NMRD Studies of monomeric and dimeric Gd3þ Complexes of Interest in magnetic resonance imaging: an integrated and theoretically self-consistent approach. J Am Chem Soc. 1996;118:9333–9346. 7. Gillis A, Gray M, Burstein D. Relaxivity and diffusion of gadolinium agents in cartilage. Magn Reson Med. 2002;48:1068–1071. 8. Kumar K, Tweedle MF. Ligand basicity and rigidity control formation of macrocyclic polyamino carboxylate complexes of gadolinium(III). Inorg Chem. 1993;32:4193–4199. 9. White DH, DeLearie LA, Moore DA, et al. The thermodynamics of complexation of lanthanide(III) DTPA-bisamide complexes and their implication for stability and solution structure. Invest Radiol. 1991; 26:S226–S228. 10. Tombach B, Heindel W. Value of 1.0-M gadolinium chelates: review of preclinical and clinical data on gadobutrol. Eur Radiol. 2002;12: 1550–1556. 11. Caravan P, Cloutier NJ, Greenfield MT, et al. The interaction of MS325 with human serum albumin and its effect on proton relaxation rates. J Am Chem Soc. 2002;124:3152–3162. 12. Lauffer RB, Parmelee DJ, Dunham SU, et al. MS-325: albumin: targeted contrast agent for MR angiography. Radiology. 1998;207: 529–538. 13. Deshpande VS, Cavagna F, Maggioni F. Comparison of gradient echo and SSFP for coronary artery MR angiography using a gadolinium based intravascular contrast agent. Invest Radiol. 2006;41:292–298. 14. de Hae¨n C, Anelli PL, Lorusso V, et al. Gadocoletic acid trisodium salt (B22956/1): a new blood pool magnetic resonance contrast agent with application in coronary angiography. Invest Radiol. 2006;41: 279–291. 15. de Hae¨n C, La Ferla R, Maggioni F. Gadobenate dimeglumine 0.5 M solution for injection (MultiHance) as contrast agent for magnetic 88 Cardiovascular Magnetic Resonance

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26. 27. 28. 29.

resonance imaging of the liver: mechanistic studies in animals. J Comp Asst Tomog. 1999;23:S169–S179. Vander Elst L, Maton F, Laurent S, et al. A multinuclear MR study of Gd-EOB-DTPA: comprehensive preclinical characterization of an organ specific MRI contrast agent. Magn Reson Med. 1997;38: 604–614. Eldredge HB, Spiller M, Chasse JM, et al. Species dependence on plasma protein binding and relaxivity of the gadolinium-based MRI contrast agent MS-325. Invest Radiol. 2006;41:229–243. Port M, Corot C, Violas X, et al. How to compare the efficiency of albumin-bound and nonalbumin-bound contrast agents in vivo: the concept of dynamic relaxivity. Invest Radiol. 2005;40:565–573. Dong Q, Hurst DR, Weinmann HJ, et al. Magnetic resonance angiography with gadomer-17: an animal study original investigation. Invest Radiol. 1998;33:699–708. Nicolle GM, Toth E, Schmitt-Willich H, et al. The impact of rigidity and water exchange on the relaxivity of a dendritic MRI contrast agent. Chem Eur J. 2002;8:1040–1048. Schnorr J, Wagner S, Abramjuk C, et al. Comparison of the iron oxide-based blood-pool contrast medium VSOP-C184 with gadopentetate dimeglumine for first-pass magnetic resonance angiography of the aorta and renal arteries in pigs. Invest Radiol. 2004;39: 546–553. Tombach B, Reimer P, Bremer C, et al. First-pass and equilibriumMRA of the aortoiliac region with a superparamagnetic iron oxide blood pool MR contrast agent (SH U 555 C): results of a human pilot study. NMR Biomed. 2004;17:500–506. Lauffer RB. Targeted relaxation enhancement agents for MRI. Magn Reson Med. 1991;22:339. Bogdanov AA, Weissleder R, Frank HW, et al. A new macromolecule as a contrast agent for MR angiography: preparation, properties, and animal studies. Radiology. 1993;187:701–706. Toth E, Van Uffelen I, Helm L, et al. Gadolinium-based linear polymer with temperature-independent proton relaxivities: a unique interplay between the water exchange and rotational contributions. Magn Reson Chem. 1998;36:S125–S134. Bertini I, Luchinat C, Parigi G. Solution NMR of Paramagnetic Molecules. Amsterdam: Elsevier; 2001. Jung CW, Jacobs P. Physical and chemical properties of superparamagnetic iron oxide MR contrast agents: ferumoxides, ferumoxtran, ferumoxsil. Magn Reson Imaging. 1995;13:661–674. Muller RN, Roch A, Colet J-M, et al. Particulate magnetic contrast agents. In: Merbach AE, Toth E, eds. Chemistry of Contrast Agents in Medical Magnetic Resonance Imaging. 2001:417–435. Tweedle MF, Wedeking P, Telser J, et al. Dependence of MR signal intensity on Gd tissue concentration over a broad dose range. Magn Reson Med. 1991;22:191–194.

55. Lanza GM, Winter P, Caruthers S, et al. Novel paramagnetic contrast agents for molecular imaging and targeted drug delivery. Curr Pharm Biotechnol. 2004;5:495–507. 56. Winter PM, Caruthers SD, Yu X, et al. Improved molecular imaging contrast agent for detection of human thrombus. Magn Reson Med. 2003;50:411–416. 57. Flacke S, Fischer S, Scott MJ, et al. Novel MRI contrast agent for molecular imaging of fibrin: implications for detecting vulnerable plaques. Circulation. 2001;104:1280–1285. 58. Yu X, Song SK, Chen J, et al. High-resolution MRI characterization of human thrombus using a novel fibrin-targeted paramagnetic nanoparticle contrast agent. Magn Reson Med. 2000;44:867–872. 59. Morawski AM, Winter PM, Crowder KC, et al. Targeted nanoparticles for quantitative imaging of sparse molecular epitopes with MRI. Magn Reson Med. 2004;51:480–486. 60. Schmieder AH, Winter PM, Caruthers SD, et al. Molecular MR imaging of melanoma angiogenesis with alphanubeta3-targeted paramagnetic nanoparticles. Magn Reson Med. 2005;53:621–627. 61. Winter PM, Morawski AM, Caruthers SD, et al. Molecular imaging of angiogenesis in early-stage atherosclerosis with alpha(v)beta3integrin-targeted nanoparticles. Circulation. 2003;108:2270–2274. 62. Winter PM, Caruthers SD, Kassner A, et al. Molecular imaging of angiogenesis in nascent Vx-2 rabbit tumors using a novel alpha(nu) beta3-targeted nanoparticle and 1.5 tesla magnetic resonance imaging. Cancer Res. 2003;63:5838–5843. 63. Anderson SA, Rader RK, Westlin WF, et al. Magnetic resonance contrast enhancement of neovasculature with alpha(v)beta(3)-targeted nanoparticles. Magn Reson Med. 2000;44:433–439. 64. Winter PM, Neubauer AM, Caruthers SD, et al. Endothelial alpha(v) beta3 integrin-targeted fumagillin nanoparticles inhibit angiogenesis in atherosclerosis. Arterioscler Thromb Vasc Biol. 2006;26:2103–2109. 65. Meding J, Urich M, Licha K, et al. Magnetic resonance imaging of atherosclerosis by targeting extracellular matrix deposition with Gadofluorine M. Contrast Media Mol Imaging. 2007;2:120–129. 66. Sirol M, Itskovich VV, Mani V, et al. Lipid-rich atherosclerotic plaques detected by gadofluorine-enhanced in vivo magnetic resonance imaging. Circulation. 2004;109:2890–2896. 67. Barkhausen J, Ebert W, Heyer C, et al. Detection of atherosclerotic plaque with Gadofluorine-enhanced magnetic resonance imaging. Circulation. 2003;108:605–609. 68. Frias JC, Williams KJ, Fisher EA, et al. Recombinant HDL-like nanoparticles: a specific contrast agent for MRI of atherosclerotic plaques. J Am Chem Soc. 2004;126:16316–16317. 69. Briley-Saebo KC, Amirbekian V, Mani V, et al. Gadolinium mixedmicelles: effect of the amphiphile on in vitro and in vivo efficacy in apolipoprotein E knockout mouse models of atherosclerosis. Magn Reson Med. 2006;56:1336–1346. 70. Mulder WJ, Strijkers GJ, Briley-Saboe KC, et al. Molecular imaging of macrophages in atherosclerotic plaques using bimodal PEG-micelles. Magn Reson Med. 2007;58:1164–1170. 71. Amirbekian V, Lipinski MJ, Briley-Saebo KC, et al. Detecting and assessing macrophages in vivo to evaluate atherosclerosis noninvasively using molecular MRI. Proc Natl Acad Sci USA. 2007;104:961–966. 72. Lancelot E, Amirbekian V, Brigger I, et al. Evaluation of matrix metalloproteinases in atherosclerosis using a novel noninvasive imaging approach. Arterioscler Thromb Vasc Biol. 2008;28:425–432. 73. Caravan P, Das B, Dumas S, et al. Collagen-targeted MRI contrast agent for molecular imaging of fibrosis. Angew Chem Int Ed Engl. 2007;46:8171–8173. 74. Helm PA, Caravan P, French BA, et al. Postinfarction myocardial scarring in mice: molecular MR imaging with use of a collagen-targeting contrast agent. Radiology. 2008;247:788–796. 75. Nahrendorf M, Sosnovik D, Chen JW, et al. Activatable magnetic resonance imaging agent reports myeloperoxidase activity in healing infarcts and noninvasively detects the antiinflammatory effects of atorvastatin on ischemia-reperfusion injury. Circulation. 2008;117: 1153–1160. 76. Corot C, Petry KG, Trivedi R, et al. Macrophage imaging in central nervous system and in carotid atherosclerotic plaque using ultrasmall superparamagnetic iron oxide in magnetic resonance imaging. Invest Radiol. 2004;39:619–625. 77. Yancy AD, Olzinski AR, Hu TC, et al. Differential uptake of ferumoxtran-10 and ferumoxytol, ultrasmall superparamagnetic iron oxide contrast agents in rabbit: critical determinants of atherosclerotic plaque labeling. J Magn Reson Imaging. 2005;21:432–442. 78. Modo M, Hoehn M, Bulte JW. Cellular MR imaging. Mol Imaging. 2005;4:143–164.

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6 CARDIOVASCULAR MAGNETIC RESONANCE CONTRAST AGENTS

30. Wedeking P, Sotak CH, Telser J, et al. Quantitative dependence of MR signal intensity on tissue concentration of Gd(HP-DO3A) in the nephrectomized rat. Magn Reson Imaging. 1992;10:97–108. 31. Stanisz GJ, Henkelman RM. Gd-DTPA relaxivity depends on macromolecular content. Magn Reson Med. 2000;44:665–667. 32. Donahue KM, Weisskoff RM, Burstein D. Water diffusion and exchange as they influence contrast enhancement. J Magn Reson Imaging. 1997;7:102–110. 33. Li X, Rooney WD, Springer Jr CS. A unified magnetic resonance imaging pharmacokinetic theory: intravascular and extracellular contrast reagents. Magn Reson Med. 2005;54:1351–1359. 34. Hazlewood CF, Chang DC, Nichols BL. Nuclear magnetic resonance transverse relaxation times of water protons in skeletal muscle. Biophys J. 1974;14:583–606. 35. McLaughlin AC, Leigh JS. Relaxation times in systems with chemical exchange. J Magn Reson. 1973;9:296–304. 36. Wright GA, Hu BS, Macovski A. Estimating oxygen saturation of blood in vivo with MR imaging at 1.5 T. J Magn Reson Imaging. 1991;1: 275–283. 37. Donahue KM, Weisskoff RM, Chesler DA, et al. Improving MR quantification of regional blood volume with intravascular T1 contrast agents: accuracy, precision, and water exchange. Magn Reson Med. 1996;36:858–867. 38. Sukstanskii AL, Yablonskiy DA. Gaussian approximation in the theory of MR signal formation in the presence of structure-specific magnetic field inhomogeneities: effects of impermeable susceptibility inclusions. J Magn Reson. 2004;167:56–67. 39. Knopp MV, Runge VM, Essig M, et al. Primary and secondary brain tumors at MR imaging: bicentric intraindividual crossover comparison of gadobenate dimeglumine and gadopentetate dimeglumine. Radiology. 2004;230:55–64. 40. Goyen M, Debatin JF. Gadobenate dimeglumine (MultiHance) for magnetic resonance angiography: review of the literature. Eur Radiol. 2003;13:N19–N27. 41. Wikstrom J, Wasser MN, Pattynama PM, et al. Gadobenate dimeglumine-enhanced magnetic resonance angiography of the pelvic arteries. Invest Radiol. 2003;38:504–515. 42. Neuwelt EA, Varallyay CG, Manninger S, et al. The potential of ferumoxytol nanoparticle magnetic resonance imaging, perfusion, and angiography in central nervous system malignancy: a pilot study. Neurosurgery. 2007;60:601–611. 43. Li W, Salanitri J, Tutton S, et al. Lower extremity deep venous thrombosis: evaluation with ferumoxytol-enhanced MR imaging and dual-contrast mechanism—preliminary experience. Radiology. 2007;242:873–881. 44. Weissleder R, Mahmood U. Molecular imaging. Radiology. 2001;219: 316–333. 45. Caravan P, Kolodziej AF, Greenwood JM, et al. EP-1242: A Fibrin Targeted Contrast Agent for Thrombus Imaging. In: 10th ISMRM Scientific Sessions. HI, USA: Honolulu; 2002. 46. Overoye-Chan K, Koerner S, Looby RJ, et al. EP-2104R: a fibrinspecific gadolinium-based MRI contrast agent for detection of thrombus. J Am Chem Soc. 2008;130:6025–6039. 47. Botnar RM, Perez AS, Witte S, et al. In vivo molecular imaging of acute and subacute thrombosis using a fibrin-binding magnetic resonance imaging contrast agent. Circulation. 2004;109:2023–2029. 48. Sirol M, Aguinaldo JG, Graham PB, et al. Fibrin-targeted contrast agent for improvement of in vivo acute thrombus detection with magnetic resonance imaging. Atherosclerosis. 2005;182:79–85. 49. Spuentrup E, Buecker A, Katoh M, et al. Molecular magnetic resonance imaging of coronary thrombosis and pulmonary emboli with a novel fibrin-targeted contrast agent. Circulation. 2005;111:1377–1382. 50. Botnar RM, Buecker A, Wiethoff AJ, et al. In vivo magnetic resonance imaging of coronary thrombosis using a fibrin-binding molecular magnetic resonance contrast agent. Circulation. 2004;110: 1463–1466. 51. Spuentrup E, Katoh M, Wiethoff AJ, et al. Molecular magnetic resonance imaging of pulmonary emboli with a fibrin-specific contrast agent. Am J Respir Crit Care Med. 2005;172:494–500. 52. Spuentrup E, Katoh M, Buecker A, et al. Molecular MR imaging of human thrombi in a swine model of pulmonary embolism using a fibrin-specific contrast agent. Invest Radiol. 2007;42:586–595. 53. Spuentrup E, Botnar RM, Wiethoff AJ, et al. MR imaging of thrombi using EP-2104R, a fibrin-specific contrast agent: initial results in patients. Eur Radiol. 2008;18:1995–2005. 54. Cyrus T, Winter PM, Caruthers SD, et al. Magnetic resonance nanoparticles for cardiovascular molecular imaging and therapy. Expert Rev Cardiovasc Ther. 2005;3:705–715.

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79. Bulte JW, Kraitchman DL. Monitoring cell therapy using iron oxide MR contrast agents. Curr Pharm Biotechnol. 2004;5:567–584. 80. Hill JM, Dick AJ, Raman VK, et al. Serial cardiac magnetic resonance imaging of injected mesenchymal stem cells. Circulation. 2003;108: 1009–1014. 81. Kraitchman DL, Heldman AW, Atalar E, et al. In vivo magnetic resonance imaging of mesenchymal stem cells in myocardial infarction. Circulation. 2003;107:2290–2293. 82. Arai T, Kofidis T, Bulte JW, et al. Dual in vivo magnetic resonance evaluation of magnetically labeled mouse embryonic stem cells and cardiac function at 1.5 t. Magn Reson Med. 2006;55:203–209. 83. Wunderbaldinger P, Josephson L, Weissleder R. Crosslinked iron oxides (CLIO): a new platform for the development of targeted MR contrast agents. Acad Radiol. 2002;9:S304–S306. 84. Moore A, Sun PZ, Cory D, et al. MRI of insulitis in autoimmune diabetes. Magn Reson Med. 2002;47:751–758. 85. Hogemann D, Josephson L, Weissleder R. Improvement of MRI probes to allow efficient detection of gene expression. Bioconjug Chem. 2000;11:941–946. 86. Schellenberger EA, Bogdanov Jr A, Hogemann D. Annexin V-CLIO: a nanoparticle for detecting apoptosis by MRI. Mol Imaging. 2002;1:102–107. 87. Sosnovik DE, Schellenberger EA, Nahrendorf M, et al. Magnetic resonance imaging of cardiomyocyte apoptosis with a novel magnetooptical nanoparticle. Magn Reson Med. 2005;54:718–724. 88. Nahrendorf M, Jaffer FA, Kelly KA, et al. Noninvasive vascular cell adhesion molecule-1 imaging identifies inflammatory activation of cells in atherosclerosis. Circulation. 2006;114:1504–1511. 89. Moller HE, Chen XJ, Saam B, et al. MRI of the lungs using hyperpolarized noble gases. Magn Reson Med. 2002;47:1029–1051. 90. Svensson J, Mansson S, Johansson E, et al. Hyperpolarized 13C MR angiography using trueFISP. Magn Reson Med. 2003;50:256–262. 91. Olsson LE, Chai CM, Axelsson O, et al. MR coronary angiography in pigs with intraarterial injections of a hyperpolarized 13C substance. Magn Reson Med. 2006;55:731–737. 92. Merritt ME, Harrison C, Storey C, et al. Hyperpolarized 13C allows a direct measure of flux through a single enzyme-catalyzed step by NMR. Proc Natl Acad Sci USA. 2007;104:19773–19777. 93. Todd DJ, Kagan A, Chibnik LB, et al. Cutaneous changes of nephrogenic systemic fibrosis: predictor of early mortality and association with gadolinium exposure. Arthritis Rheum. 2007;56:3433–3441. 94. Cowper SE. Nephrogenic systemic fibrosis: a review and exploration of the role of gadolinium. Adv Dermatol. 2007;23:131–154. 95. Cowper SE, Kuo PH, Bucala R. Nephrogenic systemic fibrosis and gadolinium exposure. Association and lessons for idiopathic fibrosing disorders. Arthritis Rheum. 2007;56:3173–3175. 96. Grobner T. Gadolinium–a specific trigger for the development of nephrogenic fibrosing dermopathy and nephrogenic systemic fibrosis? Nephrol Dial Transplant. 2006;21:1104–1108. 97. Thomsen HS. Nephrogenic systemic fibrosis: a serious late adverse reaction to gadodiamide. Eur Radiol. 2006;16:2619–2621. 98. Grobner T, Prischl FC. Gadolinium and nephrogenic systemic fibrosis. Kidney Int. 2007;72:260–264. 99. Boyd AS, Zic JA, Abraham JL. Gadolinium deposition in nephrogenic fibrosing dermopathy. J Am Acad Dermatol. 2007;56:27–30. 100. High WA, Ayers RA, Chandler J, et al. Gadolinium is detectable within the tissue of patients with nephrogenic systemic fibrosis. J Am Acad Dermatol. 2007;56:21–26. 101. High WA, Ayers RA, Cowper SE. Gadolinium is quantifiable within the tissue of patients with nephrogenic systemic fibrosis. J Am Acad Dermatol. 2007;56:710–712. 102. Saussereau E, Lacroix C, Cattaneo A, et al. Hair and fingernail gadolinium ICP-MS contents in an overdose case associated with nephrogenic systemic fibrosis. Forensic Sci Int. 2008;176:54–57. 103. Thakral C, Alhariri J, Abraham JL. Long-term retention of gadolinium in tissues from nephrogenic systemic fibrosis patient after multiple gadolinium-enhanced MRI scans: case report and implications. Contrast Media Mol Imaging. 2007;2:199–205. 104. Abraham JL, Thakral C, Skov L, et al. Dermal inorganic gadolinium concentrations: evidence for in vivo transmetallation and long-term persistence in nephrogenic systemic fibrosis. Br J Dermatol. 2008; 158:273–280.

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105. Kay J, Bazari H, Avery LL, et al. Case records of the Massachusetts General Hospital. Case 6-2008: a 46-year-old woman with renal failure and stiffness of the joints and skin. N Engl J Med. 2008;358: 827–838. 106. Kurtkoti J, Snow T, Hiremagalur B. Gadolinium and nephrogenic systemic fibrosis: association or causation. Nephrology (Carlton). 2008;. 107. Richmond H, Zwerner J, Kim Y, et al. Nephrogenic systemic fibrosis: relationship to gadolinium and response to photopheresis. Arch Dermatol. 2007;143:1025–1030. 108. Rydahl C, Thomsen HS, Marckmann P. High prevalence of nephrogenic systemic fibrosis in chronic renal failure patients exposed to gadodiamide, a gadolinium-containing magnetic resonance contrast agent. Invest Radiol. 2008;43:141–144. 109. Sieber MA, Pietsch H, Walter J, et al. A preclinical study to investigate the development of nephrogenic systemic fibrosis: a possible role for gadolinium-based contrast media. Invest Radiol. 2008;43: 65–75. 110. Leiner T, Herborn CU, Goyen M. Nephrogenic systemic fibrosis is not exclusively associated with gadodiamide. Eur Radiol. 2007;17: 1921–1923. 111. Caravan P, Lauffer RB. Contrast agents: Basic principles. In: Edelman RR, Hesselink JR, Zlatkin MB, et al, eds. Clinical Magnetic Resonance Imaging. 3rd ed. Philadelphia: Saunders; 2005:357–375. 112. Kumar K, Chang CA, Tweedle MF. Equilibrium and kinetic studies of lanthanide complexes of macrocyclic polyamino carboxylates. Inorg Chem. 1993;32:587–593. 113. Laurent S, Elst LV, Muller RN. Comparative study of the physicochemical properties of six clinical low molecular weight gadolinium contrast agents. Contrast Media Mol Imaging. 2006;1:128–137. 114. Laurent S, Elst LV, Copoix F, et al. Stability of MRI paramagnetic contrast media: a proton relaxometric protocol for transmetallation assessment. Invest Radiol. 2001;36:115–122. 115. Reilly RF. Risk for nephrogenic systemic fibrosis with gadoteridol (prohance) in patients who are on long-term hemodialysis. Clin J Am Soc Nephrol. 2008;3:747–751. 116. Kirchin MA, Runge VM. Contrast agents for magnetic resonance imaging: safety update. Top Magn Reson Imaging. 2003;14:426–435. 117. Carr JJ. Magnetic resonance contrast agents for neuroimaging. Safety issues. Neuroimaging Clin N Am. 1994;4:43–54. 118. Prince MR, Erel HE, Lent RW, et al. Gadodiamide administration causes spurious hypocalcemia. Radiology. 2003;227:639–646. 119. Emerson J, Kost G. Spurious hypocalcemia after Omniscan- or OptiMARK-enhanced magnetic resonance imaging: an algorithm for minimizing a false-positive laboratory value. Arch Pathol Lab Med. 2004;128:1151–1156. 120. Rohrer M, Bauer H, Mintorovitch J, et al. Comparison of magnetic properties of MRI contrast media solutions at different magnetic field strengths. Invest Radiol. 2005;40:715–724. 121. Tweedle MF, Hagan JJ, Kumar K, et al. Reaction of gadolinium chelates with endogenously available ions. Magn Reson Imaging. 1991;9:409–415. 122. Tweedle MF. Physicochemical properties of gadoteridol and other magnetic resonance contrast agents. Invest Radiol. 1992;27:2–6. 123. Vogler H, Platzek J, Schuhmann-Giampieri G, et al. Pre-clinical evaluation of gadobutrol: a new, neutral, extracellular contrast agent for magnetic resonance imaging. Eur J Radiol. 1995;21:1–10. 124. OptiMARK Package Insert. St. Louis: Mallinckrodt, Inc.. 125. Cavagna FM, Lorusso V, Anelli PL, et al. Preclinical profile and clinical potential of gadocoletic acid trisodium salt (B22956/1), a new intravascular contrast medium for MRI. Acad Radiol. 2002;9: S491–S494. 126. Uggeri F, Aime S, Anelli PL, et al. Novel contrast agents for magnetic resonance imaging: synthesis and characterization of the ligand BOPTA and Its Ln(III) Complexes (Ln ¼ Gd, La, Lu). X-ray Structure of Disodium (TPS-9-145337286-C-S)-[4-Carboxy-5,8,11-tris(carboxymethyl)-1-phenyl-2-oxa- 5,8,11-triazatridecan-13-oato(5-)]gadolinate(2-) in a mixture with its enantiomer. Inorg Chem. 1995;34:633–642. 127. de Hae¨n C, Gozzini L. Soluble-type hepatobiliary contrast agents for MR imaging. J Magn Reson Imaging. 1993;1993:3.

Blood Flow Velocity Assessment David Firmin

The idea of mapping measurements of blood flow onto a magnetic resonance (MR) image was first discussed in an article by Singer in 1978.1 The methods that followed could generally be categorized into time-of-flight or phase shift types, and were based on the techniques that had previously been described for nonimaging MR flow studies.2 A number of review articles have covered the subject and described the variety of methods3–5 that have been used and validated both in vitro and in vivo. The interest in flow in MR imaging has not been solely directed toward the goal of quantitative flow measurement. A large amount of effort has also been devoted to understanding the appearance of a flowing fluid on an image because this can often be indicative of the type of flow present and therefore can give important information on the diagnosis of a particular disorder. Also, the development of MR angiography techniques has required a full understanding of these effects. In 1984, soon after the development of the first clinical MR scanners, there was an increase in interest in the search for an MR method of imaging flow. Review articles were published and a number of techniques described.6–10 This chapter provides a description of and a brief historical overview of the methods that have been used to measure blood flow in the heart and great vessels.

TIME-OF-FLIGHT METHODS There are two categories of time-of-flight techniques. The first, often known as washin/washout, or flow enhancement, methods, normally rely on the saturation or partial saturation of material in a selected slice or volume being replaced by fully magnetized “high signal” spins as a result of flow (Fig. 7-1A). The second involves some form of tagging and then imaging to follow the motion of the tagged material (see Fig. 7-1B). Singer and Crooks11 adopted the first approach in an attempt to measure flow in the internal jugular veins, although quantification was questionable because of other factors affecting the flow signal. The first report to describe a tagged time-of-flight approach was by Feinberg and colleagues.10 Their method involved a variation on a dual echo spin echo sequence; the first 180 selected slice was displaced by 3 mm from the initial excitation slice, and the second was displaced by 9 mm. The first 180 selection overlapped sufficiently with the 90 selection to produce a good anatomic image. The second 180 pulse selection did not overlap with the 90 or the first 180 pulse selection and therefore produced no anatomic image,

but gave high signal from the blood that had experienced all of the preceding radiofrequency (RF) pulses (i.e., that had passed between the different selected planes). The technique was used to identify flow in the carotid and vertebral arteries of a volunteer’s neck, although flow velocities had to be within a specific range and could not be defined accurately. Methods have also been described in which the time-offlight flow movement can be visualized directly on an image.12,13 The methods involved the application of slice selection and frequency encoding in the same axis. In this way, material that had moved in this axis between selection and readout would be displaced relative to the stationary material. These techniques were therefore making use of signal misregistration, an effect that is often seen as a problem in other methods of flow imaging. Another timeof-flight approach is to saturate a band of tissue, for example, in a transverse plane, and then to follow the progress of this dark band in the coronal or sagittal plane.14 The major limitation of these saturation methods is that they are limited by the T1 of the various tissues being saturated. The contrast of the saturated blood will decrease with time, eventually making it difficult to measure accurately the distances traveled. Also, motion during the sampling gradients results in signal and thus image distortion.15 In addition, for arterial flow measurements in which cardiac gating is required, only two-dimensional (2D) images can be acquired in a reasonable time so that only limited details of the flow profile can be studied.

PHASE FLOW IMAGING METHODS Considerable knowledge had been gained on the measurement of flow from the phase of the CMR signal from the nonimaging studies after an original suggestion by Hahn in 1960 of a method of measuring the slow flow of currents in the sea.16 In 1984, the first attempts of measuring blood flow were described by van Dijk8 and Bryant and associates,9 using methods based on the theory suggested by Moran 2 years earlier.17 The imaging methods that followed fell broadly into two categories: 1. Phase contrast velocity mapping methods that mapped the phase of the signal directly to measure the flow. 2. Fourier flow imaging methods that phase encoded flow velocity to produce an image after Fourier transformation with velocity resolved on one image axis. Cardiovascular Magnetic Resonance 91

7 BLOOD FLOW VELOCITY ASSESSMENT

CHAPTER 7

Signal intensity

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

location will take up a frequency shift that depends on their position in the direction of the field gradient. When the gradient is turned off, the phase of the flowing and stationary materials can be considered equal. In the period between the positive and negative gradients, the flowing material moves away from its stationary neighbor. During the period of the negative gradient, the stationary material takes up an equal but opposite frequency shift and returns to the phase it had before the first gradient. However, the flowing fluid takes up a different frequency shift that is dependent on the distance it has moved; its final phase, therefore, also depends on this distance and hence its velocity. The relationship between the phase of the signal and flow velocity is:

A

Flow rate

f ¼ gvDAg

B

Time 1 (tagging)

Time 2 (detection)

Figure 7-1 Time-of-flight approaches to magnetic resonance flow measurement. In the saturation wash-in method (A), the signal is enhanced by an amount related to the in-flow of fully magnetized blood. In the true time-of-flight method (B), there is a time of spatial labeling or tagging, followed after a defined period by detection of the movement of the labeled blood.

Both of these methods rely on the same principles that cause flowing material to attain a phase shift that is related to its motion. Figure 7-2 shows these principles for a fluid flowing down a tube surrounded by stationary material. A bipolar gradient pulse is applied, consisting of a positive magnetic field gradient, followed a certain time later by an equal but opposite negative magnetic field gradient in the direction of the flow. During the period of the positive gradient, flowing and stationary materials in a particular

Gradient waveform

Time

Time

Magnetic field

Position

Position

Flow

Flow Signal phase Stationary Flowing material material

Stationary material

Time 1

Flowing material

Time 2

Figure 7-2 Principles of phase velocity encoding. At time 1, a positive magnetic field gradient is applied that results in an equal frequency and associated phase shift for neighboring stationary and flowing spins. At time 2, an equal but opposite magnetic field is applied. By this time, the spins in the flowing blood have separated from their original neighbors to be in a different strength of magnetic field during the gradient application. The result is that although the phase of the stationary spins will be returned to zero, the flowing spins will accumulate a phase shift proportional to the distance moved and hence the velocity. 92 Cardiovascular Magnetic Resonance

(1)

where Ag is the area of one gradient pulse (amplitude duration), D is the time between the centers of the two gradient pulses, v is the velocity, and g is the gyromagnetic ratio. Ag, D and g are all constants for a particular imaging sequence so that a quantitative measure of velocity can be determined if the phase shift can be measured. The two principal approaches of using the phase shift to produce a quantitative flow image, phase contrast velocity mapping and Fourier flow imaging, are discussed in this chapter.

Phase Contrast Velocity Mapping The early phase contrast velocity mapping methods8,9 used a spin echo sequence, which was not ideal because of problems in repeating the sequence rapidly, and signal loss as a result of shear and other more complex flows. These problems were reduced and the methods were made clinically more useful, partly by the use of a gradient echo sequence18,19 and, more importantly, by the introduction of velocity-compensated gradient waveforms.20,21 Normally, two images are acquired with different gradient waveforms in the direction of desired flow measurement. The difference in the waveforms is calculated to produce a well-defined velocity-related phase difference between the two images. A phase reconstruction is produced for each of the images, which are then subtracted pixel by pixel to produce the final velocity map. This process of subtraction removes any phase variations that are not related to flow. The velocity phase sensitivity of the final image is normally set such that the expected velocity-related phase shifts are within the range of p radians. If a larger range of velocities is present, then aliasing or wrap-around will occur, resulting in measurement of ambiguous velocities. This problem can be avoided by reducing the velocity sensitivity or potentially can be corrected by a process known as phase unwrapping (described later). Figure 7-3A shows the phase velocity images of a series of time frames from a slice just above the heart. Flow can be seen in the ascending and descending aortas, the pulmonary artery, and the superior vena cava. Flow versus time curves throughout the cardiac cycle are shown in Figure 7-3B. The stroke volume can be measured by

250 msec

40

MPA

AA MPA DA SVC

35 30

DA

300 msec

350 msec

400 msec

Flow (L/min)

SVC

25 20 15 10 5 0 –5

A

B

0

200

400

600

800

Time (msec)

Figure 7-3 A, Flow velocity images of a transverse slice at the level of the right pulmonary artery showing head/foot flow velocities at six times in the cardiac cycle. B, Plot of measured volume flow versus time for the ascending aorta (AA), descending aorta (DA), main pulmonary artery (MPA), and superior vena cava (SVC).

integrating under the aortic or pulmonary flow curve. The technique has been validated both in vitro and in vivo,22 and is now routinely used to provide useful measurements in clinical and physiologic flow studies.23

Fourier Flow Imaging Fourier flow imaging normally involves the addition of a bipolar velocity phase encoding gradient that is stepped through a range of defined amplitudes (Fig. 7-4). Because this increases the scan time by a multiple of the number of steps, it is often applied as a replacement of one of the spatial phase encoding gradients. The image is normally reconstructed and displayed with velocity information in one dimension. Stationary material is positioned in the center of the image, with faster velocities toward the edge. The method was first described by Redpath and colleagues24 in 1984; in this case, eight velocity phase encoding steps were added to a 2D imaging sequence, to image velocities in a circle of fluid-filled tubing that rotated in the image plane. Different segments of the circle, each corresponding to different velocity ranges, were seen on the eight resultant images. A year later, Feinberg and coworkers25 applied the method, both in vitro and

SPATIAL PHASE ENCODING IS REPLACED BY VELOCITY PHASE ENCODING

Spatial phase encoding gradient waveform

Velocity phase encoding gradient waveform

Figure 7-4 Gradient waveforms required for spatial and velocity phase encoding.

in vivo, and increased the velocity range and resolution by increasing the number of flow phase encoding steps. In this case, however, to maintain a tolerable scan time, only one spatial phase encoding direction was used so that there was only spatial resolution in one direction. The accuracy of the method was shown with a phantom, whereas the in vivo study, which showed the flow in the descending aorta, highlighted the problem of very high zero velocity signal from the large amount of stationary tissue imaged. In 1988, Hennig and colleagues.26 described a development of this method in which the signal from stationary tissue was saturated and the sequence was repeated much more rapidly. The issue of time precluding the use of spatial phase encoding remains a problem, although 2D radiofrequency (RF) pulses have been used successfully to locate signals within a column.27,28 More recently, Luk Pat and associates.29 showed a method of real-time Fourier velocity imaging that used an excited column to localize the signals. In vivo aortic flow waveforms were presented with a temporal resolution of only 33 msec.

IMPROVING THE ACCURACY OF PHASE CONTRAST VELOCITY MEASUREMENTS The vast majority of MR flow imaging applications have used the method of phase contrast velocity mapping. The accuracy of this method is highly dependent on such factors as flow pulsatility, velocity, and size and tortuosity of the vessel. One simple approach to improving the overall accuracy of the method is to adjust the velocity sensitivity of the sequence so that the velocity-related phase shift is close to 2p for the maximum expected velocity. Buonocore30 extended this approach by varying the velocity sensitivity during the cardiac cycle, based on the knowledge that the arterial flow velocity is high in systole but low in diastole. The accuracy can be improved further by allowing a velocity-related phase shift greater Cardiovascular Magnetic Resonance 93

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45 150 msec

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Figure 7-5 Method of phase unwrapping. A, Systolic image in which high velocities result in aliasing in both the positive and negative directions. B and C, Adjustment of the velocity window to remove aliasing in the positive and negative directions, respectively. D, The same image data after processing by the anti-aliasing algorithm.

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than 2p that will result in aliasing that can be corrected by use of a phase unwrapping algorithm (Fig. 7-5).31 Another approach to improving accuracy has been suggested by Bittoun and associates.32 The method is a combination of phase contrast velocity mapping and Fourier velocity imaging with a small number of velocity phase encoding steps. The final phase contrast velocity map is calculated from the best fit through the Fourier velocity encoded result. One potential problem with this method that can also affect conventional phase contrast velocity imaging occurs if there is any beat-to-beat variation in flow velocity. As a result of the high-velocity phase sensitivity used with this method, significant phase variations can occur, with resulting ghosting artifacts and loss of flow information. Another method that was originally described to give a measure of velocity and flow quantification was phase contrast angiography. This technique again involves acquiring two image datasets with sequences that have opposite phase velocity sensitivities, although in this case, the raw data are subtracted before reconstruction. This technique has the advantage of subtracting out signal from stationary tissue, which removes errors caused by partial volumes where voxels contained a mixture of flowing and stationary tissue. The method is, however, generally less accurate because the signal and hence the velocity measurement can be affected by factors such as in-flow enhancement and intravoxel dephasing (signal loss). In the mid-1990s, Polzin and colleagues.33 suggested a method of combining this method with phase contrast velocity mapping, which they showed to be more accurate in a number of phantom studies. The methods are yet to be fully validated in vivo, however, and are likely to be affected by problems of signal loss and motion, particularly when imaging is performed on small mobile vessels, such as the coronary arteries.34 94 Cardiovascular Magnetic Resonance

One of the most significant factors that can affect the accuracy of the flow measurement methods is flow-related signal loss. This is normally the result of loss of phase coherence within a voxel, and eventually it results in an inability to detect the encoded phase of the flow signal above the random phase of the background noise. Even if a velocity-compensated imaging sequence is used, the acceleration and even the higher orders of motion present in complex flows can result in loss of phase coherence. Figure 7-6 shows an example of a long axis image of a patient with a mitral valve stenosis in which the valve is also regurgitant. In this case, a region of blood signal is lost from the ventricle during diastole as a result of the stenosis generating complex flows and from the atrium during systole because of a regurgitant jet of flow through the valve. Partial signal loss, however, does not greatly affect the accuracy of the phase contrast velocity mapping measurement unless it is accompanied by partial volume errors. When signal loss is the result of a spread of phase within a voxel, the mean phase will be detected, although this will be affected by differential saturation effects.35 The phase contrast velocity mapping techniques are most susceptible to signal loss of one form or another, although this can normally be minimized by appropriate gradient profile design. A good way to reduce signal loss is to use a symmetrical gradient waveform that nullifies phase shifts caused by all of the odd-order derivatives of position and then to shorten the sequence as much as possible to reduce the effects of the even-order derivatives.36 Signal loss of the type described is much less of a problem with the Fourier flow imaging method. In this case, the Fourier transform is used to separate out constituent velocities. Errors can be caused by phase differences for reasons other than the velocity encoding pulses resulting in the phase map velocity values being offset from zero, even for stationary

tissues.37 This background phase offset generally varies gradually with position across the image, and it also varies with image plane orientation and other sequence parameters that affect the gradient waveforms, such as Venc. Distortion of the requested magnetic field gradients is unavoidable because of the fundamental laws of electromagnetism, and these are known in MR imaging as Maxwell, or concomitant, gradients. These background phase shifts become more significant when high-amplitude gradients are used and also when imaging is performed at lower main magnetic field strengths. However, with the latest gradient systems, Maxwell gradient effects are certainly a factor for phase velocity mapping at 1.5 Tesla. However, these velocity map offsets can be corrected precisely and automatically in software, with no user intervention required.38 A second common reason for background phase shifts is the presence of small uncorrected side effects of the gradient pulses in the magnet, known as eddy currents. These phase shifts became more of a problem with the advent of higher performance gradients, although more recently, the problem appears to have been reduced. Software is sometimes provided that allows the user to place markers identifying stationary tissues so that this background phase error can be calculated and removed from the entire image. Sometimes, however, if the phase shifts are nonlinear or if there is little signal from stationary tissue, then the only way to correct for them is to acquire an additional set of images of a large static phantom using the same sequence parameters as those used in vivo, and then to subtract out the phase errors on a pixel-by-pixel basis.39 Pixels without any signal in velocity maps have a random phase or show the phase of a weak ghost that may not be visible on the magnitude image with normal brightness settings. Particularly for poststenotic jet images, it is important to check that the magnitude image pixels of the jet are not affected by signal loss. Avoiding the inclusion of noise pixels can be problematic in regions of interest around the great vessels, and ideally, the image analysis software enables the user to set the velocity to zero for pixels whose magnitude is below a user-defined threshold. Provided that the sequence parameters are carefully chosen such that the potential errors and artifacts discussed earlier can be minimized or avoided, phase velocity mapping has been shown to be accurate and reproducible. Validation has been reported in phantoms by comparison with true measured flow and with Doppler36,40,41 in animal models by comparison with in vivo flow meter measurements.42

7 BLOOD FLOW VELOCITY ASSESSMENT

Figure 7-6 Systolic and diastolic long axis frames from cine datasets acquired in a patient with a stenotic and regurgitant mitral valve. Signal loss can be seen in the left atrium (LA) during systole and the left ventricle (LV) during diastole.

Validation has also been reported in humans by comparison with methods such as Doppler ultrasound or catheterization.43–45 Perhaps the most convincing forms of validation were those performed initially by comparing the aortic flow with the left ventricular stroke volume and later flow in both the aorta and pulmonary artery with left and right ventricular stroke volumes in normal subjects.22,46,47 For the latter, the four measurements should be the same except for small differences caused by coronary and bronchial flow, and it can be calculated that flow measurements in large vessels are accurate to within 6%. The effect of breath holding on flow measurement is another factor that could affect more recent studies. Sakuma and associates showed a significant change in both pulmonary and aortic cardiac output during a large-lungvolume breath hold.48 Conversely, flows measured during a small volume breath hold were found to be similar to those measured during normal breathing. There are other potential sources of error that have been reported. These include misalignment of the vessel with the direction of velocity encoding and misregistration of flow signal caused by flow between excitation and readout. Because of these and the other sources of errors discussed earlier, care is required when setting up the scan parameters, to minimize their effect.

Rapid Phase Flow Imaging Methods With the very rapid scanning hardware available today, it is possible to repeat a phase contrast velocity sequence so fast that low-resolution images can be acquired in 100 msec or high-resolution images can be acquired in a breath hold. The major problem is for pulsatile flow where the accuracy of the measurements and the temporal resolution can be limited if the acquisition period per cardiac cycle is too long. Also, if high spatial resolution is required, the cardiac motion of structures, such as coronary arteries, can cause blurring, with subsequent errors in flow measurement.49 For this reason, it is likely that more efficient k-space coverage methods, such as interleaved spirals, will be important. Ultra-fast flow imaging techniques have also been developed, either by combining a phase mapping type approach with imaging methods, such as single-shot echo planar and Cardiovascular Magnetic Resonance 95

Visualizing Flow and Flow Parameters The method used to visualize MR flow data has depended on the method used for acquisition. For Fourier velocity measurement, each voxel may contain a range of measured velocities, and the Fourier velocity image normally takes the form of a plot of velocity versus time or velocity versus position in one direction. Figure 7-8 shows an example in

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Figure 7-7 Real-time magnitude and flow images from the excited region containing the aorta and the superior vena cava. The plot shows the variation in flow during a period of respiratory maneuver. 96 Cardiovascular Magnetic Resonance

ECG Gate Delay 100 msec 105 msec 115 msec 125 msec 135 msec 145 msec 155 msec 165 msec 175 msec

250 mm FOV

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spiral imaging,50,51 or by imaging only one spatial dimension.52 A compromise generally has to be made in temporal or spatial resolution and probably also in the signal-tonoise ratio. However, taking into account these constraints, the methods have generally been shown to be accurate. One complication with the echo planar sequence is flow signal loss because of its inherent phase sensitivity, even when additional flow compensation is applied. However, this has been used to advantage for more qualitative flow imaging showing flow disturbances, for example.53 The one-dimensional rapid acquisition mode, real-time acquisition and velocity evaluation (RACE),52 can be used to measure flow perpendicular to the slice. The technique can be repeated rapidly throughout the cardiac cycle to give near-real-time flow information. One problem with this type of approach is that data are acquired from a projection through the patient; this means that any signal overlapping with the flow signal will combine and introduce errors to flow measurement. Several strategies have been suggested for localizing the signal to avoid this: they include spatial presaturation, projection dephasing (applying a gradient to suppress stationary tissue), and collecting a cylinder of data and multiple oblique measurements. Yang and colleagues used a 2D RF excitation scheme to excite a narrow rectangular X-section column and used only 16 echoes to spatially resolve the other dimension in high resolution.54 This approach allowed real-time flow measurements to be acquired. The authors used the method to show the effect of controlled breathing on flow in the ascending aorta and superior vena cava (Fig. 7-7).

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Figure 7-8 Series of nine Fourier velocity images at 10-msec intervals showing the velocity pulse wave propagating down the descending aorta. ECG, electrocardiogram.

which velocity images were acquired from a column of excited tissue, including the descending aorta.55 The front edge of the aortic pulse wave can be seen on successive frames as it travels down the vessel. Phase contrast velocity images contain only one velocity measure per image voxel. Historically, these have been shown with a grayscale such that flow in one direction tends toward white, flow in the other direction tends toward black, and stationary material is mid-gray, as shown previously in Figure 7-3A. When flow is measured in more than one direction, then more sophisticated methods of display can be used. Figure 7-9 shows a vector map of flow in the root of the aorta of a patient with an atherosclerotic aneurysm. This systolic image, shown alongside a pressure map (described later), shows high-velocity flow impinging on the wall of the aneurysm. An alternative type of representation would be to use the cine velocity images to calculate the path of a seed over time.56 The ability to study flow in such detail and at any site in the body is unique to MR imaging. For this reason, a large amount of interest is being generated from those who wish to understand the physiology of blood flow and its interaction with blood vessels and the cardiac chambers. Despite the relatively poor spatial resolution, a number of groups have studied methods of extracting a measure of the wall shear stress from the MR images. Both Oshinski and colleagues and Oyre and associates developed fitting methods to derive the velocity profile at subpixel distances from the vessel wall.57,58 Both groups presented expected values of stress, although it is difficult to suggest a method of validating the accuracy of these measurements. Frayne and Rutt suggested an alternative approach that potentially gave more information about the flow within a voxel that straddled the vessel wall.59 Their method used Fourier velocity encoding to distinguish the distribution of flow velocities, so that only the spatial location had to be considered. There has also been considerable interest in the possibility of deriving pressure measurements from MR images. Urchuk and coworkers considered vessel compliance and the flow

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Figure 7-9 Flow vector map showing the systolic flow pattern in the aortic root of a patient with an atherosclerotic aneurysm. The associated image shows the corresponding pressure distribution.

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Figure 7-10 Selected frames from a cine series of calculated flow pressure maps showing the variation in pressure at different times in the cardiac cycle (A to E). Each grayscale band from white to black represents a pressure gradient of 1 mm Hg. A positive pressure gradient during systole (A) reverses during the deceleration phase of diastole (D).

pulse wave to calculate the pressure waveform and showed a good correlation with catheter pressure measurements made in a pig model.60 In contrast, Yang and colleagues derived flow pressure maps from the cine phase contrast velocity maps using the Navier-Stokes equations.61 Figure 7-10 shows an example of the changing flow pressure around the aortic arch during the first half of the cardiac cycle. Figure 7-11 shows an interesting example of a flow pressure map, showing the descending aorta in a patient who underwent polyester fiber (Dacron, Dupont, Wilmington, DE) graft repair of aortic coarctation. In contrast to the rest of the aorta, no pressure gradient can be seen in the repaired region, possibly because of the reduced compliance. To fully understand, visualize, and measure flow parameters in blood vessels, a method of acquisition is required that measures velocity in three dimensions and three directions over time, with high spatial and temporal resolution. Even with the fastest gradient systems available today, this presents a significant problem in terms of acquisition time and data handling. A method has been suggested, however, using a combination of echo planar and k-space view sharing that suggests that the acquisition times can be reduced to acceptable levels.62 These types of methods are now being used to visualize and study flow patterns in much more detail.63,64

Figure 7-11 Systolic pressure map of the aortic arch in a patient with a polyester fiber (Dacron, Dupont, Wilmington, DE) repair. No pressure gradient is seen in the region of the repair. Cardiovascular Magnetic Resonance 97

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References 1. Singer JR. NMR diffusion and flow measurement and an introduction to spin phase graphing. J Phys E: Sci Instrum. 1978;11:281–291. 2. Jones DW, Child TF. NMR in flowing systems. Adv Magn Reson. 1976;8:123–148. 3. Bradley WG. Flow phenomenon in MR imaging. AJR. 1988;150: 983–994. 4. Alfidi RJ, Masaryk TJ, Haacke EM, et al. MR angiography of peripheral, carotid, and coronary arteries. AJR. 1987;149:1097. 5. Firmin DN, Dumoulin C, Mohiaddin RH. Quantiative MR flow measurement. In: Haacke EM, Potchen EJ, Gottschalk A, Siebert JE, eds. Magnetic Resonance Angiography: Concepts and Applications. St Louis: Mosby; 1993, pp. 187–219. 6. Crooks LE, Kaufman L. NMR imaging of blood flow. Brit Med Bul. 1984;40:167–169. 7. Axel L. Blood flow effects in magnetic resonance imaging. AJR. 1984;143:1157–1166. 8. van Dijk P. Direct cardiac NMR imaging of heart wall and blood flow velocity. J Comput Assist Tomogr. 1984;8:429–436. 9. Bryant DJ, Payne JA, Firmin DN, Longmore DB. Measurement of flow with NMR imaging using a gradient pulse and phase difference technique. J Comput Assist Tomogr. 1984;8:588–593. 10. Feinberg DA, Crooks LE, Hoenninger J, et al. Pulsatile blood velocity in human arteries displayed by magnetic resonance imaging. Radiology. 1984;153:177–180. 11. Singer JR, Crooks LE. Nuclear magnetic resonance blood flow measurements in the human brain. Science. 1983;221:654–656. 12. Shimizu K, Matsuda T, Sakurai T, et al. Visualisation of moving fluid: quantitative analysis of blood flow velocity using MR imaging. Radiology. 1986;159:195–199. 13. Axel L, Shimakawa A, MacFall J. A time-of-flight method of measuring flow velocity by magnetic resonance imaging. Magn Reson Imaging. 1986;4:199–205. 14. Edelman RR, Mattle HP, Kleefield J, Silver MS. Quantification of blood flow with dynamic MR imaging and presaturation bolus tracking. Radiology. 1989;171:551–556. 15. Izen SH, Haacke EM. Measuring non-constant flow in magnetic resonance imaging. IEEE Trans Med Imaging. 1990;9:450–460. 16. Hahn EL. Detection of sea-water motion by nuclear precession. Geophys Res. 1960;65:776–777. 17. Moran PR. A flow zeugmatographic interlace for NMR imaging in humans. Magn Reson Imaging. 1982;1:197–203. 18. Young IR, Bydder GM, Payne JA. Flow measurement by the development of phase differences during slice formation in MR imaging. Magn Reson Med. 1986;3:175–179. 19. Ridgway JP, Smith MA. A technique for velocity imaging using magnetic resonance imaging. Brit J Radiol. 1986;59:603–607. 20. Nayler GL, Firmin DN, Longmore DB. Blood flow imaging by cine magnetic resonance. J Comput Assist Tomogr. 1986;10:715–722. 21. Haacke EM, Lenz GW. Improving MR image quality in the presence of motion by using rephasing gradients. AJR. 1987;148:1251–1258. 22. Firmin DN, Nayler GL, Klipstein RH, et al. In vivo validation of MR velocity imaging. J Comput Assist Tomogr. 1987;11:751–756. 23. Mohiaddin RH, Pennell DJ. MR Blood flow measurement: clinical application in the heart and circulation. Cardiol Clin. 1998;16: 161–187. 24. Redpath TW, Norris DG, Jones RA, Hutchinson MS. A new method of NMR flow imaging. Phys Med Biol. 1984;29:891–895. 25. Feinberg DA, Crooks LE, Sheldon P, et al. Magnetic resonance imaging and velocity vector components of fluid flow. Magn Reson Med. 1985;2:555–566. 26. Hennig J, Mueri M, Brunner P, Friedburg H. Quantitative flow measurement with the fast Fourier flow technique. Radiology. 1988; 166:237–240. 27. Gatehouse PD, Link K, Bebbington MWP, et al. Pulse-wave and stenosis studies by cylinder excitation with Fourier velocity encoding [Abstract]. In: Proceedings of the Third Annual Meeting of the International Society of Magnetic Resonance. 1995;318. 28. Hardy CJ, Bolster Jr BD, McVeigh ER, et al. Pencil excitation with interleaved Fourier velocity encoding: NMR measurement of aortic distensibility. Magn Reson Med. 1996;35:814–819. 29. Luk Pat GT, Pauly JM, Hu BS, Nishimura DG. One-shot spatially resolved velocity imaging. Magn Reson Med. 1998;40:603–613. 30. Buonocore MH. Blood flow measurement using variable velocity encoding in the RR interval. Magn Reson Med. 1993;29:790–795.

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31. Yang GZ, Burger P, Kilner PJ, et al. Dynamic range extension of cine velocity measurements using motion registered spatio-temporal phase unwrapping. J Magn Reson Imag. 1996;6:495–502. 32. Bittoun J, Bourroul E, Jolivet O, et al. High-precision MR velocity mapping by 3D-Fourier phase encoding with a small number of encoding steps. Magn Reson Med. 1993;29:674–680. 33. Polzin JA, Alley MT, Korosec FR, et al. A complex-difference phasecontrast technique for measurement of volume flow rates. J Magn Reson Imaging. 1995;5:129–137. 34. Frayne R, Polzin JA, Mazaheri Y, Grist TM, Mistretta CA. Effect of and correction for in-plane myocardial motion on estimates of coronaryvolume flow rates. J Magn Reson Imaging. 1997;7:815–828. 35. Polzin JA, Korosec FR, Wedding KL, et al. Effects of through-plane myocardial motion on phase-difference and complex-difference measurements of absolute coronary artery flow. J Magn Reson Imaging. 1996;6:113–123. 36. Firmin DN, Nayler GL, Kilner PJ, Longmore DB. The application of phase shifts in NMR for flow measurement. Magn Reson Med. 1990;14:230–241. 37. Kilner PJ, Gatehouse PD, Firmin DN. Flow measurement by magnetic resonance: a unique asset worth optimising. J Cardiovasc Magn Reson. 2007;9:723–728. 38. Bernstein MA, Zhou XJ, Polzin JA, et al. Concomitant gradient terms in phase contrast MR: analysis and correction. Magn Reson Med. 1998;39:300–308. 39. Chernobelsky A, Shubayev O, Comeau CR, Wolff SD. Baseline correction of phase contrast images improves quantification of blood flow in the great vessels. J Cardiovasc Magn Reson. 2007;9:681–685. 40. Kilner PJ, Firmin DN, Mohiaddin RH, Underwood SR, Rees RSO, Longmore DB. Valve and great vessel stenosis: assessment with MR jet velocity mapping. Radiology. 1991;178:229–235. 41. Meier D, Maier S, Boesiger P. Quantitative flow measurements on phantoms and on blood vessels with MR. Magn Reson Med. 1988;8: 25–34. 42. Pettigrew RI, Dannels W, Galloway JR, et al. Quantitative phase-flow imaging in dogs by using standard sequences: comparison with in vivo-flow meter measurements. Am J Roentgenol. 1987;148: 411–414. 43. Kilner PJ, Manzara CC, Mohiaddin RH, et al. Magnetic resonance jet velocity mapping in mitral and aortic valve stenosis. Circulation. 1993;87:1239–1248. 44. Maier SE, Meier D, Boesiger P, Moser UT, Vieli A. Human abdominal aorta: comparative measurements of blood flow with MR imaging and multigated Doppler US. Radiology. 1989;171:487–492. 45. Van Rossum A, Sprenger KH, Peels FC, et al. In vivo validation of quantitative flow imaging in arteries and veins using magnetic resonance phase shift techniques. Eur Heart J. 1991;12:117–126. 46. Bogren HG, Klipstein RH, Firmin DN, et al. Quantitation of antegrade and retrograde blood flow in the human aorta by magnetic resonance. Am Heart J. 1989;117:1214–1222. 47. Bogren HG, Klipstein RH, Mohiaddin RH, et al. Pulmonary artery distensibility and blood flow patterns: a magnetic resonance study of normal subjects and of patients with pulmonary arterial hypertension. Am Heart J. 1989;118:990–999. 48. Sakuma H, Kawada N, Kubo H, et al. Effect of breath holding on blood flow measurement using fast velocity encoded cine MRI. Magn Reson Med. 2001;45:346–348. 49. Hofman MB, van Rossum AC, Sprenger M, Westerhof N. Assessment of flow in the right human coronary artery by magnetic resonance phase contrast velocity measurement: effects of cardiac and respiratory motion. Magn Reson Med. 1996;35:521–531. 50. Firmin DN, Klipstein RH, Hounsfield GL, et al. Echo-planar high-resolution flow velocity mapping. Mag Reson Med. 1989; 12:316–327. 51. Gatehouse PD, Firmin DN, Collins S, Longmore DB. Real time blood flow imaging by spiral scann phase velocity mapping. Magn Reson Med. 1994;31:504. 52. Mueller E, Laub G, Grauman R, Loeffler W. RACE—Real time Acquisition and Evaluation of pulsatile blood flow on a whole body MRI unit [Abstract]. In: Proceedings of the Seventh Annual Meeting of the International Society of Magnetic Resonance in Medicine. 1988;729. 53. Kose K. One shot velocity mapping using multiple spin-echo EPI and its application to turbulent flow. J Magn Reson. 1991;92:631–635.

59. Frayne R, Rutt BK. Measurement of fluid-shear rate by Fourierencoded velocity imaging. Magn Reson Med. 1995;34:378–387. 60. Urchuk SN, Fremes SE, Plewes DB. In vivo validation of MR pulse pressure measurement in an aortic flow model: preliminary results. Magn Reson Med. 1997;38:215–223. 61. Yang GZ, Kilner PJ, Wood NB, Underwood SR, Firmin DN. Computation of flow pressure fields from magnetic resonance velocity mapping. Magn Reson Med. 1996;36:520–526. 62. Firmin DN, Gatehouse PD, Yang GZ, Jhooti P, Keegan J. A 7-Dimensional echo-planar flow imaging technique using a novel k-space sampling scheme with velocity compensation [Abstract]. In: Proceedings of the Fifth Annual Meeting of the International Society of Magnetic Resonance in Medicine. 1997;118. 63. Canstein C, Cachot P, Faust A, et al. 3D MR flow analysis in realistic rapid-prototyping model systems of the thoracic aorta: comparison with in vivo data and computational fluid dynamics in identical vessel geometries. Magn Reson Med. 2008;59:535–546. 64. Bolger AF, Heiberg E, Karlsson M, et al. Transit of blood flow through the human left ventricle mapped by cardiovascular magnetic resonance. J Cardiovasc Magn Reson. 2007;9:741–747.

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54. Yang GZ, Gatehouse PD, Mohiaddin RH, et al. Zonal echo-planar flow imaging with respiratory monitoring [Abstract]. In: Proceedings of the Fifth Annual Meeting of the International Society of Magnetic Resonance in Medicine. 1997;1885. 55. Gatehouse PD, Link K, Bebbington MWP, et al. Pulse-wave and stenosis studies by cylinder excitation with Fourier velocity encoding [Abstract]. In: Proceedings of the Third Annual Meeting of the International Society of Magnetic Resonance and the Twelth Annual Meeting of the European Society for Magnetic Resonance in Medicine and Biology. 1995;318. 56. Napel S, Lee DH, Frayne R, Rutt BK. Visualizing three-dimensional flow with simulated streamlines and three-dimensional phase-contrast MR imaging. J Magn Reson Imaging. 1992;2:143–153. 57. Oshinski JN, Ku DN, Mukundan Jr S, et al. Determination of wall shear stress in the aorta with the use of MR phase velocity mapping. J Magn Reson Imaging. 1995;5:640–647. 58. Oyre S, Ringgaard S, Kozerke S, et al. Accurate noninvasive quantitation of blood flow, cross-sectional lumen vessel area and wall shear stress by three-dimensional paraboloid modeling of magnetic resonance imaging velocity data. J Am Coll Cardiol. 1998;32:128–134.

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CHAPTER 8

Special Considerations for Cardiovascular Magnetic Resonance: Safety, Electrocardiographic Setup, Monitoring, and Contraindications Jeroen J. Bax and Ernst E. van der Wall

During the last decade, cardiovascular magnetic resonance (CMR) has developed into an important diagnostic clinical tool in cardiology. Not only the anatomy of the heart but also its function, metabolism, and perfusion, as well as the coronary arteries, can be evaluated with CMR. CMR offers some special advantages over other diagnostic imaging methods. First, CMR does not use ionizing radiation. Second, the radiofrequency (RF) radiation penetrates bony structures without attenuation through relaxation parameters. Third, CMR gives additional diagnostic information about tissue characteristics. Finally, CMR provides threedimensional images or images of arbitrarily oriented slices. However, when performing CMR, particular precautions must be taken. Because CMR operates with high static and gradient magnetic fields, special safety regulations must be taken into account and certain contraindications must be considered. This chapter reviews the safety, electrocardiographic (ECG) setup, patient monitoring, and contraindications to CMR; in particular, the issue of pacemakers and implantable cardiac defibrillators (ICDs) is addressed.

SAFETY OF CARDIOVASCULAR MAGNETIC RESONANCE General Issues CMR generally takes longer than other diagnostic modalities (although the time is significantly shortened with real-time imaging1), and the confined space in which the patient is placed is rather narrow, which some patients find uncomfortable. During CMR, communication with the patient may be difficult because of interfering noise from the gradient coils. On the other hand, CMR is entirely noninvasive. Overall, the safety issues during CMR that may pose potential safety concerns2 can be summarized as follows:

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1. Biologic effects of the static magnetic field 2. Ferromagnetic attractive effects of the static magnetic field on certain devices 3. Potential effects on the relatively slowly time-varying magnetic field gradients 4. Effects of the rapidly varying RF magnetic fields, including RF power deposition concerns 5. Auditory considerations from noise from the gradients 6. Safety considerations concerning superconducting magnet systems 7. Psychological effects 8. Possible effects of the intravenous use of MR contrast agents 9. Patient safety during stress conditions The safety concerns inherent to these issues are discussed.

Biologic Effects Concerning the biologic effects of a static magnetic field, many structures in animals and humans are affected by magnetic fields. Many potential biologic effects and different magnitudes of magnetic fields have been examined, including the effect of the field on cardiac contractility and function. Gulch and colleagues concluded that static magnetic fields used in CMR do not constitute any hazard in terms of cardiac contractility.3 These magnetic fields do not increase ventricular vulnerability, as assessed by the repetitive response threshold and the ventricular fibrillation threshold.4 In one of the investigations, however, cardiac cycle length was shown to be altered.5 Numerous biologic effects on other systems have been investigated extensively, and it may be concluded that no deleterious biologic effects from static magnetic fields used in CMR have yet been established. However, as in all aspects of safety monitoring for patients, further research needs to be conducted.

The physical effect of the static magnetic field consists of a potential health hazard from the attractive effect on ferromagnetic objects. Ferromagnetic objects can be defined as those in which a strong intrinsic magnetic field can be induced when they are exposed to an external magnetic field. The existence of different kinds of scanners with different shielding makes the discussion about this topic even more crucial. When dealing with a static magnetic field, two types of physical concerns exist. First, there are concerns about forces exerted on ferromagnetic objects within, on, or distant from the patient. These forces result in rotational (torque) or translational (attractive) motion of the object. Within the human body, a ferromagnetic metallic structure may be sufficiently attracted, or have a sufficient amount of torque exerted, to create a hazardous situation. These factors should be carefully considered before subjecting a patient with a ferromagnetic implant or material to CMR, particularly if the device is located in a potentially dangerous area of the body, where movement or dislodgment of the device could injure the patient. Another potentially injurious effect is known as the projectile, or missile, effect. This refers to the fact that ferromagnetic objects have the potential to gain sufficient speed during attraction to the magnet that the accumulated kinetic energy could be injurious or even lethal if the object were to strike a patient. Numerous studies have been performed to assess the ferromagnetic qualities of various metallic implants and materials.6–10 The results indicate that patients with certain metallic implants or prostheses that are nonferromagnetic or are minimally deflected by static magnetic fields can safely undergo CMR. The literature on this topic has been extensively reviewed and compiled.8 However, there are common misconceptions about what types of objects are ferromagnetic. The most important misconception is that stainless steel is ferromagnetic, when it is not. Patients with stainless steel implants can therefore be imaged safely, except for a small number of well-described exceptions,

A

B

as discussed later. The implant will interfere locally with the images; for example, signal loss occurs around metallic prosthetic valves (Fig. 8-1) and sternal wires (Fig. 8-2) after bypass surgery, but this does not make the imaging hazardous. Non-stainless steel, which may be ferromagnetic, is not used for human implants, but is commonly used for oxygen cylinders, for example. Finally, batteries are typically attracted to the magnet, and this is one of the problems of imaging pacemakers. The second type of physical concern deals with magnetically sensitive equipment, the functioning of which may be adversely affected by the magnetic field. The most common of these is the cardiac pacemaker. Most pacemakers include a reed relay switch whereby the sensing mechanism can be bypassed and excitation in the asynchronous mode can occur. This switch is activated when a magnet of sufficient strength is held over the pacemaker.11 In addition, the function of cardiac pacemakers may be influenced by field strengths as low as 17 gauss.11 In practice, reed switch closure can be expected in all pacemakers placed in the bore of the scanner. Pacemaker and ICD function is considered again later in this chapter.

Effect of Rapidly Switched Magnetic Fields CMR exposes the patient to rapid variations of magnetic fields by the transient application of magnetic gradients during imaging. The effect of rapidly switched magnetic fields may be the induction of currents within the body or any other electrical conductor, according to Faraday’s law. The current is dependent on the time rate of change of the magnetic field (dB/dt), the cross-sectional area of the conducting tissue loop, and the conductivity of the tissue. Biologic effects of induced currents can be caused either by power deposition by the induced currents (thermal effects) or by direct effects of the current (nonthermal effects). Thermal effects as a result of switching gradients are not believed to be clinically significant.12–14 Possible

C

Figure 8-1 Cardiovascular magnetic resonance image of a patient with a Starr-Edwards prosthetic valve in the mitral position. A, Horizontal long axis. B, Vertical long axis. C, Basal short axis. The dark artifact is obvious on the images, but does not interfere with assessment of ventricular function. These images were acquired at 0.5 T with a gradient echo cine sequence and an echo time of 14 msec. Shorter echo time gradient echo sequences and spin echo sequences typically show less artifact. The Starr-Edwards valve causes the largest artifact because of the large amount of metal present in its construction. Other valves cause considerably less disturbance, especially the tissue valves. (Images courtesy of Dr. Dudley Pennell, Royal Brompton Hospital, London.) Cardiovascular Magnetic Resonance 101

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Ferromagnetism

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A

B

C

Figure 8-2 Cardiovascular magnetic resonance image of a patient with sternal wires after bypass grafting. The artifact is clearly seen (arrows) on the horizontal long axis (A) and short axis (B) gradient echo cine images, and to a much lesser extent, on the transaxial spin echo image (C). (Images courtesy of Dr. Dudley Pennell, Royal Brompton Hospital, London.)

nonthermal effects include stimulation of nerve or muscle cells. The threshold currents for nerve stimulation and ventricular fibrillation are known to be much higher than the estimated current densities induced under clinical CMR conditions. The echo planar imaging method, however, involves more rapidly changing magnetic field gradients, and peripheral muscle stimulation in humans has been reported.15 Such considerations have become more important as new technology has allowed the introduction of commercially available ultrafast gradient switching systems, and guidelines for maximum magnetic field variation are under development.

Radiofrequency Time Varying Field The transmitted RF time varying field induces electrical currents within the tissue of the patient. The majority of this power is transformed into heat within the patient’s tissue as a result of ohmic heating. The time varying magnetic gradients have the potential to cause either thermal or nonthermal biologic effects. The distinction between these two is a matter of frequency, waveform shape, and magnitude. The discussion about nonthermal effects from RF magnetic fields is controversial because of questions about the relationship between chronic exposure to electromagnetic fields over many years and the causation of cancer or developmental abnormalities. The most recent evidence suggests that proximity to power lines is not injurious.16 Of course, acute exposure of a patient to short-term RF fields for a diagnostic CMR examination is different from chronic exposure. The induced currents from RF magnetic fields are unable to cause nerve excitations. One of the difficulties faced by medicine is proving that a procedure is not injurious because of anecdotal case reports of adverse events and publication bias toward non-neutral reports.17 This issue is also faced by such well-established technology as ultrasound, where safety concerns have been raised over acoustic exposure.18 In contrast to the insignificant thermal effects caused by switched gradients, however, thermal effects as a result of 102 Cardiovascular Magnetic Resonance

RF pulses are of significant concern. The main biologic effects induced by RF fields are therefore related to the thermogenic qualities of the RF field. A general point of discussion is the appropriate safety regulations for levels of magnetic field strength in CMR imaging. Application of the fundamental law of electrostimulation is well established, both on theoretical and experimental grounds. Application of this law, in combination with Maxwell’s law, yields an equation called the fundamental law of magnetostimulation, which has the hyperbolic form of a strengthduration curve and allows an estimation of the lowest possible value of the magnetic flux density capable of stimulating nerves and muscles. Calculations have shown that the threshold for heart excitation is more than 200 times higher than for nerve and muscle stimulation, depending on pulse duration.19 However, in clinical practice, some precautions are necessary. First and most importantly, the specific absorption rate of the imaging sequence being operated is monitored by the scanner software and must be kept below limits set by such bodies as the U.S. Food and Drug Administration (FDA). Second, circumstances that could enhance the possibility of heating injury should be avoided. This includes ensuring the prevention of loops that could act as aerials within the scanner and enhance the heating effect locally. Therefore, patients should not be allowed to cross their legs (loop via the pelvis) or clasp their hands (loop via the shoulder and upper chest). The simple use of pillows prevents such problems. Other possible loops include the ECG leads, which should always be run out of the scanner parallel to the main field, and not looped across the chest. Finally, pacemaker and ICD leads make excellent aerials. In most cases, MR is contraindicated in patients with pacemakers and ICDs, although recent developments are promising (see the discussion of pacemakers and ICDs). The pacemaker lead can heat significantly during CMR and become a potential hazard (discussed later). Another consideration in patients after cardiac surgery is the effect of retained epicardial pacemaker leads. These leads can be left in place after surgery, and they may therefore act as an antenna during CMR. Recent studies have suggested that such short retained epicardial wires do not pose a significant problem.20,21

Auditory Considerations During CMR, the gradient coils and adjacent conductors produce a repetitive sound because they act essentially as loudspeakers, with current being driven through them while they are in a magnetic field. Auditory considerations should therefore be taken into account when imaging a patient. The amplitude of this noise depends on factors such as the physical configuration of the magnet, the pulse sequence type, timing specifications of the pulse sequence, and the amount of current passing through these coils.24 In general, the amplitude of the generated noise from the clinical CMR scanners remains between 65 and 95 dB. However, there have been reported instances of temporary hearing impairment as a result of CMR. Magnet-safe headphones or wax earplugs are readily available and have been shown to prevent hearing loss,25 and these are in common use. Systems combining sound attenuation with the facility to play music of the patient’s choice are also available. Research on the reduction of noise in MR scanners is ongoing, and the use of anti-noise is one area of interest.26

Superconducting System Issues Most superconducting CMR scanner systems use liquid helium. The helium maintains the magnet coils in their superconducting state. Helium achieves the gaseous state at approximately 269 C (4 K). If for any reason the temperature within the cryostat rises, or in a system quench, the helium will enter the gaseous state. This means a marked increase in volume and thereby pressure within the cryostat. A pressure-sensitive valve is designed to give way to the gaseous helium, which is always vented outside the CMR scanner room. However, it is possible that some helium gas is released into the imaging room should the system not work perfectly. Asphyxia and frostbite are potential hazards if a patient is exposed to helium vapor for a prolonged time, although there are no reports of such an occurrence in the medical community. For older scanners that still use a buffer of liquid nitrogen within the system (boils at 77 K), an oxygen monitor is recommended in the scanner room. Cryogenic dewars should be stored away from the scanner and in well-ventilated areas.

Psychological Effects Claustrophobia or other psychological problems may be encountered in up to 10% of patients undergoing CMR,27 although on average, the incidence is closer to 2% to 4%, and this can be reduced further to a small number of intractably anxious patients by the use of explanation, reassurance, and where necessary, light sedation with, for example, 2 to 5 mg intravenous (IV) diazepam.28 In addition, the development of shorter magnets

as well as open designs is proving to be helpful. Such problems are related to a variety of factors, including the restrictive dimensions of the scanner, the duration of the examination, the noise, and the ambient conditions within the magnet bore.29 Fortunately, adverse psychological effects with CMR are usually transient. In a study reported by Weinreb and colleagues,30 based on the experience of 450 patients undergoing CMR and computed tomography examinations, it was clearly shown that patients often prefer the CMR study, although CMR imaging took longer. Furthermore, the patient is placed into a confined space and there are difficulties in communicating with the patient during CMR scanning because of the noise from the gradient coils and the necessity of eliminating all extraneous RF sources from the examination room. To a certain extent, this can be avoided when the patient assumes a prone position in the scanner,31 facilitating communication with the outside surroundings. Simple maneuvers, such as using mirrors, also help in allowing the patient a clear view of the scanning room. Allowing the anxious patient to visit the scanner before the appointment gives the patient an opportunity to become familiar with the facility and staff.

SAFETY CONSIDERATIONS ASSOCIATED WITH GADOLINIUM-BASED CONTRAST AGENTS The safety profile of the contrast agents containing gadolinium currently on the market is extremely good. Gadopentetate dimeglumine (Gd-DTPA) is the best established, and its safety profile is well documented,32,33 but similar safety results have been shown with the other commercially available agents. The median lethal dose of Gd-DTPA is roughly 10 mmol/kg, which is 100 times the diagnostic dose and shows the wide safety margin that the contrast agent enjoys. Patient tolerance of this drug is also high, and the prevalence of adverse reactions is approximately 2%. Among the reactions related to the IV administration of this drug are headache, nausea, vomiting, local burning or cool sensation, and hives. There have been reported incidents of anaphylactoid reactions associated with IV injection,34 although the frequency of this appears to be approximately 1 per 100,000 doses. The safety margins with these agents appear to be considerably better than with X-ray contrast agents, although issues of nephrogenic septemic fibrosis are a concern in patients with severely impaired renal function. (See Chapter 6.) For CMR, these agents are used to increase contrast between blood and soft tissue for cine imaging for functional studies or angiography, to enhance cardiac tumors and cysts, to assess myocardial perfusion, and to examine for myocardial infiltration. In summary, FDA-registered gadolinium complexes, such as Gd-DTPA, can be safely used in patients with cardiac disorders. Multiple new CMR contrast agents are being developed and investigated. These are mainly gadolinium complexes, sometimes with novel binding molecules for special actions, but in addition, iron-based compounds are being developed. Some of these agents are retained in the Cardiovascular Magnetic Resonance 103

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Finally, the use of ECG electrodes, which are essential for cardiac gating, must be considered. Metallic ECG electrodes may cause burns during CMR,22,23 but this risk can be reduced with the use of carbon fiber electrodes, and these have now become standard.

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

vascular system and do not leak into the extravascular space. This suggests that they may have clinical utility for angiography, possibly in the coronaries, and for functional imaging.

PATIENT SAFETY DURING STRESS CONDITIONS A concern with CMR stress studies has been the ability to handle emergency situations. Patient monitoring during stress conditions is a critical issue because myocardial ischemia can be provoked in patients with coronary artery disease. Commercial equipment is available for noninvasive monitoring of blood pressure, heart rate, oxygen saturation, and other vital parameters in CMR scanners. The most crucial difference compared with conventional exercise testing outside a magnetic field is the lack of a diagnostic ECG, in particular, at high levels of stress, precluding the proper assessment of stress-induced ST-segment changes. This holds for both conventional exercise using a specially adapted bicycle ergometer and pharmacologically induced stress. Under these circumstances, only heart rate can be monitored reliably. When performing pharmacologic stress CMR (e.g., with dipyridamole, adenosine, or dobutamine), an experienced physician should be present during the examination, and appropriate treatments for complications should be in direct proximity. Dipyridamole (half-life, 30 minutes) and adenosine (half-life, 10 seconds) are both vasodilators. Both agents have similar side effects, such as bronchospasm, hypotension, arrhythmias, and bradycardia. In particular during adenosine infusion, atrioventricular heart block may develop in a small percentage of patients (0.7% to 2.8%), although this is usually asymptomatic and self-limited. When patients are symptomatic, the short physical half-life of adenosine means that heart rhythm and symptoms can be restored very quickly by halting the infusion. As a suitable antagonist to both dipyridamole and adenosine, aminophylline may be given slowly at an initial dose of 50 mg IV up to a maximum dose of 250 mg if necessary. In the case of persisting advanced heart block, 0.5 mg atropine IV should be administered up to a total dose of 3 mg. Dipyridamole and adenosine should not be administered to patients with asthma. Dobutamine (half-life, 2 minutes) is a beta-agonist leading to an increase in cardiac inotropy (contractility) and chronotropy (heart rate). Common side effects are cardiac pounding and palpitations, and less commonly, arrhythmias, such as supraventricular tachycardia and (nonsustained) ventricular tachycardia, are seen. Dobutamine can be safely administered to patients with asthma.35 The actions of dobutamine can be counteracted by IV administration of a short-acting beta-blocking agent, such as esmolol. In the case of cardiac arrest or ventricular fibrillation, the recommendations should be followed according to published guidelines, such as those proposed by the European Resuscitation Council.36 In every CMR facility, an alarm system and a written flow chart should be visually available with the necessary instructions in case of emergency. It is necessary to be able to remove the patient from the examination room quickly (preferably within 20 seconds) to an area where emergency treatment can be performed safely, away from 104 Cardiovascular Magnetic Resonance

the hazards of the magnetic field. A nonferromagnetic stretcher stored in the scanner room or a detachable scanner table is ideal. A cardiac arrest trolley must be maintained in close proximity to the scanner room, and all staff should undergo regular training in cardiopulmonary resuscitation techniques. Regular checks should be made of both the resuscitation equipment and the alarm system.

PATIENT MONITORING AND ELECTROCARDIOGRAPHIC SETUP Patient monitoring during CMR poses problems that will not be familiar to users of other technologies, such as echocardiography. Ferrous metal, which is present in most monitoring equipment, can distort the magnetic field, and such an item has the potential to become a projectile. In addition, monitoring wires that are attached to the patient, leaving the scanner, and passing to another room may act as an antenna for stray RF signals. Electrical equipment in the scanner room also can act as a source of RF noise. All of these disturbances may result in image degradation. Therefore, specific solutions to these problems have been designed. Commercially available CMR-compatible monitoring equipment, including that used to measure ECG, blood pressure, and chest wall movements, as well as for general anesthesia, has been tested in several studies.37,38 Satisfactory monitoring can be obtained and images obtained during its use can be evaluated adequately.39 For some monitoring, simple solutions work, such as that reported by Roth and associates,39 who measured arterial blood pressure outside the CMR scanner by lengthening the rubber tubing connected to a blood pressure cuff. The newest monitoring equipment eliminates the need for wires and tubes to leave the scanning room by using a microwave transmitter communicating with a slave display unit in the operating room. The CMR procedure depends on a high-quality ECG signal for routine imaging, and each manufacturer has developed its own solution to the problems posed. Fiberoptic transmission of ECG signals for gating is now commonplace, and this significantly reduces RF pulse artifacts in the ECG. Felblinger and coworkers showed that this type of system could yield signals almost free from interference,40 during both conventional41 and high gradient activity sequences, such as during echo planar imaging.42 From this signal, the authors also developed a method for respiration monitoring during MR sequences. Third-party ECG solutions are also now being incorporated into the latest generation of scanners, and these come with specific ECG recommendations for lead placement. Carbon fiber electrodes are required to eliminate the risk of burning that has been reported with standard metallic ECG electrodes.22 Typical lead placement is the result of compromise. A better signal results from widely spaced electrodes, but this results in more artifact from the gradients. In general, therefore, the leads are kept relatively close together, and on the left side, which reduces the magnetohydrodynamic effect (the effect of systolic aortic flow causing surface potentials on the ECG that distort the ST segment; Fig. 8-3). A typical

Inside magnet

Figure 8-3 The magnetohydrodynamic effect. The top trace was recorded in a patient with atrial fibrillation outside of the magnet, and the bottom trace, inside. Note the distortion of the ST segment (black arrows) caused by added potentials arising from systolic flow in the aorta. (Courtesy of Dr. Dudley Pennell, Royal Brompton Hospital, London.)

lead placement that is commonly adopted is shown in Figure 8-4. Some centers have found ECG gating using electrodes on the back to be successful, but this is not widely used. The ECG leads should not be allowed to form loops, which could present a burning hazard, and they should be braided together and brought out of the magnet while aligned parallel to the bore to reduce electrical interference. Keeping the electrical cables short is helpful, and fiberoptic conversion modules are therefore often very close to the patient’s chest. Switching between the ECG traces sometimes allows flexibility to reduce gating errors from tall T-waves or electrical interference. One thing is certain, however, and that is that time spent ensuring that the ECG is stable and working correctly at the start of the scan is time very well spent. An alternative technique to routine surface ECG recording has been described by Fischer and colleagues using vectorcardiography, and this may prove to be a useful advance and has been widely implemented.43 The system examines the three-dimensional orientation of the ECG V2

V3 V4

V1 V5 V3R

V4R V6

Figure 8-4 Typical electrode placement for cardiovascular magnetic resonance. The conventional chest leads (solid circles) are shown for comparison, and the positions of the four carbon electrocardiographic electrodes over the left chest are indicated (white circles). (Courtesy of Dr. Dudley Pennell, Royal Brompton Hospital, London.)

CONTRAINDICATIONS TO CARDIOVASCULAR MAGNETIC RESONANCE In general, there are potential hazards and artifacts of ferromagnetic and nonferromagnetic materials in CMR, for example, with neurosurgical clips and ocular implants.46–48 The (relative) contraindications of these materials for CMR are dependent on factors such as the degree of ferromagnetism, the geometry of the material, the gradient (for force), and the field strength (for torque) of the imaging magnet and many other factors.

General Contraindications to Cardiovascular Magnetic Resonance There are a number of circumstances in which CMR is better avoided. This is because of reports of death or harm that have occurred. Prominent among these are patients with Cardiovascular Magnetic Resonance 105

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Outside magnet fast atrial fibrillation

signal and uses the calculated vector of the QRS complex as a filter mechanism to ignore electrical signals, which are of similar timing and similar magnitude in the cardiac cycle, but a different vector. The system as reported identified the QRS complex correctly in 100% of cases, with 0.2% false-positive findings. In a subsequent study in normal individuals and patients with supraventricular extrasystoles, the same authors showed that vector cardiography-based triggering provided nearly 100% triggering performance during CMR examinations.44 This system represents a significant improvement for CMR in stabilizing this important gating signal. Should interpretation prove impossible for technical reasons, a standard vascular Doppler can be used to monitor heart rate during CMR. The Doppler and telemetric ECG do not contain enough ferromagnetic material to cause visible image degradation. Jorgensen and associates evaluated whether patients could be monitored during CMR with 1.5 Tesla (T) machines in a manner that complies with monitoring standards.45 The high magnetic field can interfere with normal functioning of equipment, not only monitoring equipment, but also smaller items, such as infusion pumps used for stress testing. In general, the influence of the CMR scanner on nearby equipment depends on several factors, such as the strength of the CMR magnetic field, the proximity of the equipment to the scanner, the amount of ferromagnetic material in the equipment, and the design of the electrical circuitry. Finally, simple devices, such as closed-circuit television and a two-way intercommunication system, also aid in monitoring by allowing a constant view of the patient and easy communication if the patient is in discomfort. However, the latter may be impaired during imaging because of the noise of the gradients. Because sequences are now commonly being reduced in duration to a breath hold, however, this limitation is becoming much less important.

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cerebral aneurysm clips, which have become dislodged by MR, causing fatal cerebral hemorrhage. Modern clips are not ferromagnetic and are safe, but the problem is establishing the type of clip with certainty before performing CMR. In general, therefore, CMR should be avoided in these patients unless written information is available and appropriate advice from a neurosurgical center is obtained. There have been cases of bleeding in the eye in patients with previous injury with metallic shards (usually metal workers), and again a history of this should be sought. A skull X-ray can be helpful in cases of doubt. Electronic implants are the other major problem. These may dysfunction or may be damaged by CMR, which is therefore best avoided. This applies to cochlear implants, neurostimulation systems, pacemakers/ICDs, and a number of other modern implants. In general, the main rule to observe is to determine the riskbenefit ratio for the proposed procedure. If the information can only be obtained by CMR, and if it is very important, then the risk-benefit ratio may be positive for some patients (as has been shown in some patients with pacemakers). In many clinical circumstances, however, the information required can be obtained by other means. Reference texts on the safety of specific devices are available. For CMR, a number of other specific device issues require mention. Swan-Ganz catheters and temporary pacing wires in general preclude CMR, but sternal wires and vascular clips on grafts do not present safety problems, although localized artifacts occur on the images. Three specific areas are discussed in more detail, namely, stents, valve prostheses, and pacemakers/ICDs.

Stents The first intracoronary stents were implanted in human coronary arteries in 1986 by Sigwart and colleagues.49 Recently, the indications for intracoronary stents have expanded and the use of stents has grown dramatically. One of the reasons for this development is the reduced restenosis rate compared with conventional balloon angioplasty.50–53 Because of the increased indications for intracoronary stents, the population of patients with these stents in situ who need to undergo CMR for various diagnostic reasons is rapidly expanding. Coronary stents are metallic structures that remain in situ for life. As a result, there have been concerns about stent dislodgment during CMR as well as concern about possible heating effects. Several factors determine the risk of placing these materials in a magnetic field. These factors include the ferromagnetism of the material, the strength of the static and gradient magnetic fields, the metal mass, and the geometry of the material.10,54 The most commonly used stents are made of stainless steel or tantalum, which is not ferromagnetic.55 In vitro experiments performed by Scott and Pettigrew56 evaluated and quantified the influence of magnetic fields used in clinical MR scanners on widely used coronary stents. The authors used a 1.5 T magnet and the results did not show any significant deflection of the stents used. In vivo studies have been carried out in dogs with tantalum stents in the aorta using a 1.5 T scanner.55 This study examined the CMR compatibility of tantalum stents and evaluated the feasibility of using vascular MR to evaluate the patency of the stented vessel in vivo. In this study, the animals underwent CMR repeatedly in the 106 Cardiovascular Magnetic Resonance

first 8 weeks post-implantation. No evidence was found of any stent migration after CMR immediately post-implantation in the animals. Angiographic evaluation showed no interval development of luminal narrowing or thrombus in the region of the vascular stent. CMR artifacts produced by the stent were increased with longer echo times. Further experiments by Friedrich and colleagues have shown that the common stents do not heat up during CMR.57 Large studies have subsequently confirmed the safety of stent imaging with CMR. Gerber and coworkers evaluated 111 patients with MR within 8 weeks of stent implantation (median, 8 days; range, 0 to 54 days); during follow-up, four noncardiac deaths occurred as well as three revascularization procedures.58 These events were not related to the CMR procedure. This finding agrees with widespread clinical experience in which stent imaging has been performed in the day after implantation, or soon afterward.59,60 Moreover, the feasibility of assessing changes in coronary flow reserve after percutaneous intervention with stenting was shown recently.61 Another issue regarding CMR procedures in patients with coronary stents is the image distortion caused by the ferromagnetic properties of the stent. In general, the greater the ferromagnetism of a metallic implant, the greater its magnetic susceptibility artifact. However, recent studies have shown that CMR angiography could reliably be used for noninvasive imaging and evaluation of blood flow after stent placement.62,63

Valvular Prosthesis Heart valve prostheses are all safe for CMR.64 These are the most recent recommendations, and they supersede those that suggested that pre-model 6000 series Starr-Edwards valves might cause problems during MR. This conclusion is backed up by numerous data. Several studies in 1.5 to 2.35 T static magnetic fields have shown that for a number of prosthetic valves there is no hazardous deflection during exposure of the magnetic field.65,66 Heating of small metallic implants was tested in a study reported by Davis and associates.46 The authors found no significant increase in temperature in steel and copper clips that were exposed to changing magnetic fields 6.4 times as strong as those expected to be used in the CMR scanner. As a result, there is no contraindication to CMR in patients with the currently used prosthetic valves.67–69 However, prosthetic material may lead to artifacts on CMR images. To evaluate the influence of prosthetic valves on the interpretation of CMR images and the capability of functional valve analysis, in a group of 89 patients and 100 heart valve prostheses, Bachmann and colleagues showed convincingly that all patients could be imaged with CMR without any risk and that prosthesis-induced artifacts did not interfere with image interpretation.68 In particular, physiologic valvular regurgitation was easy to differentiate from pathologic or transvalvular regurgitation. DiCesare and coworkers studied 14 patients who were surgically treated with nine biologic and seven mechanical aortic and mitral valves.69 Three classes of artifacts were distinguished and graded as minimal, moderate, or significant. The biologic valves produced minimal artifacts and the mechanical valves showed only moderate artifacts. In all 16 prosthetic valves, CMR allowed adequate semi-quantitative analysis of flow over the valve.

The number of patients with cardiac pacemakers and ICDs has increased exponentially in recent years, with 2.4 million patients in the United States having a permanent pacemaker in 2002 and more than 370,000 having an implant in 2003. Many centers consider MR absolutely contraindicated in these patients, and none of the pacemakers or ICDs has been approved by the FDA for CMR examinations. A total of 10 deaths have been attributed to MR examinations in patients with pacemakers.69 However, as is often the case, such a dogmatic approach is not entirely correct. First, the reported fatalities were poorly characterized and ECGs are not available; moreover, CMR-related deaths during physician-supervised examinations have not been reported. Second, many patients with pacemakers have safely undergone CMR.70 Therefore, by all normal rules of semantics, the presence of a pacemaker is not an absolute contraindication. However, the presence of a pacemaker is a strong relative contraindication to scanning, and such procedures still require a great deal more research. They should be undertaken only after careful evaluation of the risk-benefit ratio to the patient, and should be performed only in expert cardiovascular centers. The issues surrounding CMR of pacemakers are complex. In general, three hazardous MR interactions with pacemakers and ICDs should be considered. First, the static magnetic fields exert mechanical forces on the ferromagnetic components of the devices, including the pacemaker and shock leads; the static magnetic fields also can induce asynchronous pacing. Second, a pulsed RF field may result in oversensing or may induce currents in the leads, resulting in thermal damage at the tissue-electrode interface. Third, the gradient magnetic fields may induce voltages on leads, resulting in over- and undersensing. Combined fields may also result in device damage and failure. Various generations of pacemakers and ICDs are currently implanted in patients, and studies of the effects of MR on cardiac pacemakers, in both animal models and patients, have been reported with varying results.71 In an in vitro study by Lauck and associates, it was concluded that no disturbances arise when the systems are tested in the asynchronous mode at 0.5 T CMR under standard examination conditions with ECG-triggered imaging.72 In a study by Achenbach and colleagues, the effect of MR on pacemakers and electrodes was investigated with phantoms.73 Twenty-five electrodes were exposed in a 1.5 T scanner, with continuous registration of temperature at the tip of the electrode. Eleven pacemakers were exposed to MR and the pacemaker output was monitored. Temperature increases of up to 63.1 C were observed. Furthermore, no pacemaker malfunctions were observed in the asynchronous mode. Inhibition or rapid pacing was observed during spin echo CMR if the pacemakers were set to VVI or DDD mode. During scanning with gradient echo CMR, pacemaker function was not impaired. Conversely, Erlebacher and associates reported significant adverse effects of CMR on DDD pacemakers.74 All units paced normally in the static magnetic field, but during MR, all units malfunctioned. All malfunctions were the result of RF interference, whereas gradient and static

magnetic fields had no effects. Thus, despite magnetic field strengths adequate to close pacemaker reed switches, RF interference during MR may cause total inhibition of atrial and ventricular output in DDD pacemakers, and may also lead to dangerous atrial pacing at high rates. Gambel and coworkers studied the effect of MR in five patients with permanent cardiac pacemakers, one of whom was pacemaker dependent.75 A variety of pacing configurations was studied, but none of the patients experienced any torque or heat sensation. Four non-pacemaker-dependent patients remained in sinus rhythm throughout the MR procedure. During and after CMR, all pacemakers continued to function normally, except for one transient pause of 2 seconds toward the end of the procedure. The authors concluded that, when appropriate strategies are used, CMR may be performed with an acceptable risk-benefit ratio for the patient.75 Pennell reported four patients with pacemakers and urgent clinical problems who underwent CMR. No significant problems occurred in three patients (Fig. 8-5). However, CMR was not attempted in one patient because the pacemaker switched into full-output mode near the magnet.76 This study was unusual because all of the patients with pacemakers underwent heart scans,77 whereas most studies report the outcomes of noncardiac scans. After these preliminary data in small groups of patients, a large prospective study was performed in 54 patients undergoing a total of 62 examinations using 1.5 T CMR scanners.78 During the MR examinations, only two patients experienced mild, clinically insignificant symptoms. After the MR examinations, no loss of capture, changes in lead impedance, or battery voltages were noted. A total of 107 leads were evaluated, including 48 atrial and 59 ventricular leads. A significant change in pacing threshold was noted in 10 (9.4%) leads, and only 2 (1.9%) needed a change in programmed output. Threshold changes were not related to cardiac chamber or anatomic location. An important issue is that mid- and long-term follow-up of patients was not obtained and effects may occur at a later stage; this is of particular concern in patients with an increase in pacing thresholds. Also, the effects of CMR-related heating were not evaluated in this study. In vivo heating of pacemaker leads was evaluated recently in nine pigs undergoing CMR.79 Significant temperature increases were noted, with significant changes in impedance and minor changes in the stimulation threshold. However, histologic changes were not observed. Thus, despite the observed increases in temperature, significant tissue damage was not reported. Finally, pacemaker leads may serve as an antenna, which could result in pacing the heart during scanning at the frequency of the applied imaging pulses. This could potentially lead to hypotension and dysrhythmias. This effect was shown in experiments and in several patients while positioned in a CMR system,80–82 but this effect must be separated from excitation caused by the pulse generator. Although recent data have reported minimal effects of CMR on pacemakers, there is too little experience and there are too many types of pacemakers to allow general statements to be made about their suitability for CMR. Preferably, pacemaker-dependent patients should not be scanned, but if needed, it has been suggested to program the pacemakers in the asynchronous mode (VOO or DOO).83 Cardiovascular Magnetic Resonance 107

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Pacemakers and Implantable Cardioverter Defibrillators

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

Diastole

Patients with pacemakers should not undergo scanning unless special circumstances arise, and then only in centers with special expertise and cardiologic backup. Pacemakerdependent patients should not be scanned. Information on patients with ICDs undergoing MR examinations is scarce. In a study of the safety of devices in an animal model, Roguin and colleagues84 included 17 different ICDs and reported that 9 (53%) ICDs had interrogation or battery problems after the scan. Incidental reports showed that patients with ICDs could safely undergo CMR examination.85 Naehle and colleagues reported battery voltage declines after MR scanning86 in 18 patients with ICDs. With the rapid growth in ICD implantations, there is a clear need for more studies, in patients particularly, on the safety of ICDs in CMR.

CONCLUSION In general, CMR is safe and no long-term ill effects have been reported. Very rapidly changing gradients may induce nerve excitations that may result in muscle twitching. However, clinical scanners operate below the threshold for such effects. The threshold of excitation of the myocardium is approximately 200 times higher than that for other muscles, and so the heart will not be stimulated

108 Cardiovascular Magnetic Resonance

Systole

Figure 8-5 Images from a young patient with a pacemaker. The patient had numerous ventricular fibrillation arrests. Cardiovascular magnetic resonance imaging suggested sarcoidosis (diagnostic images not shown). The pacemaker artifact can be seen in the transaxial gradient echo cine images in the top row and in the right ventricular outflow tract cine images in the bottom row (straight arrows). Curved arrows show the artifact from the pacing lead in the apex of the right ventricle. Imaging was performed at 0.5 T with echo time of 14 msec. (Images courtesy of Dr. Dudley Pennell.)76

by the rapidly changing gradients. Most metallic implants, such as intracoronary stents, prosthetic valves, and sternal sutures, present no hazard because most materials used are nonferromagnetic. In general, patients with pacemakers and ICDs should not undergo CMR because of the unquantifiable risks. According to the recently published Council Panel Report on Clinical Indications for CMR, the following statements were made about the safety of CMR: CMR is safe and no long-term ill effects have been demonstrated. Claustrophobia occurs in about 2% of patients, but mild anxiolysis is often effective. One of the most important safety issues for CMR is the prevention of introduction into the scanner area of ferromagnetic objects which can become projectiles. Metallic implants such as hip prostheses, prosthetic heart valves, coronary stents and sternal sutures present no hazard since the materials used are not ferromagnetic (although an artefact local to the implant may be present). Care is required in patients with many cerebrovascular clips however, and specialist advice is needed for such patients. Patients with pacemakers, ICDs, retained permanent pacemaker leads and other electronic implants are not scanned, although some reports of success do exist, and there is progress towards manufacture of CMR-compatible devices.87

1. Yang PC, Kerr AB, Liu AC, et al. New real time interactive magnetic resonance imaging system complements echocardiography. J Am Coll Cardiol. 1998;32:2049–2056. 2. Kanal E, Shellock FG, Talagala L. Safety considerations in MR imaging. Radiology. 1990;176:593–606. 3. Gulch RW, Lutz O. Influence of strong static magnetic fields on heart muscle contraction. Phys Med Biol. 1986;31:763–769. 4. Doherty JU, Whitman GJR, Robinson MD, et al. Changes in cardiac excitability and vulnerability in NMR fields. Invest Radiol. 1985;20:129–135. 5. Jehenson P, Duboc D, Lavergne T, et al. Change in human cardiac rhythm induced by a 2-T static magnetic field. Radiology. 1988;166:227–230. 6. New PFJ, Rosen BR, Brady TJ, et al. Potential hazards and artifacts of ferromagnetic and nonferromagnetic surgical and dental materials and devices in nuclear magnetic resonance imaging. Radiology. 1983; 147:139–148. 7. Shellock FG, Crues JV. High-field strength MR imaging and metallic biomedical implants: an ex vivo evaluation of deflection forces. AJR. 1988;151:389–392. 8. Shellock FG. MR imaging of metallic implants and materials: a compilation of the literature. AJR. 1988;151:811–814. 9. Randall PA, Kohman LJ, Scalzetti EM, et al. Magnetic resonance imaging of prosthetic cardiac valves in vitro and in vivo. Am J Cardiol. 1988;62:973–976. 10. Teitelbaum GP, Bradley WG, Klein BD. MR imaging artefacts, ferromagnetism, and magnetic torque of intravascular filters, coils and stents. Radiology. 1988;166:657–664. 11. Pavlicek W, Geisinger M, Castle L, et al. The effects of nuclear magnetic resonance on patients with cardiac pacemakers. Radiology. 1983;147:149–153. 12. Bottomley PA, Edelstein WA. Power deposition in whole body NMR imaging. Med Phys. 1981;8:510–512. 13. Safety aspects of magnetic resonance imaging. In: Schaefer DJ, Wehrli FW, Shaw D, Kneeland BJ, eds. Biomedical Magnetic Resonance Imaging: Principles, Methodology, and Applications. New York: VCH; 1988:553. 14. Extremely low frequency (ELF) magnetic fields. In: Persson BRR, Stahlberg F, eds. Health and Safety of Clinical NMR Examinations. Boca Raton: CRC; 1989:49. 15. Cohen M, Weisskoff R, Rzedzian RR, Cantor HL. Sensory stimulation by time-varying magnetic fields. Magn Reson Med. 1990;14:409–414. 16. UK childhood cancer study investigators: exposure to powerfrequency magnetic fields and the risk of childhood cancer. Lancet. 1999;354:1925–1931. 17. Easterbrook PJ, Berlin JA, Gopalan R, Matthews DR. Publication bias in clinical research. Lancet. 1991;337:867–872. 18. Newnham JP, Evans SF, Michael CA, Stanley FJ, Landau LI. Effects of frequent ultrasound during pregnancy: a randomised controlled trial. Lancet. 1993;342:887–891. 19. Irnich W, Schmitt F. Magnetostimulation in MRI. Magn Reson Med. 1995;33:619–623. 20. Hartnell GG, Spence L, Hughes LA, Cohen MC, Saouaf R, Buff B. Safety of MR imaging in patients who have retained metallic materials after cardiac surgery. Am J Roentgenol. 1997;168:1157–1159. 21. Murphy KJ, Cohan RH, Ellis JH. MR Imaging in patients with epicardial pacemaker wires. AJR. 1999;172:727–728. 22. Boutin RD, Briggs JE, Williamson MR. Injuries associated with MR imaging: survey of safety records and methods used to screen patients for metallic foreign bodies before imaging. AJR. 1994;162:189–194. 23. Jones S, Jaffe W, Alvi R. Burns associated with electrocardiographic monitoring during magnetic resonance imaging. Burns. 1996; 22:420–421. 24. Hurwitz R, Lane SR, Bell RA, Brant-Zawadzki MN. Acoustic analysis of gradient-coil noise in MR imaging. Radiology. 1989;173:545–548. 25. Brummett RE, Talbot JM, Charuhas P. Potential hearing loss resulting from MR imaging. Radiology. 1988;169:539–540. 26. McJury M, Stewart RW, Crawford D, Toma E. The use of active noise control (ANC) to reduce acoustic noise generated during MRI scanning: some initial results. Magn Reson Imaging. 1997;15:319–322. 27. Flaherty JA, Hoskinson K. Emotional distress during magnetic resonance imaging. N Eng J Med. 1989;320:467–468. 28. Francis JM, Pennell DJ. The treatment of claustrophobia during cardiovascular magnetic resonance: use and effectiveness of mild sedation. J Cardiovasc Magn Reson. 2000;2:139–141.

29. Quirk ME, Letendre AJ, Ciottone RA, Lingley JF. Anxiety in patients undergoing MR imaging. Radiology. 1989;170:463–466. 30. Weinreb JC, Maravilla KR, Peshock R, Payne J. Magnetic resonance imaging: improving patient tolerance. AJR. 1984;143:1285–1287. 31. Hricak H, Amparo EG. Body MRI: alleviation of claustrophobia by prone positioning. Radiology. 1984;152:819. 32. Goldstein HA, Kashanian FK, Blumetti RF, et al. Safety assessment of gadopentetate dimeglumine in U.S. clinical trials. Radiology. 1990; 174:17–23. 33. Sullivan ME, Goldstein HA, Sansone KJ, et al. Hemodynamic effects of Gd-DTPA administered via rapid bolus or slow infusion: a study in dogs. AJNR. 1990;11:537–540. 34. Weiss KL. Severe anaphylactoid reaction after IV Gd-DTPA. Magn Reson Imaging. 1990;8:817–818. 35. Pennell DJ, Underwood SR, Ell PJ. Safety of dobutamine stress for thallium myocardial perfusion tomography in patients with asthma. Am J Cardiol. 1993;71:1346–1350. 36. Guidelines for advanced life support. A statement by the advanced life support working party of the European Resuscitation Council. Resuscitation. 1992;24:111–121. 37. Sellden H, de Chateau P, Ekman G, et al. Circulatory monitoring of children during anaesthesia in low-field magnetic resonance imaging. Acta Anaesthesiol Scand. 1990;34:41–43. 38. Lindberg LG, Ugnell H, Oberg PA. Monitoring of respiratory and heart rates using a fibre-optic sensor. Med Biol Eng Comput. 1992;30:533–537. 39. Roth JL, Nugent M, Gray JE, et al. Patient monitoring during magnetic resonance imaging. Anesthesiology. 1985;62:80–83. 40. Felblinger J, Boesch C. Amplitude demodulation of the electrocardiogram signal (ECG) for respiration monitoring and compensation during MR examinations. Magn Reson Med. 1997;38:129–136. 41. Felblinger J, Lehmann C, Boesch C. Electrocardiogram acquisition during MR examinations for patient monitoring and sequence triggering. Magn Reson Med. 1994;32:523–529. 42. Felblinger J, Debatin JF, Boesch C, et al. Synchronization device for electrocardiography-gated echo-planar imaging. Radiology. 1995;197:311–313. 43. Fischer SE, Wickline SA, Lorenz CH. Novel real-time R-wave detection algorithm based on the vectorcardiogram for accurate gate magnetic resonance acquisitions. Magn Reson Med. 1999;42:361–370. 44. Chia JM, Fischer SE, Wickline SA, et al. Performance of QRS detection for cardiac magnetic resonance imaging with a novel vectorcardiographic triggering method. J Magn Reson Imaging. 2000;12:678–688. 45. Jorgensen NH, Messick JM, Gray J, et al. ASA monitoring standards and magnetic resonance imaging. Anesth Analg. 1994;79:1141–1147. 46. Davis PL, Crooks L, Arakawa M, et al. Potential hazards in NMR imaging: heating effects of changing magnetic fields and RF fields on small metallic implants. AJR. 1981;137:857–860. 47. Laakman RW, Kaufman B, Han JS, et al. MR imaging in patients with metallic implants. Radiology. 1985;157:711–714. 48. Mechlin M, Thickman D, Kressel HY, et al. Magnetic resonance imaging of postoperative patients with metallic implants. AJR. 1984;143:1281–1284. 49. Sigwart U, Puel J, Mirkovitch V, et al. Intravascular stents to prevent occlusion and restenosis after transluminal angioplasty. N Engl J Med. 1987;316:701–706. 50. Fischman DL, Leon MB, Baim DS, et al. A randomized comparison of coronary stent placement and balloon angioplasty in the treatment of coronary artery disease. N Engl J Med. 1994;331:496–501. 51. Serruys PW, de Jaegere P, Kiemeneij F, et al. A comparison of balloonexpandable stent implantation with balloon angioplasty in patients with coronary artery disease. N Engl J Med. 1994;331:489–495. 52. Kimura T, Yokoi H, Nakagawa Y, et al. Three-year follow up after implantation of metallic coronary artery stents. N Engl J Med. 1996;334:561–566. 53. Carrozza JP, Kuntz RE, Levine MJ, et al. Angiographic and clinical outcome of intra-coronary stenting: immediate and longterm results from a large single-center experience. J Am Coll Cardiol. 1992;20:328–337. 54. Shellock FG, Morisoli S, Kanal E. MR procedures and biomedical implants, materials and devices: an update. Radiology. 1993;189:587–599. 55. Matsumoto AH, Teitelbaum GP, Barth KH, et al. Tantalum vascular stents: in vivo evaluation with MR imaging. Radiology. 1989;170:753–755. 56. Scott NA, Pettigrew RI. Absence or movement of coronary stents after placement in a magnetic resonance imaging field. Am J Cardiol. 1994;73:900–901.

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57. Strohm O, Kivelitz D, Gross W, et al. Safety of implantable coronary stents during H-1 magnetic resonance imaging at 1.0 and 1.5 T. J Cardiovasc Magn Reson. 2000;1:239–245. 58. Gerber TC, Fasseas P, Lennon RJ, et al. Clinical safety of magnetic resonance imaging early after coronary artery stent placement. J Am Coll Cardiol. 2003;42:1295–1298. 59. Nagel E, Hug J, Bunger S, et al. Coronary flow measurements for evaluation of patients after stent implantation. MAGMA. 1998; 6:184–185. 60. Kramer CM, Rogers WJ, Reichek N, et al. Magnetic resonance contrast enhancement versus dobutamine tagging response for assessment of myocardial viability after infarction [abstract]. J Am Coll Cardiol. 1999;33(suppl a):485A. 61. Al Saadi N, Nagel E, Gross M, et al. Improvement of myocardial perfusion reserve early after coronary intervention: assessment with cardiac magnetic resonance imaging. J Am Coll Cardiol. 2000;36: 1557–1564. 62. Duerinckx AJ, Atkinson D, Hurwitz R, et al. Coronary MR angiography after coronary stent placement. AJR. 1995;165:662–664. 63. Kotsakis A, Tan KH, Jackson G. Is MRI a safe procedure in patients with coronary stents in situ? Int J Clin Practice. 1997;51:349. 64. Shellock F. Pocket Guide to MR Procedures and Metallic Objects: Update 1998. Philadelphia, USA: Lippincott-Raven; 1998. 65. Hassler M, Le Bas JF, Wolf JE, et al. Effects of magnetic fields used in MRI on 15 prosthetic heart valves. J Radiol. 1986;67:661–666. 66. Soulen RL, Budinger TF, Higgins CB. Magnetic resonance imaging of prosthetic heart valves. Radiology. 1985;154:705–707. 67. Globits S, Higgins CB. Assessment of valvular heart disease by magnetic resonance imaging. Am Heart J. 1995;129:369–381. 68. Bachmann R, Deutsch HJ, Jungehulsing M, et al. Magnetic resonance tomography in patients with a heart valve prosthesis. Rofo. 1991; 155:499–505. 69. DiCesare E, Enrici RM, Paparoni S, et al. Low field magnetic resonance imaging in the evaluation of mechanical and biological heart valve function. Eur J Radiol. 1995;20:224–228. 70. Avery JK. Loss prevention case of the month: not my responsibility! J Tenn Med Assoc. 1988;81:523. 71. Gimbel JR, Kanal E. Can patients with implantable pacemakers safely undergo magnetic resonance imaging? J Am Coll Cardiol. 2004;43:1325–1327. 72. Lauck G, Smekal AV, Wolke S, et al. Effects of nuclear magnetic resonance imaging on cardiac pacemakers. PACE. 1995;18:1549–1555. 73. Achenbach S, Moshage W, Diem B, et al. Effects of magnetic resonance imaging on cardiac pacemakers and electrodes. Am Heart J. 1997;134:467–473.

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74. Erlebacher JA, Cahill PT, Pannizzo F, Knowles RJ. Effect of magnetic resonance imaging on DDD pacemakers. Am J Cardiol. 1986;57: 437–440. 75. Gimbel JR, Johnson D, Levine PA, Wilkoff BL. Safe performance of magnetic resonance imaging on five patients with permanent cardiac pacemakers. PACE. 1996;19:913–919. 76. Pennell . Cardiac magnetic resonance with a pacemaker in-situ: can it be done [abstract]. J Cardiovasc Magn Reson. 1999;1:72. 77. Heatlie G, Pennell DJ. Cardiovascular magnetic resonance at 0.5 T in five patients with permanent pacemakers. J Cardiovasc Magn Reson. 2007;9:15–19. 78. Martin ET, Coman JA, Shellock FG, et al. Magnetic resonance imaging and cardiac pacemaker safety at 1.5-Tesla. J Am Coll Cardiol. 2004;43:1315–1324. 79. Luechinger R, Zeijlemaker VA, Pedersen EM, et al. In vivo heating of pacemaker leads during magnetic resonance imaging. Eur Heart J. 2005;26:376–383. 80. Hayes DL, Holmes DR, Gray JE. Effect of 1.5 tesla nuclear magnetic resonance imaging scanner on implanted permanent pacemakers. J Am Coll Cardiol. 1987;10:782–786. 81. Holmes DR, Hayes DL, Gray JE, Merideth J. The effects of magnetic resonance imaging on implantable pulse generators. PACE. 1986;9:360–370. 82. Fetter J, Aram G, Holmes DR, et al. The effects of nuclear magnetic resonance imagers on external and implantable pulse generators. PACE. 1984;7:720–727. 83. Gimbel JR, Bailey SM, Tchou PJ, et al. Strategies for the safe magnetic resonance imaging of pacemaker-dependent patients. Pacing Clin Electrophysiol. 2005;28:1041–1046. 84. Roguin A, Zviman MM, Meininger GR, et al. Modern pacemaker and implantable cardioverter/defibrillator systems can be magnetic resonance imaging safe: in vitro and in vivo assessment of safety and function at 1.5 T. Circulation. 2004;110:475–482. 85. Naehle CP, Sommer T, Meyer C, et al. Strategy for safe performance of magnetic resonance imaging on a patient with implantable cardioverter defibrillator. Pacing Clin Electrophysiol. 2006;29:113–116. 86. Naehle CP, Strach K, Thomas D, et al. Magnetic resonance imaging at 1.5T in patients with implantable cardioverter-defibrillators. J Am Coll Cardiol. 2009;54:549–555. 87. Pennell DJ, Sechtem UP, Higgins CB, et al. Clinical indications for cardiovascular magnetic resonance (CMR): consensus panel report. Eur Heart J. 2004;25:1940–1965.

Special Considerations: Cardiovascular Magnetic Resonance in Infants and Children Mark A. Fogel

Cardiovascular magnetic resonance (CMR) has been in use for more than two decades and has become firmly established in the clinical evaluation of anatomy, physiology, and function in patients with congenital heart disease (CHD).1–29 Currently, CMR is used as an adjunct to other imaging modalities, such as echocardiography and invasive angiography. When requested, CMR studies are directed to clarify specific questions about the morphology and function of known anatomy rather than de novo anatomic issues.30 However, in areas such as the rings25–29 and in the assessment of left ventricular (LV) volume and mass and right ventricular (RV) volume,31,32 CMR is the established “gold standard.” It offers numerous advantages over other imaging modalities including lack of ionizing radiation, excellent soft tissue contrast, a capacity for true three-dimensional (3D) imaging of anatomy and physiology as well as function, accurate flow quantification, noninvasive labeling of the myocardium or blood (myocardial tagging), assessment of myocardial viability and perfusion, coronary imaging, and freely selectable viewing planes without limitations to “acoustic windows” or overlapping structures. These advantages and continued advances in CMR hardware, software, and imaging techniques have brought CMR into mainstream pediatric cardiology. Fast imaging with steady-state free precession (SSFP) and real-time cine,33 darkblood spin echo,34 and 3D contrast enhanced magnetic resonance angiography (CE-MRA)35 provide additional benefits to CMR. The application of CMR to CHD uses nearly all of the techniques discussed in the “adult” chapters of this text, but it nevertheless stands on its own as a separate discipline in the world of CMR for the following reasons: 1. Technical challenges: Imaging of children is more demanding than imaging of adults. Children require increased spatial resolution because of their smaller size as well as increased temporal resolution because of their higher heart rates compared with adults. Children younger than 7 to 9 years old often need to be sedated, making the children incapable of sustained breath holding. “Workarounds” to breath holding have been developed for use in imaging the pediatric patient. 2. Anatomy of CHD: The anatomy of native CHD is different from what is seen in adult cardiology, and the correct diagnosis demands a rigorous and systematic approach as well as a global appreciation of 3D thoracic anatomy. Many types of CHD rely on surgical or catheter-based therapies.

The anatomic information obtained from the CMR examination must accommodate these interventions. Thus, the physician must be knowledgeable about the pre- and postoperative anatomy of CHD. For example, the connections between the major cardiac segments (atria, ventricles, and great vessels), along with venous anatomy, can be altered in many types of CHD (e.g., transposition of the great arteries). Communications are present where they should not be (e.g., ventricular septal defect [VSD]), persistent blood vessels from fetal life may remain (e.g., patent ductus arteriosus), and cardiac structures may be markedly hypoplastic (e.g., hypoplastic left heart syndrome). In the evaluation of CHD after surgery, conduits and baffles constructed to separate the circulations have little parallel in the adult world. 3. Physiology: The unique physiology of CHD is most often a result of the altered anatomy. Evaluation of shunts plays a critical role in the assessment of many types of CHD, yet has a minor role in adults. Assessment of ventricular function in ventricles with uncommon shapes is routine in the practice of CMR in CHD. Similarly, the assessment of postoperative physiology where the surgeon has created a systemic to pulmonary artery shunt or an anastomosis between the cava and the pulmonary artery must be addressed. As with other imaging modalities, CMR has limitations and challenges in CHD. Sedation and, in some rare instances, general anesthesia to allow young patients to participate in a 45- to 60-minute scan is always a consideration. Even in older preteens or teens, cooperation may be problematic (e.g., breath holding). Although many intravascular coils, wires, stents, and clips are safe for CMR imaging (see www.mrisafety.com), they may cause image artifacts near the structure of interest. The lack of portability of CMR is disadvantageous, especially for the critically ill infant for whom transportation to the CMR suite is cumbersome. Finally, arrhythmias may not allow proper data acquisition, whereas abnormal T-waves may not allow for proper triggering. An alternative for some situations is the use of real-time CMR, in which sequential images are constructed without regard to electrocardiogram (ECG) triggering and sequences with “arrhythmia rejection.” Pacemakers remain a problem, and patients with CHD who have these devices often do not undergo CMR, although there are increasing data to suggest that CMR is safe.36–39 Cardiovascular Magnetic Resonance 111

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CHAPTER 9

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

This chapter discusses the current state of the art of CMR in CHD. For simplicity, the imaging techniques discussed are generic. The major manufacturers of CMR equipment have similar sequences with proprietary nomenclature (see Appendix I.).

GENERAL PROTOCOL FOR CARDIOVASCULAR MAGNETIC RESONANCE IN CONGENITAL HEART DISEASE The CMR procedure in patients with CHD is not standardized because unsuspected findings are common in this population. Nevertheless, certain basic principles can be formulated into a generalized protocol, as outlined in Figure 9-1. Although the specific protocol for each patient must be individualized, depending on the suspected disease process, this basic outline has been found to be efficient and comprehensive.

Anatomic Imaging After scout localizer examinations are performed, anatomic images are acquired both to survey the thoracic anatomy

GENERALIZED CMR PROTOCOL FOR CHD • Localizers • Static SSFP-axial • HASTE* • Dark blood • Coronary imaging • Cine • Velocity mapping • Gadolinium • Tagging • Viability

• Anatomy • Physiology • Perfusion • Viability

*MPR performed on stack of axial SSFP during HASTE to determine slice orientations and positions of future sequences

Figure 9-1 Generalized protocol for cardiovascular magnetic resonance imaging in patients with congenital heart disease. This is a generalized protocol and must be individualized to the patient and the specific disease. HASTE, half Fourier acquisition single-shot turbo spin echo; MPR, multiplanar reconstruction; SSFP, steady-state free precession.

and to make an “anatomic diagnosis.” These images also serve as localizers for subsequent physiologic and functional imaging. A contiguous set of 40 to 50 axial images from the diaphragm to the thoracic inlet is obtained (Fig. 9-2). The volume may extend outside the thorax for special cases, such as to the neck if arch anatomy is being evaluated (because the innominate artery may branch in the neck) or if total anomalous pulmonary venous connections are suspected (e.g., connection below the diaphragm).

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Figure 9-2 Stack of contiguous axial steady-state free precession images. A selected set of axial CMR images obtained in a patient with an atriopulmonary Fontan connection using static steady-state free precession. Images progress from superior (A) to inferior (F). Note the very dilated right atrium (RA) and the RA-to-pulmonary artery (PA) connection. Ao, aorta; LPA, left pulmonary artery; LV, left ventricle; RPA, right pulmonary artery; RV, right ventricle. 112 Cardiovascular Magnetic Resonance

Cine Cardiovascular Magnetic Resonance Cine SSFP or spoiled gradient recalled echo (GRE) acquisitions are tailored to the lesion under study. Cine CMR (Fig. 9-5) and velocity mapping are the major sequences used to delineate physiology and function in CHD. These bright-blood techniques readily show motion of the heart and blood and can detect turbulence in vessels (e.g., in the pulmonary arteries if they are stenotic) or at the valve level (e.g., aortic stenosis or regurgitation). For example, to determine LV volume overload in a patient with a VSD, a stack of LV short axis images is acquired to measure LV volume, mass, ejection fraction, stroke volume, and cardiac index. If, instead of the VSD, the LV volume overload was caused by aortic regurgitation, cine CMR of the LV outflow tract in multiple views can be used to visualize the regurgitant jet. Cine CMR can also be used to delineate the anatomy. If coarctation of the aorta is visualized on dark-blood imaging or MPR of the axial dataset, turbulence at the coarctation site should be noted on cine imaging (unless there are

multiple collateral vessels or if the coarctation is not severe). Caution must be used, however, when this signal void is used to assess the degree of stenosis (or valvular regurgitation). The size of the signal void is related to many factors, including echo time (TE). Longer TE increases the signal void size, whereas shorter TE decreases it. The size of the signal void is also a function of the direction of the stenotic jet relative to the orientation of the image voxel. Finally, SSFP imaging may underestimate the signal void seen on GRE sequences. With advances in hardware and software, cine CMR is also used to visualize valvular anatomy, often en face (Fig. 9-6). To visualize a bicuspid aortic valve, the most common type of CHD, an en face aortic valve view is obtained by setting the imaging plane to be perpendicular to the LV outflow tract in two orthogonal views at the level of the aortic valve. This typically results in the equivalent of an echocardiographic parasternal short axis orientation. Either SSFP or GRE can be used; GRE is preferred, because the high signal of the blood flowing into the imaging plane outlines the valve leaflets and gives exceptionally fine valve detail (see Fig. 9-6).

Velocity Mapping After cine imaging, blood flow data are obtained. “Retrospectively” acquired velocity mapping (compared with prospectively triggered imaging) is preferred because the former allows data to be gathered throughout the entire cardiac cycle (prospective gating does not allow for data acquisition in the 50 to 100 mg immediately before the next R-wave). Through-plane velocity mapping (flow into and out of the imaging plane) is the most useful type of velocity mapping, although in-plane velocity mapping may also be valuable (discussed later). For example, in a patient with aortic regurgitation, a velocity map across the ascending aorta measures the regurgitant fraction (the area under the curve of the reverse flow divided by the area under the curve of the forward flow). Care should be taken to place the imaging plane as close to the sinotubular junction as possible. For VSDs, through-plane velocity mapping can measure the pulmonary-to-systemic flow ratio (Qp/Qs) by placing velocity maps across the proximal aorta and proximal pulmonary artery. Measurement of Qp/Qs by velocity mapping has been validated against oximetry.19 One of the strengths of CMR velocity mapping is that it is possible to perform an independent verification to ensure data quality. In the absence of an intracardiac shunt, forward stroke volume from through-plane velocity mapping across the ascending aorta should equal forward stroke volume across the pulmonary artery. If a shunt is present, the sum of the flows from each branch pulmonary artery should equal the flow in the main pulmonary artery. LV stroke volume using cine CMR (in the absence of mitral regurgitation) should equal the aortic forward flow. Velocity mapping can also be used to assess the anatomy of valves (see Fig. 9-6). Through-plane velocity mapping can outline the leaflets of valves when imaged en face and can be successful when routine cine imaging is not. Cardiovascular Magnetic Resonance 113

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Imaging with SSFP depends on contrast between the blood pool and the surrounding tissue, which is enhanced by using a high flip angle. Therefore, the scanner is allowed to select the highest flip angle possible. One disadvantage of using “static” SSFP is that diastolic turbulence may be present on the image and may appear as an absence of a structure (signal loss), as in the case of the pulmonary arteries in a patient with a single ventricle after a Blalock-Taussig shunt. These may be very difficult to visualize because the diastolic turbulence creates a major signal loss. This drawback can be compensated for by performing dark-blood half Fourier acquisition single-shot turbo spin echo imaging or another type of dark-blood imaging. It is performed while multiplanar reconstruction (MPR) is being performed (discussed later) on the “static” SSFP images. Alternatively, a set of axial cines can be performed to visualize these structures in the hope that a phase of the cardiac cycle will show the “missing” structure. With the MPR technique, a set of contiguous images (e.g., axial) are stacked or combined to create a largevolume dataset from which software can then reconstruct any arbitrary plane. The exact slice orientation and position can therefore be used for subsequent imaging during the study. In addition, the anatomy can be inspected from multiple views from just the axial dataset. An interesting feature of this approach is that one is not restricted to planar surfaces. A “curved” cut can be created to show the important anatomic features (Fig. 9-3). A set of high-resolution, double-inversion recovery dark-blood images is obtained for the regions of interest (e.g., candy cane view of the aorta for suspected coarctation) after delineation via MPR (Figs. 9-3 and 9-4). This type of imaging is used judiciously because it requires relatively long scanning times. If there are time constraints and if there are no turbulence artifacts in the area of interest, this step can be skipped and cine imaging can be performed instead to delineate the region of interest.

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

A

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Figure 9-3 Curved multiplanar reconstruction. This patient has a cervical, circumflex right aortic arch, where the transverse aortic arch passes over the right main stem bronchus and then crosses over to the left side of the thorax posteriorly to descend on the left side of the spine. A to D, Double-inversion dark-blood axial CMR images that progress from superior (A) to inferior (D). The transverse arch crosses over to the left of the thorax posteriorly. E, When the axial double-inversion dark-blood images are stacked one atop the other, a curved plane (white line) can be drawn tracing the course of this very tortuous aortic arch that shows the entire aortic arch anatomy (F) in the reformatted image. G, Similar curved plane reconstruction from the brightblood steady-state free precession dataset. AAo, ascending aorta; DAo, descending aorta; LV, left ventricle.

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114 Cardiovascular Magnetic Resonance

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Figure 9-4 Dark-blood imaging. Three examples of cardiac pathology using double-inversion dark-blood CMR imaging. A, Patent ductus arteriosus (PDA) in a 2-week-old infant obtained while free breathing. B, Aortic coarctation in the same patient. C, Tetralogy of Fallot in a 2-day-old infant with pulmonary atresia imaged while free breathing to look for collateral vessels. A collateral vessel is seen from underneath the transverse aortic arch (TAo) connecting to the left pulmonary artery (LPA). AAo, ascending aorta; DAo, descending aorta; MPA, main pulmonary artery.

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Figure 9-5 Cine cardiovascular magnetic resonance. Steady-state free precession CMR images from cines of four-chamber (A), short axis (B), and long axis (C) images of a patient with Uhl anomaly showing an extremely dilated and thinned right ventricle (RV). D, Systolic frame from the patient in Figure 9-4 with tetralogy of Fallot and pulmonary atresia showing the ventricular septal defect (arrow) and the overriding aorta (Ao). E, Short axis cine of a patient with superoinferior ventricles and a right ventricular outflow chamber (RVOC). AAo, ascending aorta; LV, left ventricle.

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BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

Figure 9-6 Cine cardiovascular magnetic resonance and velocity mapping across a systemic semilunar valve. Phase encoded velocity maps (A and C) and gradient echo cines of a quadricuspid truncal valve (A and B) in a patient with truncus arteriosus (A and B) and a bicuspid aortic valve (C and D). Arrows point to the various leaflets. Both datasets were acquired with free breathing.

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Gadolinium-Based Anatomic Cardiovascular Magnetic Resonance The two types of 3D gadolinium techniques for anatomy (both static [Fig. 9-7] and time-resolved35 [Fig. 9-8] versions), not including late gadolinium enhancement (LGE; discussed later), are useful. If both are to be performed in the same patient, they should be separated in time (15 to 20 minutes) to allow the gadolinium of the first injection to be washed out before the second injection. These techniques not only add a special type of 3D dataset that can be rotated and manipulated in many different ways, but also can be used to image smaller vessels much more effectively than other techniques. This may also be used to create a shaded surface display or a volume-rendered 3D image (see Fig. 9-7). In addition, the gadolinium can “label” the collateral vessels. Special techniques that do not fall into this generalized protocol may be inserted at different points, depending on the lesion: Myocardial and blood tagging (Fig. 9-9): When there is a question of regional wall motion abnormalities (e.g., cardiomyopathy, as in Duchenne muscular 116 Cardiovascular Magnetic Resonance

dystrophy,40 or a single ventricle11,22,23) or mass (e.g., tumor or hypertrophic cardiomyopathy), CMR tagging can be used both qualitatively and quantitatively. With myocardial tagging, the ventricle is divided up noninvasively into “cubes of magnetization” and local deformation of the myocardium can be visualized (see Fig. 9-9). If there are concerns about a possible atrial septal defect (ASD) or VSD, a saturation band can be proscribed tagging blood on a GRE image.41 This “tagged”/dark blood can be used to visualize a shunt (see Fig. 9-9). Perfusion (see Fig. 9-8): Assessment of myocardial perfusion is also a consideration for the patient with CHD.42 These may include patients with congenital coronary anomalies (e.g., anomalous left coronary artery from the pulmonary artery) and patients who have had surgical manipulation of the coronary arteries (e.g., transposition of the great arteries after an arterial switch procedure or a Ross procedure) as well as patients with Kawasaki disease or other inflammatory disorders. A first-pass gadolinium-based method typically involves imaging with adenosine (140 mg/kg/min for 4 to 6 minutes), followed 20 minutes later with resting perfusion without vasodilator and 10 minutes later with LGE imaging (discussed later). Cine CMR dobutamine stress43 is a non– gadolinium-based option for those with impaired renal function and detects wall motion abnormalities.

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Figure 9-7 Static three-dimensional contrast-enhanced magnetic resonance angiography. A–D, A child with transposition of the great arteries (S, L, L; “corrected transposition”) with pulmonic stenosis who underwent apical left ventricular (LV) apical-to-pulmonary artery conduit (C). “Right-sided” (A and B) and “left-sided” (C and D) phases, including the right ventricle (RV) and ascending aorta (AAo). E, Maximum intensity projection image of the gadolinium injection of a patient after scimitar vein repair with a baffle (B) placed between the right pulmonary vein (RPV) and the left atrium (LA). F, Shaded surface display of the gadolinium injection of a patient with aortic coarctation (arrow) and a conduit (C) between the AAo and descending aorta (DAo). MPA, main pulmonary artery; RPA, right pulmonary artery.

Late gadolinium enhancement imaging (see Fig. 9-8): Similar to myocardial perfusion, LGE imaging (viability) has a role in CHD.42,44,45 Scar tissue preferentially accumulates gadolinium. LGE imaging takes advantage of the contrast between the gadolinium-laden scar tissue and the gadolinium-poor normal myocardium to “label” the scar tissue. Pulse sequences, first described in the mid1980s,46 have taken advantage of this property, which is unique in noninvasive, nonionizing imaging. With the development of segmented inversion recovery GRE, differences in signal intensity between normal and infarcted myocardium of up to 500% have been achieved.47 The same patients who are candidates for perfusion imaging are also candidates for LGE imaging. In addition, hearts that have undergone surgical reconstruction are candidates for LGE imaging to assess for myocardial scar tissue in both the myocardium and the areas that were reconstructed, such as the infundibulum and the pulmonary annulus in a patient after transannular patch repair of tetralogy of Fallot. LGE imaging can also be

used to delineate anatomy because surgically placed patches and valves can become bright with this technique. Coronary artery imaging45 (Fig. 9-10): Using navigator sequences,48 and for some patients, breath hold imaging, coronary artery CMR may be used to image patients with CHD who have coronary manipulation (e.g., transposition of the great arteries after arterial switch45), native disease (e.g., anomalous left coronary artery from the pulmonary artery), or acquired disease (e.g., coronary aneurysm as a result of Kawasaki disease). Both targeted 3D and whole heart coronary CMR has been successfully applied in infants and children.43,49 Two special situations that are worth mentioning occur not infrequently in pediatrics and require special protocols that use many of the techniques noted earlier as well as some that were not mentioned: Tumor/mass characterization50: Many cardiac tumors and masses can be differentiated from each other not only by where they occur in the heart, what symptoms they cause, and at what age they occur, but also Cardiovascular Magnetic Resonance 117

9 SPECIAL CONSIDERATIONS: CARDIOVASCULAR MAGNETIC RESONANCE IN INFANTS AND CHILDREN

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Figure 9-8 Late gadolinium enhancement, perfusion, and time-resolved gadolinium imaging. A, Late gadolinium enhancement (LGE) image of a patient who underwent complete atrioventricular canal repair with high signal intensity in the region of the patch (arrow). B, LGE image of a 1-year-old child with a right ventricular (RV) fibroma (F) showing how the tumor enhances with this sequence. C, Single perfusion frame of the patient with the RV fibroma showing a ring of high intensity signal in the tumor and much lower signal intensity in the middle of the tumor. D to F, Three phases of a time-resolved gadolinium injection (D, earlier phase; F, later phase) in a 5-year-old child with single ventricle physiology after Fontan completion. The patient has a left superior vena cava-to-coronary sinus that was connected to the inferior margin of the Fontan baffle (D and E). F, Later arterial and early venous phases, including the right superior vena cava connection to the right pulmonary artery (arrows).

by their characteristics on CMR. For example, a fibroma accumulates gadolinium and shows signal intensity on LGE imaging and T1-weighted imaging after gadolinium administration, whereas a lipoma shows signal intensity on noncontrast T1-weighted images and becomes signal-poor with the application of a fat saturation pulse. Tumor characterization procedures with CMR are a protocol in themselves and typically include T1- and T2-weighted images, images with fat saturation, GRE imaging (e.g., thrombus), perfusion (e.g., hemangiomas), and LGE imaging, T1-weighted images after gadolinium administration, and myocardial tagging. Functional imaging can be used to assess for effects of the tumor, such as obstruction to flow. Arrhythmogenic right ventricular cardiomyopathy: Arrhythmogenic right ventricular cardiomyopathy is characteristically associated with replacement of the RV myocardium with fatty or fibrofatty tissue.51 In its most flagrant form, the following are seen: (1) left bundle branch block tachycardia; (2) RV dilation; (3) dyskinetic RV wall motion, especially in the right ventricular outflow tract, diaphragmatic surface, and septum; and (4) RV conduction delay on ECG, inverted T-waves, and e-waves. Imaging fulfills some of the criteria set forth in the 1994 Task Force 118 Cardiovascular Magnetic Resonance

report.52 CMR has been used successfully in adults. However, in the pediatric population, there is uncertainty as to its utility.53–55 The CMR protocol includes T1-weighted imaging, cine for RV volume and function, myocardial tagging if needed to assess regional wall motion, and LGE.

TECHNICAL CONSIDERATIONS IN PEDIATRIC CARDIOVASCULAR MAGNETIC RESONANCE Because of the demands of high spatial and temporal resolution in infants and children, along with the inability of these patients to hold their breath while sedated, many adult CMR sequences must be adapted for the pediatric population.

Spatial and Temporal Resolution Small voxels needed to image infants and small children can be problematic because of the poor signal-to-noise ratio

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Figure 9-10 Navigator-gated, electrocardiogram-gated CMR image of a right coronary artery anomaly obtained with free breathing. A and B, A 3-year-old patient with a single right coronary artery (RCA) giving rise to a retroaortic left main coronary artery (LCA) with a retroaortic (benign) course and followed by bifurcation into the left circumflex (LCx) and left anterior descending (LAD) coronary arteries. C, A 1-year-old child with anomalous origin of the RCA from the left coronary cusp. Ao, aorta; PA, pulmonary artery.

(SNR). With the use of parallel imaging techniques, the SNR is further impaired, leading to poor-quality, “grainy” images. To increase the SNR, a number of strategies have been used, either in isolation or in combination, such as using multiple averages, phase oversampling, decreasing the bandwidth, and avoiding or limiting the use of parallel

imaging. These modifications come at the cost of imaging time for the benefit of image quality. At times, for the smallest voxels, even these strategies are not sufficient and the CMR imager must be satisfied with larger voxels to improve the SNR by decreasing the matrix size. Typically, when imaging the adult, a base matrix Cardiovascular Magnetic Resonance 119

9 SPECIAL CONSIDERATIONS: CARDIOVASCULAR MAGNETIC RESONANCE IN INFANTS AND CHILDREN

Figure 9-9 Myocardial and blood tagging. Diastolic (A) and systolic (B) CMR images of a “one-dimensional” tagging sequence in the fourchamber view of the patient described in Figure 9-5 with Uhl anomaly. Regional wall motion can be visualized in this manner. C and D, Myocardial “grid” tagging (spatial modulation of magnetization) in the short axis of a patient with hypoplastic left heart syndrome, again showing regional wall motion. E, A set of three images of a patient with an incomplete atrioventricular canal and an ostium primum atrial septal defect. Images with a saturation band are shown on the right side of the heart (upper right, arrow), and a saturation band is seen on the left side of the heart (arrowhead). Black blood can be seen shunting from the left atrium to the right atrium in the lower left image (arrowhead), and bright blood can be seen shunting the same way in the upper right image (lower left, arrow). LV, left ventricle; RV, right ventricle.

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

(in the frequency encoding direction) of 256 or 192 is used and this can be decreased to 128 to improve the SNR and still yield sufficient spatial resolution with a field of view of 200 mm or less (pixel size of 1.5 to 2.0 mm or less). Slice thicknesses generally are not less than 2.5 to 3.0 mm because of SNR considerations. Because heart rates can be very high in infants (not uncommonly, 140 to 170 bpm), as a general rule, for an R-R interval of less than 500 msec, 15 to 20 phases of the cardiac cycle should be acquired. For an R-R interval of 500 to 800 msec, 20 to 25 phases should be acquired. For an R-R interval of greater than 800 msec, 25 to 30 phases should be acquired. To be able to obtain this kind of temporal resolution at fast heart rates, both the repetition time (TR) and the number of segments must be low. If the TR and the number of segments are not low enough, in retrospectively gated imaging, the walls of the heart appear blurry and with a double or triple shadow, whereas in prospectively triggered imaging, the heart motion appears stilted and jerky. Adequate temporal resolution can be obtained at fast heart rates with three segments at a minimum and TR of 10 to 30 msec. In retrospective gating sequences, the CMR scanner is acquiring data continuously and recording the ECG simultaneously. After all of the lines of k-space are acquired, the computer then “bins” each line of k-space to the closest phase of the cardiac cycle it is calculating, interpolating the data as it is performing the “bin.” It is, therefore, important to recognize that the number of calculated phases should not exceed two times the measured phases (essentially the R-R interval [msec]/TR [msec], where TR is the line TR lines of k-space obtained) because there should be two measured points between each interpolated point to obtain robust data. The formula used is: 2 R-R interval ðmsecÞ=number of calculated phases ¼ TR ðmsecÞ where TR is defined as the line TR lines of k-space obtained. This is especially important in velocity mapping. At times, it is advantageous to acquire single-shot, realtime CMR in which all of the lines of k-space are acquired in one heartbeat because of respiratory motion or arrhythmia. If the heart rate is too fast, it may be advantageous to obtain the image over two heartbeats (i.e., imaged at the end of two heartbeats [2 R-R interval] instead of one.

Inability of Pediatric Patients to Hold Their Breath To allow CMR to be performed in patients who cannot hold their breath, some institutions have instituted general anesthesia during which the anesthesiologist can suspend respiration. This approach is both aggressive and invasive and adds considerable expense. At the Children’s Hospital of Philadelphia, deep sedation has been successfully used for many years without untoward effects and allows the infant to breathe freely during the CMR examination. We reserve general anesthesia for patients with cardiorespiratory compromise, those who have failed deep sedation, and those situations in which deep sedation would need to be markedly prolonged. 120 Cardiovascular Magnetic Resonance

Imaging the patient who is allowed to breathe freely often is used with the following: 1. Multiple excitations to “average out” the respiratory motion. Typically, three excitations (which increase the SNR) are needed to obtain good image quality, although more excitations may be needed for the vigorous breather. 2. Navigator-based respiratory gating techniques that monitor diaphragmatic motion with data acquisition during expiration. In addition to these strategies, placing a saturation band over the chest wall can minimize respiratory artifacts. These common strategies allow high-quality CMR images to be obtained, albeit at the cost of longer acquisition times.

Gadolinium-Based Techniques The information discussed earlier applies to all types of imaging, but the gadolinium-based techniques have additional considerations. 1. When performing “static” 3D gadolinium imaging, many CMR systems use a “bolus tracking” technique in which the major structure of interest is imaged in real time to detect the arrival of gadolinium with the static 3D sequence cued up. When the imager visualizes the gadolinium bolus arriving in this structure, the real-time sequence is stopped and the static 3D sequence is initiated. This ensures that the maximum concentration of gadolinium is in the structure of interest during imaging. In children, because of the quick circulatory time, intravenously administered gadolinium will reach the systemic circulation (e.g., aorta) much more rapidly than in adults. 2. The perfusion technique obtains images at a single-slice position at different time points in the cardiac cycle. With the high heart rates found in infants and children, it may be possible to acquire only one or two perfusion slices within a single R-R interval. If this occurs, the imager can perform the sequence over two R-R intervals without degradation of diagnostic quality and can obtain more slices (three to four) during the scan.

EXAMPLES OF CARDIOVASCULAR MAGNETIC RESONANCE IN CONGENITAL HEART DISEASE Transposition of the Great Arteries With the aorta arising from the RV and the pulmonary artery (PA) arising from the LV, the state-of-the-art surgical procedure for simple transposition of the great arteries is the arterial switch procedure, during which both great vessels are surgically transposed to arise normally and the coronary arteries are implanted above the native pulmonary (neoaortic) valve. The pulmonary arteries are typically “draped” over the ascending aorta in the Lecompte maneuver (Figs. 9-11 to 9-13), which may result in left PA stenosis. It is advantageous for the surgeon to perform a

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Figure 9-12 Cine steady-state free precession CMR imaging of a patient with transposition of the great arteries. A, The patient seen in Figure 9-11A with no left pulmonary artery (LPA) stenosis. B, The patient in Figure 911C showing a narrowed right ventricular outflow tract (RVOT). This is a diastolic frame, so no turbulence is seen. RPA, right pulmonary artery.

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Lecompte maneuver rather than making the right PA course tortuous by bringing it directly posterior and then underneath the aortic arch. For a typical assessment, after the set of axial images is obtained, both dark-blood imaging (see Fig. 9-11) and cine of the PA (see Fig. 9-12) are performed in both long axes to

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assess for stenoses, which generally can occur at the takeoff of the left PA or in the right PA between the superior vena cava (SVC) and the aorta. Cines of the RV and LV outflow tracts are used to evaluate for obstruction. In addition, cines of the four-chamber view and a set of contiguous short axis continuous cines of both ventricles are used to Cardiovascular Magnetic Resonance 121

9 SPECIAL CONSIDERATIONS: CARDIOVASCULAR MAGNETIC RESONANCE IN INFANTS AND CHILDREN

Figure 9-11 Dark-blood CMR imaging of a patient with transposition of the great arteries. The Lecompte maneuver and anatomy of the pulmonary arteries. A, “Draping” of the right (RPA) and left pulmonary arteries (LPA) over the ascending aorta (AAo) in a patient without LPA stenosis. B, Image in a patient with similar anatomy but a proximal LPA stenosis (arrow). Long axis view of the RPA (C) and LPA (D) in a patient with a narrowed right ventricular outflow tract (RVOT). DAo, descending aorta.

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

Figure 9-13 Three-dimensional volume-rendered contrast enhanced magnetic resonance angiography of a patient with transposition of the great arteries. A, The pulmonary anatomy relative to the aorta is seen in various views rotated around the superoinferior axis of the patient. Note the narrowed right ventricular outflow tract (A, C, and D). Ao, aorta; LPA, left pulmonary artery; RPA, right pulmonary artery; RVOT, right ventricular outflow tract.

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quantitate ventricular performance and to search for residual VSD. Data suggest that including RV trabeculations and papillary muscles as ventricular volume (rather than wall/mass) results in shorter analysis time and better interobserver reproducibility for ejection fraction.53 Velocity mapping across the neoaortic and neopulmonary valves is used to quantify cardiac output. This technique is used across the right and left PAs to determine flow distribution to both lungs and to assess gradients across the stenotic PA.20 Gadolinium is used for 3D reconstruction of the PA and aorta (see Fig. 9-13), and LGE44 is performed to evaluate for myocardial scarring. In the case shown in Figures 9-11C and D, 9-12A and B, and 9-13, the patient had an RV outflow tract obstruction and was being evaluated by CMR to determine whether the coronary arteries would be compromised by stent insertion. 3D reconstruction, along with coronary imaging, showed nearly 14 mm between the nearest coronary artery and the RV outflow tract. Thus, the coronary arteries were far enough away from the region of stent deployment that there would be no concerns. In general, coronary imaging can be performed successfully to show the manipulated coronary arteries in the surgical treatment of this disease.49 122 Cardiovascular Magnetic Resonance

Single Ventricles When only one usable ventricle is present to pump blood effectively and the other is hypoplastic, or when both ventricles are linked in such a way that separation of the circulations into two pumping chambers is impossible, the heart falls into the category of functional single ventricle. The aorta may be hypoplastic, and there may be other associated anomalies, such as anomalous venous connections. Single ventricles undergo staged surgical reconstruction, culminating in the Fontan procedure.56 The CMR imager must be familiar not only with the various forms of single ventricles, associated anomalies, and the physiologic and functional sequelae, but also with the various surgical reconstructive techniques to allow for the best possible medical and surgical management of these very complex patients. A comprehensive discussion of this evaluation is beyond the scope of this chapter, and the reader is referred to more specialized sources.6,8,11,12,18,21,22,24,53 As the patient with a single ventricle moves through staged reconstruction, specific targets of examination change, but the overall goal of assessing anatomy, physiology, and function remains. At all stages of surgical

the semilunar (and indirectly) atrioventricular valve, and for ventricular outflow tract obstruction (see Fig. 9-16). In the native state of the single ventricle, because much less is known about the anatomy than at other stages, it is important to perform an even more detailed anatomic assessment. Anomalous venous structures, such as a decompressing vein from the LA, the presence of a left SVC, delineation of a visceral situs for heterotaxy, the presence of an inferior vena cava, and hepatic venous drainage are all important details. In addition, because some patients may have needed resuscitation, assessment of ventricular function and valve insufficiency is also important. After the stage I procedure for hypoplastic left heart syndrome, for example, aortic arch imaging is used to evaluate the initial repair (see Fig. 9-14). Besides assessment of the distal aortic arch for obstruction and the aortic-to-pulmonary anastomosis, visualization of the aorticto-pulmonary shunt (typically, a right Blalock-Taussig shunt) or the right ventricle-to-PA shunt (i.e., Sano procedure) is important. This is generally done with doubleinversion dark-blood imaging (or CE-MRA) because the turbulence across this structure by cine generally causes

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Figure 9-14 Aortic arch, pulmonary artery, and venous pathway imaging in a patient with hypoplastic left heart syndrome. A to C, Various views of the reconstructed aorta with a three-dimensional shaded surface display with the native aorta (nAo) connected to the native main pulmonary artery (nPA). A narrowing of the aortic arch is seen (arrow). D, Two-dimensional steady-state free precession (SSFP) image of the same aortic to pulmonary anastomosis. E, Static SSFP image at the level of the right (RPA) and left pulmonary arteries (LPA) showing their size and relationship to the superior vena cava (SVC). F, Three-dimensional shaded surface display of the same connection (SVC to RPA) with the innominate vein (Inn v) in view. G, Dark-blood image of the Fontan baffle showing no clot and a widely patent systemic venous pathway. The LPA in long axis can be seen. Cardiovascular Magnetic Resonance 123

9 SPECIAL CONSIDERATIONS: CARDIOVASCULAR MAGNETIC RESONANCE IN INFANTS AND CHILDREN

reconstruction, including the single ventricle in the native state, the following is the minimum assessment: Aortic arch imaging, aimed mostly at patients with an aortic-to-pulmonary anastomosis, to assess for aortic arch obstruction (Figs. 9-14 and 9-15). Imaging of the PA6,18,24 to assess for stenosis, hypoplasia, and discontinuity (Figs. 9-14 to 9-16). Anatomic assessment of the Fontan baffle7,8 (see Figs. 9-14 to 9-16); ASD (as in a patient with hypoplastic left heart syndrome; see Fig. 9-15); ventricular outflow tract obstruction (especially in patients with a bulboventricular foramen; see Fig. 9-16); aortic-pulmonary collateral vessels; anomalous venous structures; and pulmonary or systemic venous obstruction (especially in the surgically created bidirectional Glenn shunt or the Fontan baffle itself). (see Figs. 9-14 to 9-16) Biventricular function,11,21,22,23 including regional wall motion abnormalities, ejection fraction, end-diastolic volume and mass, stroke volume, cardiac index, and atrioventricular valve regurgitant fraction (see Fig. 9-15). Velocity mapping18 to assess for cardiac index, Qp/Qs, relative flow to both lungs, and regurgitant fraction of

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

Figure 9-15 Cine, late gadolinium enhancement, and perfusion in patients with a single ventricle. Still frame steady-state free precession cines in off-axis, twochamber (A), short axis (B), and (C) long axis views of a 2-year-old child with hypoplastic left heart syndrome after hemi-Fontan procedure, imaged while free breathing. The superior vena cava to the right pulmonary artery connection can be seen in C (arrow) along with the hemi-Fontan “dam,” blocking flow from the superior vena cava into the right atrium. D, Late gadolinium enhancement in single ventricles, where patch material such as that used to reconstruct the aorta shows high signal intensity (arrow).

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signal loss. Qp/Qs is obtained by velocity mapping in the ascending aorta (over each semilunar valve) and by velocity mapping across the shunt as well as velocity maps in each PA using high velocity encoding (e.g., 400 cm/sec) and a low TE. The status of the ASD should be assessed, and because this is a volume-loaded stage, ventricular function is also a key imaging goal (see Fig. 9-15). After the bidirectional Glenn or hemi-Fontan stage (see Fig. 9-15), imaging the SVC-to-PA anastomosis is the major difference with patients undergoing the stage I procedure. This can be done with cine, dark-blood, or CE-MRA because the flows are generally low velocity. Qp/Qs uses flow in the SVC with flow in both branches of the PA as an internal check along with flow in the aorta. If a hemiFontan procedure was performed, imaging should assess whether any leak was present from the SVC-to-PA anastomosis into the atrium. Collated flow is assessed with velocity maps in the pulmonary veins. After the Fontan procedure, the most important structure to image is the entire systemic venous pathway (Fontan baffle) for obstruction or clot (see Fig. 9-16). Visualization of 124 Cardiovascular Magnetic Resonance

fenestration flow (using cine) and assessment of this structure for thrombus are important steps. Because it is known that patients who have undergone the Fontan procedure have poor ventricular function, cine imaging is an essential part of the CMR study. In addition, CE-MRA can help determine the presence of collateral arteries and assess the aortic arch. LGE imaging57,58 can be used to visualize myocardial as well as patches in both the reconstructed aorta and the atrioventricular valves (see Fig. 9-15). Perfusion may also be performed to evaluate for defects.

Coarctation of the Aorta One of the more common congenital heart lesions and a very common reason for referral for CMR, coarctation of the aorta is defined as an obstruction in the isthmus or descending aorta. It may have a number of associated abnormalities, including bicuspid aortic valve (see Fig. 9-6), complex CHD (e.g., Shone complex), double-orifice mitral valve, or a posteriorly malaligned VSD. In older children,

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Figure 9-16 Ventricular outflow obstruction, Fontan baffle, and pulmonary arteries in single ventricles. This patient with heterotaxy has a functional single left ventricle (LV) and a right ventricular outflow chamber (RVOC) with the aorta (Ao) arising from it, whereas the native pulmonary artery (PA) arises from the LV (anatomic transposition of the great vessels). The only connection between the LV and the RVOC, which connects to the Ao is through two small ventricular septal defects (VSD; arrows on A). A, Coronal steady-state free precession (SSFP) cine. B, Off-axis axial cine through the more superior VSD. Note the turbulence/dephasing through the VSD (arrow). C and D, Still frame images from in-plane velocity mapping at a velocity encoding (VENC) of 3 m/sec and 5 m/sec, respectively, of the VSD. The velocity exceeds 3 m/sec as indicated by the white signal intensity in the otherwise black jet (arrow) in C, whereas the velocity does not exceed the VENC at 5 m/sec in D. The instantaneous maximum velocity was 4.5 m/sec. E and F, Cine SSFP and X-ray angiography of the Fontan baffle and pulmonary arteries (arrows), respectively. RV, right ventricle.

multiple collateral arteries can be seen. Surgery for the coarctation can be as simple as placement of a patch over the narrowed portion or as complex as an ascending-todescending aortic conduit. The CMR procedure can provide exquisite detail of the aortic anatomy and physiology needed for diagnosis and treatment of coarctation (Fig. 9-17). An axial stack of SSFP images will show a decreased descending aortic diameter at the level of the coarctation, whereas the candy cane view with dark-blood imaging will show narrowing. It is important to obtain parallel images on both sides of the candy cane view to ensure that the narrowing is not misinterpreted by deviation of the aorta out of the imaging plane. Cine CMR is used to assess for turbulence, whereas the four-chamber and ventricular short axis stack is used to obtain the LV mass to identify possible hypertrophy. Cine CMR is also used to show aortic valve morphology and visualize a possible bicuspid aortic valve (see Fig. 9-6). Aortic stenosis or regurgitation secondary to a bicuspid valve (see Fig. 9-6) can be assessed by cine as well as phase encoded velocity mapping. Through-plane phase encoded velocity mapping across the

aortic valve is used to obtain the cardiac index, and regurgitant fraction and peak velocity is obtained for aortic regurgitation and stenosis, respectively. In addition, through-plane velocity mapping just below the coarctation and at the diaphragm can quantify collateral blood flow.57–60 In-plane velocity mapping can be used to obtain peak instantaneous velocity as well. Finally, 3D CE-MRA is used to assess for collateral arteries and to create a 3D image of the aorta and the coarctation (see Fig. 9-17).

THE FUTURE With the many ongoing advances in CMR, the future holds great promise for those with CHD. Developments include both diagnostic and therapeutic or interventional CMR (e.g., balloon valvuloplasty, stent placement), molecular imaging, and T2* techniques to evaluate the heart for myocardial iron and oxygen content as well as real-time flow assessment. Functional fetal CMR (Fig. 9-18) with real-time imaging shows the feasibility of this approach in patients Cardiovascular Magnetic Resonance 125

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BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

Figure 9-17 Cardiovascular magnetic resonance of aortic coarctation. A, Still frame image from a cine gradient echo in diastole showing narrowing of the aorta at the apex of the aortic arch. B, Accompanying in-plane phase encoded velocity map showing the jet at the coarctation site (arrow). C, Dark-blood image in a 3-yearold child with a typical coarctation of the proximal descending aorta (DAo) (arrow). D and E, Shaded surface display from a threedimensional contrast enhanced magnetic resonance angiography showing extensive collateral vessels. The coarctation is readily seen (arrow). AAo, ascending aorta.

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Figure 9-18 Functional real-time fetal cardiovascular magnetic resonance. At 36 weeks’ gestation, the patient has a diagnosis of tetralogy of Fallot and pulmonary atresia. A, CMR image from a four-chamber cine steady-state free precession (SSFP) of the heart (arrow). The mother’s anterior abdomen is at the top of the image. B, CMR image of a left ventricular outflow tract cine SSFP (arrow) with the mother’s anterior abdomen to the left of the image. The head is at the bottom. C, Cine SSFP image showing a short axis view of the heart (arrow) with the mother’s anterior abdomen to the left of the image. The head is at the bottom. 126 Cardiovascular Magnetic Resonance

CONCLUSION The use of the CMR technique for infants, children, adolescents, and adults with CHD is a unique subspecialty field

in the world of CMR. The imager must understand the complex anatomy, physiology, and function of native CHD as well as the treatments (surgical, catheter-based, medical) to perform the CMR examination accurately and efficiently. In addition to addressing the myriad of technical issues associated with imaging in pediatrics, ongoing hardware and software advances will likely lead to increasing use of CMR in the CHD population.

References 1. Didier D, Higgins CB, Fisher M, Osakai L, Silverman NH, Cheitlin MD. Congenital heart disease: gated magnetic resonance imaging in 72 patients. Radiology. 1986;158:227–235. 2. Fletcher BD, Jacobstein MD, Nelson AD, Riemenschneider TA, Alfidi RJ. Gated magnetic resonance imaging of congenital cardiac malformations. Radiology. 1984;150:137–140. 3. Higgins CB, Byrd BF, Farmer DW, Osakai L, Silverman N, Cheitlin MD. Magnetic resonance imaging in patients with congenital heart disease. Circulation. 1984;70:851–860. 4. Bank ER. Magnetic resonance of congenital cardiovascular disease. An update. Radiol Clinic North Am. 1993;31:553–572. 5. Adams R, Fellows KE, Fogel MA, Weinberg PM. Anatomic delineation of congenital heart disease with 3D magnetic resonance imaging. In: Proceedings SPIE-Medical Imaging, 1994: Physiology and Function from Multidimensional Images. 1994;2168:184–194. 6. Fogel MA, Ramaciotti C, Hubbard AM, Weinberg PW. Magnetic resonance and echocardiographic imaging of pulmonary artery size throughout stages of Fontan reconstruction. Circulation. 1994;90:2927–2936. 7. Bornemeier RA, Weinberg PM, Fogel MA. Angiographic, echocardiographic and three-dimensional magnetic resonance imaging of extracardiac conduits in congenital heart disease. Am J Cardiol. 1996;78 (6):713–717. 8. Fogel MA, Hubbard A, Weinberg PM. A simplified approach for assessment of intracardiac baffles and extracardiac conduits in congenital heart surgery with two- and three-dimensional magnetic resonance imaging. Am Heart J. 2001;142(6):1028–1036. 9. Fogel MA, Rychik J, Chin A, Hubbard A, Weinberg PM. Evaluation and follow-up of patients with left ventricular apical to aortic conduits using two and three-dimensional magnetic resonance imaging and Doppler echocardiography: a new look at an old operation. Am Heart J. 2001;141:630–636. 10. Weinberg PM, Hubbard AM, Fogel MA. Aortic arch and pulmonary artery anomalies in children. Semin Roentgenol. 1998;33(3):262–280. 11. Fogel MA, Weinberg PM, Gupta KB, et al. Mechanics of the single left ventricle: a study in ventricular-ventricular interaction II. Circulation. 1998;98:330–338. 12. Fogel MA, Weinberg PM, Rychik J, et al. Caval contribution to flow in the branch pulmonary arteries of Fontan patients using a novel application of magnetic resonance presaturation pulse. Circulation. 1999;99:1215–1221. 13. Donofrio MT, Clark BJ, Ramaciotti C, Jacobs ML, Fellows KE, Weinberg PM, et al. Regional wall motion and strain of transplanted hearts in pediatric patients using magnetic resonance tagging. Am J Physiol Regul Integr Comp Physiol. 1999;277:R1481–R1487. 14. Fogel MA, Weinberg PM, Hubbard A, Haselgrove J. Diastolic biomechanics in normal infants utilizing MRI tissue tagging. Circulation. 2000;102:218–224. 15. Eyskens B, Reybrouck T, Bagaert J, et al. Homograft insertion for pulmonary regurgitation after repair of tetralogy of Fallot improves cardiorespiratory exercise performance. Am J Cardiol. 2000;85:221–225. 16. Niezen RA, Helgbing WA, van der Wall EE, van der Geest RJ, Rebergen SA, de Roos A. Biventricular systolic function and mass studied with MRI imaging in children with pulmonary regurgitation after repair for tetralogy of Fallot. Radiology. 1996;201:135–140. 17. Rebergen SA, Chin JGJ, Ottenkamp J, van der Wall EE, de Roos A. Pulmonary regurgitaion in the late postoperative follow-up of tetralogy of Fallot: volumetric quantification by MR velocity mapping. Circulation. 1993;88:2257–2266.

18. Rebergen SA, Ottenkamp J, van der Wall EE, Chin JDJ, de Roos A. Postoperative pulmonary flow dynamics after Fontan surgery: assessment with nuclear magnetic resonance velocity mapping. J Am Coll Cardiol. 1993;21:123–131. 19. Beerbaum P, Korperich H, Barth P, Esdorn H, Gieseke J, Meyer H. Non-invasive quantification of left-to-right shunt in pediatric patients: phase-contrast cine magnetic resonance imaging compared with invasive oxymetry. Circulation. 2001;10:2476–2482. 20. Gutberlet M, Boeckel T, Hosten N, Vogel M, Kuhne T, Oellinger H, et al. Arterial switch procedure for d-transposition of the great arteries: quantitative midterm evaluation of hemodynamic changes with cine MR imaging and phase-shift velocity mapping — initial experience. Radiology. 2000;214:467–475. 21. Fogel MA, Weinberg PM, Chin AJ, Fellows KE, Hoffman EA. Late ventricular geometry and performance changes of functional single ventricle throughout staged Fontan reconstruction assessed by magnetic resonance imaging. J Am Coll Cardiol. 1996;28(1):212–221. 22. Fogel MA, Weinberg PM, Fellows KE, Hoffman EA. Study in ventricular–ventricular interaction: single right ventricles compared with systemic right ventricles in a dual chambered circulation. Circulation. 1995;92(2):219–230. 23. Fogel MA, Gupta KB, Weinberg PW, Hoffman EA. Regional wall motion and strain analysis across stages of Fontan reconstruction by magnetic resonance tagging. Am J Physiol Heart Circ Physiol. 1995;269(38):H1132–H1152. 24. Fogel MA, Weinberg PM, Hoydu A, et al. The nature of flow in the systemic venous pathway in Fontan patients utilizing magnetic resonance blood tagging. J Thorac Cardiovasc Surg. 1997;114:1032–1041. 25. Ho V, Prince M. Thoracic MR aortography: imaging techniques and strategies. Radiographics. 1998;18(2):287–309. 26. Beekman R, Hazekamp M, Sobotka M, et al. A new diagnostic approach to vascular rings and pulmonary slings: the role of MRI. Magn Reson Imaging. 1998;16(2):137–145. 27. van Son J, Julsrud P, Hagler D, et al. Imaging strategies for vascular rings. Ann Thorac Surg. 1994;57(3):604–610. 28. Didier D, Ratib O, Beghetti M, et al. Morphologic and functional evaluation of congenital heart disease by magnetic resonance imaging. J Magn Reson Imaging. 1999;10:639–655. 29. Fleenor JT, Weinberg PM, Kramer SS, Fogel M. Vascular rings and their effect on tracheal geometry. Pediatr Cardiol. 2003;24:430–435. 30. Fratz S, Hess J, Schuhbaeck A, et al. Routine clinical cardiovascular magnetic resonance in paediatric and adult congenital heart disease: patients, protocols, questions asked and contributions made. J Cardiovasc Magn Reson. 2008;10:46. 31. Mor-Avi V, Sugeng L, Weinert L, et al. Fast measurement of LV mass with real-time 3-dimensional echocardiography: comparison with MRI. Circulation. 2004;110:1814–1818. 32. Bu L, Munns S, Zhang H, et al. Rapid full volume data acquisition by real-time 3-dimensional echocardiography for assessment of LV indexes in children: a validation study compared with MRI. J Am Soc Echocardiogr. 2005;18:299–305. 33. Lee VS, Resnick D, Bundy JM, Simonetti OP, Lee P, Weinreb JC. MR evaluation in one breath hold with real-time true fast imaging with steady-state precession. Radiology. 2002;222:835–842. 34. Edelman RR, Chien D, Kim D. Fast selective black blood MR imaging. Radiology. 1991;181:655–660. 35. Finn JP, Baskaran V, Carr JC, et al. Low-dose contrast-enhanced threedimensional MR angiography with subsecond temporal resolution– initial results. Radiology. 2002;224:896–904.

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with in utero diagnosis of hypoplastic left heart syndrome to quantify ventricular volumes and cardiac output.61

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36. Martin ET, Coman JA, Shellock FG, Pulling CC, Fair R, Jenkins K. Magnetic resonance imaging and cardiac pacemaker safety at 1.5-Tesla. J Am Coll Cardiol. 2004;43:1315–1324. 37. Pulver AF, Puchalski MD, Bradley DJ, et al. Safety and imaging quality of MRI in pediatric and adult congenital heart disease patients with pacemakers. Pacing Clin Electrophysiol. 2009;32:450–456. 38. Nazarian S, Roguin A, Zviman MM, et al. Clinical utility and safety of a protocol for noncardiac and cardiac magnetic resonance imaging of patients with permanent pacemakers and implantable-cardioverter defibrillators at 1.5 Tesla. Circulation. 2006;114:1277–1284. 39. Sommer T, Naehle CP, Yang A, et al. Strategy for safe performance of extrathoracic magnetic resonance imaging at 1.5 T in the presence of cardiac pacemakers in non-pacemaker-dependent patients: a prospective study with 115 examinations. Circulation. 2006;114:1285–1292. 40. Ashford MW, Liu W, Lin SJ, et al. Occult cardiac contractile dysfunction in dystrophin-deficient children revealed by cardiac magnetic resonance strain imaging. Circulation. 2005;112:2462–2467. 41. Harris MA, Weinberg PM, Fogel MA. Cardiac magnetic resonance atrial level shunt detection utilizing presaturation tagging. Presented at the 4th World Congress of Pediatric Cardiology and Cardiac Surgery. Argentina: Buenos Aires; 2005. 42. Prakash A, Powell AJ, Krishnamurthy R, Geva T. Magnetic resonance imaging evaluation of myocardial perfusion and viability in congenital and acquired pediatric heart disease. Am J Cardiol. 2004;93:657–661. 43. Paetsch C, Jahnke A, Wahl R, et al. Comparison of dobutamine stress magnetic resonance, adenosine stress magnetic resonance, and adenosine stress magnetic resonance perfusion. Circulation. 2004;110:835–842. 44. Harris MA, Ghods S, Weinberg PM, Fogel MA. Magnetic resonance delayed enhancement for detection of fibrous tissue in postoperative pediatric patients with various forms of congenital heart disease. J Cardiovasc Magn Reson. 2005;7:157. 45. Taylor AM, Dymarkowski S, Hamaekers P, et al. MR coronary angiography and late-enhancement myocardial MR in children who underwent arterial switch surgery for transposition of great arteries. Radiology. 2005;234:542–547. 46. McNamara MT, Tscholakoff D, Revel D, et al. Differentiation of reversible and irreversible myocardial injury by MR imaging with and without gadolinium-DTPA. Radiology. 1986;158:765–769. 47. Simonetti OP, Kim RJ, Fieno DS, et al. An improved MRI technique for the visualization of myocardial injury. Radiology. 2001;218:215–223. 48. Kim WY, Danias PG, Stuber M, et al. Coronary magnetic resonance angiography for the detection of coronary stenoses. N Engl J Med. 2001;345:1863–1869.

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49. Tangcharoen T, Hegde S, Bell A, et al. Feasibility of whole-heart, steady-state free precession magnetic resonance coronary angiography (MRCA) in infants and children with congenital heart disease [abstract 219]. J Cardiovasc Magn Resonan. Abstracts from the 11th annual SCMR scientific sessions. 2008;80–81. 50. Kiaffas MG, Powell AJ, Geva T. Magnetic resonance imaging evaluation of cardiac tumor characteristics in infants and children. Am J Cardiol. 2002;89:1229–1233. 51. Basso C, Thiene G, Corrado D, Angelini A, Nava A, Valente M. Arrhythmogenic right ventricular cardiomyopathy: dysplasia, dystrophy or myocarditis. Circulation. 1996;94:983–991. 52. McKenna WJ, Thiene G, Nava A, et al. Diagnosis of arrhythmogenic right ventricular dysplasia/cardiomyopathy. Task Force of the Working Group Myocardial and Pericardial Disease of the European Society of Cardiology and of the Scientific Council on Cardiomyopathies of the International Society and Federation of Cardiology. Br Heart J. 1994;71:215–218. 53. Winter MM, Bernink FJP, Groenink M, et al. Evaluating the systemic right ventricle by CMR: the importance of consistent and reproducible delineation of the cavity. J Cardiovasc Magn Reson. 2008;10:40. 54. Fogel MA, Weinberg PM, Rhodes L. Usefulness of magnetic resonance imaging for the diagnosis of right ventricular dysplasia in children. 2006;97(8):1232–1237. 55. Midiri M, Finazzo M, Brancato M, et al. Arrhythmogenic right ventricular dysplasia: MR features. Eur Radiol. 1997;7:307–312. 56. Fontan F, Baudet E. Surgical repair of tricuspid atresia. Thorax. 1971;26:240–248. 57. Fogel MA, ed. Ventricular Function and Blood Flow in Congenital Heart Disease. Malden MA: Blackwell Futura; 2005. 58. Fogel MA. Cardiac magnetic resonance of single ventricles. J Cardiovasc Magn Reson. 2006;8(4):661–670. 59. Harris M, Johnson T, Weinberg P, Fogel M. Delayed enhancement cardiovascular magnetic resonance identifies fibrous tissue in children after congenital heart surgery. J Thorac Cardiovasc Surg. 2007; 133:676–681. 60. Steffens JC, Bourne MW, Sakuma H, O’Sullivan M, Higgins CB. Quantification of collateral blood flow in coarctation of the aorta by velocity encoded cine magnetic resonance imaging. Circulation. 1994; 90:937–943. 61. Fogel MA, Wilson DR, Flake A, et al. A new method of functional assessment of the fetal heart using a novel application of “real time” cardiac magnetic resonance imaging. Fetal Diagn Ther. 2005;20:475–480.

Use of Navigator Echoes in Cardiovascular Magnetic Resonance and Factors Affecting Their Implementation David Firmin and Jennifer Keegan

Respiration has been shown to be an important factor influencing the quality of cardiovascular magnetic resonance (CMR) images. In addition to cardiac motion, which can be addressed reasonably well by electrocardiographic (ECG) triggering, respiratory motion moves the position and distorts the shape of the heart by several millimeters between inspiration and expiration. In 1991, Atkinson and Edelman1 showed the detrimental effects of breathing on the quality of cardiac studies by showing improved detail (fast low angle shot) in breath hold segmented fast gradient echo images compared with conventional nonbreath-hold images. Although breath holding produces images that are free of respiratory motion artifact, it is not without problems. The breath hold position may vary from one breath hold scan to the next, giving rise to misregistration effects, and it may also vary during the breath hold period itself,2 resulting in image blurring and artifacts. In addition, the scan parameters are limited by the need to perform imaging within the duration of a comfortable breath hold period, and for a number of patients, this period may be very short. An alternative to breath holding is to monitor respiratory motion throughout the data acquisition period and to correct the data for that motion, either in real time or through postprocessing, with the efficacy of both techniques being strongly dependent on the accuracy of the method of motion assessment. During normal tidal respiration, the superior-inferior (SI) motion of the diaphragm is approximately four to five times the anterior-posterior motion of the chest wall,3 and consequently, diaphragm motion is the most sensitive measure of respiratory motion. In 1989, Ehman and Felmlee4 were the first to introduce navigator echoes for measuring the displacement of a moving structure and to demonstrate their use in determining diaphragm motion during respiration. The navigator echo is the signal from a column of material oriented perpendicular to the direction of the motion to be monitored. On Fourier transformation, this signal results in a welldefined edge of the moving structure. The navigator echoes may be interleaved with the imaging sequence and consequently enable the motion to be determined throughout the data acquisition period.

In CMR, there have been a number of developments in the use of navigator measurement to reduce the problems of respiratory motion. This chapter discusses these developments, considers the various choices that have been studied in the implementation of navigators, and discusses their importance. There are many variables in the application and use of navigator echoes, and although there have been some attempts to study these, it is unlikely that we are close to optimizing their application.

USE OF NAVIGATOR INFORMATION There are two distinct ways of using navigator echoes to reduce the problems of respiratory motion in CMR, which are multiple breath holding with feedback and free breathing methods. The first of these uses the navigator information to provide visual feedback on the diaphragm position to subjects to allow them to hold their breath at the same point repeatedly.5 The second uses the navigator echo measurement as an input to some form of respiratory gating algorithm while the patient breathes normally. Figure 10-1 shows actual respiratory trace data in a subject when performing multiple breath holds and when free breathing. In both cases, a navigator acceptance window, typically 5 mm wide, is defined, and all data acquired when the navigator is outside of this window are ignored. The resulting image therefore consists of data acquired over a narrow range of respiratory positions. The respiratory or scan efficiency is defined as the percentage of ECG triggers that fall within the navigator acceptance window and is a measure of the data rejection rate, which in turn determines the overall scan duration. As the navigator acceptance window is reduced, the rejection rate increases and the scan efficiency decreases. Figure 10-2 shows the residual diaphragm displacements that occurred during data acquisitions performed during conventional breath holding, breath holding with navigator feedback, and navigator free breathing in normal subjects.6 Both navigator techniques result in images acquired over a reduced range of Cardiovascular Magnetic Resonance 129

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CHAPTER 10

Standard deviation

0

±0.7 mm

±0.7mm

30

−20

20 −40

−60

A

0

Time (s)

60

0

Diaphragm displacement (mm)

Diaphragm displacement (mm)

±2.4 mm

Diaphragm displacement (mm)

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

5mm NE acceptance window

10

0

−10 BH

−20

LED

−20

+ −40

B Figure 10-1 Navigator echo respiratory trace data during multiple breath holding with navigator feedback (A) and free breathing (B). In each case, the shaded region shows the position of a 5-mm navigator acceptance window outside of which data is rejected. (Adapted from Taylor AM, Keegan J, Jhooti P, Gatehouse PD, Firmin DN, Pennell DJ. Differences between normal subjects and patients with coronary artery disease for three different MR coronary angiography respiratory suppression techniques. J Magn Reson Imaging. 1999;9:786–793, with permission.)

diaphragm displacement compared with those acquired using repeated conventional breath holding. In addition, they allow a longer overall scan time. This allows for averaging of data to improve the signal-to-noise ratio, increasing the k-space coverage for improved spatial resolution and increasing the temporal resolution by reducing the number of individual image views acquired per cardiac cycle.

Multiple Breath Hold Methods Wang and colleagues7 were the first to show the use of a respiratory feedback monitor to reduce misregistration artifacts in consecutive breath hold segmented fast gradient echo coronary artery images and to show improved image quality from averaging scans acquired over multiple breath holds. When used in informed healthy volunteers, this technique has been shown to produce good results with reasonable scan efficiency.8 However, a period of training is required, and the process can be problematic, particularly with patients who have difficulty holding their breath because of a combination of illness and anxiety.6 Although 130 Cardiovascular Magnetic Resonance

FR

−30 Figure 10-2 Mean diaphragm displacement in 17 normal subjects with conventional breath holding (open circles), breath holding with navigator feedback (closed circles), and free breathing (plus marks). The navigator-controlled studies used a 5-mm navigator acceptance window. (Adapted from Taylor AM, Keegan J, Jhooti P, Gatehouse PD, Firmin DN, Pennell DJ. Differences between normal subjects and patients with coronary artery disease for three different MR coronary angiography respiratory suppression techniques. J Magn Reson Imaging. 1999; 9:786–793, with permission.)

it might be expected that breath holding with respiratory feedback would enable the completion of a cardiac study much more quickly than when using the free breathing methods described later, because of the time required for training and the required rest periods between breath holds, the overall examination times are longer than anticipated. In fact, in a group of patients with coronary artery disease, there was no significant difference between the overall examination times with the two techniques,6 although the same study showed that, in a group of normal healthy subjects, multiple breath holding resulted in a time reduction of 20%.

Free Breathing Methods Free breathing methods require very little cooperation from the patient. The main disadvantage is the potential for respiratory drift, which can cause considerably reduced scan efficiency.9 Recently, therefore, most effort has gone toward improving scan efficiency with this approach. Much of the early work used retrospective respiratory gating.10 With this method, data acquisition was oversampled, typically by a factor of five, and then sorted retrospectively so that the final image was constructed from data acquired over the narrowest possible range of respiratory

much greater scan efficiency than other methods, while retaining scan quality (Table 10-1 and Figure 10-3). An alternative method, initially developed by Sinkus and Bornert to address general respiratory motion16 and more recently applied to imaging of the coronary arteries, used a tailored acceptance window through k-space as opposed to phase encode ordering to obtain a very similar result.17 Both of these phase ordering or windowing techniques use a predefined navigator acceptance window, and scan efficiency is reduced when the respiratory pattern changes during study acquisition. This has been more recently addressed by Jhooti and colleagues, who developed a technique that combines the benefits of phase ordering with an automatic window selection that enables the highly efficient acquisition of high-quality coronary artery images without the need for a predefined acceptance window.18 Three-dimensional motion-adapted gating19 is a similar technique that yields images comparable to standard prospective navigator gating, with significantly improved scan efficiency.20

NAVIGATOR ECHO IMPLEMENTATION Method of Column Selection Two methods have been used for the generation of a navigator echo. With the spin echo technique, a spin echo signal is generated from the column of material formed by the intersection of two planes, one excited by a 90 radiofrequency (RF) pulse and the other by a 180 RF pulse. The column cross-section may be either rectangular or rhombic, depending on the orientation of the two planes. This approach is very robust and produces an extremely welldefined column. However, it cannot be repeated rapidly, and care must be taken to ensure that the column selection planes do not impinge on the region of interest. The alternative approach is to use a selective twodimensional (2D) RF pulse to excite a column of approximately circular cross-section.21 Although this technique is much more sensitive to factors such as shimming errors, which can potentially cause blurring and distortion of the column, with a reduced flip angle, it can be repeated more rapidly and the navigator artifact is less extensive.

Table 10-1 Image Quality Scores and Scan Efficiencies{ for Three-Dimensional Magnetic Resonance Angiography* Image Quality Mean Score Scan Efficiency

Phase Ordered

ARA

DVA{

RRG

4.4 72 ( 11.6)

4.7 48 ( 11.5)

6.6 72 ( 11.6)

6.8 20

*Mean image quality scores (1 ¼ excellent, 10 ¼ very poor) and scan efficiencies{ ( SD) for data acquired using phase ordering, an accept/reject algorithm (ARA), the diminishing variance algorithm (DVA), and retrospective respiratory gating (RRG) in 15 subjects. { Scan time for the DVA technique is set to that of the phase ordered technique. (Adapted from Jhooti P, Keegan J, Gatehouse PD, et al. 3D coronary artery imaging with phase reordering for improved scan efficiency. Magn Reson Med. 1999; 41: 555.)

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positions. In 1995, Hofman and associates11 showed that, using this approach, the image quality of three-dimensional (3D) coronary acquisitions was improved over those acquired with multiple averages. However, scan efficiency was poor (20% for an oversampling ratio of five), and although the final image was constructed from the narrowest respiratory window possible, the range was highly dependent on the subject’s breathing pattern during the long acquisition period and was often still unacceptably high. After the introduction of prospective control techniques, navigators have most commonly been used with a simple accept-reject algorithm where data are acquired or not, depending on whether the navigator measurement is within a predefined acceptance window. Oshinski and coworkers12 were the first investigators to show high-quality coronary artery images with such an approach. The problem with this method, however, is that for reasonably high scan quality, a narrow acceptance window of 5 mm or less is required, and this generally results in relatively poor scan efficiency. In addition, as noted earlier, many subjects and patients undergo a drift in diaphragm position over time,9 such that the predefined acceptance window becomes less and less suitable as the study progresses. The diminishing variance algorithm overcomes this problem because it does not use a predefined acceptance window.13 With this method, one complete scan is acquired and the navigator positions are saved for each line of data. At the end of the initial scan, the most frequent diaphragm position during that scan is determined, and a process of reacquiring lines of data that were acquired with diaphragm positions furthest offset from this position begins. As time progresses, the range of diaphragm positions for the data making up the final set is considerably reduced. In addition to the lack of requirement of an acceptance window, this method has the advantage that an image can be reconstructed at any time after the initial dataset is complete. Another alternative to the simple accept-reject algorithm that can improve both image quality and scan efficiency is to use a k-space ordering that depends on diaphragm position. Two similar approaches have been suggested, based on the finding that the center of k-space appears to be more sensitive to motion than the edges.14 Jhooti and colleagues developed a phase encode ordered method that used a dual acceptance window of 5 mm for the center of k-space and 10 mm for the outer regions.15 This approach allowed

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

A

C DVA (scan eff. = 73%)

Phase reordered (scan eff. = 73%)

B

Figure 10-3 A single slice from a three-dimensional dataset showing a long section of the right coronary artery. The phase ordered images (A) are of comparable quality to those acquired with the acceptreject algorithm (B) and better than those acquired with both the diminishing variance algorithm (C) and retrospective respiratory gating (D). Scan efficiency is also significantly higher for phase ordering than for both the acceptreject algorithm and retrospective respiratory gating techniques. (Adapted from Jhooti P, Keegan J, Gatehouse PD, Collins S, Rowe A, Taylor AM, Firmin DN. 3D coronary artery imaging with phase reordering for improved scan efficiency. Magn Reson Med 1999. 41:555–562, with permission.)

D ARA (scan eff. = 49%)

RRG (scan eff. = 20%)

Both methods are used routinely for research studies on coronary imaging, without any reported problems.

Correction Factors In CMR, navigator echoes are most frequently used to measure the position of the diaphragm. However, the motion of the heart is not straightforward, and only the inferior border that sits on the diaphragm will move to the same extent, whereas superiorly, the relative motion will be reduced. This was first studied by Wang and associates,3 who measured the displacement of the right coronary artery root, the origin of the left anterior descending artery, and the superior and inferior margins of the heart relative to the diaphragm in 10 healthy subjects. For the right coronary artery origin, the mean ( SD) relative displacement (or correction factor) was 0.57 ( 0.26). McConnell and coworkers22 first used this correction factor to track the position of the imaging slice during breath holding and showed improved image registration relative to untracked acquisitions. In free breathing studies, the correction factor was first applied by Danias and colleagues,23 who showed that tracked image quality was maintained as the navigator acceptance window increased from 3 mm to 7 mm, whereas in untracked images, it decreased significantly. This technique, called real-time prospective slice following, 132 Cardiovascular Magnetic Resonance

or slice tracking, is now used routinely for both 2D and 3D methods of acquisition. Of note, however, is the relatively high standard deviation of the correction factor noted earlier, which reflects considerable intersubject variation in the degree of cardiac motion with respiration. This was also observed by Danias and coworkers, who used real-time 2D echo planar imaging to study the SI motion of the heart as a function of navigator position.24 The accuracy of slice-following techniques will obviously depend on the accuracy of the correction factor implemented. In 1997, Taylor and colleagues25 showed how a subject-specific factor could be measured rapidly with end-inspiratory and end-expiratory breath hold scans before the coronary imaging protocol. Figure 10-4 shows the relationship between the motion of the right hemi-diaphragm and the coronary ostia measured in one subject, with the slope of the graph giving the correction factor. Figure 10-5 shows two examples of subjects with very different correction factors, showing how a wider acceptance window can be used, thus improving scan efficiency. The need for a subject-specific correction factor has further been confirmed in 3D coronary angiography, where its use was found to yield optimal image quality.26 In 2002, Keegan and associates further developed this area of work by studying the variability of correction factors in the SI, anterior-posterior, and right-left directions for both breath holding and free breathing.27 The study concluded that subject variability in correction

Downward coronary displacement (mm)

y = –0.08 – (0.45 x) r = 0.99

4

8

12

16 0

10

20

30

Downward diaphragm displacement (mm) Figure 10-4 Plot of superior-inferior right coronary artery displacement against superior-inferior diaphragm displacement for a single subject. The gradient of the linear regression line is the subjectspecific correction factor. (Modified from Taylor AM, Keegan J, Jhooti P, Firmin DN, Pennell DJ. Calculation of a subject-specific adaptive motion-correction factor for improved real-time navigator echo-gated magnetic resonance coronary angiography. J Cardiovasc Magn Reson. 1999; 1:131–138, with permission.)

factors, together with within-subject differences between breath holding and free breathing, is such that slice following should be performed with subject-specific factors determined from free breathing acquisitions. An additional or alternative approach to the real-time slice following described earlier is to use a postprocessing adaptive motion correction technique4 to correct an image retrospectively for movement occurring during data acquisition. This technique, which can be used to correct a 2D acquisition for in-plane displacement or a 3D acquisition for inplane and through-plane displacement, may not appear to be an attractive option initially, but it has the advantages of allowing the correction factor to be optimized for each individual patient and provides an alternative approach to those centers with scanners that do not have a real-time decision making capability. This approach has been implemented with both segmented gradient echo28 and interleaved spiral29 coronary artery acquisitions, with promising results.

Column Positioning The degree of diaphragm motion detected by the navigator echo is dependent on the positioning of the navigator column. The dome of the right hemi-diaphragm is higher than that of the left, and the two move coherently with respiration, but to differing degrees.30 Motion of the diaphragm is also greater posteriorly than anteriorly (anterior and dome excursions are 56% and 79%, respectively, of posterior excursions), and at the level of the dome, it is greater laterally than medially.31 The correction factor implemented in real-time slice following or postprocessing adaptive motion correction, as described earlier, is strongly dependent on the positioning of the navigator column and further

supports the use of a subject-specific factor, as described in the previous section. McConnell and colleagues32 studied the effects of varying the navigator location on the image quality of coronary artery studies. Navigators were positioned through the dome of the right hemi-diaphragm, through the posterior portion of the left hemi-diaphragm, through the anterior and posterior left ventricular walls, and through the anterior left ventricular wall, as shown in Figure 10-6. The advantage of the latter navigator position is that it would eliminate the need for a correction factor, as described in the previous section, relating the navigator-echomeasured displacement to the coronary artery motion. The results are summarized in Table 10-2 and show no significant differences in the image quality scores obtained with varying navigator location. There was a tendency for the anterior left ventricular wall navigator scans to be longer in duration, but the difference did not reach statistical significance. One of the problems with monitoring the heart itself is the complex anatomy, making it more difficult to find a suitable position for the navigator column. More sophisticated methods of positioning the column may further improve this method of monitoring cardiac motion.

MULTIPLE COLUMN ORIENTATIONS There is a linear relationship between the SI and anteriorposterior motions of the heart, with the SI motion being approximately five times that of the anterior-posterior motion.3 For this reason, the real-time slice-following methods first used by McConnell and colleagues22 and by Danias and associates23 included a correction for anterior-posterior motion of the heart, assuming it to be equal to 20% of the SI motion. Unfortunately, there is not always such a strong relationship between the directions of motion of the heart with respiration. Sachs and colleagues showed this by using three navigators to measure the SI, anterior-posterior, and right-left motions of the heart.33 Figure 10-7 shows an example from this study illustrating the scatter of SI, right-left, and anteriorposterior measurements, made over a period of approximately 10 minutes. The group went on to compare the use of one, two, and three navigators for imaging the right coronary artery and showed an improvement when multiple directions of motion were considered. This improvement in image quality, however, must be offset against the main disadvantage, which is that scan efficiency is reduced, potentially introducing more problems associated with long-term drift in the breathing pattern. A more recent study by Jahnke and coworkers used a new cross-correlation-based approach and showed the potential advantage of combining three orthogonal navigators.34

Navigator Timing Navigator timing is one of the more important parameters; however, flexibility to alter this is often limited by the computing architecture of the scanner being used (discussed Cardiovascular Magnetic Resonance 133

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0

6 mm 16 mm

0.70

0.00

1.00

CF

0.25

0.00

1.00

6 mm

CF

16 mm

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

A

B Figure 10-5 Right coronary artery origin images acquired with navigator acceptance windows of 6 mm and 16 mm in subjects with subjectspecific correction factors (CFs) of 0.7 (A) and 0.25 (B). For both subjects, images were also acquired with CFs of 0 and 1. In the absence of slice following (CF ¼ 0), image quality is reduced as the navigator acceptance window increases from 6 mm to 16 mm. When slice following with a subject-specific CF is used, however, image quality is maintained. (Modified from Taylor AM, Keegan J, Jhooti P, Firmin DN, Pennell DJ. Calculation of a subject-specific adaptive motion-correction factor for improved real-time navigator echo-gated magnetic resonance coronary angiography. J Cardiovasc Magn Reson. 1999; 1:131–138, with permission.)

134 Cardiovascular Magnetic Resonance

C

B

D

Figure 10-6 Navigator column locations positioned on tranverse, coronal and sagittal pilot images: (A) through the dome of the right hemidiaphragm, (B) through the posterior left hemi-diaphragm, (C) through both anterior and posterior left ventricular walls and (D) through the anterior left ventricular wall.

Table 10-2 Image Quality Scores, Registration Errors, and Total Scan Times* Parameter Image Quality Score (0–4) Registration error (mm) Craniocaudal Anteroposterior Total Scan Time{ (sec)

Right Diaphragm Navigator

Left Diaphragm Navigator

Left Ventricle Navigator

Anterior LV Wall Navigator

2.3 0.1

2.3 0.1

2.4 0.1

2.2 0.2

0.5 0.1 0.3 0.1 294 28

0.4 0.1 0.3 0.1 314 30

0.6 0.1 0.3 0.1 342 62

0.4 0.1 0.4 0.1 427 111

*Image quality scores (0 ¼ very poor, 4 ¼ excellent) registration errors, and total scan times for different navigator column positions during free breathing MR coronary angiography. There were no significant differences between the navigator column locations. Data are presented as mean standard error of the mean (SEM); LV ¼ left ventricle { Total scan time is the time from start to finish for 6 scans. Adapted from McConnell MV, Khasgiwala VC, Savord BJ, et al. Comparison of respiratory suppression methods and navigator locations for MR coronary angiography. Am J Roentgenol. 1997;168:1369.

Figure 10-7 Anterior-posterior (A/P; A) and right-left (R/L; B) navigator echo measurements as a function of superior-inferior navigator echo measurements in a healthy subject. (Data provided by Todd Sachs, Stanford University.)

80 60

R/L

A/P

40

A

0

S/I

later). Figure 10-8 shows the three main alternatives: (1) pre-, (2) pre- and post-, and (3) navigators repeated regularly throughout the cardiac cycle. A simple prenavigator provides the highest scan efficiency when a navigator acceptance window is used, but may not be reliable if there is a sudden change in breathing between the navigator measurement and image data acquisition. Pre- and

150

B

0

S/I

150

postnavigators overcome this problem, but of course, they also reduce scan efficiency. In our experience, the use of prenavigators only produces acceptable results for free breathing studies, whereas multiple breath hold acquisitions certainly require both pre- and postnavigators. An important factor that depends on the computer hardware and architecture is the time required after the navigator Cardiovascular Magnetic Resonance 135

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A

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

D. acq. PrePre- and postD. acq.

D. acq.

D. acq.

D. acq.

Rep. Figure 10-8 Timing of the navigators for pre-, pre- and post-, and repeated (Rep.) navigator echo-controlled data acquisition (D. acq.).

acquisition before the start of the imaging sequence. Particularly if prenavigators only are being used, the longer this interval, the greater the potential for errors caused by respiratory motion. Also, for ECG R-wave-triggered scans, this may also have implications for the minimum gating delay that can be obtained and for short gating delay or cine scans, post-only navigators can be used as an alternative. This approach was recently implemented in left ventricular function studies, where it was found that image quality in a group of patients with heart failure was significantly improved over conventional breath hold scans.35 Repeated navigators allow for improved cine or multislice imaging and also provide some potential for estimating internavigator respiratory motion. The potential problem with this is that the navigator signal-to-noise ratio could be reduced and this may affect the accuracy of navigator edge detection. In addition, as the time for navigator output increases, the time for imaging decreases and the number of phases or slices that can be acquired is reduced.

Precision of Navigator Measurement Commonly, a spatial resolution of 1 mm is used along the navigator echo column, for example, having a field of view of 512 mm and sampling 512 points on the navigator readout. However, the precision of the measurement is dependent to a large extent on the signal-to-noise ratio of the navigator measurement. The most important factor affecting the signal-to-noise ratio is the coil arrangement. If, for example, a single coil is used for imaging and navigator detection, it must be large enough to cover both the imaging area of interest and the region of navigator detection. On the other hand, if phased array coils are used, it is possible to position one coil specifically for navigator detection, possibly over the region of the right diaphragm. Another important factor in the precision of the measurement is the quality of the edge on the navigator trace. To obtain a well-defined edge of the diaphragm, for example, it is important to have a reasonably small column cross-section and to position it through the dome of the diaphragm, so that the column is perpendicular to the diaphragm edge, rather than more posteriorly, where motion is greatest. 136 Cardiovascular Magnetic Resonance

Finally, the diaphragm edge may be detected by edge detection, correlation, or least squares fit algorithms. For rapid tracking (repetition time < 100 msec) or narrower columns, the signal-to-noise ratio in the diaphragm trace could be too poor for simple edge detection methods to succeed. Of the remaining two techniques, the least squares fit method has been shown to be more resistant to the effects of noise and to the diaphragm profile deformation that occurs during respiration than the correlation method and therefore would be the technique of choice.36 However, most navigator techniques acquire only one or two navigators per cardiac cycle, and in such cases, the signalto-noise ratios are usually relatively high and edge detection algorithms are generally adequate.

MORE RECENT APPROACHES Other Forms of Navigators As has been mentioned, there are problems with the conventional navigators that have been described earlier because they generally do not give a direct measure of the respiratory-related motion of the heart and they cannot be implemented simply and efficiently to give a measure of this motion in 3D.37 A number of ingenious alternative approaches have been described. In 2003, Nguyen and colleagues38 described a method that selectively excited the epicardial fat, followed by a rapid readout scheme that instantaneously gave three 1D images of its position in the x, y, and z directions. Tested on six subjects, the method showed a slight improvement over conventional navigators; however, the authors noted a number of problems that would need to be resolved before it could be in routine use. In the same year, another method of rapidly localizing heart signals for measurement of its position was suggested by Pai and Wen, who used a phase contrast angiographic approach to selectively image the flowing blood in the heart chambers.39 These blood signals were used to define the heart position in the SI and anterior-posterior directions. Despite the potential advantages of truly tracking the heart position, this method does not appear to have been developed further or used. Subsequently, Stehning and associates used radial imaging for “self-navigation.” They developed an interleaved 3D radial acquisition modified in such a way that the first readout was always in the SI direction.40 This readout could then be reconstructed every cardiac cycle to give an SI projection for motion extraction. In this work, the authors showed improved definition compared with conventional navigators when imaging a moving phantom and similar image quality on initial in vivo coronary scans. Hardy and colleagues41 used cross-correlation of low-resolution real-time 2D spiral coronary artery images to accept or reject images for averaging. This adaptive averaging technique was extended to highresolution segmented acquisitions by cross-correlating subimages reconstructed from individual data segments. Breathing autocorrection with spiral interleaves (BACSPIN) is a similar technique42 that involves the acquisition of a multislice spiral dataset during breath holding, followed by repeated acquisition of the same slices during free breathing.

Motion Models Because of the complexity of cardiac motion and the difficulty in extracting measures to correct for it, there has been an interest in developing modeling methods to assist with this process. Manke and coworkers compared onedimensional (1D) translation (SI direction), 3D translation, and 3D affine transformation motion models in a group of healthy subjects.44 By using an elastic image registration algorithm on 3D coronary images acquired at different breath hold positions, the superiority of the 3D translational model over the 1D translational model was shown. The authors also used a fast model-based image registration to extract motion information from a time series of lowresolution 3D images. This was used in conjunction with conventional navigators to calibrate a respiratory motion model that allowed the prediction of affine transformation parameters, including 3D translation, rotation, scale, and shear motion from the navigator measurements, and was shown with coronary artery imaging.45 McLeish and colleagues46 acquired images at a number of breath hold positions and studied the accuracy of both rigid and nonrigid registration methods in registering the other breath hold images with those acquired at end-expiration. They used principal component analysis to produce patient-specific statistical motion models and suggest how this could be used to assist motion correction in CMR. In 2005, Nehrke and Bornert described their study in which they used a patient-specific model to control the acquisition.47 Initially, multiple low-resolution 3D images were preacquired during free breathing. An affine model of the respiratory-related cardiac motion was then extracted from these images, and this was steered by real-time navigators to control the high-resolution acquisition. The method was demonstrated in both phantoms and volunteers when a 20-mm navigator acceptance window was used. The results in the volunteers showed slightly inferior image quality to images acquired simply with a 5-mm navigator acceptance window. However, scan efficiency was considerably higher

and image quality was also considerably better than for those acquired simply with a 20-mm acceptance window. A more recent study showed that the residual coronary artery motion observed using affine navigators with a 10mm acceptance window is similar to that observed with conventional navigator gating with a 5-mm window and that observed using a single breath hold.48

Computer Architecture The computer architecture of modern CMR scanners can be very complex, generally incorporating three main computers. The host computer runs the user interface and allows connection to the image database, the reconstruction computer is a dedicated rapid processor for reconstruction of the CMR image data, and the scan computer allows control and adjustment of parameters associated with the scanning sequence. The architecture of these computers can significantly affect the potential and usefulness of navigator echoes. On many systems, for example, the navigator signal must be reconstructed and processed on the reconstruction computer, but the measurement made must be passed through the host to control the parameters on the scan computer. This arrangement inevitably adds a variable and unknown delay that is dependent on other tasks being performed by the host operating system. To overcome this, either a direct and rapid link is required, allowing transfer of data from the reconstruction computer to the scan computer, or the scan computer itself must be capable of acquiring and reconstructing the navigator data, so that no data transfer is required. Newer scanners are generally being designed with rapid acquisition, processing, and control in mind.

CONCLUSION Navigator echo has been shown to be an important method for monitoring respiration that has been used for defining the position of the heart, enabling improved coronary and other cardiac imaging. The limited number of studies and the many parameters and variables involved in their use suggest that an optimal method of application may not yet have been developed. With future system development, there will be minimal cost, in imaging time or other factors, involved in obtaining this positional information. Therefore, it would seem worthwhile to collect and use it where appropriate. The methods are not robust, probably because of their relative lack of sophistication. One of the major advantages of this technique is that navigators allow images to be acquired during free respiration, eliminating the need for patient cooperation. They also allow longer acquisition times, enabling higher spatial and temporal resolution and increasing the potential for more sophisticated techniques, such as detailed flow imaging.49,50 A balance must be maintained, however, and imaging time should not be increased so much that increased respiratory drift cancels any potential benefit.

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Each heavily undersampled free breathing spiral interleaf is compared with the same interleaf acquired during breath holding, and those that match closely are incorporated into a multi-slice, multi-average dataset. This technique has been shown in a group of six healthy volunteers where increased signal-to-noise ratio has been achieved with minimal motion blurring. An alternate approach to respiratory motion correction that has been applied to spiral imaging is to acquire a low-resolution 3D dataset on the fat resonance immediately before a high-resolution interleaf on the water resonance. Cross-correlation of a selected region of interest of the low-resolution fat images from beat to beat is then used to determine the x, y, and z translations of that region and can be used to correct the next high-resolution interleaf retrospectively. This approach has been shown in darkblood coronary vessel wall imaging, where high-quality images have been obtained without the need for a gating window.43

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24. Danias PG, Stuber M, Botnar RM, Kissinger KV, Edelman RR, Manning WJ. Relationship between motion of coronary arteries and diaphragm during free breathing: lessons from real-time MR imaging. Am J Roentgenol. 1999;172:1061–1065. 25. Taylor AM, Keegan J, Jhooti P, Firmin DN, Pennell DJ. Calculation of a subject-specific adaptive motion-correction factor for improved realtime navigator echo-gated magnetic resonance coronary angiography. J Cardiovasc Magn Reson. 1999;1:131–138. 26. Nagel E, Bornstedt A, Schnackenburg B, Hug J, Oswald H, Fleck E. Optimisation of real-time adaptive navigator correction for 3D magnetic resonance coronary angiography. Magn Reson Med. 1999;42: 408–411. 27. Keegan J, Gatehouse P, Yang GZ, Firmin D. Coronary artery motion with the respiratory cycle during breath-holding and free-breathing: implications for slice-followed coronary artery imaging. Magn Reson Med. 2002;47:476–481. 28. Wang Y, Ehman RL. Retrospective adaptive motion correction for navigator-gated 3D coronary MR angiography. J Magn Reson Imaging. 2000;11:208–214. 29. Keegan J, Gatehouse PD, Yang GZ, Firmin DN. Adaptive motion correction of interleaved spiral images and velocity maps: implications for coronary imaging. Proceedings of the 15th Annual Meeting of the ESMRMB. MAGMA. 1998;6(suppl 1):67. 30. Korin HW, Ehman RL, Riederer SJ, Felmlee JP, Grimm RC. Respiratory kinematics of the upper abdominal organs: a quantitative study. Magn Reson Med. 1992;23:172–178. 31. Gierada DS, Curtin JJ, Erickson SJ, Prost RW, Strandt JA, Goodman LR. Diaphragmatic motion: fast gradient recalled echo MR imaging in healthy subjects. Radiology. 1995;194:879–884. 32. McConnell MV, Khasgiwala VC, Savord BJ, et al. Comparison of respiratory suppression methods and navigator locations for MR coronary angiography. AJR Am J Roentgenol. 1997;168:1369–1375. 33. Sachs TS, Meyer CH, Pauly JM, Hu BS, Nishimura DG, Macovski A. The real-time interactive 3D-DVA for robust coronary MRA. IEEE Trans Med Imag. 2000;19:73–79. 34. Jahnke C, Nehrke K, Paetsch I, et al. Improved bulk myocardial motion suppression for navigator-gated coronary magnetic resonance imaging. J Magn Reson Imaging. 2007;26:780–786. 35. Bellenger NG, Gatehouse PD, Rajappan K, Keegan J, Firmin DN, Pennell DJ. Left ventricular quantification in heart failure by cardiovascular MR using prospective respiratory navigator gating: comparison with breath-hold acquisition. J Magn Reson Imaging. 2000;11:411–417. 36. Wang Y, Grimm RC, Felmlee JP, Riederer SJ, Ehman RL. Algorithms for extracting motion information from navigator echoes. Magn Reson Med. 1996;36:117–123. 37. Nehrke K, Bornert P, Manke D, Bock JC. Free-breathing cardiac MR imaging: study of implications of respiratory motion: initial results. Radiology. 2001;220:810–815. 38. Nguyen TD, Nuval A, Mulukutla S, Wang Y. Direct monitoring of coronary artery motion with cardiac fat navigator echoes. Magn Reson Med. 2003;50:235–241. 39. Pai VM, Wen H. Rapid-motion-perception based cardiac navigators: using the high flow blood volume as a marker for the position of the heart. J Cardiovasc Magn Reson. 2003;5:531–543. 40. Stehning C, Bornert P, Nehrke K, Eggers H, Stuber M. Free-breathing whole-heart coronary MRA with 3D radial SSFP and self-navigated image reconstruction. Magn Reson Med. 2005;54:476–480. 41. Hardy CJ, Saranathan M, Zhu Y, Darrow RD. Coronary angiography by real-time MRI with adaptive averaging. Magn Reson Med. 2000;44:940–946. 42. Hardy CJ, Zhao L, Zong X, Saranathan M, Yucel EK. Coronary MR angiography: respiratory motion correction with BACSPIN. J Magn Reson Imaging. 2003;17:170–176. 43. Keegan J, Gatehouse PD, Yang GZ, Firmin DN. Non-model-based correction of respiratory motion using beat-to-beat 3D spiral fat-selective imaging. J Magn Reson Imaging. 2007;26:624–629. 44. Manke D, Rosch P, Nehrke K, Bornert P, Dossel O. Model evaluation and calibration for prospective respiratory motion correction in coronary MR angiography based on 3D image registration. IEEE Trans Med Imaging. 2002;21:1132–1141. 45. Manke D, Nehrke K, Bornert P. Novel prospective respiratory motion approach for free breathing coronary angiography using patientadapted affine motion model. Magn Reson Med. 2003;50:122–131.

49. Nagel E, Bornstedt A, Hug J, Schnackenburg B, Wellnhofer E, Fleck E. Noninvasive determination of coronary blood flow velocity with magnetic resonance imaging: comparison of breath-hold and navigator techniques with intravascular ultrasound. Magn Reson Med. 1999;41: 544–549. 50. Keegan J, Gatehouse P, Mohiaddin RH, Yang GZ, Firmin D. Comparison of spiral and FLASH phase velocity mapping, with and without breath-holding, for the assessment of right and left coronary artery blood flow velocity. J Magn Reson Imaging. 2004;20:953–960.

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46. McLeish K, Hill DL, Atkinson D, Blackall JM, Razavi R. A study of the motion and deformation of the heart due to respiration. IEEE Trans Med Imaging. 2002;21:1142–1150. 47. Nehrke K, Bornert P. Prospective correction of affine motion for arbitrary MR sequences on a clinical scanner. Magn Reson Med. 2005;54: 1130–1138. 48. Fischer RW, Botnar RM, Nehrke K, Boesiger P, Manning WJ, Peters DC. Analysis of residual coronary artery motion for breath hold and navigator approaches using real-time coronary MRI. Magn Reson Med. 2006;55:612–618.

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

CHAPTER 11

Normal Cardiac Anatomy, Orientation, and Function Michael L. Chuang, Warren J. Manning, and Ronald M. Peshock

Cardiovascular magnetic resonance (CMR) can be used to obtain images of the heart in any plane. Thus, to define normal anatomy and function, it is important to define standard imaging planes in order to develop knowledge of normal anatomy, anatomic variants, and potential artifacts. Standard CMR planes have evolved from other imaging modalities, including body computed tomography (CT) imaging, echocardiography, and X-ray contrast angiography, and consistent nomenclature across imaging modalities is important1 for accurate and unambiguous communication. The problem is often one of determining the appropriate plane as rapidly as possible to make the diagnosis. As with most other cardiac imaging techniques, it is important to know as much as possible regarding the clinical question prior to determining the protocol. All examinations therefore should be planned to answer a specific clinical question. The basic imaging planes can be grouped into planes oriented with respect to the heart, such as horizontal and vertical long axis and short axis, and planes oriented with respect to the major axes of the body, such as the transaxial, sagittal, and coronal planes. Cardiac-oriented planes are essential for evaluation of cardiac chamber size and function and are familiar from other cardiac imaging techniques. With CMR, these planes can be positioned very accurately. As shown in Figure 11-1A, a breath hold scout image in the coronal or sagittal plane is the usual starting point. An axial scout (Fig. 11-1B) is used to define the vertical long axis (also known as the two-chamber view, Fig. 11-1C). The horizontal long axis (Fig. 11-1D), which depicts both atria and both ventricles but is slightly different from the true four-chamber view, is then planned and is followed by the short axis (Fig. 11-1E and F), which can be used to generate the left ventricular (LV) outflow tract view (Fig. 11-1G), which is similar to the parasternal long axis view of echocardiography. The main structures of normal cardiac anatomy in the coronal, axial, and sagittal planes are shown for spin echo sequences in Figures 11-2A to J. There are many atlases of cross-sectional anatomy by CMR that can be helpful2 and web sites (e.g., www.scmr.org) with interactive learning of the cross-sectional anatomy can be very useful teaching aids. It is recommended that the reader refer to these for further details. From the standpoint of tissue characterization, the spin echo images typically permit the differentiation of fat (white) from muscle (intermediate gray). Black regions in spin echo CMR studies may represent several tissues or materials, including air, bone, fibrous 140 Cardiovascular Magnetic Resonance

tissue, metal, or rapidly moving blood. It is important to note that if fluid moves relatively slowly (for example, in an aneurysm), its signal intensity will increase, which can mimic more solid tissue such as thrombus or muscle. The placement of imaging planes, slice thickness, and in-plane resolution are determined by the size of the structure of interest. Presaturation bands can be added to remove specific artifacts. Other preparatory prepulses can be applied to emphasize or deemphasize the signal contribution of specific tissues. For example, in the evaluation of arrhythmogenic right ventricular cardiomyopathy, highspatial resolution spin echo images of the anterior right ventricular (RV) wall that are free from respiratory artifact are needed. This can be achieved by using a surface coil to improve signal-to-noise ratio compared with the body coil and thus facilitate higher spatial resolution. Breath hold, double inversion recovery spin echo techniques can also be very effective in removing respiratory and flow artifacts. Imaging planes oriented with respect to the principal axes of the body are particularly useful in the evaluation of the aorta, pericardium, RV free wall, and paracardiac masses. Coronal images can also be quite useful because they present tomographic information in an orientation similar to that of the chest X-ray, which is familiar to most clinicians (see Fig. 11-2A and B). In general, axial planes are also useful because they are familiar to the clinician from knowledge of CT (see Fig. 11-2C–H). Specific vascular structures of interest that can be evaluated well with axial imaging include the thoracic aorta and its branches, the pulmonary artery and veins, and the superior vena cava (see Fig. 11-2C and D). Axial images through the heart can be particularly useful in the evaluation of the pericardium and RV free wall (see Fig. 11-2E–H). They are of limited value in the assessment of myocardial wall thickness and chamber size because of the variable orientation of the heart relative to the principal axes of the body. Sagittal images are in general the least familiar to clinicians and are often more difficult to interpret (see Fig. 11-2I and J). However, sagittal images are useful in depicting the RV outflow tract and are therefore helpful in the evaluation of patients with congenital heart disease and RV cardiomyopathy. Oblique sagittal planes are useful in the evaluation of the thoracic aorta, and these planes can be easily defined from the transaxial images, especially if three-point plane definition is available by using the arch and lower ascending and descending aorta as the reference points (see Fig. 11-2K and L). Black-blood images (see Fig. 11-2M–P) oriented

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B

D

E

F

G H Figure 11-1 For legend see next page.

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Figure 11-1 A, Scout image 1, coronal: typical breath hold CMR image used to begin study (alternatively, a sagittal image could be used). The white line indicates the location of an axial image used to locate the mitral valve plane and interventricular septum. B, Scout image 2, axial: Typical breath hold image obtained to set up a vertical long axis (VLA) image. The white line indicates the position of the VLA imaging plane and is drawn to pass through the middle of the mitral valve and the ventricular apex. C, Vertical long axis image from breath hold steady-state free precession (SSFP) cine CMR oriented as described above. The white line indicates the position of the horizontal long axis (HLA) imaging plane and is selected to pass through the ventricular apex and between the attachment points of the mitral valve. D, Horizontal long axis breath hold SSFP cine CMR image oriented based on the prescription in panel C. The white lines indicate the positions of a stack of images oriented in the left ventricular (LV) short axis orientation, which will be obtained next. E, A representative end-diastolic breath hold SSFP cine LV short axis image from the middle of the stack shown in panel D. The white line perpendicular to the imaginary line between the insertion points of the right ventricular (RV) free wall is used to select the imaging plane to obtain a four-chamber view of the heart. F, Basal image in the LV short axis orientation showing an oblique view of the aortic valve. The white line shows the orientation of the LV outflow tract (LVOT) view. G, An LVOT view at end-diastole. This imaging plane is comparable to the parasternal long axis view of transthoracic echocardiography. H, Four-chamber breath hold SSFP cine CMR image at end-diastole. This view is similar to the HLA view, but note that typically, less of the aortic outflow tract is seen.

along the functional axes introduced in Figure 11-1 can be useful in the definition and tissue characterization of intracardiac and paracardiac masses. In addition, depiction of these planes with double inversion recovery black-blood imaging is useful for characterization of valvular disease and the coronary artery wall.3–5 Myocardial function is typically assessed by using cine steady-state free precession (SSFP) imaging, which has largely supplanted the older cine segmented gradient echo methods. SSFP provides improved contrast between blood pool and myocardium,6 particularly in the presence of impaired ventricular function, as it is dependent mainly on T1/T2 ratio rather than inflow of unsaturated protons. Both SSFP and gradient echo cine methods depict blood as bright (white), while muscle is an intermediate gray, and air, bone, fibrous tissue, and metal are dark. The main findings using bright-blood cine CMR of the heart are shown in Figures 11-1C–H and 11-2Q–T. These cines are typically used to assess myocardial and valve function. The LV outflow tract view, for example, is used to visualize the mitral and aortic valves (see Fig. 11-1G). One advantage of CMR is the ability to precisely position the long axis planes to avoid the foreshortening that can occur in contrast X-ray ventriculography or echocardiography (Fig. 11-1G and H). Short axis views are planned from the long axis views to span the entire LV. The short axis views in Figure 11-2R–T are useful in the evaluation of ventricular size and regional function. Using the same orientation to obtain views of the atria can be useful for assessing atrial masses as well as chamber size and function. To most accurately assess LV function and size, it is important to obtain correctly oriented images that encompass the entire LV throughout the cardiac cycle. To obtain true LV short axis views, we go through the following steps. From an axial scout image, the vertical long axis or twochamber view is obtained. The horizontal long axis (HLA) view is planned from the two-chamber view, ensuring that the imaging plane passes through the apex of the LV and through the center of the mitral valve annulus. A stack of short axis images is planned from the HLA view (Fig. 111D), with the short axis planes perpendicular to an imaginary line passing through the LV apex distally and midway between the visualized portions of the mitral valve annulus basally. It is important to plan the short axis stack so that it extends just distal to the apex and slightly above the base of the LV to ensure coverage of the entire LV. Failure to do so results in an incomplete dataset of suboptimal use for quantitative volumetric LV measures. The stack of images obtained as described above will be oriented in the LV short 142 Cardiovascular Magnetic Resonance

axis orientation. Though the long axis of the RV and LV are not parallel, from a practical perspective, the LV short axis dataset is used to calculate and assess RV volumes and systolic function. Coronary artery CMR requires yet another set of imaging planes to plane the coronary arteries in tomographic slices in the atrioventricular groove and axial planes when the targeted slab approach is used. Whole heart methods, which encompass the entire coronary tree in a 3D axial volume, yield a volumetric dataset similar to that obtained by ECG-gated cardiac CT. This subject is discussed in more detail in Chapter 21.

ANATOMIC VARIANTS Given the ability to obtain images in many planes, it is important to be aware of normal structures and anatomic variants that can complicate interpretation of studies. Several potential confusing features have been described: Prominence of the lateral border of the right atrial wall (Fig. 11-3A). This structure is a prominence of the trabeculae carneae and crista terminalis and does not represent an atrial mass.7 Lipomatous hypertrophy of the atrial septum (Fig. 113B). Fat deposition in the atrial septum is occasionally seen, particularly in the elderly. This process spares the region of the fossa ovalis, leading to the characteristic “dumbbell” appearance.8,9 It is benign but is associated with atrial arrhythmias. Severe and extensive lipomatous hypertrophy may extend outside the heart.10 Imaging with and without fat saturation readily characterizes this abnormality. Superior pericardial recess (Fig. 11-3C). The pericardium normally extends up the ascending aorta, and this space may contain fluid. This recess can be mistaken for aortic dissection or potentially an anomalous coronary vessel.

COMMON ARTIFACTS A number of CMR artifacts can complicate image interpretation. These artifacts relate primarily to several features of CMR. The acquisition time is often relatively long in comparison to physiologic processes, which leads to cardiac and respiratory motion artifacts. These motion artifacts must be recognized and minimized at the acquisition stage if possible. Also, because the strength of the local magnetic

Pulmonary artery Left ventricle

RVOT Right atrium

Left ventricular apex

A

B Ascending aorta

Superior vena cava

Superior vena cava

Main pulmonary artery

Transverse aortic arch Trachea

C

Descending aorta

D Pericardium

Right atrial appendage

Pericardium Epicardial fat

Right ventricular outflow tract

Tricuspid valve

RV RA

LV LA

Left atrium

Mitral valve Pulmonary vein

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E

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Inferior vena cava

Right atrium Coronary sinus

G Figure 11-2

Coronary sinus

H

For legend see page 145.

Cardiovascular Magnetic Resonance 143

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Ascending aorta

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Ascending aorta

Right pulmonary artery

Pericardium

Left atrium

RVOT Pericardium Epicardial fat

Right atrium

Right ventricular wall

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I

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Ascending aorta

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L

RCA origin Aortic valve leaflets PDA

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N Left ventricular apex

Anterior AV groove

O Figure 11-2 For legend see page 145.

144 Cardiovascular Magnetic Resonance

Posterior AV groove

P

Descending aorta

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11 NORMAL CARDIAC ANATOMY, ORIENTATION, AND FUNCTION

Q

Figure 11-2 A, Coronal breath hold double inversion recovery, spin echo CMR images. Fat is white, myocardium is intermediate gray in intensity, and blood is dark. The slice is positioned anteriorly and cuts through the RV, the RV outflow tract (RVOT), the interventricular septum, and the LV apex. B, Coronal image positioned more posteriorly than the image in panel A. This shows the right atrium, LV, ascending aorta, and pulmonary artery. The aortic valve leaflets are also seen. C, Transverse conventional gated spin echo image at the left of the transverse aortic arch. The trachea and superior vena cava are also demonstrated. D, Transverse conventional gated spin echo image at the level of the main pulmonary artery. The views in panels C and D are useful in the evaluation of possible aortic dissection. E, Transverse conventional spin echo image at the level of the aortic valve. F, Transverse conventional spin echo image at the level of the interatrial septum. The pericardium and epicardial fat are clearly demonstrated. This view can be useful in evaluating atrial masses and pericardial disease. G, Transverse conventional spin echo image at the level of the coronary sinus. The RV wall, epicardial fat, and pericardium are also demonstrated. This view can be helpful in evaluating patients for constrictive pericarditis and RV dysplasia. H, Transverse conventional spin echo image at the level of the entrance of the inferior vena cava into the right atrium. I, Sagittal conventional spin echo image obtained through the ascending aorta. The pericardium is clearly demonstrated. This view can be helpful in the evaluation of the ascending aorta and pericardium. J, Sagittal conventional spin echo image obtained through the RVOT. This view can be helpful in evaluating the pericardium, RVOT, and RV free wall. K, Transverse breath hold double inversion recovery, spin echo images obtained at the level of the transverse portion of the aortic arch (left) and the main pulmonary artery (right). The white line indicates the position of a parasagittal oblique plane used to obtain a “candy cane” view of the aorta (next panel). L, Parasagittal view of the aorta. The ascending aorta, transverse aorta, and descending aorta are seen in a single slice. The vessels to the head and neck are also well seen. This view can be helpful in the evaluation of aortic disease. M, Long axis view using breath hold double inversion recovery technique. This image is comparable with the parasternal long axis view in transthoracic echocardiography. Both the RV and LV are well demonstrated. The origin of the right coronary artery (RCA) is seen in the fat of the anterior atrioventricular groove. The aortic valve leaflets are also well seen. This view can be useful in the evaluation of hypertrophic cardiomyopathy with septal asymmetry. N, Short axis view using breath hold double inversion recovery technique. The LV and RV walls are well demonstrated. In this image the posterior descending artery (PDA) is also seen in cross section in the posterior interventricular groove. O, Four-chamber view using breath hold double inversion recovery technique. P, Vertical long axis or two-chamber view using breath hold double inversion recovery technique. Q, End-diastolic image in the HLA orientation from an steady-state free precession (SSFP) breath hold cine CMR sequence. The white lines indicate the locations of short axis imaging planes in subsequent panels R-T, all of which were obtained by using breath hold cine SSFP imaging). R, Basal RV and LV. The left anterior descending coronary artery is seen in the anterior interventricular groove. S, End-diastolic short axis image obtained at the mid level of the left ventricle. T, End-systolic image at the same imaging level as in panel S. LA, left atrium, LV, left ventricle; RA, right atrium, RV, right ventricle.

field determines the position of an object in a CMR image, alterations in the local magnetic field can shift the image position of the structure. Therefore, metal on or in the body can alter the local magnetic field, leading to distortion and local signal loss. Finally, hydrogen nuclei in fat experience a slightly different magnetic field in comparison with hydrogen nuclei in water molecules because of the local

chemical environment. This chemical shift is used in CMR spectroscopy to differentiate one compound from another. However, in CMR, this results in what is known as a chemical shift artifact at the interface of water and fatty tissues. This artifact results from sharing of fat and water components within a pixel, leading to signal cancellation. Specific examples are given for each type of artifact. Cardiovascular Magnetic Resonance 145

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A

Figure 11-3 A, Transverse gradient echo CMR image at the level of the aortic valve obtained by using respirator gating with a navigator echo. A right atrial ridge is noted in the lateral wall of the right atrium (arrow). This finding is normal and should not be mistaken for a right atrial mass. B, Single frame from horizontal long axis steady-state free precession cine CMR. Lipomatous hypertrophy of the atrial septum is demonstrated (arrow). There is fatty infiltration of the septum that does not involve the region of the fossa ovalis, resulting in the typical “dumbbell” appearance. C, Oblique double inversion recovery breath hold CMR image obtained at the level of the right pulmonary artery. The extension of the pericardial space both anterior and posterior to the ascending aorta is demonstrated (arrows). The pericardial recess should not be mistaken for evidence of aortic dissection.

B

C

Cardiac Motion Artifacts Except for single-shot echo planar imaging (EPI) or other realtime imaging approaches, CMR requires gating to the electrocardiogram (ECG) or peripheral pulse. Inaccurate cardiac gating can result in ghosting and other artifacts (Fig. 11-4A and B). Focused efforts to confirm accurate QRS detection are well worth the additional time and effort with regards to image quality. Vectorcardiographic techniques have been implemented to take advantage of the difference in the normal vector and the vector of the artifact from the magnetohydrodynamic effect to improve ECG gating.11 Surprisingly good-quality cine images can be obtained in patients with atrial fibrillation, which may be related to the relatively consistent length of systole relative to changes in heart rate.12 In contrast, bigeminy (trigeminy, etc.) often results in poor-quality cine images, in that every other beat is activated differently, resulting in combining data from two different activation patterns. Some CMR systems provide arrhythmia rejection in an attempt to reduce these effects; however, use of these tools generally results in increased scan time because of rejection of cardiac cycles.

Respiratory Motion Artifacts Respiration is associated with marked displacement of the heart. Motion in the craniocaudal direction is on the order 146 Cardiovascular Magnetic Resonance

of 1 to 1.5 cm in normal individuals.13 This motion can result in image degradation with ghosting and blurring, particularly in patients with inconsistent respiratory patterns. Strategies to reduce respiratory artifact include the use of sustained breath hold, presaturation of the highintensity signal from fat in the chest wall, and the use of free breathing with respiratory gating. Respiratory gating may be accomplished by using a thoracic bellows or by tracking the diaphragm position by using a navigator echo.14 These methods accept cardiac cycles only during some portion (typically end-expiration) of the respiratory cycle. It can substantially improve image quality and can be useful in coronary imaging without breath hold15 and in patients with heart failure.16 All respiratory gating methods increase total scan time.

Metal Artifact Pieces of metal outside or inside the body alter the local magnetic field and can result in artifacts (Fig. 11-4C–I). Patients must be screened carefully for the presence of metal, but despite vigilance, objects that are common in the hospital might still go with the patient into the scanner. Figure 11-4C shows an artifact related to a safety pin on the patient’s gown. Note that signal loss and distortion are present in both the fast spin echo and gradient echo images. The severity of the artifact is larger in the gradient echo

B

C

D

E

F

G

H

11 NORMAL CARDIAC ANATOMY, ORIENTATION, AND FUNCTION

Figure 11-4

A

For legend see next page.

Cardiovascular Magnetic Resonance 147

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

I

J

K

Figure 11-4 A, Artifacts due to respiration and poor gating. In this gated spin echo CMR image there is mottling of the ventricular wall and loss of edge sharpness. B, The same image as in panel A, but with the window and level adjusted to accentuate the artifact. There are ghosts of the chest wall related to respiratory motion and additional artifact over the heart as a result of poor ECG gating. C, Metal artifact. The upper images were obtained with a safety pin present on the subject’s gown. The resultant signal void is very evident. The bottom row shows corresponding images after removal of the safety pin. Distortion from metal artifact is markedly more prominent in the gradient echo images (right column) than in the spin echo images (left column). D, Plain-film X-ray showing sternal wires (dashed arrow) and metallic coronary artery bypass graft (CABG) markers (solid arrow) in a patient with prior CABG surgery. E, Artifact from sternal wires (dashed arrow) and CABG markers (solid arrow) on T1-weighted spin echo CMR imaging. F, Signal voids (arrows) in two views of a bioprosthetic aortic valve replacement by breath hold cine steady-state free precession imaging. The artifact results from the nonorganic struts. G, Metal in bileaflet mitral valve prosthesis produces signal voids (arrows). H, There is minimal artifact from the tricuspid (dashed arrow) and mitral (solid arrow) annuloplasty rings. I, Metal artifact from a coronary artery stent in the left anterior descending coronary artery (arrow) seen on a scout image. J, Chemical-shift artifact. The image on the left is done with a relatively short signal acquisition time (wide bandwidth). The image on the right is done with a longer signal acquisition time (narrow bandwidth). This display accentuates the effect of the difference in chemical shift of water and fat, creating the artifactual space between the aortic wall at fat (arrow). K, Chemical-shift artifact in echo-planar imaging (EPI). In EPI, the chemical shift artifact occurs in the frequency-encoding direction (right to left in these images). The image on the left is obtained by using a multishot EPI sequence with a relatively short EPI acquisition with each shot. The chemical shift artifact is indicated by the white line in the posterior chest wall. The image on the right is obtained by using fewer shots with a longer EPI acquisition. The chemical shift is larger, as indicated by the longer white line posteriorly. The image is degraded by superimposition of anterior subcutaneous fat onto the heart. This problem can be addressed by adding fat saturation to the sequence.

images, severely compromising interpretation of the RV and septum. Figure 11-4D and E shows the artifacts related to sternal wires and coronary artery bypass graft markers. Figure 11-4F shows the artifact related to a bioprosthetic aortic valve while Figure 11-4G is a mechanical bileaflet mitral valve prosthesis. Figure 11-4H shows the minimal artifacts associated with mitral and tricuspid annuloplasty rings, and Figure 11-4I depicts artifact from a stent in the left anterior descending coronary artery. 148 Cardiovascular Magnetic Resonance

Chemical Shift Artifact A chemical shift artifact occurs because the hydrogen nuclei in fat experience a slightly different magnetic field than hydrogen nuclei in water because of the different local chemical environment (Fig. 11-4J,K).17 This process results in displacement of the fat signal in the frequency-encoding direction relative to water and is accentuated with narrow bandwidth sequences, which can present a diagnostic

NORMAL CARDIAC SYSTOLIC AND DIASTOLIC FUNCTION The management of cardiovascular disease is critically dependent on the assessment of cardiac function. Thus, every cardiac imaging technique has been used to assess systolic and diastolic function. There is an extensive body of evidence to indicate that CMR provides highly accurate and reproducible assessments of global and regional cardiac function, and CMR is often considered as the gold standard for the noninvasive evaluation of cardiac function,19 a standard by which other noninvasive methods are validated. An important consideration in determining ventricular function is the temporal resolution or frame rate of the cine CMR sequence. A frame rate of at least 25 frames/sec (i.e., temporal resolution of 40 msec/frame) is required to accurately identify end-systole. Historically, X-ray left ventriculography has been obtained at a frame rate of 30 frames/ sec (temporal resolution of 33 msec). The frame rate in echocardiography is dependent on the speed of ultrasound in the body and the distance of the heart from the transducer but typically is at least 30 frames/sec (temporal resolution of 33 msec or better). With modern CMR scanners with high-performance gradients, frame rates over 30 frames/sec for breath hold sequences are routine. Further, the apparent frame rate can be increased by using viewsharing techniques that reconstruct intermediate images by combining recent, but previously acquired, k-space data with selectively updated current data.20 Finally, partially parallel imaging methods (e.g., SENSE, SMASH, GRAPPA) are now routinely used in clinical imaging to decrease acquisition time and increase frame rate.21 These are described in greater detail in Chapter 3. Real-time CMR approaches have inferior spatial and temporal (50 to 70 msec) resolution but may be preferred for patients with frequent arrhythmias and in examining respiratory changes in interventricular septal motion (e.g., respiratory variation with tamponade or constriction).22 While the CMR cine loop appears to display a single cardiac cycle, the image data are generally acquired over multiple heartbeats. As a result, image quality can be markedly degraded in the presence of arrhythmias (see above) or unreliable ECG triggering or gating. Cardiac function is typically assessed during breath holding, to minimize bulk cardiac translation; so if the goal is to assess the effect of respiration on chamber size or function (as in evaluating

for possible constriction, for example), then non–breath hold real-time methods22 may be more appropriate, as such techniques acquire images during a fraction of a single heartbeat and not as a composite over multiple cardiac cycles as with usual segmented k-space sequences. However, the trade-off with real-time CMR is a decrease in both spatial and temporal resolution.

Left Ventricle Assessment of LV function includes global and regional function. Assessment of global LV function is based on measuring changes in chamber volumes. These changes can be estimated from unidimensional (linear) measurements as in echocardiography, but with CMR, more accurate measures of chamber volume can be made by using three-dimensional methods. The two-dimensional methods (e.g., the area-length method or biplane angiographic formulas) have no advantages over echocardiography and are not widely used.23–25 More accurate measures, particularly in deformed ventricles, which do not fit common geometric formula–based models, can be obtained by using the summation of disks method. (In the cardiac imaging literature, this method is often referred to as the Simpson’s rule method. The mathematical Simpson’s rule is a fourth-order polynomial approximation for numerical integration;26 failure to distinguish between the mathematical and medical definitions of Simpson’s rule leads to confusion between clinicians and engineers or medical physicists. The term modified Simpson’s rule is sometimes used in the echocardiography literature to refer to a simplified summation of disks method.) With the summation of disks method, short axis images are obtained that span the entire ventricle; the cross-sectional area in each slice is measured, multiplied by the slice thickness (and interslice gap if applicable), and summed over the entire ventricle.27 This approach is highly accurate and reproducible and is widely used both clinically and in research.26 Regional LV systolic function can be assessed both qualitatively (“eyeball” method similar to echocardiography) and quantitatively. As shown in Figure 11-5, the standard long axis, four-chamber, two-chamber, and short axis views can be mapped onto the 17-segment American Heart Association1 model for qualitative assessment of wall thickening. Wall thickening can also be determined quantitatively by using centerline or other methods, and commercial software is available for such analyses. Myocardial tagging methods28–30 can be used to quantify myocardial contractility using strain without the need to identify endocardial or epicardial borders explicitly. However, software for tracking deformation of the tag lines over time nonetheless requires human interaction, prolonging analysis times. The harmonic phase (HARP) technique was proposed to obviate the human postprocessing issue by computing strain based on local spatial frequency of the tag lines.31 With appropriate bandpass filtering and subsequent transformation of the k-space signal, strain can be extracted automatically. An alternative method, displacement encoding with stimulated echoes (DENSE),32 combines a stimulated echo with bipolar gradient to encode displacement, from which strain can be calculated. These methods are detailed in Chapter 13. Cardiovascular Magnetic Resonance 149

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problem in spin echo imaging of patients with suspected aortic dissection.18 It can be addressed by using a wider bandwidth sequence or repeating the sequence with frequency encoding in the alternate direction, which will result in changing the orientation of the artifact and thus help exclude the presence of an aortic dissection. In echo planar images, chemical shift effects lead to artifacts displaced in the phase-encoding direction. As shown in Figure 11-4K, this effect can be minimized using multishot EPI techniques. In single-shot EPI with long acquisition times, the chemical shift effects can be quite large. For this reason, single-shot echo planar images often employ fat saturation to suppress this artifact.

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

Regional LV function

Ap-AS

Mid- Bs-AS AS

BsSept

Ap Ap-Inf/ Lat

Bs-AS

BsMidInf/Lat Inf/Lat

Bs-Ant

BsLat

Bs-Inf

Bs-Inf/Lat

MidAS

Mid-Ant

Figure 11-5 A, Cine CMR images in each of the standard planes typically used for scoring wall thickening and segmental function. The segments visualized correspond to those depicted in the standard American Heart Association (AHA) 17-segment model. B, AHA 17segment model of the LV (2). Ant, anterior; Ap, apical; As, anteroseptal; Inf; inferior; Lat, lateral; Sept, septal; Bs, basal.

Apex Ap-Sept

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Right Ventricle Historically, measurement of RV volumes has been largely qualitative. This situation is due to the lack of standard geometric models for the RV. A distinct advantage of CMR is that the same summation of disks approach may be readily applied to the RV. Studies in ventricular casts have shown an excellent correlation between CMR and displacement measurements.33

Stroke Volume Stroke volume is the amount of blood ejected from the heart with each cardiac cycle. It can be readily calculated by subtracting the end-systolic volume from the end-diastolic volume. Multiplying the stroke volume by the heart rate yields the ventricular cardiac output, typically reported in liters per minute. Studies comparing CMR cardiac output with invasive thermodilution methods have shown good correlation.34–37 The cardiac output that is determined in this way from the stroke volume is the same as the cardiac output that is determined in the catheterization laboratory. In the setting of aortic or mitral regurgitation, however, part of this volume does not result in the net delivery of blood to the periphery. In this setting, the cardiac output based on the summation of disks LV stroke volume is greater than the forward flow (which can be measured by using flow techniques (see Chapter 37), but the regurgitant volume can be determined by subtracting the forward flow from the apparent stroke volume.38

Left Ventricular Mass CMR is similarly considered the gold standard for the assessment of in vivo LV mass and is often used for validation of other methods.39 The mass of the LV wall can be estimated 150 Cardiovascular Magnetic Resonance

by measuring the volume of the myocardium and multiplying it by the specific gravity of myocardium, 1.05 g/mL. CMR, in conjunction with a volumetric method such as summation of disks, provides accurate estimates of myocardial mass over a broad range of heart sizes and deformed ventricles both in animals and in man.40–42 In cadaver heart studies, linear correlation analysis demonstrated a correlation coefficient of 0.99 with a standard error of 6.8 g. Intraobserver, interobserver, and interstudy variability are excellent. Notably, LV mass that is determined by using linear measurements and cubed-power geometric formulas, such as the echocardiographic Penn formula,43 generally overestimate volumetric mass even when the linear measurements are made from CMR images.44 The volumetric summation of disks approach has been used to determine RV mass with good results as well.45–46

EFFECT OF IMAGING SEQUENCE AND MAGNETIC FIELD STRENGTH ON VENTRICULAR VOLUMES AND MASS AND IMPLICATIONS FOR REFERENCE STANDARDS Compared with cine gradient echo cine sequences, cine SSFP imaging provides superior contrast between blood pool and myocardium. Perhaps owing to this difference in delineation of endocardial contours, ventricular volumes and stroke volume by SSFP imaging are slightly but systematically greater than corresponding values measured using cine gradient echo cine methods.47 Conversely, LV

imaging sequences is used. Table 11-1 shows cine gradient echo LV reference (“normal”) values for healthy adults (free of any history of hypertension and cardiac disease) from the primarily Caucasian Framingham population and Dallas Heart Study populations. Table 11-2 shows cine SSFPbased LV reference values from Framingham among adults strictly free of hypertension or clinical cardiovascular disease. Table 11-3 presents RV reference values, obtained by using cine gradient echo imaging, from the MESA cohort for different ethnic groups. Higher-field-strength CMR systems (i.e., 3 Tesla[T]) offer signal-to-noise advantages over “conventional” 1.5Tesla systems and are well established for neurologic imaging. As the installed base of 3-T systems increases, there has been increasing interest in 3-Tesla CMR (see Chapter 13). With respect to cardiac size and function, the magnetic field strength of the scanner (i.e., 1.5-T versus 3-Tesla) does not appear to have any significant effect on measured ventricular volumes and mass, but the systematic difference

Table 11-1 Breath Hold Cine First Gradient Echo CMR Reference Values for the Left Ventricle Based on Community-Dwelling Adult Subjects, Aged 57 9 Years and Strictly Free of Cardiovascular Disease and Any History of Hypertension, Drawn from the Framingham Heart Study Offspring Cohort and Dallas Heart Study Raw LV EDV (mL) LV ESV (mL) LVM (g) LV EF LV EDV/HT (mL/m) LV ESV/HT (mL/m) LVM/HT (g/m) LV EDV/BSA (mL/m2) LV ESV/BSA (mL/m2) LVM/BSA (mLm2)

Men (n ¼ 63) 115 36 155 0.69 0.70 (DHS) 66 21 89 58 54 (DHS) 18 78

Men 95th Percentile 169 65 201 0.59* 0.55{ 94 36 114 80 31 95

Women (n ¼ 79) 84 25 103 0.70 0.75 (DHS) 52 16 64 50 49 (DHS) 15 61

Women 95th Percentile 117 41 134 0.60* 0.61{ 70 25 82 66 24 75

*Fifth percentile (lower limit) for ejection fraction. { Below the 2.5th percentile of the DHS. BSA, body surface area; DHS, Dallas Heart Study data; EDV, end-diastolic volume; EF, ejection fraction; ESV, end-systolic volume; HT, height; LV, left ventricle; LVM, left ventricular mass. Source: Adapted from CJ Salton et al. J Am Coll Cardiol. 2002;39:1055; Chung et al. Circulation. 2006;113:1597.

Table 11-2 Breath Hold Cine Steady-State Free Precession CMR Reference Values for the Left Ventricle Based on Community-Dwelling Adult Subjects, Aged 61 8 Years and Strictly Free of Cardiovascular Disease and Any History of Hypertension, Drawn from the Framingham Heart Study Offspring Cohort Raw LV EDV (mL) LV ESV (mL) LV SV (mL) LVM (g) LV EF LV EDV/HT (mL/m) LV ESV/HT (mL/m) LVM/HT (g/m) LV EDV/BSA (mL/m2) LV ESV/BSA (mL/m2) LVM/BSA (g/m2)

Men (n ¼ 239) 144 26 54 14 93 17 123 22 0.65 0.05 81 14 29 8 70 12 71 12 25 7 61 10

Men 95th Percentile

Women (n ¼ 367)

196 78 126 167 0.55* 109 43 93 95 38 79

106 19 35 10 71 11 81 14 0.67 0.05 65 10 22 6 50 8 61 9 20.5 47 7

Women 95th Percentile 143 55 93 109 0.57* 86 33 66 78 30 60

*Fifth percentile (lower limit) for ejection fraction. BSA, body surface area; EDV, end-diastolic volume; EF, ejection fraction; ESV, end-systolic volume; HT, height; LV, left ventricle; LVM, left ventricular mass; SV, stroke volume. Source: Adapted from Salton CJ et al. Circulation. 2006;114: II-669.

Cardiovascular Magnetic Resonance 151

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mass with SSFP CMR is smaller than CMR gradient echo mass. This has implications for what constitutes normalrange ventricular volumes and mass, as the reference values must be imaging sequence specific. Gradient echo reference values for LV volumes and mass have been published based on data from the community-dwelling Framingham Heart Study.48 The Multiethnic Study of Atherosclerosis (MESA) and Dallas Heart group have also published RV49 and LV50,51 reference values based on gradient echo imaging with a slightly higher LV ejection fraction demonstrated in women.51 Reports on SSFP-based reference values for RV volumes and mass also demonstrate increased RV ejection fraction in women.52 Other publications have proposed reference values partitioned by age as well as sex, using SSFP imaging.53–55 These are useful, though the data are somewhat limited by small sample sizes or scanning of younger, athletically fit populations. Despite greater cavity volumes by SSFP as compared with gradient echo methods, LV ejection fraction is similar regardless of which of these

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

Table 11-3 Breath Hold Cine Gradient Echo CMR Reference Values for the Right Ventricle Based on Community-Dwelling Adult Subjects, Aged 61 10 Years, Drawn from the Multi-Ethnic Study of Atherosclerosis (MESA) cohort Raw RV RV RV RV RV RV RV RV RV RV

EDV (mL) ESV (mL) SV (mL) EF EDV/HT (mL/m) ESV/HT (mL/m) SV/HT (mL/m) EDV/BSA (mL/m2) ESV/BSA (mL/m2) SV/BSA (mL/m2)

Men (n ¼ 219) 142 31 54 17 88 22 0.62 0.1 73 14 28 8 45 10 82 16 31 9 51 11

Men 95th Percentile 201 85 125 0.50* 98 43 63 101 48 70

Women (n ¼ 268)

Women 95th Percentile

110 24 35 13 75 18 0.69 0.1 65 11 21 7 44 8 69 14 22 8 47 10

155 57 106 0.58* 83 33 58 95 34 63

*Fifth percentile (lower limit) for ejection fraction. BSA, body surface area; EDV, end-diastolic volume; EF, ejection fraction; ESV, end-systolic volume; HT, height; LVM, left ventricular mass; RV, right ventricle; SV, stroke volume. Source: Adapted from Tandri H et al. Am J Cardiol. 2006;98:1660.

between gradient echo and SSFP imaging sequences is preserved regardless of field strength.56,57 It should be noted that SSFP imaging at 3-T is more subject to artifacts than imaging at 1.5-T; this results in part from greater sensitivity to field inhomogeneities, longer T1, and limitations on radiofrequency power deposition.58 Cardiac SSFP imaging at 3-T at present requires more careful sequence optimization and understanding of how to identify and correct artifacts than CMR at 1.5-T.

Aortic Flow An important feature of CMR is its sensitivity to motion, and this feature can be harnessed to allow the measurements of velocity and vessel flow without the constraints of Doppler methods. Although there is potential for artifacts, CMR allows accurate and reproducible measurement of vessel flow in vivo.37,59–61 The approach for the measurement of aortic flow is straightforward. The coronal scout image is used, and the imaging plane is placed perpendicular to the direction of flow several centimeters above the aortic valve, typically at approximately the level of the bifurcation of the main pulmonary artery (Fig. 11-6A). The pulse sequence uses bipolar gradients oriented in the expected direction of flow, so tissues, such as blood, moving through the imaging plane accrue a nonzero net phase, while stationary tissues gain and lose phase equally for zero net phase change. The phase change corresponds to velocity. The velocity encoding value (VENC) should be chosen to be above the anticipated maximum velocity (approximately 1.5 to 2 m/sec in normal individuals and higher in patients with aortic valve disease). Setting the VENC too low can result in aliasing, in which phase change exceeds þ180 , so there is an abrupt discontinuity in apparent velocity, also known colloquially as wrap-around. For example, a phase change of, for example, þ270 will be misinterpreted as a phase of 90 , owing to wraparound. The reconstructed images are typically presented as a set of magnitude images that are used to determine the crosssectional area of the aorta in each frame (Fig. 11-6B). There is also a set of phase-encoded (velocity map) images in 152 Cardiovascular Magnetic Resonance

which the gray scale indicates the velocity of motion in each voxel (Fig. 11-6C). Velocity is measured at each voxel across the vessel, integrated over the cross-sectional area of the vessel, and then integrated over the cardiac cycle (Fig. 11-6D). The integrated flow across the slice at each point in time can be graphed. Note that some retrograde flow is normal in the ascending aorta during early diastole, owing to closure of the aortic valve, diastolic ascent of the base of the heart, and diastolic coronary flow. The difference between LV stroke volume and aortic systolic forward flow can also be used to quantify mitral regurgitation.

Pulmonary Artery Flow Pulmonary artery flow can be measured by using techniques similar to those for aortic flow. This is particularly valuable in evaluation of patients with left-to-right intracardiac shunting to determine the ratio of pulmonary to systemic flow (Qp/Qs). Data demonstrated a very good correlation with invasive techniques.62,63 Depending on the pulmonary artery orientation, the location of the flow images can be planned from the axial and/or sagittal scout with a perpendicular plane positioned several centimeters distal to the pulmonary valve. The velocity profile is then integrated over the cross-sectional area of the artery over time to determine the volume flow per cardiac cycle in a manner similar to that used in the ascending aorta. The difference between RV stroke volume and pulmonary systolic forward flow can also be used to quantify tricuspid regurgitation.

NORMAL VALVULAR FUNCTION Assessment of valve function involves evaluation of morphology, motion, competence, and effects on ventricular function. Imaging cardiac valve morphology poses significant problems for CMR.64,65 First, the normal valve is a thin, fibrous structure often less than 1 mm thick, leading to potential for partial volume effects. Second, it is

11 NORMAL CARDIAC ANATOMY, ORIENTATION, AND FUNCTION

B

A

C Scan time 00:02:51 PHILIPS MR 1.5 T SESSION INFORMATION: Q-Flow: AorticFlow (not validated).

Flux results (slice 1) ml/s 400 350

RESULTS SUMMARY: Heart rate : RR-interval :

300 250

60 bpm 1000 ms

(from heart rate)

ANALYSIS RESULTS: slice 1 Vessel 1

200 150 100 50 0 0 100 200 300 400 500 600 700 800 9001000

time (ms)

RR-interval: 1000 ms (from heart rate)

Stroke volume (ml) Forward flow vol. (ml) Backward flow vol. (ml) Regurgitant fract. (%) Abs. stroke volume (ml) Mean flux (ml/s) Stroke distance (cm) Mean velocity (cm/s)

73.5 73.8 0.3 0.4 74.1 73.5 9.5 9.5

Vessel 1, slice 1

Q-Flow: AorticFlow (not validated) Flux: Peak vel: 12.82 cm/s 7.19 ml/s Mean vel: 0.94 cm/s Max. vel: 12.82 cm/s Area: Min vel: –6.96 cm/s 7.65 cm2 Pixels: 557 pixels Vel stddev: 3.38 cm/s

300 cm/s 200 100 0 –100 –200

FFE/M SI 1 Ph 1/000 ms

D Figure 11-6

PCA/P SI 1 Ph 1/000 ms PCv FH 300 cm/s

–300 cm/s

For legend see next page. Cardiovascular Magnetic Resonance 153

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

Figure 11-6 A, Coronal scout image for measuring aortic flow. The white line indicates the anatomic position of a flow sequence. The plane is positioned at the level of the pulmonary artery well above the aortic valve and perpendicular to the walls of the aorta. B, Magnitude reconstruction from the flow sequence positioned in panel A. The imaging plane is transverse and positioned at the level of the bifurcation of the main pulmonary artery. The ascending aorta is seen anteriorly (arrow). C, Velocity map reconstruction from the same flow sequence as shown in panel B. The gray scale in this image indicates the velocity of motion toward the head as bright (solid arrow, ascending aorta) and away as dark (dashed arrow, descending aorta). D, Typical dataset for semiautomated analysis. The area of the ascending aorta is determined in each frame and the velocity over the area is integrated to calculate flow volume per frame. The top left subpanel shows a graph of flow volume over the cardiac cycle in the ascending aorta. The forward stroke volume, shown in the results listed in upper right subpanel, is calculated by integrating the flow over the cardiac cycle.

constantly in motion. In most cine CMR, the image is acquired over a number of cardiac cycles, requiring that the valve return to the same position with each cycle. With breath hold imaging, one can obtain very high-resolution images, indicating that normal valve motion appears to be quite reproducible over a limited number of cycles when the effects of respiratory motion are removed. However, vegetations and other valve pathology are characterized by chaotic valve motion, which is less reproducible, leading to loss of signal and motion artifacts. Third, valve pathology frequently involves fibrosis and calcification, both of which are characterized by loss of signal on CMR, making it difficult to detect, particularly in spin echo imaging, Last, regions of turbulence are associated with loss of signal on CMR, which may lead to overestimation of the extent of abnormality on gradient echo imaging. In spite of these concerns, CMR can be used to obtain high-resolution images of valves using bright-blood SSFP techniques and/

or black-blood breath hold, double inversion recovery techniques.66 An example of a normal valve image obtained by using the latter technique is shown in Figure 11-7A. The LVOT view in Figure 11-7B shows aortic and mitral valves by SSFP imaging, and Figure 11-7C is an en face view of a normal trileaflet aortic valve; this frame from the SSFP cine loop shows the closed valve at end-diastole. Figure 11-7D is another normal aortic valve open during early systole. Figure 11-7E is an en face view of a sclerotic, mildly stenosed aortic valve during early systole. Figure 11-7F shows a dark signal void due to turbulent flow of aortic regurgitation. Imaging the valvular abnormality is only one part of the evaluation of the patient with valvular disease. It is essential to quantify the functional severity of the lesion and to determine its impact on ventricular size and function. Full details of assessing valvular abnormalities are given in Chapter 37, but some general comments are useful. As

A

B

C

D

Figure 11-7 A, Double inversion recovery spin echo CMR long axis image. The aortic valve leaflets are demonstrated (arrow). B, An LVOTview image acquired using breath hold steady-state free precession (SSFP) CMR imaging. C, A normal trileaflet aortic valve depicted en face at end diastole (valve closed) by SSFP CMR imaging. D, An open, normal trileaflet aortic valve seen during early systole. continued 154 Cardiovascular Magnetic Resonance

F

Figure 11-7 cont’d E, An en face view of a sclerotic, mildly stenosed, trileaflet aortic valve also during early systole. Note the small opening and deformed leaflets as compared with panel D. F, Aortic regurgitation is visualized qualitatively in the dephasing jet (arrow) in this SSFP image.

was noted earlier, CMR is highly accurate in determining ventricular dimensions and volumes. In addition, phase contrast quantitative flow techniques provide effective means for determining velocity and blood flow. Thus, in addition to demonstrating that valvular disease is present, CMR can be used to quantify the degree of dysfunction and to determine its effect on ventricular size and function. Valve pressure gradients estimated by using CMR correlate well with ultrasound measurements.67,68 In addition, measurements of aortic valve area and cardiac output by CMR agree with measurements of valve area at catheterization. Regurgitant jets are generally well visualized, owing to turbulence (dephasing) creating dark regions of signal loss in gradient echo cines, but it is hazardous to estimate even qualitative severity of regurgitation from the size of the region of signal loss on cine CMR, as the apparent size of the jet is highly dependent on the details of the particular imaging sequence used. Specifically, although SSFP imaging generally provides excellent depiction of myocardial and valve anatomy, the size of dephasing jets is smaller (see Fig. 11-7F) than that with gradient echo sequences. Regurgitant jets can be underestimated or overlooked entirely based on visual assessment alone. Quantitative flow techniques are more useful in determining the regurgitant volume, for example, by measuring retrograde flow in diastole in the aorta.69–71 Interestingly, measures of chamber volume and cardiac output by CMR in patients with atrial fibrillation agree well with invasive measures.12 Mitral regurgitation

has also been studied using quantitative measures.72,73 Interrogation of the aortic and pulmonic valves is relatively straightforward using velocity-encoding methods as described above, but the application of this method to the mitral and tricuspid valves is not necessarily straightforward, owing to the through-plane translation (10 to 20 mm in a normal heart) of the base of the ventricles and thus the mitral and tricuspid annuli. Assessment of mitral regurgitation is more reliably achieved by calculating the difference between stroke volume, by summation of disks method, and net aortic forward flow, by phase-encoded velocity mapping, as long as there is no intracardiac shunt. The presence of a prosthetic valve is not a contraindication for CMR74 except in the case of probable valve dehiscence,75 although mechanical valves (see Fig. 11-4G) or the struts of bioprosthetic valves will results in artifacts (see Fig. 11-4F). Similarly, rings used for valve repair may produce local artifacts (see Fig. 11-4H).

CONCLUSION CMR can be used to clearly delineate cardiac anatomy and to assess function. As with any imaging technique, it is important to have a strategy for imaging and be familiar with the normal anatomy and potential artifacts. When this knowledge is in hand, CMR can be a very effective tool in the evaluation of patients with cardiovascular disease.

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4. Fayad ZA, Fuster V, Fallon JT, et al. Noninvasive in vivo human coronary artery lumen and wall imaging using black-blood magnetic resonance imaging. Circulation. 2000;102:506–510. 5. Botnar RM, Stuber M, Kissinger KV, et al. Non-invasive coronary vessel wall and plaque imaging with magnetic resonance imaging. Circulation. 2000;102:2582–2587. 6. Thiele H, Nagel E, Paetsch I, et al. Functional cardiac MR imaging with steady-state free precession (SSFP) significantly improves endocardial border delineation without contrast agents. J Magn Reson Imag. 2001;14:362–367. Cardiovascular Magnetic Resonance 155

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32. Aletras AH, Ding S, Balaban RS, Wen H. DENSE: Displacement encoding with stimulated echoes in cardiac functional MRI. J Magn Reson. 1999;137:247–252. 33. Boxt LM, Katz J, Kolb T, et al. Direct quantitation of right and left ventricular volumes with nuclear magnetic resonance imaging in patients with primary pulmonary hypertension. J Am Coll Cardiol. 1992;19:1508–1515. 34. Utz JA, Herfkens RJ, Heinsimer JA, et al. Cine MR determination of left ventricular ejection fraction. AJR Am J Roentgenol. 1987;148:839–843. 35. Culham JAG, Vince DJ. Cardiac output by MR imaging: an experimental study comparing right ventricle and left ventricle with thermodilution. J Can Assoc Radiol. 1988;39:247–249. 36. Hunter GJ, Hamberg LM, Weisskoff RM, et al. Measurement of stroke volume and cardiac output within a single breath hold with echoplanar MR imaging. J Magn Reson Imaging. 1994;4:51–58. 37. Hundley WG, Li HF, Hillis LD, et al. Quantitation of cardiac output with velocity-encoded, phase-difference magnetic resonance imaging. Am J Cardiol. 1995;75:1250–1255. 38. Hundley WG, Li HF, Willard JE, et al. Magnetic resonance imaging assessment of the severity of mitral regurgitation: a comparison with invasive techniques. Circulation. 1995;92:1151–1158. 39. Takeuchi M, Nishikage T, Mor-Avi V, et al. Measurement of left ventricular mass by real-time three-dimensional echocardiography: validation against magnetic resonance and comparison with twodimensional and M-mode measurements. J Am Soc Echo. 2008;21: 1001–1005. 40. Franco F, Dubois SK, Peshock RM, Shohet RV. Magnetic resonance imaging accurately estimates cardiac mass in a transgenic mouse model of cardiac hypertrophy. Am J Physiol. 1998;43:H679–H683. 41. Maddahi J, Crues J, Berman DS, et al. Noninvasive quantification of left ventricular myocardial mass by gated proton nuclear magnetic resonance imaging. J Am Coll Cardiol. 1987;10:682–692. 42. Katz J, Milliken MC, Stray-Gundersen J, et al. Estimation of human myocardial mass with MR imaging. Radiology. 1988;169:495–498. 43. Devereux RB, Reichek N. Echocardiographic determination of left ventricular mass in man: anatomic validation of the method. Circulation. 1977;55:613–618. 44. Bottini PB, Carr AA, Prisant LM, et al. Magnetic resonance imaging compared to echocardiography to assess left ventricular mass in the hypertensive patient. Am J Hyperten. 1995;8:221–228. 45. Katz J, Whang J, Boxt LM, Barst RJ. Estimation of right ventricular mass in normal subjects and in patients with primary pulmonary hypertension by nuclear magnetic resonance imaging. J Am Coll Cardiol. 1993;21:1475–1481. 46. McDonald KM, Parrish T, Wennberg P, et al. Rapid, accurate and simultaneous noninvasive assessment of right and left ventricular mass with nuclear magnetic resonance imaging using the snapshot gradient method. J Am Coll Cardiol. 1992;19:1601–1607. 47. Moon JCC, Lorenz CH, Francis JM, Smith GC, Pennell DJ. Breathhold FLASH and FISP cardiovascular MR imaging: left ventricular volume differences and reproducibility. Radiology. 2002;223:789–797. 48. Salton CJ, Chuang ML, O’Donnell CJ, et al. Gender differences and normal left ventricular anatomy in an adult population free of hypertension: a cardiovascular magnetic resonance study of the Framingham Heart Study offspring cohort. J Am Coll Cardiol. 2002;39:1055–1060. 49. Tandri H, Daya SK, Nasir K, et al. Normal reference values for the adult right ventricle by magnetic resonance imaging. Am J Cardiol. 2006;98:1660–1664. 50. Natori S, Lai S, Finn JP, et al. Cardiovascular function in multi-ethnic study of atherosclerosis: normal values by age, sex and ethnicity. AJR Am J Roengtenol. 2006;186:S357–S365. 51. Chung AK, Das SR, Leonard D, et al. Women have higher left ventricular ejection fractions than men independent of differences in left ventricular volume: the Dallas Heart Study. Circulation. 2006;113: 1597–1604. 52. Salton CJ, Gona P, Chuang ML, et al. Gender-specific reference values for left ventricular anatomy and systolic function using modern cardiovascular magnetic resonance methods in a longitudinally followed normotensive population: the Framingham Heart Study. Circulation. 2006;114(suppl S):669. 53. George KP, Birch KM, Pennell DJ, Myerson SG. Magnetic-resonanceimaging-derived indices for the normalization of left ventricular morphology by body size. Magn Reson Imaging. 2009;27:207–213. 54. Maceira AM, Prasad SK, Khan M, et al. Reference right ventricular systolic and diastolic function normalized to age, gender and body surface area from steady-state free precession cardiovascular magnetic resonance. Eur Heart J. 2006;27:2879–2888.

65. Globits S, Higgins CB. Assessment of valvular heart disease by magnetic resonance imaging. Am Heart J. 1995;129:369. 66. Arai AE, Epstein FH, Bove KE, Wolff SD. Visualization of aortic valve leaflets using black blood MRI. J Magn Reson Imaging. 1999;10: 771–777. 67. Kilner PJ, Firmin DN, Rees RSO, et al. Valve and great vessel stenosis: assessment with magnetic resonance jet velocity mapping. Radiology. 1991;178:229–235. 68. Eichenberger AC, Jenni R, von Shulthess GK. Aortic valve pressure gradients in patients with aortic valve stenosis: quantification with velocityencoded cine MR imaging. AJR Am J Roentgenol. 1993;160:971–977. 69. Dulce MC, Mostbeck GH, O’Sullivan M, et al. Severity of aortic regurgitation: interstudy reproducibility of measurements with velocityencoded cine MR imaging. Radiology. 1992;185:235–240. 70. Honda N, Machida K, Hashimoto M, et al. Aortic regurgitation: quantitation with MR imaging velocity mapping. Radiology. 1993; 185:189–194. 71. Walker PG, Oyre S, Pedersen EM, et al. A new control volume method for calculating valvular regurgitation. Circulation. 1995;92:579–586. 72. Fujita N, Chazouilleres AE, Hartiala JJ, et al. Quantification of mitral regurgitation by velocity-encoded cine nuclear magnetic resonance imaging. J Am Coll Cardiol. 1994;23:951–958. 73. Hundley WG, Li HF, Willard JE, et al. Magnetic resonance imaging assessment of the severity of mitral regurgitation: a comparison with invasive techniques. Circulation. 1995;92:1151–1158. 74. Deutsch HJ, Bachmann R, Sechtern U, et al. Regurgitant flow in cardiac valve prostheses: diagnostic value of gradient echo nuclear magnetic resonance imaging in reference to transesophageal twodimensional color doppler echocardiography. J Am Coll Cardiol. 1992;19:1500–1507. 75. Shellock FG, Kanal E. Magnetic Resonance: Bioeffects, Safety, and Patient Management. 2nd ed. Philadelphia: Lippincott-Raven; 1996.

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55. Maceira AM, Prasad SK, Khan M, Pennell DJ. Normalized left ventricular systolic and diastolic function by steady state free precession cardiovascular magnetic resonance. J Cardiovasc Magn Reson. 2006;8:415–426. 56. Hudsmith LE, Petersen SE, Tyler DJ, et al. Determination of cardiac volumes and mass with FLASH and SSFP cine sequences at 1.5 vs. 3 Tesla: a validation study. J Magn Reson Imag. 2006;24:312–318. 57. Grothues F, Boenigk H, Graessner J, Kanowski M, Klein HU. Balanced steady-state free precession vs. segmented fast low-angle shot for the evaluation of ventricular volumes, mass and function and 3 Tesla. J Magn Reson Imag. 2007;26:392–400. 58. Scha¨r M, Kozerke S, Fischer SE, Boesiger P. Cardiac SSFP imaging at 3 Tesla. Magn Reson Med. 2004;51:799–806. 59. Stahlberg F, Sondergaard L, Thomsen C, Henriksen O. Quantification of complex flow using MR phase imaging: a study of parameters influencing the phase/velocity relation. Magn Reson Imaging. 1992;10: 13–23. 60. Tang C, Blatter DD, Parker DL. Accuracy of phase-contrast flow measurements in the presence of partial volume effects. J Magn Reson Imaging. 1993;3:377–385. 61. Buonocore MH, Bogren H. Factors influencing the accuracy and precision of velocity-encoded phase imaging. Magn Reson Med. 1992;26:141–154. 62. Brenner LD, Caputo GR, Mostbeck G, et al. Quantification of left to right atrial shunts with velocity-encoded cine nuclear magnetic resonance imaging. J Am Coll Cardiol. 1992;20:1246–1250. 63. Hundley WG, Li HE, Lange RA, et al. Assessment of left-to-right intracardiac shunting by velocity-encoded, phase-difference magnetic resonance imaging: a comparison with oximetric and indicator dilution techniques. Circulation. 1995;91:2955–2960. 64. Duerinckx AJ, Higgins CB. Valvular heart disease. Radiol Clin North Am. 1994;32:613–630.

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CHAPTER 12

Comprehensive Cardiovascular Magnetic Resonance in Coronary Artery Disease Sven Plein

Cardiovascular magnetic resonance (CMR) imaging is a highly versatile imaging modality, capable of assessing several manifestations of coronary artery disease (CAD).1,2 Furthermore, CMR data are acquired in freely definable and highly reproducible imaging planes, so the different components of a CMR study can be readily registered and correlated. These unique characteristics of CMR have led to the expectation that the technology will provide a comprehensive assessment of CAD in a single imaging study. However, until recently, the realization of such comprehensive CMR studies has been hindered by long image acquisition times and a lack of studies demonstrating the clinical effectiveness of CMR. In the past years, these limitations have been largely overcome. Several CMR methods have now been shown to provide clinically relevant and often complementary information in CAD patients.3–12 At the same time, advances in technology and acquisition techniques have shortened scan durations for most CMR methods13–15 More data can therefore be acquired within tolerable examination times, and protocols integrating several CMR methods have become feasible. It is expected that such comprehensive protocols will further improve the clinical utility and diagnostic accuracy of CMR in the management of patients with CAD. In this chapter, the most relevant CMR methods for the assessment of CAD are briefly discussed, with particular emphasis on their acquisition times and suitability for a single-session CMR study. This is followed by a discussion of current comprehensive protocols for typical clinical scenarios.

CARDIOVASCULAR MAGNETIC RESONANCE IN CORONARY ARTERY DISEASE The following measurements are of relevance to the management of most patients with CAD: cardiac morphology, global and regional myocardial function, myocardial ischemia, viability status of dysfunctional myocardium, and the presence of flow-limiting coronary stenoses. CMR can provide data in all of these aspects of CAD, and acquisition times have been reduced for most of the relevant CMR methods (Table 12-1). 158 Cardiovascular Magnetic Resonance

Morphology and Function CMR can accurately assess cardiac morphology, global and regional cardiac function in normal as well as deformed ventricles.16,17 Cine imaging forms an essential component of any CMR study in CAD. With the introduction of steadystate free precession (SSFP) pulse sequences, data acquisition times have been reduced to around 5 minutes for acquisition of a stack of 10 to 12 two-dimensional (2D) short axis sections.13,14 Combined with spatial or spatiotemporal undersampling methods, acquisition times can be reduced further, albeit at the expense of signal-to-noise ratio. These acceleration methods can also be employed to achieve either 2D real-time or three-dimensional (3D) whole heart acquisition.15,18 Combined with myocardial tagging, cine CMR imaging can provide a further detailed assessment of regional cardiac function in similar acquisition times.19

Ischemia Myocardial ischemia as the principal manifestation of CAD can be detected by two different CMR techniques: contrastenhanced myocardial perfusion and dobutamine-stress MR (DSMR). The place of myocardial perfusion imaging in a comprehensive CMR study, and its duration, are determined mainly by the need for patient preparation and monitoring, the length of the administration of the stress agent (3 to 4 minutes for adenosine as the preferred pharmacologic stress agent),20,21 and some time to allow for hemodynamics and symptoms to recover after the study. Approximately 10 minutes should therefore be considered for an adenosine stress perfusion acquisition. With the use of a variety of acquisition protocols, the accuracy of CMR in the published literature is at least comparable to that of nuclear scintigraphy.3,8,10,22–25 Dobutamine stress CMR (DCMR) detects inducible wall motion abnormalities with accuracy at least equal to that of stress echocardiography (See Chapter 15.)9,26,27 The duration of a DCMR study is determined primarily by the need to increase the dose of dobutamine incrementally in steps of 5 to 10 mg/kg/min up to 30 or 40 mg/kg/min. Data acquisition itself, usually using 2D SSFP pulse sequences and acquiring several orthogonal views, is rapidly accomplished at each stress level. Recently, 3D cine imaging has been proposed to

Indication

CMR method

Pulse sequence

Wall motion Ischemia

Cine function or tagging DSMR Perfusion LGE DSMR Angiography

SSFP SSFP T1 GRE IR GRE SSFP T2 prep, GRE or SSFP, navigator or breath hold Velocity encoded GRE

Viability Coronary imaging

Flow

Acquisition time/slice

Total acquisition time

5-15 sec 5-15 sec <300 ms 10-15 sec 5-15 sec n/a

3-5 min 15 min (for stress administration) 10 min (for stress administration) 10 min 15 min (included in DCMR ischemia) 1–40 min

20 sec

n/a

DSMR, dobutamine stress MR; SSFP, steady-state free precession; LGE, late gadolinium enhancement CMR; GRE, gradient recalled echo; IR, inversion recovery; T2 prep, T2 prepared.

cover the entire heart at each stress level.28 Including setup and recovery times, around 20 minutes should be scheduled for a complete dobutamine stress study.

Viability Morphologic parameters such as wall thickness and wall thickening, in particular in response to low-dose inotropic stimulation, can provide information on the likelihood of functional recovery.29 This information can be obtained in the early stress stages of a DCMR protocol at no additional scan time. Late gadolinium enhancement CMR (LGE) images of myocardial scar using current segmented inversion recovery gradient echo pulse sequences can be obtained in a breath hold and a complete LGE study covering the whole heart in multiple 2D sections can be accomplished in no more than 10 minutes.5,30 Three-dimensional approaches for LGE have recently been presented that provide full cardiac coverage in a single breath hold.31 In integrating LGE into a comprehensive CMR study, the key consideration concerns the necessary delay between contrast injection and data acquisition. This time can be used to acquire other image data.

predictable, but only a fraction of the data can be acquired within a breath hold in comparison with navigator-gated techniques so that the achievable signal-to-noise ratio and resolution are lower.

COMPREHENSIVE CMR ASSESSMENT OF CORONARY ARTERY DISEASE General Considerations Definitions The term comprehensive suggests that a CMR protocol provides all or at least the majority of the relevant information for a patient presenting with known or suspected CAD. In practice, imaging protocols will always be targeted to answer a particular clinical question and might not be comprehensive in a general sense. Because CMR protocols will vary from patient to patient, depending on the clinical indication, a universal comprehensive CMR protocol cannot be prescribed. The protocols suggested below should therefore be seen as general guidance providing a framework that requires adjustments according to the clinical questions posed.

Coronary Artery Imaging

Selection of Methods

Coronary artery MR angiography (MRA) delineates the anatomy of the coronary arteries and detects at least proximal coronary stenosis.4,16,32 Acquisition times are more variable than for other CMR methods. Free-breathing techniques based on respiratory navigators allow acquisition of large high-resolution 3D datasets aligned with individual coronary arteries (targeted approach) or covering the entire heart (whole heart approach) but are also the most timeconsuming. Nominal acquisition times for the currently favored whole heart acquisitions are in the order of 5 to 10 minutes, giving a total acquisition time of 10 to 20 minutes at an average acquisition efficiency of 50%.33 However, in individual patients, efficiency can be much lower, so scan times may exceed 30 minutes. This variability in the acquisition times makes it challenging to incorporate navigator-gated MR coronary angiography into a comprehensive protocol. Breath hold acquisitions are more

Testament to the versatility of CMR, a comprehensive CMR protocol for CAD can be achieved in several different ways. One of the principal choices concerns the type of stress agent used. Inotropic stress can be used for assessment of viability by stimulation of functional reserve and for ischemia detection through inducible wall motion abnormalities. Vasodilator stress is particularly suited for myocardial perfusion imaging and can be usefully combined with LGE for viability assessment. In a head-to-head comparison, DCMR was found to be more accurate than stress perfusion CMR to detect ischemia.34 However, for comprehensive CMR protocols, most authors have favored adenosine-based perfusion imaging over DCMR. Several reasons may account for this preference: First, as the initial step in the ischemic cascade, perfusion defects occur before wall motion abnormalities, making perfusion imaging theoretically more sensitive than DCMR to detect ischemia. Cardiovascular Magnetic Resonance 159

12 COMPREHENSIVE CARDIOVASCULAR MAGNETIC RESONANCE IN CORONARY ARTERY DISEASE

Table 12-1 CMR Methods Available for the Assessment of CAD

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

Second, vasodilator stress is associated with fewer complications than inotropic stress, which is of particular relevance in the MR environment.35,36 Third, the administration protocol for vasodilators is less timeconsuming and hemodynamic effects recover more quickly than they do after inotropic stress.

Contrast Agent Delivery In many comprehensive CMR protocols, contrast agents are delivered in several separate boluses, for example, to acquire two perfusion measurements and LGE data. Such a split dose regime requires some consideration: 1. If two perfusion studies are acquired within one imaging session, they should be separated in time as much as possible to reduce the effects of residual contrast on the latter acquisition. Acquiring the stress perfusion data first is usually preferred. 2. While first pass perfusion images are most commonly obtained during the bolus injection of 0.05 to 0.1 mmol/kg of a gadolinium-based contrast agent, LGE data are best acquired 10 to 15 minutes after a dose of 0.1 to 0.3 mmol/kg of contrast agent. In the current literature, these higher doses are often obtained by giving additional contrast boluses following perfusion studies. However, how splitting the contrast delivery affects the LGE images has not been studied. 3. The intervals between the various contrast injections for perfusion imaging and LGE can be used to acquire cine or coronary images. The presence of the residual contrast agent will have an effect on those acquisitions that need to be considered in the pulse sequence design and image interpretation.

Comprehensive CMR Protocols Comprehensive CMR imaging protocols have been mainly applied to two clinical scenarios: the detection of CAD and the assessment of viability. Table 12-2 lists some of the studies that have integrated multiple CMR methods.

Detection of CAD One of the first reports of a comprehensive CMR protocol came from Kramer and coworkers.37 Twenty-seven patients were studied after a first myocardial infarction. The protocol included tagged cine imaging of regional LV function, coronary flow assessment, and resting perfusion. Twentythree patients completed the protocol in a mean scan time of 46 5 minutes. In 2000, Sensky and coworkers presented a comprehensive study incorporating cine imaging at rest and during low-dose DCMR as well as myocardial perfusion imaging at rest and during adenosine stress.38 This protocol was applied to six patients with CAD and was completed in a mean acquisition time of 49 6 minutes, but the small sample size prevented further evaluation. With the advent of improved LGE methods, later reports have tended to avoid using two different stress agents in a single imaging session. In 2003, Kwong and coworkers used CMR to detect the presence of acute coronary syndrome (ACS) in 161 consecutive patients presenting to the emergency room with chest pain.39 The CMR protocol commenced with perfusion imaging at rest, followed by cine imaging. An additional bolus of contrast agent was given after the perfusion study to provide a total of 0.2 mmol/kg for subsequent LGE. The protocol was completed in 38 12 minutes. The sensitivity and specificity for detecting ACS by CMR were 84% and 85%, respectively. Figure 12-1 shows an example from this study. A subsequent study by Ingkanisorn was performed in a further 135 patients presenting to the emergency room with chest pain.40 The CMR protocol consisted of resting cine imaging, adenosine stress perfusion, and LGE. The duration of the CMR studies was approximately 45 minutes (personal communication). Of the three CMR components, stress perfusion abnormalities showed the best diagnostic yield, with a reported 100% sensitivity and 93% specificity to detect CAD. LGE and resting wall motion were highly specific (97% and 96%, respectively) but relatively insensitive in this context (55% and 70%, respectively). An abnormal CMR result of any of the components added significant prognostic value in predicting future diagnosis of CAD, myocardial infarction (MI), or death over clinical risk factors.

Table 12-2 Overview of Comprehensive CMR Protocols in the Literature Rest Function Kramer 1997 Sensky 2000 Kwong 2003 Plein 2002, 2004 Chiu 2003 Fenchel 2005 Bodi 2005 Ingkanisorn 2006

Rest Perfusion

Stress Perfusion

LGE

Low Dose DCMR

Coronary CMR

3 3 1 1 2 2 3 -

4 4 3 3 2

3 5 4 4 4 3

2 2 -

2 (coronary flow) 3 -

1 (tagging) 1 2 2 1 1 1 1

Imaging Time (min) 46 49 38 63 Less than 90 50 60 45

The numbers indicate the order in which the components were acquired in each study. The imaging time is the mean time quoted for completion of the whole study (where available). DCMR, dobutamine stress MR; LGE, late gadolinium enhancement CMR.

160 Cardiovascular Magnetic Resonance

B

C

Figure 12-1 Image example. CMR images from a 68-year-old female who presented to the emergency department with chest discomfort for 2 hours. She had no risk factors for CAD. Her electrocardiogram showed nonspecific changes, and troponin levels were normal. CMR was performed within 6 hours of presentation. A, Cine images show a wall motion abnormality in the anteroseptal wall (arrow in systolic frame). B, Resting perfusion shows a matching perfusion defect (arrow). C, LGE shows no scar in the dysfunctional segments (arrow). A diagnosis of anteroseptal ischemia at rest with preserved viability can be made on the basis of these images with a recommendation for urgent coronary angiography and revascularization. Images courtesy of R. Kwong, Brigham and Women’s Hospital, Boston, MA.

Two studies have incorporated both resting and stress perfusion imaging into single-session protocols and combined these with LGE and functional cine imaging.41,42 Both administered an additional bolus of 0.1 mmol/kg of a gadolinium-based contrast agent after the second perfusion study in preparation for LGE. Chiu and coworkers studied patients with acute coronary syndromes, completing resting cine imaging, rest and stress perfusion imaging, and LGE in less than 90 minutes.41 Fenchel and coworkers used a similar array of acquisitions with a mean acquisition time of 50 4 minutes.42,43 The most comprehensive protocol to date has been presented by Plein and coworkers.44,45 The protocol commenced with resting perfusion, which was followed by cine imaging, a targeted coronary MRA of the right coronary artery, stress perfusion assessment, and MRA of the left coronary system, and concluded with LGE. The rest perfusion study was used to aid in the identification of perfusion artifacts. In a feasibility study, the protocol was applied to 10 patients with CAD and was completed in an average scan time of 63 minutes.44 An image example from this work is given in Figure 12-2. In a subsequent clinical study, the same protocol was applied to 72 patients who presented with non-ST elevation acute coronary syndrome.45 The average scan time was similar to that of the feasibility study at 62 8 minutes. In this study, the diagnostic accuracy of the four individual CMR methods (perfusion, function, viability, and coronaries) to predict the presence of CAD on subsequent coronary angiography was measured individually and by combining two or more of the methods. Similar to the study by Ingkanisorn and coworkers, perfusion was the most accurate individual predictor of the presence of CAD (sensitivity 87%, specificity 83%) followed by coronary MRA (sensitivity 84%, specificity 75%). Wall motion and LGE were specific (75% and 83%, respectively) but

less sensitive (68% and 57%, respectively) to detect CAD. The study showed further that a combined analysis of two or more CMR methods increased the diagnostic yield of the study. The inclusion of coronary MRA significantly increased the predictive power of a logistic regression model containing perfusion alone (w2 (8.29), p ¼ 0.004). The inclusion of wall motion or late enhancement did not add any incremental predictive power to the model. The best overall diagnostic accuracy was achieved when the observers had access to all CMR data in combination with a sensitivity of 96.4%, specificity of 83.3%, negative predictive value of 83.3%, and positive predictive value of 96.4%. Figure 12-3 illustrates the sensitivities and specificities for the separate and combined analysis of CMR methods in the studies by Ingkanisorn and coworkers40 and Plein and coworkers.44 In both studies, the combined analysis of several CMR methods yielded very high diagnostic accuracy. However, because both studies were acquired in patients with acute coronary syndromes, these data might not be applicable to other patient groups. In summary, a comprehensive CMR study for the detection of CAD should include cine imaging at rest, adenosine or dipyridamole stress perfusion, and LGE (Fig. 12-4). Such a study can be completed in approximately 40 minutes. There is early evidence that the combined analysis of these methods improves the diagnostic accuracy and prognostic value over the analysis of individual CMR methods. Resting perfusion imaging may be added at little additional time expense to aid in the identification of perfusion artifacts and to allow semiquantitative measurements of rest-stress perfusion indexes. A coronary MRA study can also be incorporated and might improve the sensitivity and negative predictive value of the study in particular, but at present, it will add considerable scan time to the protocol. Cardiovascular Magnetic Resonance 161

12 COMPREHENSIVE CARDIOVASCULAR MAGNETIC RESONANCE IN CORONARY ARTERY DISEASE

A

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

A

B

C

D

E

F

G

H

162 Cardiovascular Magnetic Resonance

Figure 12-2 Image example. Patient presenting with a recent history of increasing central chest pain and previous myocardial infarction. Her electrocardiogram showed inferior Q-waves and T-wave inversion. Troponin levels were raised, consistent with a diagnosis of NSTEMI (non-ST elevation myocardial infarction). The patient underwent CMR within 48 hours from presentation prior to coronary angiography. The CMR images show akinesia of the inferior wall (arrow in panel A) with evidence of scar on the LGE images (B). There is an inferior perfusion deficit (full arrows) present at rest (C) and stress (D). In addition, there is subendocardial anteroseptal inducible ischemia that is present only on the stress image (dotted arrows). Coronary MRA shows a lesion in the mid LAD (arrow in panel E) and an occluded RCA (arrow in panel G). The CMR study thus suggests chronic inferior myocardial infarction with probable chronic occlusion of the RCA and anteroseptal ischemia. Revascularization to the LAD was recommended. The subsequent X-ray angiogram confirmed the CMR findings (arrows in panels F and H).

100 80 60 40 20 0 Coronaries Perfusion

A

Wall motion

LGE

Combined

LGE

Combined

SPECIFICITY 100 80 60 40 20 0 Coronaries Perfusion

B

Wall motion

Ingkanisorn

Plein

Figure 12-3 Sensitivity (A) and specificity (B) of individual CMR methods and combined analysis from studies by Ingkanisorn40 and Plein.44 The study by Ingkanisorn did not include coronary CMR imaging.

Several authors have compared or combined CMR methods to delineate myocardial viability and predict recovery of dysfunctional myocardium. In 2000, Lauerma and coworkers described a protocol that used cine imaging at rest, followed by perfusion and cine imaging during low-dose dobutamine infusion.46 After a 5-minute delay, LGE images were obtained by using a now outdated technique. In 10 patients, the ability to predict recovery of dysfunctional segments at 6 months was assessed. CMR was completed in 50 10 minutes. Lowdose dobutamine alone yielded a sensitivity of 79% and a specificity of 93%. LGE yielded 62% and 84%, respectively, but these results cannot be compared with the improved modern LGE methods. Perfusion imaging showed a sensitivity of 97% with a sensitivity of 91%, while the combination of both parameters yielded 97% and 96%, respectively. Kaandorp and colleagues performed a head-to-head comparison between LGE and low-dose dobutamine CMR.47 The protocol commenced with cine imaging at rest, followed by bolus injection of contrast and LGE and finally low-dose dobutamine cine imaging. The study time was not recorded. The authors found direct correlations between the transmural extent of scar tissue and contractile reserve during low-dose dobutamine. Interestingly, agreement between the two methods was not good in the segments with an intermediate extent of scar tissue (50 to 75%), with 42% of segments showing contractile reserve during low-dose dobutamine stimulation and 58% of segments showing no response. Although no follow-up after revascularization was carried out in this study, the authors speculate that low-dose DSMR may allow a further characterization of this intermediate-viability group on

Top-up bolus contrast agent

5

10

15

20

25

30

35

40

45

50

Late enhancement

Resting perfusion

Coronary imaging

Resting function

Perfusion

Scout images 0

Stress infusion

Additional imaging

Contrast agent bolus

Contrast agent bolus

55

60

Time (min) Figure 12-4 Suggested comprehensive CMR protocol for detection of CAD. The width of the blocks indicates the approximate duration of data acquisition. Dark orange blocks indicate that the component should be included in all comprehensive studies; light orange blocks indicate optional components. Following scout images to determine the main cardiac axes, scanning commences with a stress perfusion study. This is followed by assessment of cardiac function at rest. Optionally, coronary images and a rest perfusion study may next be acquired. Immediately following the last perfusion study, a top-up dose of contrast agent should be given to achieve the total required dose for LGE. The delay of 10 to 15 minutes from this injection to the LGE acquisition can be used for further imaging, such as further coronary angiography or coronary flow measurements. The protocol concludes with LGE. Cardiovascular Magnetic Resonance 163

12 COMPREHENSIVE CARDIOVASCULAR MAGNETIC RESONANCE IN CORONARY ARTERY DISEASE

Viability Assessment

SENSITIVITY

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

late-enhancement CMR and may be a useful addition to an imaging protocol in these patients. A similar study by Wellnhofer and colleagues compared DCMR and LGE in predicting functional recovery.48 In 29 patients, LGE and DCMR were performed. No detailed information about the study protocol or duration is provided. DCMR predicted functional recovery better than LGE, and LGE added complementary information to DCMR but not vice versa. The most comprehensive study to date has been presented by Bodi and coworkers, who assessed four CMRderived myocardial viability indexes in 40 patients one week after a first ST-segment elevation MI.49 All had angiographically proven open infarct-related arteries. Cine imaging at rest and during low-dose dobutamine stress, resting perfusion with a bolus of 0.1 mmol/kg of gadolinium, and LGE were carried out. No additional contrast was administered for LGE. The acquisition time for this protocol was 60 12 minutes, and all patients completed the entire study. The authors found widely different ability to predict segmental functional recovery for the four parameters tested. Wall thickness on cine images greater than 5.5 mm and normal perfusion were the most sensitive (95% and 80%, respectively) and wall thickening with low-dose dobutamine greater than 2 mm as well as transmural extent of scar less than 50% on LGE images the most specific (93% each). In a multivariate analysis, LGE had the highest predictive power, followed by DCMR. On the basis of the multivariate analysis, the authors generated a comprehensive score combining the four variables. No functional recovery was seen if LGE showed transmural necrosis and all other indexes were negative, 17% when only wall thickness and/or perfusion were positive, and 45% when response to dobutamine was detected. If LGE showed nontransmural necrosis, the rate of functional recovery was 77% in the absence of response to dobutamine and 95% in its presence. Figures 12-5 and 12-6 show examples from this study, highlighting the ability of different CMR methods to predict the likelihood of functional recovery after myocardial infarction. In summary, despite some controversy, the current evidence is in favor of using LGE as the principal method for viability assessment. Figure 12-7 illustrates a possible imaging protocol. During the administration of the contrast agent for LGE, resting perfusion images can be acquired. Resting cine imaging of wall motion can then be performed until LGE is carried out at an appropriate time delay of 10 to 15 minutes from the contrast injection. In patients with intermediate transmurality of scar, low-dose dobutamine imaging can be considered for further viability characterization.

Analysis of Comprehensive CMR Studies The wealth of data provided by a comprehensive CMR study poses several analytic challenges. In which order should the data be reviewed? How should the results of the different study components be combined? Should the complete study be read as abnormal if one component is 164 Cardiovascular Magnetic Resonance

abnormal or only if all components show matching abnormalities? How should the study be interpreted if results from the different CMR methods are discrepant? The answers to these questions depend on the expected diagnostic accuracy of each CMR method and the clinical context in which the test is undertaken. By varying the analysis algorithm, the achievable sensitivity, specificity, and predictive value of the combined analysis can be modified. For example, if the scan is requested for the noninvasive screening of patients at risk of CAD, a high negative predictive value will be paramount. A combined CMR analysis should thus interpret the study as abnormal if any one of the components shows an abnormality. If the clinical question is whether a patient with dysfunctional myocardium will benefit from revascularization, a higher positive predictive value may be more relevant, and the study should be interpreted as positive only if several measurements of viability are normal.

Detection of Coronary Artery Disease Fuster and Kim35 have suggested an analysis algorithm for the detection of CAD from a CMR study comprising rest and stress perfusion and LGE. Such algorithms become more complex when cine and coronary imaging are also included. Figure 12-8 illustrates a potential pathway for the analysis of such comprehensive data, which is based on the suggestion of Fuster and Kim. Several assumptions have been made in generating this algorithm: LGE is highly specific but (unless in the presence of previous MI) not sensitive to detect CAD. Perfusion is sensitive and specific as long as the study is of good quality. Coronary MRA at present is less accurate than perfusion imaging but can be used for confirmation if other results are inconclusive. A positive LGE result with a typical ischemic pattern should thus lead to a diagnosis of CAD. If the LGE study is negative or shows a pattern that is not consistent with CAD, perfusion as the next most accurate method should be reviewed. If the stress perfusion study can be confidently interpreted as abnormal with a clearly normal rest perfusion study, a diagnosis of CAD can be made. If both perfusion studies are positive with a normal LGE study, the perfusion images should be interpreted as showing an artifact. Alternatively, coronary CMR, if available, and cine images can be reviewed at this point to increase the sensitivity and negative predictive value of the analysis.

Viability Studies If LGE shows no enhanced myocardium, the tissue is viable, and functional recovery after revascularization is likely, providing the cause of dysfunction is CAD. Conversely, complete transmural scar suggests that functional recovery cannot be expected. For intermediate degrees of scar (e.g., between 25% and 75% of the myocardium), other methods may be used to aid the analysis.50 If available, low-dose DCMR images should initially be reviewed, and functional improvement should be regarded as a strong predictor of recovery after revascularization. Other available data, such as first pass perfusion and wall thickness, can also be used if the findings on LGE and DCMR are inconclusive (Fig. 12-9).

Diastole

B

Systole

C

Perfusion

D

LGE

E

Diastole 6 months

F

Systole 6 months

Figure 12-5 Image example. CMR acquired in a patient within 6 hours of a reperfused anterior myocardial infarction. Cine images in diastole (A) and systole (B) acquired during low-dose dobutamine infusion show preserved wall thickness but impaired wall thickening in the anteroseptal segments. C, Resting perfusion images demonstrate a matching full-thickness perfusion defect. D, LGE image shows a transmural scar in the same territory. All CMR components suggest that revascularization is unlikely to improve segmental cardiac function. E, F, The follow up cine images at 6 months show that there is now thinning of the infarct area with persistent impairment of wall thickening. Images courtesy of Vicente Bodı´, Hospital Clı´nico y Universitario de Valencia, Valencia, Spain.

Cardiovascular Magnetic Resonance 165

12 COMPREHENSIVE CARDIOVASCULAR MAGNETIC RESONANCE IN CORONARY ARTERY DISEASE

A

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

A

Diastole

B

C

Perfusion

D

E

Diastole 6 months

F

Systole

LGE

Systole 6 months

Figure 12-6 Image example. CMR acquired in a further patient within 6 hours of anterior myocardial infarction. A, B, Cine images during low-dose dobutamine infusion (diastole in panel A, systole in panel B) show preserved wall thickness and moderately impaired wall thickening. C, Resting perfusion images show no regional perfusion defect. D, The LGE image shows a small subendocardial scar. CMR analysis suggests that segmental cardiac function is likely to improve following revascularization, with all CMR methods showing consistent findings. E, F, Follow-up cine images 6 months after percutaneous coronary revascularization to the LAD artery show that regional cardiac function has improved. Images courtesy of Vicente Bodı´, Hospital Clı´nico y Universitario de Valencia, Valencia, Spain.

166 Cardiovascular Magnetic Resonance

Top-up bolus contrast injection

0

5

10

15

20

25

Stress function

Late enhancement

Resting function

Scout images

Resting perfusion

Contrast agent bolus

Dobutamine infusion 30

35

40

45

Time (min)

LGE +

–

–

–

Rest perfusion

+

Regional WM Coronary MRA

CAD

+

Intermediate transmurality

Stress perfusion +

LGE

– –

+

Low-dose DCMR

+

– No CAD

Figure 12-8 Analysis of comprehensive CMR data for the detection of CAD. LGE images should be reviewed first. If these show a typical ischemic pattern, a diagnosis of CAD can be made. If the LGE study is negative or shows a pattern that is not consistent with CAD, stress perfusion images should be reviewed. If the stress perfusion study is abnormal, the rest perfusion study, if available, can be reviewed to rule out an image artifact. A perfusion defect at stress with a normal rest study will lead to a diagnosis of CAD. If both perfusion studies show the same defect and the LGE study is normal, the perfusion images should be interpreted as showing an artifact. If available, coronary MRA and regional wall motion can be included in the analysis at this point. WM, wall motion.

CONCLUSION CMR today allows a comprehensive assessment of CAD, providing a wide range of clinically relevant data. Early results suggest that the combined acquisition and analysis of several CMR methods increases the diagnostic accuracy of a comprehensive CMR study. For the detection of CAD, such a study should incorporate cine imaging at rest, perfusion imaging during vasodilator stress, and LGE. This protocol can be completed in less than 1 hour. The wide range of data provided by such a study could become a cost-effective alternative to the

Recovery

–

Wth < 5.5 mm Rest perfusion

+

No recovery

Figure 12-9 Analysis of comprehensive CMR data for viability studies. LGE images should be reviewed first. If there is no evidence of scar tissue, the myocardium is viable, and functional recovery after revascularization can be expected. Transmural scar on LGE suggests that functional recovery is unlikely. For intermediate degrees of scar, other methods can be used to aid the analysis. If available, low-dose DCMR images should initially be reviewed with functional improvement on stress images being a strong predictor of recovery after revascularization. Other available data, such as first pass perfusion and wall thickness, can also be used if the findings on LGE and DCMR are inconclusive. Wth, wall thickness.

current use of multiple other investigations.51 For viability assessment, the combination of several morphologic and functional CMR methods into a 45-minute protocol provides a highly accurate tool to determine the potential for recovery of myocardial contractility.

FUTURE DIRECTIONS Further increases in data acquisition speed and image quality can be expected in the coming years. These should be of particular benefit to comprehensive CMR protocols. In Cardiovascular Magnetic Resonance 167

12 COMPREHENSIVE CARDIOVASCULAR MAGNETIC RESONANCE IN CORONARY ARTERY DISEASE

Figure 12-7 Comprehensive CMR protocol for viability assessment. The width of the blocks indicates approximate duration of data acquisition. Dark orange blocks indicate that the component should be included in all comprehensive studies; light orange blocks indicate optional components. Following scout images to determine the main cardiac axes, scanning may commence with a rest perfusion study. Immediately following this, a top-up dose of contrast agent should be given to achieve the total required dose for LGE. The delay of 10 to 15 minutes from this injection to the LGE acquisition can be used to acquire cine images of cardiac function at rest. LGE should then be acquired. If intermediate scar is present on LGE images, a dobutamine infusion may be commenced, and further cine images may be acquired during low-dose dobutamine infusion.

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

particular, the development of 3D methods for cine and viability imaging will further shorten scan times and permit comprehensive assessment in shorter scan times. So far, comprehensive CMR imaging in CAD has been evaluated only in small single-center studies in dedicated research institutions. A variety of imaging protocols with important differences have been used in those studies. Larger clinical trials and multicenter studies will be needed to further

determine the clinical role of comprehensive CMR protocols and to establish the optimal combination of CMR methods in different clinical scenarios. Standardized protocols for image acquisition and analysis need to be developed and validated in large-scale clinical trials. Finally, the cost effectiveness of comprehensive CMR in comparison with established and other evolving management strategies needs to be determined.

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19. 20. 21.

22.

23. 24. 25.

26.

27. 28. 29. 30. 31.

32. 33. 34.

35. 36.

space broad-use linear acquisition speed-up technique (k-t BLAST). J Magn Reson Imaging. 2008;27:510–515. Castillo E, Lima JA, Bluemke DA. Regional myocardial function: advances in MR imaging and analysis. Radiographics. 2003;23 Spec No: S127–S140. Pennell DJ. Cardiovascular magnetic resonance and the role of adenosine pharmacologic stress. Am J Cardiol. 2004;94(suppl): 26D–32D. Hilleman DE, Lucas Jr BD, Mohiuddin SM, Holmberg MJ. Costminimization analysis of intravenous adenosine and dipyridamole in thallous chloride TI 201 SPECT myocardial perfusion imaging. Ann Pharmacother. 1997;31:974–979. Ishida N, Sakuma H, Motoyasu M, et al. Noninfarcted myocardium: correlation between dynamic first-pass contrast-enhanced myocardial MR imaging and quantitative coronary angiography. Radiology. 2003;229:209–216. Manning WJ, Atkinson DJ, Grossman W, et al. First-pass nuclear magnetic resonance imaging studies using gadolinium-DTPA in patients with coronary artery disease. J Am Coll Cardiol. 1991;18:959–965. Wilke N, Jerosch-Herold M, Wang Y, et al. Myocardial perfusion reserve: assessment with multisection, quantitative, first-pass MR imaging. Radiology. 1997;204:373–384. Schwitter J, Wacker CM, van Rossum AC, et al. MR-IMPACT: comparison of perfusion-cardiac magnetic resonance with single-photon emission computed tomography for the detection of coronary artery disease in a multicentre, multivendor, randomized trial. Eur Heart J. 2008;29:480–489. Hundley WG, Hamilton CA, Thomas MS, et al. Utility of fast cine magnetic resonance imaging and display for the detection of myocardial ischemia in patients not well suited for second harmonic stress echocardiography. Circulation. 1999;100:1697–1702. Hundley WG, Morgan TM, Neagle CM, Hamilton CA, Rerkpattanapipat P, Link KM. Magnetic resonance imaging determination of cardiac prognosis. Circulation. 2002;106:2328–2333. Jahnke C, Paetsch I, Gebker R, Bornstedt A, Fleck E, Nagel E. Accelerated 4D dobutamine stress MR imaging with k-t BLAST: feasibility and diagnostic performance. Radiology. 2006;241:718–728. Heusch G, Schulz R. Hibernating myocardium: a review. J Mol Cell Cardiol. 1996;12:2359–2372. Simonetti OP, Kim RJ, Fieno DS, et al. An improved MR imaging technique for the visualization of myocardial infarction. Radiology. 2001;218:215–223. Peters DC, Appelbaum EA, Nezafat R, Dokhan B, Han Y, Kissinger KV, Goddu B, Manning WJ, et al. Left ventricular infarct size, peri-infarct zone, and papillary scar measurements: a comparison of high-resolution 3D and conventional 2D late gadolinium enhancement cardiac MR. J Magn Reson Imaging. 2009;30:794–800. Manning WJ, Li W, Edelman RR. A preliminary report comparing magnetic resonance coronary angiography with conventional angiography. N Engl J Med. 1993;328:828–832. Sakuma H, Ichikawa Y, Suzawa N, et al. Assessment of coronary arteries with total study time of less than 30 minutes by using whole-heart coronary MR angiography. Radiology. 2005;237:316–321. Paetsch I, Jahnke C, Wahl A, et al. Comparison of dobutamine stress magnetic resonance, adenosine stress magnetic resonance, and adenosine stress magnetic resonance perfusion. Circulation. 2004;110: 835–842. Fuster V, Kim RJ. Frontiers in cardiovascular magnetic resonance. Circulation. 2005;112:135–144. Nagel E, Lorenz C, Baer F, et al. Stress cardiovascular magnetic resonance: consensus panel report. J Cardiovasc Magn Reson. 2001;3: 267–281.

45. Plein S, Greenwood JP, Ridgway JP, Cranny G, Ball SG, Sivananthan MU. Assessment of non–ST-segment elevation acute coronary syndromes with cardiac magnetic resonance imaging. J Am Coll Cardiol. 2004;44:2173–2181. 46. Lauerma K, Niemi P, Ha¨nninen H, et al. Multimodality MR imaging assessment of myocardial viability: combination of first-pass and late contrast enhancement to wall motion dynamics and comparison with FDG PET-initial experience. Radiology. 2000;217:729–736. 47. Kaandorp TA, Bax JJ, Schuijf JD, et al. Head-to-head comparison between contrast-enhanced magnetic resonance imaging and dobutamine magnetic resonance imaging in men with ischemic cardiomyopathy. Am J Cardiol. 2004;93:1461–1464. 48. Wellnhofer E, Olariu A, Klein C, et al. Magnetic resonance low-dose dobutamine test is superior to scar quantification for the prediction of functional recovery. Circulation. 2004;109:2172–2174. 49. Bodi V, Sanchis J, Lo´pez-Lereu MP, et al. Usefulness of a comprehensive cardiovascular magnetic resonance imaging assessment for predicting recovery of left ventricular wall motion in the setting of myocardial stunning. J Am Coll Cardiol. 2005;46:1747–1752. 50. Van der Wall EE, Bax JJ. Letter regarding article by Selvanayagam et al, “Value of delayed-enhancement cardiovascular magnetic resonance imaging in predicting myocardial viability after surgical revascularization”. Circulation. 2005;111:e286. 51. Hunink MG, Kuntz KM, Fleischmann KE, Brady TJ. Noninvasive imaging for the diagnosis of coronary artery disease: focusing the development of new diagnostic technology. Ann Intern Med. 1999;131:673–680.

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37. Kramer CM, Rogers WJ, Geskin G, et al. Usefulness of magnetic resonance imaging early after acute myocardial infarction. Am J Cardiol. 1997;80:690–695. 38. Sensky PR, Jivan A, Hudson NM, et al. Coronary artery disease: combined stress MR imaging protocol-one-stop evaluation of myocardial perfusion and function. Radiology. 2000;215:608–614. 39. Kwong RY, Schussheim AE, Rekhraj S, et al. Detecting acute coronary syndrome in the emergency department with cardiac magnetic resonance imaging. Circulation. 2003;107:531–537. 40. Ingkanisorn WP, Kwong RY, Bohme NS, et al. Prognosis of negative adenosine stress magnetic resonance in patients presenting to an emergency department with chest pain. J Am Coll Cardiol. 2006;47: 1427–1432. 41. Chiu CW, So NMC, Lam WWM, Chan KY, Sanderson JE. Combined first-pass perfusion and viability study at MR imaging in patients with non–ST segment-elevation acute coronary syndromes: feasibility study. Radiology. 2003;226:717–722. 42. Fenchel M, Helber U, Kramer U, et al. Detection of regional myocardial perfusion deficit using rest and stress perfusion MRI: a feasibility study. AJR Am J Roentgenol. 2005;185:627–635. 43. Fenchel M, Franow A, Stauder NI, et al. Myocardial perfusion after angioplasty in patients suspected of having single-vessel coronary artery disease: improvement detected at rest-stress first-pass perfusion MR imaging—initial experience. Radiology. 2005;37:67–74. 44. Plein S, Ridgway JP, Jones TR, Bloomer TN, Sivananthan MU. Coronary artery disease: assessment with a comprehensive MR imaging protocol-initial results. Radiology. 2002;225:300–307.

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CHAPTER 13

High Field Cardiovascular Magnetic Resonance Grigorios Korosoglou and Matthias Stuber

In theory, the signal-to-noise ratio (SNR) is directly related to the strength (B0) of the static magnetic field. Thus, an improved SNR can be expected for cardiovascular magnetic resonance (CMR) by using magnets with higher magnetic field strength, of which 3 tesla (3 T) systems are now available for clinical use. A higher SNR can be exploited to provide higher spatial resolution/detail, greater temporal resolution, abbreviated scanning times, or a combination of the above. This will ultimately result in an improved diagnostic value of CMR and lead to new areas of innovation and research. However, at present, the limited availability of these higher field systems equipped for cardiac applications together with other potential impediments1 including increased susceptibility artifacts at tissue borders, reduced T2*,2,3 increased T1, radiofrequency (RF) field distortions,4 and changed tissue dielectric constants or body dielectric resonances5,6 are limiting factors. Furthermore, at higher field strengths, increased RF deposition may remove flexibility for general sequence design and reliable R-wave triggering becomes more challenging, owing to the amplified magnetohydrodynamic effects.7 Figure 13-1 displays electrocardiographic (ECG) traces of the same individual recorded at four different field strengths. The gradually increasing artifact on the T-wave can readily be seen as a function of increasing B0. Such an augmented T-wave could be misinterpreted as an R-wave by conventional R-wave detection algorithms, thereby compromising image quality of any ECG-triggered CMR acquisition. Therefore, adapted ECG modules and R-wave detection algorithms are a necessity for successful cardiovascular imaging at high field strength. Despite the above-mentioned challenges, tremendous progress in hardware and the adaptation of cardiac-specific software have made it possible to implement CMR on commercial whole body 3 T systems, and these systems are expected to become the dominant CMR platform in the not-too-distant future. Contemporary results obtained with this new technology are discussed in this chapter.

APPLICATIONS Functional Imaging At 1.5 T, cine steady-state free precession (SSFP)8 techniques have been adopted by most of the clinical and research CMR centers for functional cardiac imaging. 170 Cardiovascular Magnetic Resonance

Regional and global systolic function, ventricular volume, and mass are measures that are qualitatively and quantitatively assessed.9 By using SSFP techniques at 1.5 T, a major advantage in SNR and contrast-to-noise ratio (CNR) has been obtained and the resulting improved endocardial border definition facilitated automated or semiautomated contouring.10 However, a short repetition time (TR) is crucial to minimize flow artifacts and signal voids, owing to local field inhomogeneities. Although these challenges have been successfully addressed by the vendors at 1.5 T, an increase in general DB0 effects and TR that may be governed by the maximum RF deposition are among the impediments for cine SSFP approaches at 3 T. However, careful adaptation of the RF excitation angles, higher-order shimming, and modifications of the imaging protocol to the changed physical boundary conditions have been implemented for long axis, four-chamber, and short axis functional imaging11 (Fig. 13-2). A high visual contrast between the ventricular blood pool and the surrounding myocardium is obtained. Simultaneously, excellent separation of the epicardium and the lungs will support automated delineation. It is noteworthy that no major susceptibility artifacts are observed on either of the images obtained at different anatomic levels and in different spatial orientations. These images were obtained with a body send and a six-element phased array cardiac receive coil. To minimize adverse effects of B0 inhomogeneity, localized higher-order shimming was utilized.12 In accord with these findings, studies demonstrate that imaging at 3 T using two-dimensional (2D) cine SSFP provides substantially higher SNR and CNR compared with imaging at 1.5 T, offering overall good image quality.13,14 Michaely and colleagues15 also reported a similar SNR gain for functional cardiac cine image at 3 T. In this study, spoiled gradient echo sequences seemed to benefit more from the higher field strength, being less affected by artifacts in comparison with SSFP pulse sequences. Although comparison studies with 1.5 T have not been reported, dobutamine stress CMR at 3 T is also possible.16

Myocardial Tagging CMR myocardial tagging has been shown to be a useful tool for the assessment of local myocardial motion in healthy and diseased states.17 Fading of the tags limits access to diastolic myocardial motion for most of the contemporary tagging sequences. This fading is T1 dependent. At higher

MHD

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Figure 13-1 Electrocardiogram (ECG) traces acquired in the same individual at four different field strengths (0.0 T, 0.5 T, 1.5 T, 3.0 T). Note the gradually increasing artifact on the T-wave of the ECG, which can be attributed to the amplified magnetohydrodynamic effect at higher field strength. R-wave triggering is therefore one of the challenges at 3 T. This necessitates sophisticated ECG hardware and R-wave detection algorithms for successful CMR. Source: Courtesy of Stefan Fischer.

RV

field strength, however, T1 of myocardium is increased, resulting in prolonged tag persistence. Together with the increase in SNR at higher field strength, this attenuated fading results in a higher temporal resolution and more accurate identification of the tagged lines, which facilitates identification of subtle changes in myocardial motion during the cardiac cycle. In Figure 13-3, three CSPAMM18 line-tagged cine frames out of a series with a temporal resolution of 50 frames per R-R interval are displayed. These images were acquired by using spiral imaging (8-msec spiral readouts) in conjunction with spectral spatial excitation for fat suppression and ramped flip angles for constant tagging CNR. Data have confirmed the increase of CNR at 3 T for myocardial tagging techniques.13,14,19 Adenosine myocardial stress tagging has been implemented at 3 T in combination with stress perfusion.20 Together with the increase in SNR, this has been attributed to a prolonged T1 of the myocardium at 3 T. In contrast to conventional tagging, strain-encoded CMR imaging uses tag surfaces that are not orthogonal but parallel to the image plane, combined with out-of-plane phaseencoding gradients in the perpendicular section-select direction.21 Implementation of strain encoding on higher field systems may increase the spatial and temporal resolution of this technique, thereby enhancing its sensitivity to detect ischemic disorders in patients with coronary artery disease.22

RV

LV

LV

LV

A

B

D

E

C

Figure 13-2 Steady-state free precession (SSFP) functional images acquired with parallel imaging (SENSE) and an acceleration factor of 2. To minimize adverse effects of B0 inhomogeneity, higher-order shimming was used. A, Long axis. B, four chamber. C–E, short axis at different anatomic levels. LV, left ventricle; RV, right ventricle. Cardiovascular Magnetic Resonance 171

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End-Diastole

End-Systole

Late Diastole

A

B

C

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Figure 13-3 Short axis (A–C) and four chamber (D–F) spiral CSPAMM myocardial tagging images acquired at 3 T. Myocardial tagging supports the quantification of local myocardial motion, including stress, strain, rotation shear, and so on. Consistent with the increased muscle T1 at 3 T, the persistence of the tags is prolonged, facilitating an accurate wall motion for systole and late diastole (see panels C and F ).

Coronary Artery Cardiovascular Magnetic Resonance The first international multicenter trial of coronary artery CMR demonstrated the efficacy of coronary artery CMR for the noninvasive assessment of significant multivessel coronary artery disease.23 However, visualization of the more distal segments was not possible, owing to their small caliber. It is likely that higher spatial coronary artery CMR approaching 500 mm may be required. The higher SNR gained by the higher field strength would allow for a further improved spatial resolution, which will support isotropic resolution and improved visualization of both proximal and distal segments. However, potential challenges have to be considered at higher field strength, as has already been mentioned. Therefore, a first study investigating the feasibility of in vivo human coronary artery CMR at higher field strength was performed24 using the 1.5 T CMR methodology25 adopted with only minor modifications to account for the changed physical boundary conditions at 3 T. Vector ECG technology was used,7 and a body send coil was used together with a prototype six-element cardiac synergy surface receive coil. High-resolution imaging was performed by using a three-dimensional (3D) segmented k-space gradient echo imaging sequence using a T2Prep and fat saturation for endogenous contrast enhancement26 and a 2D selective 172 Cardiovascular Magnetic Resonance

real-time navigator for respiratory motion suppression during free breathing.27 To minimize the sensitivity of the navigator 2D selective RF pulse to DB0 and T2* effects, the number of turns in k-space was reduced (from N ¼ 12 at 1.5 T25) to N ¼ 8 to shorten the excitation duration. To suppress contamination of the navigator by spurious signal originating from higher-order components (aliasing) concentric with the central 2D selective excitation, a local receive coil positioned over the sternum was selectively employed for navigator signal reception. Ten RF excitations (TR ¼ 8.2 msec; echo time ¼ 2.4 msec) were performed for imaging during each R-R interval (acquisition window ¼ 82 msec). The RF excitation angles were constant (25 ) and were adapted to account for the prolonged blood T1 at 3.0 T (1600 msec).2 A heart rate–specific individual diastolic trigger delay28 was used to minimize intrinsic myocardial motion during the acquisition interval. In-plane resolution was 0.6 to 0.7 0.6 to 1.0 mm,2 and the slice thickness was 3 mm. Coronary artery CMR studies were completed within an hour with uniform image quality and average scan duration of approximately 7 minutes. No major susceptibility artifacts were seen in the vicinity of the left or the right proximal to mid-coronary arterial segments. The image of the left coronary system shown in Figure 13-4A was acquired with a voxel size of 0.6 0.6 3 mm, and the image of the right coronary artery (RCA) shown in Figure 13-4B was acquired with a voxel size of

A

Ao

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Figure 13-4 Coronary artery CMR of the left and right coronary arterial system acquired during free breathing using a T2Prep navigator gated and corrected three-dimensional segmented k-space gradient echo imaging sequence (image resolution: 0.6 0.6 3 mm). The images were acquired with a body send coil and a prototype six-element synergy surface receive coil. A, A left coronary system including the left main (LM), 8 cm of the left anterior descending (LAD) with distal branching vessels (dashed arrow) and 4 cm of the left coronary circumflex. B, A 9-cm segment of the right coronary artery together with the left main (LM), the left coronary circumflex, and some branching vessels (dashed arrow). Ao, aorta; LV, left ventricle; PA, pulmonary artery; RV, right ventricle. Source: Stuber M, Botnar RM, Fischer SE, et al. Preliminary report on in vivo coronary MRA at 3 Tesla in humans. Magn Reson Med. 2002;48(3):425–429.

0.7 1.0 3 mm. In these images, a considerable visual contrast between the coronary blood pool and the surrounding tissue is apparent. Smaller-diameter branching vessels are seen on both images (broken arrows). Further studies demonstrated the feasibility of a radial sampling technique to acquire 3D coronary artery CMR at 3 T within a single breath hold. The combination of a radial k-space acquisition, parallel imaging, and interleaved sliding window reconstruction provided good SNR and high coronary vessel sharpness.29 In a head-to-head comparison of coronary vessel imaging at 3 T versus 1.5 T, SNR, CNR, and vessel sharpness were reported to be substantially higher at 3 T.30,31 Similarly, in another 3.0 T study, fat suppression fast gradient echo sequences32 provided high spatial resolution for the delineation of coronary arteries with good interobserver variability for quantitative vessel sharpness and CNR measurements. In a small clinical study of Figure 13-5 Coronary artery CMR of the right coronary artery obtained with a conventional T2Prep (A) and an adiabatic T2Prep (B) for contrast enhancement between the myocardium and the blood pool. By using an adiabatic T2Prep, local signal variations can be substantially minimized.

patients with suspected coronary artery disease, the SNR gain at 3.0 T (versus 1.5 T) did not translate into improved accuracy for the detection of coronary lesions.33 This has been attributed to the presence of intrinsic and extrinsic motion artifacts, to susceptibility-related local magnetic field variations, and to off-resonance effects, which linearly increase at the higher magnetic fields. However, the use of adiabatic refocusing T2 preparation sequences to suppress artifacts originating from B1 inhomogeneity can provide significantly improved definition of coronary arteries at 3 T34 (Fig. 13-5). The potential of this sequence to improve the sensitivity of coronary artery CMR for the detection of high-grade lesions remains to be evaluated in further clinical studies. Data with gadolinium contrast appear particularly promising.35 Finally, the very preliminary data on very high field (7.0 T) whole body scanners capable of coronary artery imaging show promise.36

Conventional T2 Prep

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Coronary Artery Wall Imaging Several studies at 1.5 T have already demonstrated ability to visualize the coronary vessel wall in both healthy and diseased states,37–39 including visualization of a positive arterial remodeling in patients with suspected coronary artery disease at 1.5 T.40 However, the ultimate goal would be not only the coronary vessel wall thickness measurement but also the differentiation between plaque components and consequently the identification of vulnerable plaque.37 Still, this necessitates a significantly improved spatial resolution. At higher field strength, a further step in this direction is expected. Owing to the relatively small size of the coronary vessel wall (1 mm), the central location within the thorax, and constant movement due to intrinsic and extrinsic myocardial motion, coronary vessel wall imaging is among the technically most challenging topics in CMR. At 3 T, navigator-gated and real-time motion corrected cross-sectional views of the RCA were obtained in healthy adults by Botnar and colleagues.38 The CMR pulse sequence consisted of a dual-inversion black-blood prepulse with the inversion time (TI)41 adapted for the prolonged T1 of blood (700 msec at 60 beats per minute, dual inversion, and data collection every other heartbeat) at 3 T. The 2D segmented k-space gradient echo imaging sequence was immediately preceded by the 2D selective navigator pulse, a frequency selective fat suppression prepulse, and a navigator-restore prepulse42 for optimization of navigator performance in the presence of a nonselective 180 prepulse. Other imaging parameters included echo time ¼ 2.3 msec, TR ¼ 8 msec, bandwidth ¼ 135 Hz, flip angle ¼ 30 , 10 excitations per shot, NSA ¼ 2, FOV ¼ 360 mm, image matrix ¼ 512 512, resulting in an inplane resolution of 0.7 0.7 mm and a slice thickness of 5 mm. Adjacent to the dual-inversion fast spin echo scout scan in parallel to the RCA shown in Figure 13-6A, an excellent depiction of the coronary vessel acquired at 3 T can be seen in Figure 13-6B. The prolonged inversion time of the dual-inversion prepulse yields an adequate suppression of the left and right ventricular intracavitary blood-pool.

Although a relatively SNR inefficient black-blood segmented k-space gradient echo sequence was used, these data demonstrate the feasibility of 3 T coronary artery vessel wall CMR using a free breathing black-blood 2D fast gradient echo technique with ECG triggering and navigator gating. These results have been confirmed by others using a single breath hold43 and using navigator gating.43,44 Particularly the use of adiabatic pulses for both inversion and reinversion44 has been reported to be robust to offresonance and RF inhomogeneity. However, substantial improvements in resolution and image quality are still required in order to evaluate the composition of coronary atherosclerotic plaques.

Perfusion and Late Gadolinium Enhancement Both myocardial stress perfusion and late gadolinium enhancement (LGE) CMR require administration of extracellular gadolinium-based contrast agents. Because gradient echo techniques with short echo times have been applied successfully for both modalities at 1.5 T, no major impediments are expected from a sequence standpoint at 3 T. In vitro studies demonstrate a higher gadolinium contrast enhancement at higher field strengths.45 In a study involving healthy subjects, the assessment of myocardial perfusion and viability at 3 T substantially increased SNR, CNR, and overall image quality compared with imaging at 1.5 T.13 In accord with these findings, the use of 3 T MR imaging was reported to significantly increase contrast in first pass myocardial perfusion imaging for different gadolinium doses compared with 1.5 T.46 This suggests that lower doses of contrast may be used at 3 T. Comparison adenosine stress perfusion CMR studies of 1.5 T and 3 T suggest an accuracy benefit for 3 T.47 Multicenter clinical studies will be needed to verify the added value of 3-T LGE CMR for the more accurate detection of scarred myocardial tissue. As an alternative to contrast perfusion for assessment of ischemia, noncontrast 3 T stress with myocardial blood oxygenation level dependent imaging compares favorably with positron emission tomography (PET).48

RCA wall

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174 Cardiovascular Magnetic Resonance

B

Figure 13-6 Coronary vessel wall CMR showing a cross section of the proximal right coronary artery (RCA). The localizer (A) was acquired with a dual-inversion fast spin echo sequence using navigator technology during free breathing. The coronary vessel wall CMR (B) was acquired during free breathing by using a dual-inversion navigator-gated and corrected two-dimensional segmented k-space gradient echo imaging sequence. Source: Images courtesy Rene M. Botnar, Guy’s and St. Thomas Hospital, London.

CMR spectroscopy allows the noninvasive evaluation of lipid and creatine metabolism in cardiac tissue,49,50 which may aid in the differentiation of viable myocardium from scarred tissue in ischemic heart disease. High field whole body CMR systems offer increased spectral resolution. However, shorter T2* values and increased field inhomogeneities at 3 T would necessitate higher-order shimming algorithms for the accurate acquisition of cardiac proton spectra. Schar and colleagues have implemented navigator gating and volume tracking for cardiac proton spectroscopy at 3 T incorporating local field map-based shimming. Volume tracking allowed the suppression of tissue outside the planned voxel and provided the acquisition of highquality spectra during free breathing.12

Parallel Imaging Parallel imaging techniques such as simultaneous acquisition of spatial harmonics51 or sensitivity encoding for fast MRI (SENSE)52 are particularly interesting at higher field strengths and for selected applications. While the theoretically doubled SNR at 3 T necessitates an approximately four time signal averaging at 1.5 T (to obtain similar SNR), this twofold SNR at 3 T could be traded for abbreviated scanning times using parallel imaging with an acceleration factor approaching four. In areas in which SNR is not necessarily the limiting factor (e.g., functional cardiac imaging), this helps to shorten breath hold durations and will be most beneficial in patients who cannot tolerate prolonged breath holds. In this context, SENSE has been reported to contribute to scan time reduction for coronary artery CMR.53 Shortening of the scanning time would make coronary artery CMR better suited for integration as part of a comprehensive cardiac examination. Furthermore, combining dual-stack 3D coronary artery CMR with parallel imaging54 has been shown to improve coronary artery delineation by providing shorter temporal delays between navigator, T2 preparation, and the actual image acquisitions. This approach may considerably facilitate the simultaneous acquisition of high-resolution images of the left and right coronary system. However, to the present date, a further improved resolution for both coronary artery CMR and coronary vessel wall CMR may have a higher priority than abbreviated scanning times with subsequent SNR loss as intrinsically coupled with parallel imaging. Shortened T2* values and an increase in △B0 consistent with an increased B0 will be among the limitations for techniques with prolonged signal readout (spiral techniques, EPI techniques).2 However, by using parallel imaging, echo trains or readout times can be shortened, which may support optimized image quality using such techniques at higher field strength. Sodickson and colleagues also demonstrated that parallel imaging may be employed beneficially for first pass contrast agent bolus injection techniques.55 Hereby, accelerated data acquisition results in a more time-effective signal sampling during the first pass. This also remains to be revisited at higher field strength. More recently, higher numbers of receiver channels (up to 32) became commercially available on different vendors’

CONCLUSION With state-of-the-art 3 T hardware, software, and methodology, high-quality images can be acquired in all the major areas of contemporary CMR. At some centers, 3 T CMR is already the primary scanner for clinical cases.60 Cardiovascular Magnetic Resonance 175

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Spectroscopy

platforms. The development of such large coil arrays is ongoing, but impressive early results suggest these will be quickly adopted by the CMR community.13 With the availability of increasing numbers of 3 T CMR scanners, it is likely that for the reasons reviewed above, 3 T CMR will evolve to be the dominant CMR platform. Ongoing studies are needed to optimize hardware components including body send coils, surface receive coils, coil arrays for parallel imaging, higher-order shimming algorithms, and sophisticated ECG triggering hardware to avoid adverse effects of the amplified magneto hydrodynamic effect on R-wave triggering. Cardiac-specific software with segmented k-space gradient echo imaging sequences (including SSFP and EPI), fast spin echo sequences and prepulses such as fat saturation, inversion, dual-inversion and T2Prep will support the majority of the contemporary cardiovascular applications. However, many of these prepulses or combinations thereof will undoubtedly benefit from a re-design of the RF pulses with the aim to reduce SAR and to minimize the adverse effects of B1 inhomogeneity. For high-resolution applications with prolonged equilibrium scanning times, navigator technology should be available. The expected theoretical SNR benefit at higher field strength and its consequences are obviously most attractive for enhanced spatial resolution, enhanced temporal resolution, and abbreviated scanning times or combinations thereof. However, potential challenges as mentioned in the introductory section of this chapter exist, and their individual effects on image quality remain to be studied for each application. With the higher field strength, RF deposition increases; therefore, certain limitations will apply for sequences with large RF excitation angles or continuous RF excitations over a prolonged period of time. Therefore, SSFP or fast spin echo sequences need to be adapted accordingly. For coronary artery and coronary vessel wall CMR, data can be acquired only in a narrow time window within each cardiac cycle. Although this makes CMR data acquisition rather inefficient, it offers more flexibility for high field sequence design, since RF deposition is minimized for both applications. While the increase in T1 may lead to a reduced steadystate magnetization with subsequent relative loss in SNR, this effect can in turn also be used to improve the efficacy of myocardial tagging or spin-labeling techniques, for example, which will be able to take advantage not only of the enhanced SNR at 3 T but also of the prolonged T1.56 Shortened T2* may be useful for blood oxygenation measurements57 or perfusion imaging. Together with navigator technology, it may be possible to implement cardiac spectroscopy during free breathing58 while the enhanced SNR will support spectroscopic data acquisition in relatively small and well-defined volumes. If extrinsic and intrinsic myocardial motion is sufficiently constrained, diffusion imaging together with fiber tracking may allow for a more detailed insight into fiber architecture and the orientation of myocytes.59

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

After further refinement of imaging sequences and adaptation of hardware components (larger coil arrays), the specific clinical benefits associated with the higher field strength will have to be defined as a next step. The areas

in which 3 T outperforms 1.5 T remain to be identified in direct comparisons, and the outcome of patient studies will ultimately define the value of 3 T CMR in clinical applications.

References 1. Wen H, Denison TJ, Singerman RW, Balaban RS. The intrinsic signalto-noise ratio in human cardiac imaging at 1.5, 3, and 4 T. J Magn Reson. 1997;125(1):65–71. 2. Noeske R, Seifert F, Rhein KH, Rinneberg H. Human cardiac imaging at 3 T using phased array coils. Magn Reson Med. 2000;44(6):978–982. 3. Atalay MK, Poncelet BP, Kantor HL, Brady TJ, Weisskoff RM. Cardiac susceptibility artifacts arising from the heart-lung interface. Magn Reson Med. 2001;45(2):341–345. 4. Dougherty L, Connick TJ, Mizsei G. Cardiac imaging at 4 Tesla. Magn Reson Med. 2001;45(1):176–178. 5. Wen H, Jaffer FA, Denison TJ, Duewell S, Chesnick AS, Balaban RS. The evaluation of dielectric resonators containing H2O or D2O as RF coils for high-field MR imaging and spectroscopy. J Magn Reson B. 1996;110(2):117–123. 6. Bottomley PA, Andrew ER. RF magnetic field penetration, phase shift and power dissipation in biological tissue: implications for NMR imaging. Phys Med Biol. 1978;23(4):630–643. 7. Fischer SE, Wickline SA, Lorenz CH. Novel real-time R-wave detection algorithm based on the vectorcardiogram for accurate gated magnetic resonance acquisitions. Magn Reson Med. 1999;42(2):361–370. 8. Oppelt A. FISP: a new fast MRI sequence. Electromedica. 1986;54:15–18. 9. Salton CJ, Chuang ML, O’Donnell CJ, et al. Gender differences and normal left ventricular anatomy in an adult population free of hypertension: a cardiovascular magnetic resonance study of the Framingham Heart Study Offspring cohort. J Am Coll Cardiol. 2002;39 (6):1055–1060. 10. Heid O, True FISP. Cardiac fluoroscopy. Proceedings of the International Society for Magnetic Resonance in Medicine (abstract). 1997;1:320. 11. Schar M, Kozerke S, Fischer SE, Boesiger P. Cardiac SSFP imaging at 3 Tesla. Magn Reson Med. 2004;799–806. 12. Schar M, Kozerke S, Boesiger P. Navigator gating and volume tracking for double-triggered cardiac proton spectroscopy at 3 Tesla. Magn Reson Med. 2004;51(6):1091–1095. 13. Gutberlet M, Noeske R, Schwinge K, Freyhardt P, Felix R, Niendorf T. Comprehensive cardiac magnetic resonance imaging at 3.0 Tesla: feasibility and implications for clinical applications. Invest Radiol. 2006;41(2):154–167. 14. Gutberlet M, Schwinge K, Freyhardt P, et al. Influence of high magnetic field strengths and parallel acquisition strategies on image quality in cardiac 2D CINE magnetic resonance imaging: comparison of 1.5 T vs. 3.0 T. Eur Radiol. 2005;15(8):1586–1597. 15. Michaely HJ, Nael K, Schoenberg SO, et al. Analysis of cardiac function—comparison between 1.5 Tesla and 3.0 Tesla cardiac cine magnetic resonance imaging: preliminary experience. Invest Radiol. 2006;41(2):133–140. 16. Kelle S, Hamdan A, Schnackenburg B, et al. Dobutamine stress cardiovascular magnetic resonance at 3 Tesla. J Cardiovasc Magn Reson. 2008;10:44. 17. McVeigh ER. MRI of myocardial function: motion tracking techniques. Magn Reson Imaging. 1996;14(2):137–150. 18. Fischer SE, McKinnon GC, Scheidegger MB, Prins W, Meier D, Boesiger P. True myocardial motion tracking. Magn Reson Med. 1994;31(4):401–413. 19. Valeti VU, Chun W, Potter DD, et al. Myocardial tagging and strain analysis at 3 Tesla: comparison with 1.5 Tesla imaging. J Magn Reson Imaging. 2006;23:477–480. 20. Thomas D, Strach K, Meyer C, et al. Combined myocardial stress perfusion imaging and myocardial stress tagging for detection of coronary artery disease at 3 Tesla. J Cardiovasc Magn Reson. 2008;10:59. 21. Garot J, Lima JA, Gerber BL, et al. Spatially resolved imaging of myocardial function with strain-encoded MR: comparison with delayed contrast-enhanced MR imaging after myocardial infarction. Radiology. 2004;233(2):596–602.

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22. Pan LFA, Spooner A, Weiss RG, Stuber M, Osman NF. Strain encoded imaging for detection of regional dysfunction in patients with myocardial infarction at 3T. JCMR. 2006;535:295–296. 23. Kim WY, Danias PG, Stuber M, et al. Coronary magnetic resonance angiography for the detection of coronary stenoses. N Engl J Med. 2001;345(26): 1863–1869. 24. Stuber M, Botnar RM, Fischer SE, et al. Preliminary report on in vivo coronary MRA at 3 Tesla in humans. Magn Reson Med. 2002;48 (3):425–429. 25. Stuber M, Botnar RM, Danias PG, et al. Double oblique free-breathing high-resolution 3D coronary MRA. J Am Coll Cardiol. 1999;34 (2):524–531. 26. Botnar RM, Stuber M, Danias PG, Kissinger KV, Manning WJ. Improved coronary artery definition with T2-weighted free-breathing 3D-coronary MRA. Circulation. 1999;99(24): 3139–3148. 27. McConnell MV, Khasgiwala VC, Savord BJ, et al. Comparison of respiratory suppression methods and navigator locations for MR coronary angiography. AJR Am J Roentgenol. 1997;168(5):1369–1375. 28. Kim WY, Stuber M, Kissinger KV, Andersen NT, Manning WJ, Botnar RM. Impact of bulk cardiac motion on right coronary MR angiography and vessel wall imaging. J Magn Reson Imaging. 2001; 14 (4):383–390. 29. Bi X, Park J, Larson AC, Zhang Q, Simonetti O, Li D. Contrastenhanced 4D radial coronary artery imaging at 3.0 T within a single breath-hold. Magn Reson Med. 2005;54(2): 470–475. 30. Bi X, Li D. Coronary arteries at 3.0 T: contrast-enhanced magnetization-prepared three-dimensional breathhold MR angiography. J Magn Reson Imaging. 2005;21(2):133–139. 31. Nayak KS, Cunningham CH, Santos JM, Pauly JM. Real-time cardiac MRI at 3 tesla. Magn Reson Med. 2004;51(4):655–660. 32. Kaul MG, Stork A, Bansmann PM, et al. Evaluation of balanced steady-state free precession (TrueFISP) and K-space segmented gradient echo sequences for 3D coronary MR angiography with navigator gating at 3 Tesla. Rofo. 2004;176(11):1560–1565. 33. Sommer T, Hackenbroch M, Hofer U, et al. Coronary MR angiography at 3.0 T versus that at 1.5 T: initial results in patients suspected of having coronary artery disease. Radiology. 2005;234(3):718–725. 34. Nezafat R, Stuber M, Ouwerkerk R, Gharib AM, Desai MY, Pettigrew RI. B(1)-insensitive T(2) preparation for improved coronary magnetic resonance angiography at 3 T. Magn Reson Med. 2006;55(4):858–864. 35. Yang Q, Li K, Liu X, et al. Contrast-enhanced whole-heart coronary magnetic resonance angiography at 3 T: a comparative study with x-ray angiography in a single center. J Am Coll Cardiol. 2009;54:69–76. 36. van Elderen SGC, Webb AG, Versluis M, et al. In vivo human coronary magnetic resonance angiography at 7 Tesla. J Cardiovasc Magn Reson. 2009;(suppl 1):36. 37. Fayad ZA, Fuster V, Fallon JT, et al. Noninvasive in vivo human coronary artery lumen and wall imaging using black-blood magnetic resonance imaging. Circulation. 2000;102(5):506–510. 38. Botnar RM, Stuber M, Lamerichs R, Smink J, Fischer SE, Manning WJ. Initial experiences with in-vivo right coronary artery human MR vessel wall imaging at 3 Tesla. J Cardiovasc Magn Reson. 2003;5:589–594. 39. Yeon SB, Sabir A, Clouse M, et al. Delayed-enhancement cardiovascular magnetic resonance coronary artery wall imaging. J Am Coll Cardiol. 2007;50:441–447. 40. Kim WY, Stuber M, Bornert P, Kissinger KV, Manning WJ, Botnar RM. Three-dimensional black-blood cardiac magnetic resonance coronary vessel wall imaging detects positive arterial remodeling in patients with nonsignificant coronary artery disease. Circulation. 2002;106 (3):296–299. 41. Fleckenstein JL, et al. Fast short-tau inversion-recovery MR imaging. Radiology. 1991;179(2):499–504. 42. Stuber M, Botnar RM, Spuentrup E, Kissinger KV, Manning WJ. Three-dimensional high-resolution fast spin-echo coronary magnetic resonance angiography. Magn Reson Med. 2001;45(2): 206–211.

51. Sodickson DK, Manning WJ. Simultaneous acquisition of spatial harmonics (SMASH): fast imaging with radiofrequency coil arrays. Magn Reson Med. 1997;38(4):591–603. 52. Pruessmann KP, Weiger M, Scheidegger MB, Boesiger P. SENSE: sensitivity encoding for fast MRI. Magn Reson Med. 1999;42(5):952–962. 53. Huber ME, Kozerke S, Pruessmann KP, Smink J, Boesiger P. Sensitivity-encoded coronary MRA at 3T. Magn Reson Med. 2004;52 (2):221–227. 54. Huber ME, Kozerke S, Boesiger P. Improved artery delineation in dualstack coronary magnetic resonance angiography using parallel imaging at 3 T. J Magn Reson Imaging. 2005;21(4):443–448. 55. Sodickson DK, McKenzie CA, Li W, Wolff S, Manning WJ, Edelman RR. Contrast-enhanced 3D MR angiography with simultaneous acquisition of spatial harmonics: a pilot study. Radiology. 2000;217(1):284–289. 56. Wang J, Alsop DC, Li L, et al. Comparison of quantitative perfusion imaging using arterial spin labeling at 1.5 and 4.0 Tesla. Magn Reson Med. 2002;48(2):242–254. 57. Li D, Waight DJ, Wang Y. In vivo correlation between blood T2* and oxygen saturation. J Magn Reson Imaging. 1998;8(6):1236–1239. 58. Kozerke S, Schar M, Lamb HJ, Boesiger P. Volume tracking cardiac 31P spectroscopy. Magn Reson Med. 2002;48(2):380–384. 59. Dou J, Reese TG, Tseng WY, Wedeen VJ. Cardiac diffusion MRI without motion effects. Magn Reson Med. 2002;48(1):105–114. 60. Rajaram M, Seabra LF, Abdullah SM, et al. 3T cardiac magnetic resonance performs well as the primary scanner in a clinical setting: our initial experience at a tertiary care center. J Cardiovasc Magn Reson. 2009;11(suppl 1):86.

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43. Koktzoglou I, Simonetti O, Li D. Coronary artery wall imaging: initial experience at 3 Tesla. J Magn Reson Imaging. 2005;21(2):128–132. 44. Priest AN, Bansmann PM, Kaul MG, Stork A, Adam G. Magnetic resonance imaging of the coronary vessel wall at 3 T using an obliquely oriented reinversion slab with adiabatic pulses. Magn Reson Med. 2005;54(5):1115–1122. 45. Rinck PA, Muller RN. Field strength and dose dependence of contrast enhancement by gadolinium-based MR contrast agents. Eur Radiol. 1999;9(5):998–1004. 46. Araoz PA, Glockner JF, McGee KP, et al. 3 Tesla MR imaging provides improved contrast in first-pass myocardial perfusion imaging over a range of gadolinium doses. J Cardiovasc Magn Reson. 2005;7 (3):559–564. 47. Cheng ASH, Pegg TJ, Karamitsos TD, et al. Cardiovascular magnetic resonance perfusion imaging at 3-Tesla for the detection of coronary artery disease: a comparison with 1.5-Tesla. J Am Coll Cardiol. 2007;49:2440–2449. 48. Karamitsos TD, Leccisotti L, Arnold JR, et al. Relationship between regional myocardial oxygenation and perfusion in patients with coronary artery disease: insights from cardiovascular magnetic resonance and positron emission tomography. Circulation Cardiovasc Imaging. 2009;0:109, 860148. 49. den Hollander JA, Evanochko WT, Pohost GM. Observation of cardiac lipids in humans by localized 1H magnetic resonance spectroscopic imaging. Magn Reson Med. 1994;32(2):175–180. 50. Bottomley PA, Weiss RG. Non-invasive magnetic-resonance detection of creatine depletion in non-viable infarcted myocardium. Lancet. 1998;351(9104):714–718.

Assessment of Cardiac Function Alicia M. Maceira, Nicholas G. Bellenger, and Dudley J. Pennell

The accurate and reproducible assessment of cardiac function is a fundamental aim of noninvasive cardiac imaging. It forms the foundation upon which much of the assessment and management of myocardial dysfunction, ischemia, viability, remodeling, valvular, and other cardiac disorders are based. In this chapter, we will discuss the importance of the measurement of global cardiac function, compare techniques, and provide a practical step-by-step guide to its assessment by cardiovascular magnetic resonance (CMR).

THE POPULATION IMPACT OF CARDIAC DYSFUNCTION Cardiac dysfunction can result from a broad spectrum of organ-specific and multisystem disorders. The defining property that all these disorders have in common is impairment of the ventricle’s ability to eject blood, which, in its broadest sense and omitting any discussion of semantics, is known as heart failure. Heart failure is common, with approximately 1.5% to 2.0% of the population below 65 years old being affected, rising to 6% to 10% of those older than 65 years.1 It afflicts 4.8 million people in the United States, with 400,000 to 700,000 new cases developing each year. It is the leading cause of hospital admission in people older than 65 years of age, and despite aggressive treatment, 40,000 patients die of heart failure each year.2 It is also an enormous consumer of health care budgets in the Western world, and treatments to prevent its occurring, to slow its progression, or to prevent repeated hospitalizations can have an important economic impact.

THE IMPORTANCE OF MEASURING CARDIAC FUNCTION A single assessment of cardiac function can provide important diagnostic and prognostic information, whether in the setting of post-infarction recovery,3 left ventricular (LV) hypertrophy,4 or chronic heart failure.5,6 As well as improving symptoms and quality of life, treatment of heart failure aims to decrease the likelihood of disease progression that could result in costly hospital admissions and ultimately death. To allow early evaluation and alteration of an individual patient’s management, serial studies need to be performed using a technique that not only is accurate but also has good interstudy reproducibility. This principle also applies in considering study populations of trial therapies.

TECHNIQUES FOR ASSESSING CARDIAC FUNCTION Bedside clinical assessment of cardiac function is generally poor,7 and electrocardiogram findings are nonspecific, although the presence of an entirely normal electrocardiogram (ECG) has a 95% likelihood of normal systolic function.8 In the search for a better assessment, it is worth considering the ideal imaging technique. This would provide a noninvasive, accurate, and reproducible assessment of cardiac function without exposure to ionizing radiation. It would be widely available and would be time- and cost-effective. Although no technique meets these ideals, there are clear differences between modalities that merit discussion.

Echocardiography Echocardiography is a widely available but less than ideal imaging technique for quantifying cardiac function, as the image acquisition is operator and acoustic window dependent.9 The quantification of ventricular function is limited by a priori geometric assumptions that may provide a reasonable assessment in the normal ventricle but are less reliable in remodeled hearts, owing to complex irregular shape changes.10,11 M-mode echocardiography was developed in the early 1970s and was immediately applied in practice for LV function assessment, because of its simple algorithm and noninvasiveness. In this technique, mid-LV diameters are measured in the minor axis (Fig. 14-1), and volumes are obtained by cubing these values (thereby cubing the errors).12–14 Functional estimates, such as fractional shortening and ejection fraction (EF), are then derived from these diameters and volumes, respectively. This method assumes that a single view is representative of all the myocardial segments and that contraction is uniform throughout the ventricle. Two-dimensional (2D) echocardiography, with the ability of imaging the heart in tomographic views, considerably improved the accuracy of LV volume measurement, providing the opportunity to derive the cardiac function from cardiac volumes by the arealength and Simpson’s method of discs (Fig. 14-2).12 This, however, relies on good visualization of the entire endocardial border, which is frequently not possible, although several software-based algorithms for automatic endocardial border detection and on-line calculation of LV volume have been developed.15 For example, in a multicenter study that required good-quality echocardiograms as an entry Cardiovascular Magnetic Resonance 181

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CW RV IVS ESD

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Figure 14-1 M-mode echo recording at the midventricular level in a patient with tricuspid regurgitation showing paradoxical systolic motion of the interventricular septum and right ventricle. CW, chest wall; EDD, end-diastolic dimension; ESD, end-systolic dimension; IVS, interventricular septum; LV, left ventricle; PVW, posterior ventricular wall; RV, right ventricle.

practical difficulties of quantifying global function by echo are underlined by the fact that in the “real world,” it is often simply estimated by the clinician performing the imaging. This is highly subjective but can be clinically valid with experience.16 Recently, there have been two important advances in echocardiography. On one hand, contrast echocardiography has been shown to allow improved assessment of LV volumes and LVEF17,18 with low interobserver variability.19 On the other hand, three-dimensional (3D) echocardiography has emerged as a more accurate and reproducible approach to LV quantitation by removing the need for geometric assumptions about the LV shape.20 Volumetric analysis of real-time 3D echocardiographic data allows fast dynamic measurement of LV volumes.21 But this technique needs a stable cardiac rhythm and constant cardiac function during image acquisition, and issues of acoustic windows and practicality in clinical practice still remain to be answered as more experience is gained.

Nuclear Cardiology

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Figure 14-2 Echo measurement by Simpson’s method of discs. This represents a more accurate measurement than M-mode or arealength assessment but relies on good endocardial border definition. A, annulus; LA, left atrium; RA, right atrium; RV, right ventricle.

criterion, the echo core laboratory was unable to perform a confident 2D analysis in 31% of patients.14 Thus, being the most frequently used modality in clinical practice, echocardiography has gained little acceptance in clinical trials, owing to its moderate reproducibility and accuracy to define left ventricular ejection fraction (LVEF), because of geometric assumptions, poor acoustic windows, and inadequate discrimination of the endocardial border. Overall, the 182 Cardiovascular Magnetic Resonance

Nuclear cardiology with radionuclide ventriculography is commonly used to measure ventricular function by measuring the LVEF, but it has relatively low spatial and temporal resolution, and preparation and scanning times are relatively prolonged.22 In addition, ventricular volumes are difficult to measure and are rarely performed clinically (though they are used occasionally for research),3,23 and ventricular mass cannot be obtained. The use of gated perfusion single photon emission computed tomography (SPECT) has allowed the development of 3D solutions to global and regional ventricular function,24 and this is now achieving widespread use, especially in the United States. This is useful when perfusion needs to be assessed, and it adds prognostic value to the perfusion assessment,25 but it is not being performed solely to assess ventricular function. This technique has good reproducibility,26 and ventricular volumes are reported as reliable,27 but there is concern over their accuracy in both small and large ventricles because of the limited spatial resolution and the problems of assigning a ventricular border in areas of transmural infarction and thinning where counts are very low.28 It has been reported that Tc-99m sestamibi ECG-gated SPECT performed 3 weeks after acute myocardial infarction can predict ventricular remodeling,29 but contrast echocardiography is more accurate than gated SPECT for estimation of LV remodeling.30 Recently, ECG-gated F-18-fluorodeoxyglucose positron emission tomography (PET) has been shown to be reasonably accurate for measurement of cardiac function,31 but this technique is time consuming and is performed only in viability studies. For these nuclear cardiology techniques, the need for repeated radionuclide doses in follow-up studies is problematic, especially for research, in which radiation exposure must be justified in a milieu of competing technologies and public pressure in general to limit radiation burdens.

Computed Tomography Electron beam tomography has been used in the past to study both function and perfusion, but technical questions prevent its clinical role. Multidetector computed tomography (MDCT)

Cardiovascular Magnetic Resonance CMR has some fundamental advantages over other imaging techniques, which have fueled the growing enthusiasm for its use in clinical practice and research. CMR offers accurate and reproducible tomographic, static, or cine images

of high spatial and temporal resolution in any desired plane without exposure to contrast agents or ionizing radiation. As such, true two- and four-chamber and short axis views can be easily and rapidly acquired to allow a visual, qualitative assessment of function, similar to that of echocardiography. The main advantage of CMR, however, lies in its quantitative accuracy and reproducibility. There are two main methods for measuring the enddiastolic and end-systolic volumes. The earliest semiquantitative method was an adaptation of the echo area-length method, in which the volume of the left ventricle is assumed to constitute an ellipsoid. With a single apical four-chamber view, the area (A) of the LV endocardial border can be traced, as well as the length (L) from the apex to the mitral annulus. The volume (V) for systole and diastole is then easily calculated: V ¼ 0.85A2/L. A more accurate adaptation uses both long axis planes (vertical and horizontal long axis) to measure two perpendicular areas and lengths: V ¼ (0.85 area A area B)/smaller length. While this offers a simple and time-efficient volume analysis,38 it suffers from the same limitations that echocardiography has: the need for geometric assumptions and the inability to take into account regional differences in wall motion. An example of a distorted heart post-infarction is shown in Figure 14-3. A better method of measuring volumes and thereby function is by the use of Simpson’s rule. A stack of contiguous tomographic slices are acquired that encompass the

Figure 14-3 A patient with ischemic heart disease in whom the left ventricle (LV) no longer conforms to geometric assumptions in either diastole (A, C) or systole (B, D) frames. Both the VLA (A, B) and the HLA (C, D) views are illustrated.

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of the heart is a rapidly developing technique that is used mainly to evaluate the coronary arteries. It can also be used for evaluating global and regional ventricular function, with good agreement with echocardiography,32,33 cineventriculography,34 SPECT,35 and cine magnetic resonance imaging.36 Postprocessing tools allow fast and semiautomatic determination of LV function parameters from MDCT data in analogy to known CMR evaluation approaches. It cannot be considered a first-line modality, owing to the exposure to radiation, the use of potentially nephrotoxic contrast, and the low temporal resolution, but it may be used for accessory dynamic information in patients undergoing computed tomography coronary angiography. Notably, no normal reference ranges have been established so far, and CMR normal values are for comparison. MDCT reproducibility for volumes and function is lower than that with CMR, with an interobserver variability of 4.2% for end-diastolic volume, 5.4% for end-systolic volume, and 4.1% for LVEF.37

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Figure 14-4 The end-diastolic and end-systolic slices from multiple contiguous short axis cines that encompass the left ventricle (LV), from base to apex, in a patient with ventricular dilatation and hypertrophy from chronic aortic regurgitation. The epicardial and endocardial borders are traced, and the summation values are shown. Note that there is one more image at end-diastole11 than at end-systole,10 reflecting the need to allow for the systolic descent of the atrioventricular ring, as described in the text. Note also the ingress of the LV outflow tract in the most basal end-diastolic image, where there is no ventricular mass, and the open ends of the LV myocardial horseshoe are joined together to form the appropriate volume.

entire left ventricle. The ventricular volume is equal to the sum of the endocardial areas multiplied by the distance between the centers of each slice (Fig. 14-4). The volumes that are obtained by this method are independent of geometric assumptions and dimensionally accurate.39,40 At one time, it was common to use a stack of transverse images for this measurement, but while this makes it easy to define the mitral valve plane and thereby the true base of the LV,41 it is also subject to considerable partial volume effects, especially in the inferior wall. For this

Long axis of left ventricle

reason, more recently, short axis slices have been employed, and nearly all sites specializing in CMR have now adopted this practice (Fig. 14-5). The question has been raised as to whether the right ventricular (RV) volumes should be measured in the axial orientation, as this appears to have better interobserver and intraobserver reproducibility.42 Yet the interstudy reproducibility of RV measurements in the short axis orientation is good,43 and in practice, this orientation allows the LV and RV dimensions to be measured simultaneously.

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Figure 14-5 Schematic of the left ventricular short axis slices that encompass the entire left ventricle. Each slice is acquired as a cine. A1, end-diastolic endocardial area; A2, end-systolic endocardial area; A3, myocardium.

Steady-state free precession (SSFP) sequences rephase the transverse magnetization that undergoes dephasing during phase encoding and readout between radio-frequency pulses; therefore, imaging occurs when all transverse and longitudinal magnetization components are at steady state. SSFP cine imaging eliminates blood saturation artifacts and makes the cines independent of inflow enhancement and based on the ratio of T2 to T1. This results in substantially improved blood-myocardium contrast, especially in the long axis planes,46 which may allow easier delineation of the endocardial borders. At the epicardial border, fat-myocardium delineation is also improved. This sequence runs at its best with ultrafast gradients, as a very short TR is required to reduce the sensitivity of the sequence to movement artifact. Because of these characteristics, LV end-diastolic and end-systolic volumes are larger and LV mass is smaller with SSFP than with GRE.47 In terms of reproducibility, only EF shows differences favoring SSFP, while SSFP and GRE are equal for volumes and mass reproducibility. In the past, improved performance was achieved mainly through improvements in gradient hardware, but nowadays, further developments in this direction are limited, owing to physiologic constraints such as the risk of peripheral nerve stimulation. The introduction of parallel imaging provides alternative means for increasing acquisition speed. By using information from multiple receiver coils, images can be reconstructed from a sparsely sampled set of data, allowing for twofold to threefold acceleration of the imaging process. However, further increases in acquisition speed are difficult to achieve for current clinical field strengths and typical fields of view. Recently developed methods, k-t BLAST (broad-use linear acquisition speed-up technique) and k-t SENSE (sensitivity encoding), have been proposed that significantly improve the performance of dynamic imaging, taking into account the similarity of image information at different time points during a dynamic series. By using these methods, improved temporal resolution (fivefold to eightfold acceleration) or improved spatial resolution for a given amount of acquisition can be achieved in

ACCURACY AND REPRODUCIBILITY OF CMR It is now widely accepted that CMR offers the reference standard for the noninvasive assessment of cardiac function, being both accurate55,56 (Figs. 14-6 to 14-9) and reproducible in normal as well as abnormal ventricles.57–61 Much validation work was done by using conventional nonbreath-hold cine CMR on older scanners, however, and currently, breath hold sequences are routinely employed in

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Differences Between Steady-State Free Precession and Gradient Echo

cardiac imaging.48 Another recently developed method is UNFOLD (unaliasing by Fourier-encoding the overlaps using the temporal dimension),49 which works by forcing aliased signals to behave in specific ways through time, so unwanted signals are detected and removed, thus reducing the amount of acquired data. This method can in theory be used with k-t SENSE and k-t BLAST to accelerate the acquisition and/or to suppress artifacts in free breathing. All these techniques are very promising but still in development, and no specific data are available. In addition to the LV, it is important to remember the right ventricle, as its function is also known to be an important determinant of prognosis, both in coronary artery disease50 and in heart failure,51 congenital heart disease, and pulmonary disease.52 Global RV function is difficult to assess adequately by echocardiography, while radionuclide ventriculography suffers from assumptions concerning projection of overlapping structures unless research techniques such as first pass techniques with ultra-short half-life isotopes are used. CMR does not experience such problems, and RV function and mass are well characterized.53,54 The RV is discussed in greater detail in Chapters 11 and 28.

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In the past, to achieve full 3D coverage of the ventricle using conventional free breathing gradient echo cine sequences, a total scanning time of 30 minutes or more was required, but on modern scanners with fast imaging, a single cine can be acquired in just one breath hold of about 8 to 10 seconds, allowing the whole stack of images to be acquired in 5 to 10 minutes.14 This also has the considerable additional advantage of reducing breathing and movement artifact. Moreover, real-time steady-state free precession (SSFP) imaging can acquire all the ventricular slices in just one breath hold with acceptable accuracy and image quality.44 In patients who are unable to hold their breath consistently or who are orthopneic, solutions using the same sequence with more signal averages or combined with navigator echo imaging have been shown to be successful, during free breathing,45 with a slight increase in the acquisition duration.

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nearly all centers. More recent studies have been performed to examine whether this has affected the quality of the results, and these show that the two techniques have similar reproducibility.62–65 This is illustrated in Figure 14-10, which shows the intraobserver, interobserver, and interstudy variability for volume and functional assessment by breath hold CMR in dilated and normal ventricles, compared with conventional cine CMR.65 186 Cardiovascular Magnetic Resonance

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Figure 14-7 Validation of right and left ventricular stroke volumes (LVSV) in vivo using CMR ventricular volume analysis. There is an excellent agreement between the stroke volumes (SV) of both ventricles, which is strong evidence that both measurements are accurate, because they are equivalent in vivo in the absence of valve regurgitation or shunting. Source: Data from Longmore DB, Klipstein RH, Underwood SR, et al. Dimensional accuracy of magnetic resonance in studies of the heart. Lancet. 1985; 1:1360–1362.

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Figure 14-9 Validation of right ventricular (RV) mass with comparison of CMR-derived mass against postmortem bovine hearts. Source: Data from Katz J, Whang J, Boxt LM, Barst RJ. Estimation of right ventricular mass in normal subjects and in patients with primary pulmonary hypertension by nuclear magnetic resonance imaging. J Am Coll Cardiol. 1993;21:1475–1481.

CMR has been shown to have higher reproducibility than 2D echocardiography for LV end-systolic volume (4.4% to 9.2% versus 13.7% to 20.3%, p <.001), LVEF (2.4% to 7.3% versus 8.6% to 19.4%, p < 0.001), and LV mass measurements (2.8% to 4.8% versus 11.6% to 15.7%, p < 0.001), and this higher reproducibility allows for sample size reductions of 55% to 93%.66 The excellent reproducibility of CMR versus echocardiography can be illustrated by considering the sample size required for a drug trial designed to show a 10-g decrease in LV mass with antihypertensive treatment. A direct comparison of CMR with echocardiography for reproducibility has shown that for an 80% power and a p value of 0.05, the sample size required would be 505 patients with 2D echo but only 14 patients with CMR.67 Similarly, we have found that to show a 5% difference in EF with a 90% power and a p value of 0.05 would require only 7 normal subjects or 5 patients with dilated ventricles.66 Similarly, CMR measurements of RV function parameters both in healthy subjects and in patients show good interstudy reproducibility (4.2% to 7.8% for RVEDV, 8.1% to 18.1% for RV end-systolic volume, 4.3% to 10.4% for RV ejection fraction and 7.8% to 9.4% for RV mass), which was lower than that for the left ventricle for all measures but only significantly for EF.68 This reproducibility has significant implications for research and in particular for pharmaceutical companies, for which CMR offers a more cost- and time-effective research tool.

A PRACTICAL GUIDE TO FUNCTIONAL CMR In modern medicine, a balance must be struck between the information that can be gained from an investigation and the resources it demands. The following protocol is designed to be as efficient as possible in gaining the volumetric data

Intraobserver % variability

12 10 8 6 4 2 0 Bellenger Bellenger Bellenger normal LVH DCM

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EF Mass

Figure 14-10 The intraobserver (A), interobserver (B), and interstudy (C) percentage variability for end-diastolic volume (EDV), end-systolic volume (ESV), ejection fraction (EF), and left ventricular mass. Results using breath hold gradient echo from our center in patients with heart failure and dilated ventricles (DCM) and left ventricular hypertrophy (LVH) are compared with breath hold gradient echo in normals (Bogaert) and traditional, slower gradient echo cine imaging (Semelka).65 Overall, the results are very similar (probably favoring the breath hold imaging), both between techniques and between different population groups.

from the ventricles of the heart22 (See also Chapter 11). Figure 14-11 illustrates the sequence of pilot images used to achieve imaging in the long axis of the left ventricle and thereby the true short axis. A coronal pilot is first taken and used to acquire transverse pilots, which show both the mitral

valve and the apex of the LV. By taking a plane through the center of the mitral valve (halfway between the back end of the septum and the back end of the lateral wall) and the tip of the apex, the vertical long axis (VLA) is acquired. This VLA is used to plan the horizontal long axis (HLA), by again using a plane through the center of the mitral valve (halfway between the back of the anterior and inferior walls) and the tip of the apex. It should be noted that it is common to find centers describing planes that are parallel to the septum for the VLA and parallel to the inferior wall for the HLA, but these are not correct, as they are likely to lead to the long axis plane not passing through the center of the basal ring of the LV, and they may lead to an offset from the tip of the apex, which is also undesirable. This can lead to problems planning the short axis cuts to adequately cover the full extent of left and right ventricles; in addition, it reduces the reproducibility of the short axis plane positioning for repeated studies. Finally, the short axis slices are placed on the HLA to encompass the heart. To achieve the most reliable results, which are the most reproducible, attention to detail is required. First, if the short axis cuts are to be acquired by using a breath hold cine sequence, then the VLA and HLA must also have been acquired by using a breath hold and at end-expiration. Second, in using breath hold techniques, it is more reproducible to ask the patient to hold the breath at end-expiration rather than elsewhere in the respiratory cycle; this applies to the pilots as well as to the cines.69 Third, the first short axis plane should be placed at the base of the heart, covering the most basal portion of the left and right ventricles just forward of the atrioventricular (AV) ring, and it should be placed on the end-diastolic HLA image. Finally, further short axis planes should then be planned to move apically from this plane until the apex is encompassed. While it is possible to acquire the VLA and HLA as single pilot images instead of cines, there is little practical merit in this, as the time for two breath hold cines is small, and the contraction pattern in these two planes is very useful during qualitative assessment of ventricular function. Nowadays, with the development of 3D postprocessing software solutions for analyzing the cines, these long axis cines are mandatory. Also, full 3D analysis of atria as well as ventricles is simple and practical with automated analysis, in which case cines encompassing the entire heart should be acquired. Following are some technical tips: The average Fast gradient echo breath hold that is required to acquire 18 phases with a phase-encoded grouping (PEG) of 28 is approximately 10 sec. The PEG size and the number of phases acquired are related through the repetition time (TR) of the sequence. Modern CMR scanners with faster gradients allow a shorter TR and echo time (TE), which improve this compromise, so the breath hold length is no longer generally a problem for patients. If necessary, increasing the PEG can reduce the breath hold, but a compromise is reached, as fewer phases will be captured with a higher PEG. The number of phases should be at least 12 to 15 to give adequate information on wall motion and cover end systole. Decreasing the field-of-view will also reduce the breath hold time but may result in some wrap-around occurring at the edges of the image. If this remains remote from the heart, it may be considered an acceptable compromise. Typically, the best results are seen when the time between cine phases is 30 msec or less, and this yields a cine with approximately 25 frames in clinical practice. Cardiovascular Magnetic Resonance 187

14 ASSESSMENT OF CARDIAC FUNCTION

14

ISCHEMIC HEART DISEASE

B

A

E

C

D

Figure 14-11 Pilot images used to achieve the true short axis of the left ventricle. A coronal pilot (A) is first acquired and used to pilot the transverse image (B). The vertical long axis (C) is obtained from this, and subsequently, the horizontal long axis (D) is obtained, upon which a stack of short axis cines are placed. One representative end-diastolic and end-systolic image from a single cine is shown.

The ECG gating can be prospective or retrospective. In prospective gating, images are acquired during systole and the beginning of diastole, while with retrospective gating, images from the whole cardiac cycle are obtained. Although the breath hold with prospective gating is shorter, it is advisable to use retrospective gating, as with modern 3D automated analysis of volumes, diastolic function can be assessed. Also, in patients with arrhythmia, retrospective gating generally yields better image quality. A slice thickness of 8 mm at 1.5 T provides adequate spatial resolution without overly increasing the number of slices and thereby the analysis time. It also limits partial volume effects. A 2-mm slice gap is commonly used to allow easy calculation of volumes, as the center of each slice is then 1 cm apart. There are no formal studies to aid in the choice of slice thickness, but 8 mm is a reasonable consensus. Some centers prefer thinner slices, but it is important to maintain acceptable signalto-noise ratio and image quality. 3D imaging may eventually allow more partitions with thinner slices. Analysis of the short axis slices is relatively straightforward, provided that the quality of the images is reasonable (mainly depending on accurate ECG gating and good breath holding). The main source of error is in separating the ventricles from the atria. Identifying this basal slice is made more difficult by the through plane descent of the atrioventricular ring in systole, which is 188 Cardiovascular Magnetic Resonance

usually about 1 cm. This makes the placement of the first, most basal short axis slice very important. By ensuring that this basal slice is carefully positioned on the end-diastolic, end-expiratory breath hold HLA image just forward of the atrioventricular ring, the first short axis cine will by definition contain end-diastolic volume and mass within both ventricles. However, at end-systole, the basal slice will include only atrium, owing to descent of the AV ring, and in general the systolic area in the basal cine is not included in the analysis of the systolic volume. In general, the next slice down contains both end-diastolic and end-systolic volume. An alternative approach to this rigorous approach is to oversample with short axis slices into the atrium and attempt to retrospectively differentiate ventricle from atrium on the basis of the degree of descent of the AV ring in the long axis images and whether the chamber dilates or contracts in systole. In general, we prefer not to oversample but to ensure that the first basal slice is acquired correctly, as this leads to a reproducible approach and is more time-efficient for both acquisition and analysis and because oversampling relies more heavily on good image quality to differentiate atrium from ventricle. Papillary muscles and endocardial trabeculae should be excluded from the LV volume and included in the LV mass. Although there is no clear consensus at present, LV mass is usually taken from the end-diastolic images.

OTHER CMR MEASURES OF GLOBAL FUNCTION: BRIEF SYNOPSIS Systolic Function There are other ways to derive important functional information from the heart besides volumetry alone. For example, flow in the major vessels is easily measured with great accuracy by using velocity mapping.71 If the aortic flow is measured over a complete cardiac cycle, this represents the LV stroke volume. The RV stroke volume can likewise be found from flow in the pulmonary artery. Cardiac output can be easily derived by using the following formula: Cardiac output ¼ stroke volume heart rate. The measurement of stroke volume is useful in valve disease72 and for determining the need for surgery in cardiovascular shunting73 and is noninvasive and quantitative.74 The peak flow velocity and

Table 14-1 Normal SSFP Left Ventricular Volumes, Systolic Function, and Mass (Absolute and Indexed to Body Surface Area) by Age Decile (Mean, 95% Confidence Interval) in Males and Females 20–29 Years

30–39 Years

MALES EDV (mL), SD 21 ESV (mL), SD 11 SV (mL), SD 14 EF (%), SD 4.5 Mass (g), SD 20

Absolute Values 167 (126–208) 163 (121–204) 58 (35–80) 56 (33–78) 109 (81–137) 107 (79–135) 65 (57–74) 66 (57–75) 148 (109–186) 147 (109–185)

EDV/BSA (mL/m2), SD 9.0 ESV/BSA (mL/m2), SD 5.5 SV/BSA (mL/m2), SD 6.1 EF/BSA (%/m2), SD 3.3 Mass/BSA (g/m2), SD 8.5

Indexed to Body Surface Area 86 (68–103) 83 (66–101) 30 (19–41) 29 (18–39) 56 (44–68) 55 (43–67) 34 (28–40) 34 (28–40) 76 (59–93) 75 (59–92)

FEMALES EDV (mL), SD 21 ESV (mL), SD 9.5 SV (mL), SD 14 EF (%), SD 4.6 Mass (g), SD 18

Absolute Values 139 (99–179) 135 (94–175) 48 (29–66) 45 (27–64) 91 (63–119) 89 (61–117) 66 (56–75) 66 (57–75) 105 (69–141) 106 (70–142)

EDV/BSA (mL/m2), SD 8.7 ESV/BSA (mL/m2), SD 4.7 SV/BSA (mL/m2), SD 6.2 EF/BSA (%/m2), SD 4.7 Mass/BSA (g/m2), SD 7.5

Indexed to Body Surface Area 82 (65–99) 79 (62–96) 28 (19–37) 27 (17–36) 54 (42–66) 53 (40–65) 39 (30–48) 39 (30–49) 62 (47–77) 62 (47–77)

40–49 Years

50–59 Years

60–69 Years

70–79 Years

159 (117–200) 54 (31–76) 105 (77–133) 66 (58–75) 146 (108–185)

154 (113–196) 51 (29–74) 103 (75–131) 67 (58–76) 146 (107–184)

150 (109–191) 49 (27–72) 101 (73–129) 67 (58–76) 145 (107–183)

146 (105–187) 47 (25–70) 99 (71–127) 68 (59–77) 144 (106–183)

81 27 54 34 75

(64–99) (17–38) (42–66) (28–40) (58–91)

130 (90–171) 43 (25–62) 87 (59–115) 67 (58–76) 107 (71–143) 76 25 51 40 63

(59–93) (16–34) (39–63) (30–49) (48–77)

79 26 53 34 74

(62–97) (15–37) (41–65) (28–41) (57–91)

126 (86–166) 41 (22–59) 85 (57–113) 68 (59–77) 108 (72–144) 73 24 50 40 63

(56–90) (14–33) (37–62) (31–49) (48–78)

77 25 52 34 73

(60–95) (14–36) (40–64) (28–41) (57–90)

122 (82–162) 39 (20–57) 83 (56–111) 69 (60–78) 109 (73–145) 70 22 48 40 63

(53–87) (13–31) (36–60) (31–49) (48–78)

75 24 51 34 73

(58–93) (13–35) (39–63) (28–41) (56–89)

118 (77–158) 36 (18–55) 81 (54–109) 69 (60–78) 110 (74–146) 67 21 47 40 63

(50–84) (12–30) (34–59) (31–49) (49–78)

BSA, body surface area; EDV, end-diastolic volume; EF, ejection fraction; ESV, end-systolic volume; SD, standard deviation; SSFP, steady state free precession; SV, stroke volume.

Cardiovascular Magnetic Resonance 189

14 ASSESSMENT OF CARDIAC FUNCTION

with a 3D analytical approach will be available to fully take into account the motion in systole of the atrioventricular ring, by incorporating VLA and HLA cines to define accurately the mitral valve position at each phase in the cardiac cycle. A fully robust solution for the postprocessing is the holy grail of this field, but there is every reason to be optimistic that it will soon be solved.

There are unpublished data (B. Cowan, personal communication) showing that LV mass by CMR varies by a small amount from end-diastole to end-systole, and this may be due to expulsion of intramyocardial blood into the venous system. The reproducibility of LV diastolic volumes is in general better than at end-systole because the volumes are larger, and this is probably as good a reason as any for working from the end-diastolic images to determine mass, but in addition, if the routine above is followed, there may be doubt as to whether LV mass is present in the most basal LV slice at end-systole if its quality is less than ideal, but by definition, LV mass is always present at end-diastole in the most basal slice. Normal reference ranges have been reported for LV70 dimensions and systolic function. This is important because most LV and RV parameters are dependent on gender, age, and body surface area. The reference data are shown in Tables 14-1 (LV) and 14-2 (RV) and Figure 14-12 (LV). For the very best interstudy reproducibility for followup studies and drug trials, it is necessary to go the extra mile in the analysis and always have the first set of cines with the regions of interest (ROIs) on screen alongside the follow-up cines during the analysis. At least a printout on paper of the first study and the ROIs used for analysis is required for comparison. For this reason, we routinely print out the diastolic and systolic ROIs in a systematic way for every patient on whom volumes are analyzed. As new approaches are being generated regularly, these recommendations may change in the future. In particular, automated post-processing of the cines

ISCHEMIC HEART DISEASE

Table 14-2 Normal SSFP Right Ventricular Volumes, Systolic Function, and Mass (Absolute and Indexed to Body Surface Area) by Age Decile (Mean, 95% Confidence Interval) in Males and Females 20–29 Years

30–39 Years

MALES EDV (mL), SD 25.4 ESV (mL), SD 15.2 SV (mL), SD 17.4 EF (%), SD 6.5 Mass (g), SD 14.4

Absolute Values 177 (127–227) 171 (121–221) 68 (38–98) 64 (34–94) 108 (74–143) 108 (74–142) 61 (48–74) 63 (50–76) 70 (42–99) 69 (40–97)

EDV/BSA (mL/m2), SD 11.7 ESV/BSA (mL/m2), SD 7.4 SV/BSA (mL/m2), SD 8.2 EF/BSA (%/m2), SD 4 Mass/BSA, (g/m2), SD 6.8

Indexed to Body Surface Area 91 (68–114) 88 (65–111) 35 (21–50) 33 (18–47) 56 (40–72) 55 (39–71) 32 (24–40) 32 (25–40) 36 (23–50) 35 (22–49)

FEMALES EDV (mL), SD 21.6 ESV (mL), SD 13.3 SV (mL), SD 13.1 EF (%), SD 6 Mass (g), SD 10.6

Absolute Values 142 (100–184) 136 (94–178) 55 (29–82) 51 (25–77) 87 (61–112) 85 (59–111) 61 (49–73) 63 (51–75) 54 (33–74) 51 (31–72)

EDV/BSA (mL/m2), SD 9.4 ESV/BSA (mL/m2), SD 6.6 SV/BSA (mL/m2), SD 6.1 EF/BSA (%/m2), SD 5.2 Mass/BSA (g/m2), SD 5.2

Indexed to Body Surface Area 84 (65–102) 80 (61–98) 32 (20–45) 30 (17–43) 51 (39–63) 50 (38–62) 37 (27–47) 38 (27–48) 32 (22–42) 30 (20–40)

40–49 Years

50–59 Years

60–69 Years

70–79 Years

166 (116–216) 59 (29–89) 107 (73–141) 65 (52–77) 67 (39–95)

160 (111–210) 55 (25–85) 106 (72–140) 66 (53–79) 65 (37–94)

155 (105–205) 50 (20–80) 105 (71–139) 68 (55–81) 63 (35–92)

150 (100–200) 46 (16–76) 104 (70–138) 70 (57–83) 62 (33–90)

85 30 55 33 34

(62–108) (16–45) (39–71) (25–41) (21–48)

130 (87–172) 46 (20–72) 84 (58–109) 65 (53–77) 49 (28–70) 76 27 49 38 29

(57–94) (14–40) (37–61) (28–49) (19–39)

82 28 54 34 33

(59–105) (13–42) (38–70) (26–42) (20–46)

124 (81–166) 42 (15–68) 82 (56–108) 67 (55–79) 47 (26–68) 72 24 48 39 27

(53–90) (11–37) (36–60) (29–49) (17–37)

79 25 53 35 32

(56–101) (11–40) (37–69) (27–42) (19–45)

117 (75–160) 37 (11–63) 80 (55–106) 69 (57–81) 45 (24–66) 68 21 46 40 26

(49–86) (8–34) (34–58) (30–50) (16–36)

75 23 52 35 31

(52–98) (8–37) (36–69) (27–43) (18–44)

111 (69–153) 32 (6–58) 79 (53–105) 71 (59–83) 43 (22–63) 64 19 45 41 24

(45–82) (6–32) (33–57) (31–51) (14–35)

BSA, body surface area; EDV, end-diastolic volume; EF, ejection fraction; ESV, end-systolic volume; SD, standard deviation; SSFP, steady state free precession; SV, stroke volume.

acceleration, which are derived from the flow curves, have been used to determine the burden of ventricular ischemia during dobutamine stress (see Chapter 15).75

Diastolic Function For diastolic function, velocity mapping of the transmittal E wave agrees with that of echocardiography but tends to underestimate the A wave, owing to beat-to-beat variability that occurs in the diastolic period.76 A decrease in the E/A ratio, however, can be demonstrated by CMR in cases of reduced ventricular compliance. Cine CMR also allows direct visualization of the abnormal diastolic filling in dilated ventricles with inflow directed not toward the apex but toward the free wall, giving rise to a well developed circular flow pattern turning back toward the septum and outflow tract, persisting through diastole.77 With 3D automated analysis of cine images, it is possible to measure volumes over a complete cardiac cycle, obtaining a volume curve whose first derivative is the flow curve; normal values of LV diastolic function with this method have been described and are shown in Figure 14-13.70 With the same approach, RV diastolic function curves have been produced and are shown in Chapter 28. Studies of myocardial velocity in diastole both at rest and during dobutamine have also been performed by using CMR, but experience with this technique is very limited.78 Tagging in diastole has also been examined, but again more experience is needed.79 More details of the assessment of diastolic function can be found in Chapter 5. 190 Cardiovascular Magnetic Resonance

Regional Function Cine CMR sequence allows a qualitative assessment of regional cardiac function in the same way as with echocardiography, but with improved image quality and a lower loss of nonvisualized segments. In addition, it is possible to image routinely in the true long and short axis of the heart without compromises that result from restricted angulation caused by awkward acoustic windows. This wall motion analysis can be performed at rest, with low dose dobutamine for detection of viable myocardium80,81 and high-dose dobutamine for detection of ischemia.82,83 Studies using real-time CMR are now being published; these too show that CMR is superior to dobutamine stress echocardiography in patients with limited acoustic access.84,85 Several methods based on cine CMR have been suggested to provide quantitative assessment of wall motion and wall thickening.86–91 In reality, however, myocardial dynamics are more complicated than simple thickening and 2D motion because of a complex interaction of contraction, expansion, twisting, and through-plane motion. This is best measured by the process of tagging.92 Briefly, tagging quantifies local deformation parameters and accurately measures several determinants of local loading conditions, such as wall thickness and curvature. These combined with systolic blood pressure give a measure of local loading, and by using the relation between local loading and deformation, a measure of local performance is produced. Tagging is discussed in more detail in Chapter 5.

+95% CI Mean –95% CI

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30

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30

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20

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30

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50 40 20

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90 80

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Figure 14-12 Left ventricular (LV) volumes, mass, and function in systole normalized to age and body surface area for males and females.

Cardiovascular Magnetic Resonance 191

14 ASSESSMENT OF CARDIAC FUNCTION

MALES

LV active peak filling rate/BSA - males (mL/s/m2)

LV early peak filling rate/BSA - males (mL/s/m2)

700 600 500 400 300

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Age (years) 8 7 6 5 4 3 2 1 0

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LV active peak filling rate/BSA - females (mL/s/m2)

FEMALES LV early peak filling rate/BSA - females (mL/s/m2)

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ISCHEMIC HEART DISEASE

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Age (years) Figure 14-13 Left ventricular (LV) function in diastole normalized to age and body surface area for males and females.

192 Cardiovascular Magnetic Resonance

CMR offers a rapid, accurate, and reproducible assessment of cardiac function that is free of geometric assumptions and is truly noninvasive. At present, CMR analysis of cardiac function can be slow, and experience is required, but automated border detection systems that incorporate both the long and short axis views, and thereby overcome the difficulties of basal slice differentiation, are already in use.

Aiding improved measures is the use of improved cine sequences such as SSFP, parallel imaging, and possibly intravascular contrast agents.93 A true single breath hold 3D volume acquisition should be achievable, and this will greatly speed up the acquisition. As CMR continues to become faster, it also becomes more cost-effective for this function, and this simple technique will become a major tool in clinical and research practice.

References 1. Schocken DD, Arrieta MI, Leaverton PE, Ross EA. Prevalence and mortality rate of congestive heart failure in the United States. J Am Coll Cardiol. 1992;20:301–306. 2. Anonymous. From the Centers for Disease Control and Prevention. Mortality from congestive heart failure: United States, 1980–1990. JAMA. 1994;271:813–814. 3. White HD, Norris RM, Brown MA, Brandt PW, Whitlock RM, Wild CJ. Left ventricular end-systolic volume as the major determinant of survival after recovery from myocardial infarction. Circulation. 1987;76:44–51. 4. Dunn FG, Pringle SD. Sudden cardiac death, ventricular arrhythmias and hypertensive left ventricular hypertrophy. J Hypertens. 1993;11:1003–1010. 5. Levy D, Garrison RJ, Savage DD, Kannel WB, Castelli WP. Prognostic implications of echocardiographically determined left ventricular mass in the Framingham heart study. N Engl J Med. 1990;332:1561–1566. 6. Ghali JK, Liao Y, Cooper RS. Influence of left ventricular geometric patterns on prognosis in patients with or without coronary artery disease. J Am Coll Cardiol. 1998;31:1635–1640. 7. Marantz PR, Tobin JN, Wassertheil-Smoller S, et al. The relationship between left ventricular systolic function and congestive heart failure diagnosed by clinical criteria. Circulation. 1988;77:607–618. 8. O’Keefe JH, Zinsmeister AR, Gibbons RJ. Value of normal electrocardiographic findings in predicting resting left ventricular function in patients with chest pain and suspected coronary artery disease. Am J Med. 1989;86:658–662. 9. Allison JD, Flickinger FW, Wright CJ, et al. Measurement of left ventricular mass in hypertrophic cardiomyopathy using MRI: comparison with echo. Magn Reson Imaging. 1993;11:329–334. 10. Kronik G, Slany J, Mosslacher H. Comparative value of eight M-mode echocardiographic formulas for determining left ventricular stroke volume. Circulation. 1979;60:1308–1316. 11. Teichholz LE, Kreulen T, Herman MV, Gorlin R. Problems in echocardiographic volume determinations: echocardiographic-angiographic correlations in the presence or absence of asynergy. Am J Cardiol. 1976;37:7–11. 12. Bellenger NG, Pennell DJ. Comparison of nuclear and non-nuclear techniques for the combined assessment of perfusion and function. In: Germano G, Berman DS, eds. Clinical Gated Cardiac SPECT. New York: Futura; 1999:349–377. 13. Bellenger NG, Marcus N, Davies LC, Yacoub M, Banner N, Pennell DJ. Left ventricular function and mass after orthotopic heart transplantation: a comparison of cardiovascular magnetic resonance with echocardiography. J Heart Lung Transpl. 2000;19:444–452. 14. Bellenger NG, Burgess M, Ray SG, et al., on behalf of the CHRISTMAS steering committee and investigators. Comparison of left ventricular ejection fraction and volumes in heart failure by two-dimensional echocardiography, radionuclide ventriculography and cardiovascular magnetic resonance: are they interchangeable? Eur Heart J. 2000; 21:1387–1396. 15. Spencer KT, Bednarz J, Mor-Avi V, DeCara J, Lang RM. Automated endocardial border detection and evaluation of left ventricular function from contrast-enhanced images using modified acoustic quantification. J Am Soc Echocardiogr. 2002;15:777–781. 16. Amico AF, Lichtenberg GS, Reisner SA, et al. Superiority of visual versus computerized echocardiography estimation of radionuclide left ventricular ejection fraction. Am Heart J. 1989;118:1259–1265. 17. Thomson H, Basmadjian AJ, Rainbird AJ, et al. Contrast echocardiography improves the accuracy and reproducibility of left ventricular remodeling measurements. J Am Coll Cardiol. 2001;38:867–875.

18. Malm S, Frigstad S, Sagberg E, Larsson H, Skjaerpe T. Accurate and reproducible measurement of left ventricular volume and ejection fraction by contrast echocardiography: a comparison with magnetic resonance imaging. J Am Coll Cardiol. 2004;44:1030–1035. 19. Hoffmann R, von Bardeleben S, ten Cate F, et al. Assessment of systolic left ventricular function: a multi-centre comparison of cineventriculography, cardiac magnetic resonance imaging, unenhanced and contrastenhanced echocardiography. Eur Heart J. 2005;26: 607–616. 20. Chuang ML, Hibberd MG, Salton CJ, et al. Importance of imaging method over imaging modality in noninvasive determination of left ventricular volumes and ejection fraction: assessment by two- and three-dimensional echocardiography and magnetic resonance imaging. J Am Coll Cardiol. 2000;35:477–484. 21. Corsi C, Lang RM, Veronesi F, et al. Volumetric quantification of global and regional left ventricular function from real-time three-dimensional echocardiographic images. Circulation. 2005;111:1161–1170. 22. Bellenger NG, Francis JM, Davies CL, Coats A, Pennell DJ. Establishment and performance of a magnetic resonance cardiac function clinic. J Cardiovasc Magn Reson. 2000;2:15–22. 23. Gaudron P, Eilles C, Kugler I, Ertl G. Progressive left ventricular dysfunction and remodeling after myocardial infarction: potential mechanisms and early predictors. Circulation. 1993;87:755–763. 24. Lapeyre AC, Klodas E, Rogers PJ, Sinak LJ, Hammell TC, O’Connor MK, Gibbons RJ. Quantitation of regional ejection fractions using gated tomographic imaging with 99mTc-sestamibi. Chest. 2005;127:778–786. 25. Sharir T, Germano G, Kavanagh PB, et al. Incremental prognostic value of post-stress left ventricular ejection fraction and volume by gated myocardial perfusion single photon emission computed tomography. Circulation. 1999;100:1035–1042. 26. Johnson LL, Verdesca SA, Aude WY, et al. Postischemic stunning can affect left ventricular ejection fraction and regional wall motion on post-stress gated sestamibi. J Am Coll Cardiol. 1997;30:1641–1648. 27. Iskandrian AE, Germano G, van Decker W, et al. Validation of left ventricular volume measurements by gated SPECT 99mTc-labeled sestamibi imaging. J Nucl Cardiol. 1998;5:574–578. 28. Anagnostopoulos C, Gunning MG, Pennell DJ, Laney R, Proukakis H, Underwood SR. Resting regional myocardial motion and thickening assessed by ECG-gated Tc99m-MIBI emission tomography and by magnetic resonance imaging. Eur J Nucl Med. 1996;23:909–916. 29. Lipiecki J, Cachin F, Durel N, de Tauriac O, Ponsonnaille J, Maublant J. Influence of infarct-zone viability detected by rest Tc99m sestamibi gated SPECT on left ventricular remodeling after acute myocardial infarction treated by percutaneous transluminal coronary angioplasty in the acute phase. J Nucl Cardiol. 2004;11:673–681. 30. Lim TK, Burden L, Janardhanan R, et al. Contrast echocardiography versus gated single photon emission computed tomography for the assessment of parameters of left ventricular remodeling after acute myocardial infarction. J Am Soc Echocardiogr. 2006;19:280–284. 31. Freiberg J, Hove JD, Kofoed KF, et al. Absolute quantitation of left ventricular wall and cavity parameters using ECG-gated PET. J Nucl Cardiol. 2004;11:38–46. 32. Heuschmid M, Rothfuss JK, Schroeder S, et al. Assessment of left ventricular myocardial function using 16-slice multidetector-row computed tomography: comparison with magnetic resonance imaging and echocardiography. Eur Radiol. 2006;16:551–559. 33. Lessick J, Mutlak D, Rispler S, et al. Comparison of multidetector computed tomography versus echocardiography for assessing regional left ventricular function. Am J Cardiol. 2005;96:1011–1015. 34. Juergens KU, Grude M, Fallenberg EM, et al. Using ECG-gated multidetector CT to evaluate global left ventricular myocardial function in Cardiovascular Magnetic Resonance 193

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57. Semelka RC, Tomei E, Wagner S, et al. Normal left ventricular dimensions and function: interstudy reproducibility of measurements with cine MR imaging. Radiology. 1990;174:763–768. 58. Semelka RC, Tomei E, Wagner S, et al. Interstudy reproducibility of dimensional and functional measurements between cine magnetic resonance imaging studies in the morphologically abnormal left ventricle. Am Heart J. 1990;119:1367–1373. 59. Pattynama PM, Lamb HJ, van der Velde EA, van der Wall EE, De Roos A. Left ventricular measurements with cine and spin-echo MR imaging: a study of reproducibility with variance component analysis. Radiology. 1993;187:261–268. 60. Shapiro EP, Rogers WJ, Beyar R, et al. Determination of left ventricular mass by MRI in hearts deformed by acute infarction. Circulation. 1989;79:706–711. 61. Lorenz CH, Walker ES, Morgan VL, Graham TP, Klein SS. Normal human right and left ventricular mass, systolic function and gender differences by cine magnetic resonance imaging. J Cardiovasc Magn Reson. 1999;1:7–21. 62. Bogaert JG, Bosmans HT, Rademakers FE, et al. Left ventricular quantification with breath hold MR imaging: comparison with echocardiography. MAGMA. 1995;3:5–12. 63. Sakuma H, Fujita N, Foo TKF. Evaluation of left ventricular volume and mass with breath hold cine MR imaging. Radiology. 1993; 188:377–380. 64. Bloomgarden DC, Fayad ZA, Ferrari VA, Chin B, St John Sutton M, Axel L. Global cardiac function using fast breath-hold MRI: validation of new acquisition and analysis techniques. Magn Reson Med. 1997;37:683–692. 65. Bellenger NG, Davies LC, Francis JM, Marcus NJ, Pennell DJ. Reduction in sample size for studies of remodelling in heart failure by the use of cardiovascular magnetic resonance. J Cardiovasc Magn Reson. 2000;2:271–278. 66. Grothues F, Smith GC, Moon JCC, et al. Comparison of interstudy reproducibility of cardiovascular magnetic resonance with two-dimensional echocardiography in normal subjects and in patients with heart failure or left ventricular hypertrophy. Am J Cardiol. 2002;90:29–34. 67. Bottini PB, Carr AA, Prisant M, Flickinger FW, Allison JD, Gottdiener JS. Magnetic resonance imaging compared to echocardiography to assess left ventricular mass in the hypertensive patient. Am J Hypertens. 1995;8:221–228. 68. Grothues F, Moon JC, Bellenger MG, Smith GC, Klein HU, Pennell DJ. Interstudy reproducibility of right ventricular volumes, function, and mass with cardiovascular magnetic resonance. Am Heart J. 2004; 147:218–223. 69. Taylor AM, Jhooti P, Wiesmann F, Keegan J, Firmin DN, Pennell DJ. MR navigator-echo monitoring of temporal changes in diaphragm position: implications for MR coronary angiography. J Magn Reson Imaging. 1997;7:629–636. 70. Maceira AM, Prasad SK, Khan M, Pennell DP. Normalized left ventricular systolic and diastolic function by steady state free precession cardiovascular magnetic resonance. J Cardiovasc Magn Reson. 2006; 8:417–426. 71. Mohiaddin RH, Longmore DB. Functional aspects of cardiovascular magnetic resonance imaging: techniques and application. Circulation. 1993;88:264–281. 72. Underwood SR, Klipstein RH, Firmin DN, et al. Magnetic resonance assessment of aortic and mitral regurgitation. Br Heart J. 1986;56:455–462. 73. Taylor AM, Stables RH, Poole-Wilson PA, Pennell DJ. Definitive clinical assessment of atrial septal defect by magnetic resonance imaging. J Cardiovasc Magn Reson. 1999;1:43–47. 74. Hundley WG, Li HF, Lange RA, et al. Assessment of left-to-right intracardiac shunting by velocity-encoded, phase-difference magnetic resonance imaging: a comparison with oximetric and indicator dilution techniques. Circulation. 1995;91:2955–2960. 75. Pennell DJ, Firmin DN, Burger P, et al. Assessment of magnetic resonance velocity mapping of global ventricular function during dobutamine infusion in coronary artery disease. Br Heart J. 1995;74: 163–170. 76. Karwatowski SP, Brecker SJ, Yang GZ, et al. Mitral valve flow measured with cine MR velocity mapping in patients with ischemic heart disease: comparison with Doppler echocardiography. J Magn Reson Imaging. 1995;5:89–92. 77. Mohiaddin R, Hasegawa M. Flow pattern in the dilated ischaemic left ventricle studied by MRI in healthy volunteers and in patients with myocardial infarction. J Magn Reson Imaging. 1995;5:493–498.

86. van Rugge FP, van der Wall EE, Spanjersberg SJ, et al. Magnetic resonance imaging during dobutamine stress for detection and localization of coronary artery disease: quantitative wall motion analysis using a modification of the centerline method. Circulation. 1994; 90:127–138. 87. Sechtem U, Sommerhoff BA, Markiewicz W, White RD, Cheitlin MD, Higgins CB. Regional left ventricular wall thickening by magnetic resonance imaging: evaluation in normal persons and patients with global and regional dysfunction. Am J Cardiol. 1987;59:145–151. 88. Pflugfelder PW, Sechtem U, White RD, et al. Quantification of regional myocardial function by rapid (cine) MRI. Am J Radiol. 1988;150:525. 89. van Rugge FP, van der Wall EE, Spanjersberg SJ, et al. Magnetic resonance imaging during dobutamine stress for detection and localization of coronary artery disease: quantitative wall motion analysis using a modification of the centerline method. Circulation. 1994;90: 127–138. 90. van Rugge FP, Holman ER, van der Wall EE, et al. Quantitation of global and regional left ventricular function by cine magnetic resonance imaging during dobutamine stress in normal human subjects. Eur Heart J. 1993;14:456–463. 91. Buser PT, Auffermann W, Holt WW, et al. Non-invasive evaluation of the global left ventricular function using cine MRI. J Am Coll Cardiol. 1989;13:1294. 92. Kramer C, Rogers WJ, Theobald TM, Power TP, Petruolo S, Reichek N. Remote noninfarcted region dysfunction soon after first anterior myocardial infarction: a magnetic resonance tagging study. Circulation. 1996;94:660–666. 93. Taylor AM, Panting JR, Keegan J, et al. Use of the intravascular contrast agent NC100150 Injection in spin echo and gradient echo imaging of the heart. J Magn Reson Imaging. 1999;1:23–32.

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78. Karwatowski SP, Mohiaddin RH, Yang GZ, Firmin DN, St John Sutton MG, Underwood SR. Regional myocardial velocity imaged by magnetic resonance in patients with ischaemic heart disease. Br Heart J. 1994;72:332–338. 79. Rademakers FE, Buchalter MB, Rogers WJ, et al. Dissociation between left ventricular untwisting and filling: accentuation by catecholamines. Circulation. 1992;85:1572–1581. 80. Baer FM, Voth E, Schneider CA, Theissen P, Schicha H, Sechtem U. Comparison of low-dose dobutamine-gradient-echo magnetic resonance imaging and positron emission tomography with [18F]fluorodeoxyglucose in patients with chronic coronary artery disease: a functional and morphological approach to the detection of residual myocardial viability. Circulation. 1995;91:1006–1015. 81. Dendale PA, Franken PR, Waldman GJ, et al. Low-dosage dobutamine magnetic resonance imaging as an alternative to echocardiography in the detection of viable myocardium after acute infarction. Am Heart J. 1995;130:134–140. 82. Pennell DJ, Underwood SR, Manzara CC, et al. Magnetic resonance imaging during dobutamine stress in coronary artery disease. Am J Cardiol. 1992;70:34–40. 83. Nagel E, Lehmkuhl HB, Bocksch W, et al. Noninvasive diagnosis of ischemia induced wall motion abnormalities with the use of high dose dobutamine stress MRI: comparison with dobutamine stress echocardiography. Circulation. 1999;99:763–770. 84. Yang PC, Kerr AB, Liu AC, et al. New real time interactive magnetic resonance imaging system complements echocardiography. J Am Coll Cardiol. 1998;32:2049–2056. 85. Hundley WG, Hamilton CA, Thaomas MS, et al. Utility of fast cine magnetic resonance imaging and display for the detection of myocardial ischemia in patients not well suited for second harmonic stress echocardiography. Circulation. 1999;100:1697–1702.

ISCHEMIC HEART DISEASE

CHAPTER 15

Wall Motion Stress Cardiovascular Magnetic Resonance: Ischemia, Viability, and Prognosis Thomas F. Walsh, Chirapa Puntawangkoon, Craig A. Hamilton, and W. Gregory Hundley

Master and Oppenheimer described the first stress test in 1929.1 Subsequently, numerous methods for inducing and identifying myocardial ischemia have been developed, including those that incorporate electrocardiography or in combination with imaging, including echocardiography, myocardial scintigraphy, and, most recently, cardiovascular magnetic resonance (CMR). For over two decades,2 stress CMR has been used to identify inducible ischemia,3 to measure contractile reserve,4 and to identify those at risk of future myocardial infarction and cardiac death.5,6 In this chapter, we will review the utility of CMR wall motion stress testing, including the pharmacologic agents used during the procedure, the evolution of the procedure from dobutamine stress echocardiography (DSE), the performance and safety of a dobutamine stress CMR (DCMR) examination, and the efficacy of DCMR in identifying inducible ischemia, myocardial viability, and cardiac prognosis. Data comparing DCMR with perfusion stress CMR will also be provided, along with an introduction to physiologic/exercise stress CMR.

INTRAVENOUS DOBUTAMINE AND ATROPINE Developed in 1975 by Tuttle and Mills,7 dobutamine is a synthetic beta-1-selective catecholamine agonist that binds to myocyte beta-1-receptors and increases intracellular calcium concentrations and myocardial contraction.8 In 1984, dobutamine was first described as a stress-testing agent.9 In addition to enhancing myocardial contractility, it increases heart rate by promoting peripheral vasodilation and a reflex tachycardia.8 Because of rapid metabolism by catechol-o-methyltransferase, dobutamine’s onset of action and half-life are both approximately 2 minutes; inactive metabolites are readily excreted by the kidneys and liver.10 At low doses, less than 10 mg/kg/min, myocardial contractility is augmented and minor peripheral vasodilation occurs.11 As the dose increases (20 to 40 mg/kg/min), heart rate and myocardial oxygen demand are increased along with myocardial work.12 In the presence of a flow-limiting epicardial stenosis, intravenous dobutamine promotes an oxygen supply/ demand mismatch that induces a left ventricular (LV) wall motion abnormality.13,14 196 Cardiovascular Magnetic Resonance

Intravenous dobutamine increases myocardial oxygen demand in a fashion similar to exercise. Thus, it is useful for stress testing in patients whose exercise ability is limited owing to peripheral vascular disease, physical incapacitation, pulmonary disease, or chronic deconditioning.15 Other benefits of dobutamine include its tolerance for peripheral vein infusion and its effectiveness in both perfusion and wall motion imaging.16 Side effects of dobutamine at stress test doses include chest pain, ventricular ectopy, dyspnea, nausea, elevated blood pressure, and arrhythmias. Dobutamine is contraindicated in patients with an obstructive cardiomyopathy, severe aortic stenosis, second- and third-degree heart block, sudden death syndrome, tachyarrhythmias, and previous episodes of dobutamine hypersensitivity. Although dobutamine exhibits positive inotropic activity, its chronotropic response can be suboptimal. Studies using changes in LV wall motion to identify inducible ischemia exhibit heightened sensitivity in appreciating flow-limiting epicardial luminal narrowing when the heart rate response during testing exceeds 85% of the maximum age-predicted heart rate response (MPHRR ¼ 220 age in years).17 For some patients, the addition of atropine (0.5 to 2 mg) may be needed to achieve an adequate heart rate response. Atropine is a naturally occurring alkaloid that is a competitive antagonist of muscarinic cholinergic receptors. Atropine increases heart rate by inhibiting vagal tone18 and is particularly useful in patients taking beta-blockers or rate-limiting calcium antagonists or those possessing a high parasympathetic drive. Atropine is administered in increments of 0.5 mg up to a total of 2 mg. Side effects of atropine include dry mouth, tachycardia, and hallucination.10 Abnormally high heart rates after atropine administration are best treated with beta-blockers.

SAFETY PROFILE OF DOBUTAMINE AND ATROPINE STRESS TESTING Safety is a primary concern for physicians and health care providers who administer dobutamine/atropine stress exams.16,19,20 The safety profile of dobutamine-atropine

Study

Year

Modality

Picano Geleijnse Garcia Hamilton Kuijpers Wahl

1994 1997 1999 2001 2004 2004

DSE DSE DSE DCMR DCMR DCMR

Number of Patients

Minor Events

Major Events

Deaths

2799 2246 325 469 400 1000

78% 71% 57% 67% 71% 64%

5% 5% 21% 0% 3% 6%

0 0 1 0 0 0

Major complications included death, ventricular fibrillation, supraventricular tachycardia, atrioventricular block, acute myocardial infarction, rupture of the left ventricular free wall or ventricular septal defects, transient ischemic attacks, or severe symptomatic hypotension. Minor complications included nausea/vomiting, anxiety, hypotension, and atropine poisoning with hallucinations lasting several hours in the absence of myocardial ischemia. Adapted from Mandapaka S, Hundley WG. Dobutamine cardiovascular resonance: a review. J Magn Reson Imag. 2006;24(3):499–512.

stress has been reported in six large studies, three each utilizing DSE and DCMR (Table 15-1).21–25 Major complications among more than 5000 DSE tests included death, ventricular fibrillation, sustained ventricular tachycardia, complete atrioventricular block, acute myocardial infarction, cardiac rupture/ventricular septal defect, transient ischemic attack, and severe hypotension. Minor complications included nausea, anxiety, and atropine poisoning with hallucinations lasting several hours. The rates of major and minor complications ranged from 0.1% to 5%, and from 57% to 78%, respectively. The incidences of major and minor side effects with DCMR are similar26–28 except that there were no episodes of death during DCMR in these studies (see Table 15-1). The frequency of other major events was 0% to 6%. The percentage of patients sustaining minor events was 64% to 71%. Major events are usually associated with continued administration of pharmacologic stress in the setting of concurrent myocardial ischemia. For this reason, it is important to identify ischemia promptly and to discontinue the stress protocol when ischemia is recognized.29 In the studies reported above, many (56%) of the patients that developed side effects had known coronary artery disease (CAD) and a diminished resting LV ejection fraction (LVEF).30 For this reason, a high level of scrutiny is used in reviewing image data for ischemia in the setting of reduced LVEF. Medications used for testing and emergency medications (Table 15-2) for treatment of potential cardiac complications need to be readily available to improve the probability of a successful outcome should complications occur during DCMR. Practice drills responding to major complications are suggested.

DOBUTAMINE STRESS ECHOCARDIOGRAPHY Since 1977, transthoracic echocardiography has been used to acquire images before, during, and after exercise (treadmill and bicycle) stress and pharmacologic stress.31 During DSE, two-dimensional (2D) images are acquired in four standard imaging planes: parasternal short axis (PSA), parasternal long axis (PLA), apical four-chamber (A4C), and apical two-chamber (A2C) views.32,33 A common dobutamine infusion protocol

initiates dobutamine at 10 mg/kg/min for 3 minutes, followed by 10-mg/kg/min infusion increments in 3-minute stages to 50 mg/kg/min. Echocardiographic images (PLA, PSA, A4C, A2C) are obtained prior to dobutamine and during the last minute of each stage. If the patient is not able to achieve 85% of MPHRR, supplemental atropine is administered in 0.5-mg increments each minute (up to 2.0 mg) until the target heart rate is reached.34 Throughout the test, the four echocardiographic planes are reviewed to monitor for ischemia. During testing, LV myocardial wall motion is defined as follows: 1 ¼ normal, 2 ¼ hypokinesis, 3 ¼ akinesis, and 4 ¼ dyskinesis. Myocardial ischemia is defined as an increase in wall motion score during testing or the persistence of severe hypokinesis during testing.35,36 The scoring system is determined by using a 17-segment model.37,38 A DSE is interrupted when there is evidence of ischemia, severe hypertension or hypotension, dysrhythmias, unstable angina, or unexpected neurologic findings.21 DSE is useful for assessing risk of future myocardial infarction, death, and the need for coronary intervention.38–40 For many years, it was the modality of choice for risk stratification of perioperative cardiovascular events during noncardiac surgery.41,42 However, DSE accuracy is dependent on adequate acoustic windows, which are often suboptimal, owing to a large body habitus, prior cardiothoracic surgery, obstructive airway disease, or advanced age.17,43,44 In addition, there is a high interobserver variability for the interpretation of the images. Consequently, the accuracy of the test depends heavily on the proficiency of the sonographer and the physician.29,45

DOBUTAMINE STRESS CARDIOVASCULAR MAGNETIC RESONANCE CMR is well suited to overcome many of the limitations of DSE. It has no acoustic window limitations and therefore can obtain comprehensive visualization of the LV myocardium regardless of body habitus and age.46–48 In addition, image acquisition can be standardized, and there is less reliance on an individual technologist’s technique for acquiring high-quality images.

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15 WALL MOTION STRESS CARDIOVASCULAR MAGNETIC RESONANCE: ISCHEMIA, VIABILITY, AND PROGNOSIS

Table 15-1 Safety Profile of Select Dobutamine Stress Echocardiography (DSE) and Dobutamine Stress Cardiovascular Magnetic Resonance (DCMR) Studies

ISCHEMIC HEART DISEASE

Table 15-2 Common On-Site Medications for CMR Stress Tests Stress Agent Dobutamine Atropine Adenosine

Stimulates b1 receptors to increase myocardial contractility and stroke volume; also acts as a peripheral vasodilator Anticholinergic that inhibits acetylcholine at the parasympathetic neuroeffector junction, blocking vagal effects on the sinoatrial and atrioventricular nodes; enhances conduction through the atrioventricular node and increases the heart rate Peripheral and central vasodilator; increases heart rate and blood flow

Treatment for ventricular dysrhythmia Lidocaine Decreases depolarization, automaticity, and excitability in the ventricles; can suppress ventricular dysrhythmia Procainamide Decreases excitability, conduction velocity, automaticity, and membrane responsiveness with prolonged refractory period Treatment for Supraventricular Tachycardia Metoprolol Diltiazem Verapamil

b1-antagonist that decreases myocardial contractility, heart rate, cardiac output, and blood pressure and reduces myocardial oxygen demand Calcium channel antagonist; reduces myocardial contractility and SA node conductor Calcium channel antagonist; reduces myocardial contractility and oxygen demand; decreases AV nodal conduction rate; and dilates coronary arteries and arterioles

Treatment for Chest Pain Nitroglycerin Aspirin Heparin Morphine sulfate

Reduces cardiac oxygen demand by decreasing left ventricular end-diastolic pressure (preload) and systemic vascular resistance (afterload) Analgesic/antipyretic that impedes clotting by blocking prostaglandin synthesis Anticoagulant that accelerates the formation of antithrombin III–thrombin complex and deactivates thrombin, preventing the conversion of fibrinogen to fibrin Opium derivative with primary effects on the central nervous system (CNS) and organs containing smooth muscle

Treatment of Allergic Reaction or Anaphylaxis Diphenhydramine Hydrocortisone Anxiolytic Diazepam Midazolam

Prevents histamine-mediated responses Corticosteroid that decreases inflammation, mainly by stabilizing leukocyte lysosomal membranes and suppressing the immune responses Benzodiazepine with relatively long half-life that works through the immune system to exert a calming effect Water-soluble benzodiazepine with a short half-life exerts calming effect at low doses and a CNS depressant effect at high doses

Reversal Agents Flumazenil Naloxone

Benzodiazepine receptor antagonist that can be used to reverse the effects of midazolam or diazepam Narcotic antagonist that can be used to reverse the effects of morphine sulfate

Other Furosemide Albuterol (nebulizer) Aminophylline

Loop diuretic for treatment of pulmonary congestion due to intravascular volume overload Inhaled beta-agonist that causes bronchodilation. Important for emergent treatment of bronchospasm Relaxes smooth muscle of the bronchial airways and pulmonary blood vessels for treatment of bronchospasms

DCMR stress tests are commonly performed on short, 1.5-T closed bore systems, with bore diameters ranging from 55 to 70 cm. Participants are asked to refrain from food and drink for several hours prior to the test, and beta-blockers are held the morning of the test. Electrocardiogram (ECG) interpretation within the magnetic for ST segment changes is difficult, owing to ECG distortion from the magnetohydrodynamic effect of pulsatile blood. Therefore, a resting supine 12-lead ECG is first performed outside of the magnet room. Intravenous access is then established, and a phased-array cardiac surface coil, brachial blood pressure cuff, pulse oximetry monitor, and respiratory gating belt are attached (Fig. 15-1).49 Present throughout testing are a physician and nurse, who continuously monitor the patient’s heart rhythm, respiratory rate, oxygen saturation, and blood pressure. The protocol is similar to that described above for DSE in that dobutamine infusions are started at 5 (if viability is being assessed) or 10 mg/kg/min followed by 3-minute stages of 10-mg/kg/ min increments to 50 mg/kg/min. If the patient’s target heart 198 Cardiovascular Magnetic Resonance

rate (85% MPHRR) is not obtained, atropine is added at 0.5-mg/min increments (up to 2 mg total) to augment the heart rate response (Fig. 15-2). Cine gradient recalled echo (GRE) or steady-state free precession (SSFP) bright-blood images (Table 15-3) are obtained in three apical views (horizontal, vertical, and apical long axis) and at least three short axis (the base, middle portion, and apex) LV planes at each level of pharmacologic stress (Fig. 15-3). If summation of disks will be used to calculate LV volumes (See Chapter 11.), short slices spanning the LV from base to the apex can be acquired.50 These views are obtained at rest, with each dobutamine stage, and after 10 minutes of recovery to confirm that and induced LV wall motion abnormality has resolved. After the patient is removed from the magnet room, a supine 12-lead ECG is repeated. The DCMR test is stopped when there is development of ischemia, as evidenced by a new wall motion abnormality (Figs. 15-4 and 15-5), a greater than 40-mmHg fall in systolic blood pressure, significant ventricular or atrial arrhythmias, or achievement of the 85% MPHRR.46,51–53

Scanner

Preparation and

recovery area

Figure 15-2 Dobutamine infusion protocol for cardiovascular stress testing. As shown, three apical and three short axis views are acquired at baseline (B) over a 3-minute time period. Also, heart rate (HR) and blood pressure (BP) are collected. Next, dobutamine is infused at 10 mg/kg/min, and the three apical and three short axis views are repeated. At peak stress, a combination of dobutamine and atropine (At) are infused to achieve 85% of the maximum predicted heart rate response for age. At this peak stress level, three apical views and three short axis views are repeated. Upon completion of the peak stress image acquisitions, recovery images (R) are obtained prior to patient discharge. The total scan time for an individual with a negative stress test averages 25 minutes using this protocol. Ox, oxygen saturation.

• Review H & P • Baseline ECG • HR, BP, Ox.

Emphasize target HR 85% MPHRR for age Three apical views and three short axis views HR BP

Patient prep

B

HR BP

10

HR BP

HR BP

20 + At

Total scan time ~25 min

R

Dobutamine (μg/kg/min)

30 min

3 min

3 min

3 min

3 min

Table 15-3 Advantages and Disadvantages of Various Cine Bright-Blood CMR Techniques Technique

Advantage

Disadvantage

GRE (breath hold) GRE (respiration triggered) SSFP SENSE Real-time

Reliable ↑ heart rate No breath hold Apical views Rapid acquisition Rapid acquisition

Breath hold ¼12 sec 45-sec scan Artifact ↑ heart rate Unknown reliability Noise, low spatial and temporal resolution

GRE, gradient recalled echo; SENSE, sensitivity encoding; SSFP, steady state free precession.

Cardiovascular Magnetic Resonance 199

15 WALL MOTION STRESS CARDIOVASCULAR MAGNETIC RESONANCE: ISCHEMIA, VIABILITY, AND PROGNOSIS

Figure 15-1 Images highlighting the preparation/ recovery area (left) as well as the scanning facility (right). Note that by utilizing a preparation and recovery area, cardiovascular magnetic resonance imaging center patient throughput is enhanced, as the scanning facility is not used for time-consuming, non–scan-related tasks such as insertion of intravenous catheters or placement of special monitoring equipment.

ISCHEMIC HEART DISEASE

Coronal locator

Long axis

Long axis

Anterior chest Antero septum LV Posterior wall Ao

LA

Four-chamber Apical short axis Septum

RV

Middle short axis

LV

Lateral wall

RA LA LV

Basal short axis

Two-chamber

RV 60°

Inferior LV wall LA

Anterior wall

Figure 15-3 Imaging acquisition strategy for obtaining three apical and three short axis views during dobutamine stress. As shown at the top left, a long axis plane (dashed line, top left panel) is obtained from a coronal locating plane. This long axis or apical three-chamber view is similar to the parasternal long axis plane acquired during the transthoracic echocardiography (middle panel top). Next, three short axis views (apical, middle, and basal) are obtained as plotted off the apical long axis or three-chamber view. As shown in the bottom panel (labeled “Basal short axis”), 60 rotations are utilized to obtain image acquisitions in the four-chamber and two-chamber planes of the left ventricle. Ao, aorta; LA, left atrium; LV, left ventricle; RA, right atrium; RV, right ventricle.

Long axis

Ap-AS

Four-chamber

Ap-Post

Two-chamber

Ap-Lat

Ap-Inf

Ap-Ant

Mid-AS Mid-Post

Mid-Sept Mid-Lat

Mid-Inf

Mid-Ant

Bs-AS

Bs-Sept

Bs-Inf

Bs-Ant

Bs-Post

Basal short axis

Ap-Sept

Middle short axis

Bs-Ant

Mid-Ant

Bs-Lat Bs-AS Bs-Post Bs-Sept

Bs-Inf

Bs-Lat

Mid-AS

Mid-Lat Mid-Post

Mid-Sept Mid-Inf

200 Cardiovascular Magnetic Resonance

Apical short axis

Ap-Ant Ap-Lat Ap-AS

Ap-Post

Ap-Sept Ap-Inf

Figure 15-4 Delineation of orthogonal left ventricular myocardial segments. Long axis, four-chamber, and twochamber views are displayed in the top series of images. Apical, middle, and basal segments are identified on each of the corresponding left ventricular myocardial walls. The orthogonal display of the same myocardial segments in short axis planes is displayed in the bottom row of images. Recognition of ischemia during dobutamine stress is achieved when an observed wall motion abnormality occurs during dobutamine/atropine infusions in two orthogonal planes. Ant, anterior; Ap, apical; AS, anterior septum; Bs, basal; Inf, inferior; Lat, lateral; Post, posterior.

End diastole

End systole

Middle short axis Normal response Apical view

Middle short axis Inducible ischemia Apical view

UTILITY OF DOBUTAMINE CARDIOVASCULAR MAGNETIC RESONANCE FOR IDENTIFYING INDUCIBLE ISCHEMIA (EARLY STUDIES) DCMR to detect LV wall motion abnormalities indicative of ischemia has been employed for nearly two decades.11,54 In 1992, Pennell and colleagues reported the first clinical (Fig. 15-6) use of DCMR.55 In their study, 25 patients were evaluated with DCMR, dobutamine thallium-201 single photon emission computed tomography (SPECT), and Xray coronary angiography. Study subjects received dobutamine infusions of up to 20 mg/kg/min and underwent cine GRE CMR. Of the 22 patients with significant CAD on angiography, 20 (91%) had LV wall motion abnormalities on CMR, and 21 (95%) had reversible ischemia on dobutamine SPECT. DCMR sensitivity for CAD was 91% and agreement with dobutamine SPECT was 96%. Subsequently, multiple studies have examined the utility of DCMR for a variety of clinical situations. Baer and colleagues reported on the utility of DCMR for individuals with no previous infarction or resting LV wall motion but with high grade (70% diameter) coronary artery stenoses.56 Twenty-eight subjects underwent DCMR (5, 10, 15, and 20 mg/kg/min at 1.5 T). Segmental wall motion analysis was performed in basal, midventricular, and apical short axis planes as well as in two transverse planes. Each of 72 segments was related to a coronary artery territory and was graded as normal, hypokinetic, or akinetic. Nearly

90% of epicardial coronary artery narrowings were correctly identified in the left anterior descending (LAD), 62% in the left circumflex (LCX), and 78% in the right coronary artery (RCA) territories. Single-vessel and multivessel disease were identified with 73% and 100% sensitivity, respectively. Comparison of DCMR and SPECT for the detection of CAD was examined in another study by Baer and colleagues.57 In 35 consecutive patients, sensitivity for CAD was 84% and 87% for CMR and SPECT, respectively. This comparison between the two techniques indicated that the efficacy of DCMR is maintained when compared with a well-established clinical tool. Van Rugge and colleagues assessed the efficacy of quantitative58 and qualitative59 measures of LV wall motion during DCMR for identification and localization of myocardial ischemia in patients with CAD. For the qualitative study of 45 patients, including 37 with CAD, 30 (81%) had an LV wall motion abnormality during DCMR with a specificity of 100%. These values were superior to those from exercise electrocardiography (70, or 63%) or DSE (51, or 63%). Sensitivity for detection of single-, double-, and triplevessel disease was 75%, 80%, and 100%, respectively. In the quantitative portion of the study, 39 patients with previously demonstrated LV wall motion abnormalities and 10 healthy subjects were assessed by using a short axis stress cine analyzed with a modified centerline method. The LV was divided into 100 cords with stress wall motion preselected as abnormal if four or more adjacent cords showed systolic wall thickening below 2 standard deviations of that obtained from the healthy subjects. This quantified approach had a sensitivity for detecting single-, double-, and triplevessel CAD of 88%, 91%, and 100%, respectively. Sensitivity Cardiovascular Magnetic Resonance 201

15 WALL MOTION STRESS CARDIOVASCULAR MAGNETIC RESONANCE: ISCHEMIA, VIABILITY, AND PROGNOSIS

Figure 15-5 Examples of normal contractility as well as inducible ischemia during dobutamine stress. In the top series of images, a middle short axis plane and an apical fourchamber plane are demonstrated at end diastole in a sequential series of images proceeding from diastole (far left images) to end systole (far right images). As shown in the normal situation, left ventricular cavity dimensions diminish symmetrically in all myocardial segments. In the bottom series of images with evidence of inducible ischemia, one notices that in both the middle short axis and the apical three-chamber view, the middle and apical segments of the septum do not thicken in systole. As shown by the white arrowheads, the failure to thicken in the same segments in two orthogonal planes is a region consistent with inducible ischemia.

20 µg

Max

Rest

10 µg

20 µg

Max

End diastole

10 µg

End systole

ISCHEMIC HEART DISEASE

Rest

Figure 15-6 Positive DCMR. Top, DCMR cine images (short axis and two-chamber view) of a patient with chest pain and suspected CAD. At rest and under increasing dobutamine, there is normal regional function. However, at maximum stress, an inferior and inferolateral wall motion abnormality is elicited. Bottom, X-ray coronary angiography images demonstrating single-vessel LCX disease (black arrows). Source: From Paetsch I, Jahnke C, Fleck E, Nagel E. Current clinical applications of stress wall motion analysis with cardiac magnetic resonance imaging. Eur J Echocardiogr. 2005;Oct;6(5):317–326, with permission.

for detecting stenosis or specific abnormalities in the LAD, RCA, and LCX was 75%, 87%, and 63%, respectively. These studies demonstrated the utility of DCMR using both quantitative and qualitative methods. Each of the early studies was relatively small, was performed in a single institution, and enrolled participants in whom the investigators were aware of the extent of coronary artery luminal narrowings prior to their undergoing DCMR. LV wall motion was assessed at baseline and at peak stress and not assessed continuously throughout the course of stress testing. Infusions were terminated prematurely when patients developed chest pain. Nevertheless, these studies effectively demonstrated the feasibility and diagnostic utility of DCMR.

DOBUTAMINE CARDIOVASCULAR MAGNETIC RESONANCE (CONTEMPORARY STUDIES) Over the last decade, advances in computer software and CMR hardware have made it possible to more easily perform DCMR in patients with suspected CAD. In 1999, Nagel and colleagues published the first study using highdose DCMR (up to 40 mg/kg/min) and atropine. Comparison was made with DSE.60 In 208 subjects (61 women, 147 men) who underwent DCMR (5, 10, 20, 30, and 40 202 Cardiovascular Magnetic Resonance

mg/kg/min during 3-minute stages with atropine added as an adjunct if necessary to achieve 85% of MPHRR for age) and DSE before X-ray coronary angiography. Regional LV wall motion was visually assessed by using a 16-segment model with CAD (50% diameter stenosis) present in 107 patients. Eighteen patients could not be examined by DSE, owing to poor image quality, and 18 could not be examined by DCMR (6 because of large body mass and 11 because of claustrophobia). Only four patients failed to reach the target heart rate. Overall, the diagnostic accuracy of DCMR was superior to that of DSE (p < 0.05), as the specificity increased from 70% to 86% and sensitivity increased from 74% to 86% (Table 15-4). While Nagel and colleagues directly compared DSE and DCMR, Hundley and coworkers used DCMR to study

Table 15-4 Comparison of DSE and DCMR as Compared with X-Ray Coronary Angiography Sensitivity Specificity Positive predicting value Negative predicting value Accuracy

DSE

DCMR

p

74% 70% 81% 61% 73%

86% 86% 91% 78% 86%

<0.05 <0.05 <0.05 <0.05 <0.005

Source: Adapted from Nagel et al. Noninvasive diagnosis of ischemiainduced wall motion abnormalities with the use of high-dose dobutamine stress MRI. Circulation. 1999;16;730–732, with permission).

89

93 91 80

80

81

84 82

80

75

79 72

68 62 50

Sensitivity

Specificity

Accuracy

PPV

NPV

Figure 15-7 Sensitivity, specificity, diagnostic accuracy, and positive and negative predictive values (PPV and NPV, respectively) of high-dose DCMR for CAD (>50% diameter stenoses). Orange bars ¼ left anterior descending coronary artery; peach bars ¼ left circumflex coronary artery, blue bars ¼ right coronary artery. (From Wahl A, Paetsch I, Roethemeyer S, Klein C, Fleck E, Nagel E. High dose dobutamine-atropine stress cardiovascular MR imaging after coronary revascularization in patients with wall motion abnormalities at rest. Radiology. 2004;233:210–216, with permission)

coronary artery stenoses were similar to or higher than those observed with other noninvasive imaging techniques. A meta-analysis of studies comparing DCMR with X-ray coronary angiography in 754 patients63 who underwent DCMR yielded sensitivity of 83% and specificity of 86% for identified greater than 50% coronary arterial luminal narrowings. From these data, it was concluded that DCMR demonstrates a high overall sensitivity and specificity for the diagnosis of CAD in patients with high disease prevalence. Initial 3-T DCMR data are now available. Kelle and colleagues reported on 30 consecutive patients undergoing 3-T DCMR using breath hold GRE with three LV short axis and three LV long axis images obtained during each 3-minute stage of dobutamine up to 40 mg/kg/min. Gadolinium (0.1 mmol/ kg) was administered before the rest acquisitions to improve image quality.63a DCMR was successful in 27 patients (90%), and 68% of 22 subjects who underwent angiography had significant CAD. Patient-based sensitivity and specificity were 80% and 86%, respectively, and accuracy was 82%.

Table 15-5 Sensitivity and Specificity of DCMR for Detection of 50% Coronary Arterial Luminal Narrowings Author 55

Pennell Van Rugge59 Van Rugge58 Baer57 Nagel60 Hundley17 Wahl61 Gebker89a Summary

Year

Dose (mg/kg/min)

Number of Patients

Sensitivity

Specificity

1992 1993 1994 1994 1999 1999 2004 2008

20 20 20 20 40 þ atropine 40 þ atropine 40 þ atropine 40

25 45 39 35 208 163 160 455 1130

91 81 91 84 86 83 89 85* 86**

100 80 86 83 84 82* 84**

*70%. **Weighted average.

Cardiovascular Magnetic Resonance 203

15 WALL MOTION STRESS CARDIOVASCULAR MAGNETIC RESONANCE: ISCHEMIA, VIABILITY, AND PROGNOSIS

163 patients with inadequate second harmonic DSE.43 Ten percent of patients undergoing DSE had five to eight myocardial segments not visualized with DSE, and 90% had eight or more. Continuous cines were obtained in three apical (horizontal, vertical, and apical long axis) views and three LV short axis (base, mid, and apex) views during 5minute dobutamine infusions of 5, 10, 20, and 40 mg/kg/ min to achieve 85% MPHRR (with atropine administered in 0.3-mg/min increments if the target heart rate was not reached). Inducible ischemia was evident in 36 patients and absent in 103 patients. While 29 patients developed chest discomfort, only 8% developed new wall motion abnormalities concurrently. Among patients who later underwent X-ray coronary angiography, the sensitivity and specificity for detecting CAD (>50% diameter stenosis) were both 83%. The authors concluded that DCMR provides a useful mechanism to diagnose inducible ischemia in patients who are not well suited for DSE. DSE can be difficult to interpret when patients have resting wall motion abnormalities with many false positive results. In many of the aforementioned studies, patients with a prior myocardial infarction or resting LV wall motion abnormalities were excluded. Wahl and colleagues performed high-dose DCMR in 160 patients with previously documented resting LV wall motion abnormalities.61 The subjects had prior revascularization and were difficult to assess with DSE. The sensitivity and specificity of DCMR for detecting CAD (50% diameter stenosis) were 89% and 84%, respectively (Fig. 15-7). The sensitivity of detecting luminal narrowing of one-, two-, or three-vessel CAD was 87%, 91%, and 100%, respectively. There was a single, nonfatal, major complication that required external defibrillation. This study demonstrates that high-dose DCMR can be useful even in patients with previously documented LV wall motion abnormalities and prior coronary revascularization. Many of the previous studies were unable to utilize single breath hold techniques and consequently led to longer exam times. Paetsch and colleagues were among the first to use cine SSFP sequences with parallel imaging for faster data acquisition. Sensitivity and specificity were 89% and 80%, respectively.62 These bright-blood imaging techniques allow the acquisition of multiple slice positions during a single breath hold, and therefore reduce scan time by a factor of 3 to 4 (see Fig. 15-6). As shown in Table 15-5, the sensitivity and specificity of DCMR for identifying greater than 50%

Several studies have examined the utility of DCMR for identifying LV myocardial segments that, owing to myocardial stunning or hibernation, display abnormal LV wall motion at rest and will improve in contractility after coronary arterial revascularization procedures. Early studies examining contractile reserve compared DCMR with DSE and radionuclide studies. Development of tagging has better enabled quantitative assessment of myocardial viability. Several studies have utilized tagging to quantitatively assess myocardial viability, including movement of individual layers, circumferential shortening, and wall thickening. All of these represent important determinants of potential recovery after vascularization. Dendale and colleagues were among the first to explore whether low-dose DCMR represents a reasonable alternative64 by performing low-dose DCMR during the second week after myocardial infarction in patients without risk of serious adverse events. In comparison with DSE, there was similar efficacy and overall accuracy in prediction of recovery of contractile function (79% with DCMR and 83% with DSE). This study was limited by the lack of quantitative analysis of the DCMR images. Quantitative analysis was performed by Saito and colleagues in a study comparing DSE with tagged DCMR images.65 With the use of software that automatically traced all intersecting points, the LV wall motion of patients with ischemic heart disease and healthy subjects was analyzed. Viability of the myocardium was subsequently determined by a quantitative analysis of the improvement of the regional wall motion. The sensitivity of low-dose DCMR with tagging was 76% compared with 66% for DSE, while the specificity of DCMR and DSE was 86% and 100%,

respectively. Overall accuracy of DCMR and DSE was 78% and 72%, respectively. The authors concluded that quantitative evaluation of LV wall motion in every segment was possible. Thus, DCMR with tagging was considered superior to DSE for evaluation of myocardial viability. Low-dose DCMR has also been compared with metabolic assessments of viability obtained during radionuclide studies. Baer and colleagues66 reported that DCMR provided more accurate information regarding viability than resting LV end-diastolic wall thickness. When compared with positron emission tomography (PET) with 18Ffluorodeoxyglucose, DCMR was found to be superior for residual metabolic activity (sensitivity: 81%; specificity: 95%, positive predictive value: 96%) than PET (sensitivity: 72%; specificity: 89%, positive predictive value: 91%). In a small study, Sayad and coworkers67 demonstrated that tissue tagging along with DCMR predicts viability of hibernating and stunned myocardium. Among 10 patients with resting segmental LV wall motion abnormalities referred for coronary revascularization, quantitative DCMR had a sensitivity of 89%, a specificity of 93%, a negative predictive value of 82%, and a positive predictive value of 96% for recovery of segmental function after mechanical revascularization. There was an excellent correlation between LV endsystolic wall thickness at peak dobutamine infusion and after revascularization (Fig. 15-8). By using a combination of three apical views, all myocardial segments were completely assessed. Myocardial segments with a resting end-systolic wall thickness of less than 7 mm were unlikely to demonstrate contractile reserve with dobutamine or recovery of function after mechanical revascularization. After a myocardial infarction, the transmural extent of necrosis and decrease in LV ejection fraction can vary considerably. To determine the contribution of residual subepicardial viable myocardium to global LV function, Bogaert and colleagues studied 12 patients with single-vessel disease after successful reperfusion after their first transmural anterior myocardial infarction with DCMR tagging (Fig. 15-9).68 Measured parameters included regionally quantified fiber strains, wall thickening, and ejection fraction at 1 week

Contractile reserve in abnormal segment N = 43

Correlation between end-systolic thickness at peak dobutamine and post-revascularization N = 43

20 20 End-systolic thickness postrevascularization (mm)

End-systolic thickness (mm)

ISCHEMIC HEART DISEASE

UTILITY OF DOBUTAMINE CARDIOVASCULAR MAGNETIC RESONANCE IN IDENTIFYING MYOCARDIAL VIABILITY

16 12 8 4 0

y = 1.08x − 0.31 r = 0.95

10 5 0

Rest

A

15

Peak

Post

Dobutamine revascularization

0

B

5

10

15

20

End-systolic thickness at peak dobutamine (mm)

Figure 15-8 A, Plot of end-systolic wall thickening at rest, at peak dobutamine, and after revascularization in all 43 abnormal segments. Squares represent the three segments that had contractile reserve with dobutamine but no recovery of function after revascularization. Triangles represent the segment without contractile reserve but which improved after revascularization. No segment with end-systolic wall thickness less than 7 mm had contractile reserve or improved after revascularization. B, Regression line for end-systolic wall thickening at peak dobutamine and after revascularization. Source: From Sayad DE, Willett DL, Hundley WG, Grayburn PA, Peshock RM. Dobutamine magnetic resonance imaging with myocardial tagging quantitatively predicts improvement in regional function after revascularization. Am J Cardiol. 1998 Nov 1;82(9):1149–1151, A10, with permission. 204 Cardiovascular Magnetic Resonance

B

C

D

Figure 15-9 Short-axis basal views at baseline (rest) and peak dobutamine dose (40 mg) before and after tagging. Diastolic phase of a normal left ventricle at rest without tagging (A) and with tagging (B). C, Early systolic phase of the same left ventricle at peak dobutamine dose. Wall contraction pattern (wall thickening) appears to be normal. D, Same phase as panel C. Preservation of tagging lines at the anterior wall of the left ventricle. There is akinesia of three segments: septal, anterior-septal, and anterior (arrows), which indicate myocardial ischemia. CAG (coronary angiography) showed significant stenosis in the left anterior descending coronary artery. Source: From Kuijpers, et al. Circulation. 2004, with permission.

and 3 months after infarction. Normal and shear strains were calculated on the basis of displacement of node points along the subepicardial and subendocardial contours from end diastole to end systole. Improved subepicardial fiber shortening during dobutamine administration between week 1 and the 3-month mark was associated with an improved regional LV wall motion and global LVEF. Consequently, the functional recovery of viable subepicardial regions was attributed as a mechanism of late improvement in regional and global ejection fraction after transmural infarction. This study confirmed the necessity of early reperfusion, as restoration of flow in the infarct related vessel appears to preserve fibers in the subepicardial and lateral border zone of a transmural infarct (Fig. 15-10). Whereas Bogaert and coworkers studied mathematically quantified fiber strains, Geskin and colleagues examined percent intramyocardial circumferential segment shortening in a group of 20 patients after their first reperfused myocardial infarction.69 A normal response to dobutamine administration was defined as a 5% or greater increase in percent intramyocardial segment shortening. Using low-dose dobutamine, they demonstrated that an increase in circumferential shortening in resting dysfunctional segments during dobutamine infusion predicted the myocardial functional recovery at 8 weeks after myocardial infarction. The investigators noted that a normal increase in shortening elicited by dobutamine within dysfunctional midwall and subepicardium predicted greater functional recovery at 8 weeks after infarction, but the response within the subendocardium was not predictive. These results confirm that the response of intramyocardial function to low-dose DCMR after reperfused infarction can be quantified with CMR tagging. While wall thickening as measured by DCMR is one method of assessing contractility, scar quantification by late gadolinium enhancement (LGE) imaging represents another predictor of systolic function after revascularization. Wellnhofer and colleagues examined 29 patients with CAD with DCMR and LGE.70 Similar efficacy was noted for segments with no (0%) LGE and segments with extensive (>50% transmural) LGE. However, for segments with an intermediate (1% to 49%) LGE pattern, low-dose DCMR was superior to LGE for identifying segments that demonstrated an improvement in contractility after revascularization (Fig. 15-11). The reason for this finding may

relate to DCMR as a functional test, while LGE assesses underlying cardiac structure and is superior for localizing and quantifying scar. The authors concluded that LGE and DCMR provide complementary information. We commonly perform both tests in patients who are referred for viability testing.

PROGNOSIS While being able to diagnose transient myocardial ischemia and viability has important short-term implications, the long-term predictive value of DCMR has also been established. Hundley and colleagues reported on the utility of DCMR for cardiac prognosis in 279 patients who underwent DCMR and were followed for 20 months for the occurrence of unstable angina, congestive heart failure, coronary arterial revascularization, cardiac death, and death attributable to any cause.71 Reduced LVEF and DCMR evidence of inducible myocardial ischemia were associated with future myocardial infarction and cardiac death (Fig. 15-12). By using a multivariate analysis, a low LVEF, inducible ischemia, and contractile reserve were associated with infarction and cardiac death, independent of the presence of traditional risk factors for coronary arteriosclerosis or myocardial infarction. These findings have also been confirmed in a female population.71a Jahnke and colleagues reported on the prognostic utility of DCMR,72 comparing the prognostic value of adenosine perfusion CMR and DCMR wall motion in 461 patients. The prognostic utility of the presence or absence of dobutamine-induced LV wall motion abnormalities was similar to that identified during adenosine perfusion CMR. The 3-year event-free survival rate was 99.2% for patients with normal adenosine stress CMR perfusion or DCMR. DCMR has been also been used to determine preoperative cardiovascular risk in cardiac and noncardiac surgery patients. Rerkpattanapipat and colleagues studied a group of 102 patients for preoperative risk evaluation for noncardiac surgery.73 Congestive heart failure, myocardial infarction, and death were the major events that were assessed. Six patients experienced cardiac events, including death (N ¼ 4), nonfatal myocardial infarction (N ¼ 1), and congestive heart failure (N ¼ 1) during surgery or the postoperative period. A positive DCMR was independently associated with major perioperative cardiac events. Cardiovascular Magnetic Resonance 205

15 WALL MOTION STRESS CARDIOVASCULAR MAGNETIC RESONANCE: ISCHEMIA, VIABILITY, AND PROGNOSIS

A

20 µg

Max

End diastole

10 µg

End systole

ISCHEMIC HEART DISEASE

Rest

Figure 15-10 Contractile reserve in left circumflex and right coronary artery disease. The format is the same as in Figure 15-6. At rest, a severely hypokinetic inferolateral segment was found (white arrows). Under low-dose dobutamine infusion up to 20 mg/kg/min, a substantial and gradual improvement in the contractile response of the inferolateral segment could be demonstrated, though at maximum stress, both the inferolateral and inferior segments became akinetic. The “biphasic profile of contraction” of the inferolateral segment was compatible with viable myocardium. On angiogram, there is subtotal occlusion of the distal portion of a marginal branch (arrowheads) and high-grade stenosis of the distal left circumflex as well as a high-grade stenosis of the distal right coronary artery (RCA) (arrows). The open arrows demonstrate the development of small coronary collaterals of the RCA to compensate for the chronically reduced blood supply to the inferolateral wall. Source: From Paetsch I, Jahnke C, Fleck E, Nagel E. Current clinical applications of stress wall motion analysis with cardiac magnetic resonance imaging. Eur J Echocardiogr. 2005 Oct;6(5):317–326, with permission.

TISSUE TAGGING DURING DOBUTAMINE CARDIOVASCULAR MAGNETIC RESONANCE The application of tissue tagging (using radiofrequency pulses to create regular dark bands of selective saturation at the onset of image acquisition) helps visualization of wall motion with a high degree of precision.74–76 With this technique, it is possible to calculate the local Lagrangian strain tensor, which describes the motion around a given point in the tissue as it transverses space and time. Different tagging schemes have been developed, including lines,77 grids,78 and radial stripes.79 For the quantitative analysis using tissue 206 Cardiovascular Magnetic Resonance

tagging, tag intersection points are used to assess contraction (by measures of thickening or strain) and relaxation. To date, tissue tagging has been the most widely reported form of quantitative analysis of LV systolic thickening associated with myocardial ischemia and contractile reserve. Kuijpers and colleagues reported on the qualitative assessment of tagged images in comparison with nontagged cine CMR80 among 194 patients with chest pain. The researchers used DCMR with and without myocardial tagging at three LV short axis levels. Patients with new wall motion abnormalities underwent X-ray coronary angiography. Tagged DCMR images detected new wall motion abnormalities in 68 patients, compared with only 58 patients without tagging; X-ray coronary angiography data corresponded well (p ¼ 0.002) (see Fig. 15-9) . This was the first study to suggest that high-dose DCMR with qualitative assessment of tag deformation may detect inducible

100

N = 86

N = 53

N = 51

N = 45

N = 53

No

<25

25–49

50–74

≥75

80

(%)

60

40

20

0 LGE (% transmurality) Prevalence Specificity

Cardiac death or MI after dobutamine/atropine MRI LVEF ≥ 40%; (–) ischemia

Proportion hard event-free

1.0 0.9

LVEF ≥40%; (+) ischemia

0.8

LVEF < 40%

0.7 0.6 0.5 0

6

12

18

24

30

36

Time (months) Figure 15-12 Kaplan-Meyer survival curve indicative of cardiovascular events. In 279 individuals undergoing dobutamine stress magnetic resonance imaging for chest pain, the proportion of individuals free of cardiac death or myocardial infarction (y-axis) and the three-year follow-up (x-axis) are shown. As noted, individuals with a left ventricular ejection fraction greater than 40% and no inducible ischemia exhibited fewer cardiac events compared to those individuals with inducible ischemia or those with a left ventricular ejection fraction (LVEF) less than 40%.

LV wall motion abnormalities with greater accuracy than is seen in images acquired without tagging. A unique application of tissue tagging includes the potential assessment of LV diastolic function. Paetsch and colleagues evaluated the diastolic parameters from myocardial tagging of 20 patients with low- and high-dose DCMR to identify patients with significant CAD.81 Diastolic (velocity

Sensitivity Correct predictions

of circumferential lengthening and diastolic rotation velocity) and systolic (circumferential shortening, systolic rotation, and systolic rotation velocity) parameters during low-dose DCMR changed significantly in patients without CAD but not in those with CAD (p < 0.05). Identification of patients without and with CAD was possible by using the diastolic parameter “time to peak untwist” (Fig. 15-13). Thus, the results of this study showed that myocardial tagging helps in a quantitative analysis of systolic and diastolic function during both low- and high-dose dobutamine stress. Another potential development for the use of tissue tagging during DCMR includes the use of automated analyses of the images with harmonic phase magnetic resonance (HARP). Image inspection with HARP has the possibility to reduce tag analysis times substantially. Developed by Osman and colleagues,82,83 this technique concentrates on the Fourier transformation of tagged images, which is directly related to the motion of the tag lines. The HARP technique can be used with any tagging strategy, provided that the tag pattern is planar and tag lines are placed a uniform distance from one another. In addition to LV wall motion and thickening, myocardial strain can be analyzed. Selecting tag filter specifications for the images requires up to 20 minutes, but once the filter has been set, a full quantitative analysis of the data typically takes less than 3 minutes. This represents a substantial time savings over manual analysis, which can require up to 8 hours. Further research is proceeding in this area to provide users with a mechanism to obtain quantitative tag data in near real time. By standardizing acquisitions of tagged images, one can preset the filter and provide very efficient myocardial strain mapping. By extracting HARP images directly from the raw data using faster microprocessors, the possibility of on-line detailed quantitative assessment of 2D myocardial Cardiovascular Magnetic Resonance 207

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Figure 15-11 Transmurality of late gadolinium enhancement (LGE): subgroup analysis. Bars refer to the prevalence of recovery and sensitivity, specificity, and percentage of correct predictions by DCMR and are subgrouped with respect to LGE (cutoff: 25%). The specificity of DCMR remains high irrespective of the extent of LGE. The test retains a high sensitivity in 25% to 49% LGE. Source: From Wellnhofer E, Olariu A, Nagel E, et al. Magnetic resonance low-dose dobutamine test is superior to SCAR quantification for the prediction of functional recovery. Circulation. 2004 May 11;109 (18):2172–2174. with permission.

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End diastole

End systole

End systole

A

B

C Figure 15-13 Representative end-diastolic and end-systolic apical images of a patient without significant coronary artery disease. Row A, Rest. Row B, Low-dose dobutamine stress. Row C, High-dose dobutamine stress. Note the increase in systolic rotation and endocardial motion from rest to low-dose stress. Source: From Paetsch I, Foll D, Kaluza A, et al. Magnetic resonance stress tagging in ischemic heart disease. Am J Physiol Heart Circ Physiol. 2005; 288:H2708-H2714, with permission.

strains for use during stress testing may become a reality in the near future. Pan and colleagues84 have proposed a fast and semiautomatic method for tracking three-dimensional (3D) cardiac images. Using a material mesh model that is built to represent a collection of material points inside the LV wall, the phase time-invariance property of material points is then used to track mesh points. With a series of nine time frame CMR images, researchers were able to initialize settings, build the mesh, and track the 3D images in approximately 10 minutes. Eventually, more complex mesh algorithms will be developed that will allow 3D calculation of regional and global wall motion in real time. 208 Cardiovascular Magnetic Resonance

ADENOSINE AND DIPYRIDAMOLE AS WALL MOTION STRESS AGENTS DURING STRESS CARDIOVASCULAR MAGNETIC RESONANCE Dipyridamole and adenosine exert their effect through coronary vasodilation, producing up to fivefold increases in coronary artery blood flow. Dipyridamole raises the interstitial levels of adenosine by inhibiting adenosine

Perfusion had the highest accuracy values for coronary stenoses less than 75% (cutoff: 59%), whereas wall motion abnormalities exhibited the highest accuracy for coronary stenoses greater than 75% (cutoff: 84%) (p < 0.001).89 The authors concluded that during dipyridamole CMR stress testing, perfusion and wall motion abnormalities exhibit a similar diagnostic accuracy, with perfusion showing higher sensitivity, particularly in the detection of moderate stenoses, and wall motion showing higher specificity. Both studies demonstrated that adenosine and dipyridamole for the detection of stress-inducible wall motion abnormalities are inducible only in patients with more severe coronary artery stenosis (>75%). Finally, Gebker and coworkers examined the value of adding myocardial perfusion during DCMR in 455 consecutive patients who had perfusion CMR images acquired in three LV short axis views at rest and at peak dobutamineatropine stress.89a Combined perfusion and wall motion methodology improved sensitivity for CAD (70% diameter stenosis) from 85% to 91% but provided lower specificity (70% versus 82%), resulting in a similar diagnostic accuracy of 84% for both approaches.

LEFT VENTRICULAR WALL MOTION DURING EXERCISE STRESS CARDIOVASCULAR MAGNETIC RESONANCE Potential advantages of physiologic stress include reduction in stress protocol duration, acquisition of data regarding functional capacity and hemodynamic response, and the ability to extract information related to cardiac prognosis on the basis of exercise duration. In the past, exercise stress CMR using supine leg ergometry was limited by motion artifacts associated with relatively long image acquisition times. Faster CMR sequences have become available that allow for rapid image acquisition after exercise stress. Roest and colleagues studied 16 healthy subjects using bicycle exercise in the supine position on a magnetic resonance–compatible bicycle ergometer (Fig. 15-14).90 Respiratory motion artifacts were minimized by brief end-expiratory breath holding for all LV short axis images. Stroke volume and ejection fraction increased in both ventricles, while the end-systolic volume of both ventricles decreased. End-diastolic volume remained the same. Similar results were obtained for GRE and with standard echo planar CMR. This study demonstrates that exercise CMR can be used to assess physiologic changes in both the left and right ventricles. While other studies have examined supine bicycle ergometry stress, Rerkpattanapipat and coworkers demonstrated the feasibility of CMR in combination with upright treadmill exercise in 27 patients. Cine CMR images were obtained less than 1 minute post-exercise.91 Ischemia (>70% diameter stenosis) was identified in 14 patients, with good overall sensitivity (79%) and specificity (85%). By placing the treadmill approximately 20 feet away from the CMR table and having patients exercise immediately before imaging, the authors were able to counteract one of the previous limitations of stress exercise CMR in that images were acquired within 60 to 90 seconds after Cardiovascular Magnetic Resonance 209

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breakdown via the adenosine deaminase pathway, and it inhibits the facilitated uptake of adenosine into cells. Normally, there is also a mild rise in heart rate and fall in blood pressure. Because caffeine and theophylline block adenosine binding and can significantly reduce adenosine’s vasodilatory effects, it is recommended to have patients avoid use of these substances prior to adenosine or dipyridamole administration. Dipyridamole can be infused at 0.142 mg/kg/min for 4 minutes with maximum result achieved 3-4 minutes after infusion.85 The safety profile of dipyridamole is favorable, with a well-known and low complication rate. Side effects include chest pain, shortness of breath, flushing, and dizziness, all of which can be reversed with aminophylline administration.86 Fatal side effects are very rare; however, owing to the potential for bronchospasm, dipyridamole is contraindicated in patients with asthma. Transient side effects occur more frequently with adenosine infusion, but reversal with aminophylline is not needed, owing to the very short half-life (10 seconds). Adenosine is administered at 140 mg/kg/min over 4 to 6 minutes and carries an increased risk of first-degree atrioventricular block.62 Therefore, adenosine is contraindicated in patients with sick sinus syndrome or greater than first-degree block. Because a wall motion abnormality would generally require a “steal” phenomenon with decreased flow, the use of vasodilation wall motion stress would be expected to have a lower sensitivity than DCMR or vasodilator stress CMR would. Available data confirm this. Pennell and coworkers were the first to report on the use of dipyridamole with CMR in the assessment of LV wall motion.87 In 40 patients, dipyridamole was infused at a dose of 0.56 mg/ kg along with a 10-mg bolus after 10 minutes. Compared to thallium tomography, the sensitivity CMR systolic dysfunction was 67%.88 Not surprisingly, there was an apparent inability to detect smaller areas of ischemia. The procedure was well tolerated overall, and only mild side effects (flushing, chest pain, and headache) were common. Paetsch and colleagues compared DCMR with adenosine perfusion CMR for identifying LV wall motion abnormalities indicative of CAD.62 In 79 patients with suspected or known CAD who were scheduled for X-ray coronary angiography, segmental perfusion was graded by the transmural extent of inducible perfusion deficits (<25%, 25% to 50%, 51% to 75%, and >75%). Fifty-three patients (67%) had CAD (50% coronary arterial diameter stenosis). The sensitivity of adenosine wall motion analyses and adenosine-induced perfusion defects were 89%, 40%, and 91% with corresponding specificity of 80%, 96%, and 62%, respectively. Adenosine-induced wall motion abnormalities were seen only in segments with more than 75% transmural perfusion deficit. These findings suggest that DCMR is superior to adenosine stress perfusion and cine CMR for detecting CAD. Another comparison study by Pingitore and colleagues involved a head-to-head comparison in 94 subjects who underwent perfusion and function CMR during dipyridamole. Wall motion and the perfusion reserve index exhibited similar diagnostic accuracy (86% for both) for identifying greater than 50% coronary artery diameter reduction. Wall motion analyses had a higher specificity (96% versus 66%, p < 0.01) and lower sensitivity (82% versus 93%, p < 0.05) than the perfusion reserve index.

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Exercise bike attached to the scanner

Treadmill positioned outside the scan room

Figure 15-14 Methods of exercise-induced stress in patients undergoing wall motion analyses during magnetic resonance stress testing. Left, An individual pedals a nonferromagnetic electronically braked bike (Loda, the Netherlands). Right, An individual performs maximal exercise stress according to a protocol on a treadmill located outside the scanning room. Immediately after exercise, the participant is placed on the cardiovascular magnetic resonance imaging table, and images of left ventricular wall motion are collected within 60 seconds of exercise cessation.

cessation of treadmill exercise. Comprehensive coverage of the myocardium was not achieved in all subjects. Further investigations in this area are needed to determine the clinical utility of exercise-induced LV wall motion abnormalities for the assessment of patients with suspected CAD. A similar approach, but with the treadmill positioned in the magnet room, was used by Jekie and colleagues in 20 healthy subjects (see Fig. 15-14).91a Real-time, short axis cines were acquired at seven levels within 45 seconds of exercise cessation. This was followed by perfusion CMR within 60 seconds of exercise cessation.

CONCLUSION Inducible LV wall motion abnormalities during stress CMR has been found useful for identifying patients with

CAD and cardiac prognosis. Improvement in resting LV regional wall motion with low-dose DCMR accurately identifies viable myocardium that will improve contractility after mechanical revascularization and appears to be superior to LGE methods for intermediate LGE values (1% to 50% transmural hyperenhancement). Wall motion assessment using dobutamine is safe and superior to using adenosine for the detection of regional LV abnormalities. DCMR is superior to DSE primarily because of the ability to acquire high-resolution images in any tomographic plane without limitations imposed by body habitus or acoustic window. With new tagging techniques, qualitative and quantitative assessments of LV wall motion and strain are possible. While not yet perfected for routine clinical use, these approaches may soon provide realtime 3D determinations of regional and global LV function during DCMR.

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74. Garot J, Bluemke DA, Osman NF, et al. Fast determination of regional myocardial strain fields from tagged cardiac images using harmonic phase MRI. Circulation. 2000;101(9):981–988. 75. Ryf S, Spiegel MA, Gerber M, Boesiger P. Myocardial tagging with 3D-CSPAMM. J Magn Reson Imaging. 2002;16(3):320–325. 76. Pan L, Lima JA, Osman NF. Fast tracking of cardiac motion using 3D-HARP. Inf Process Med Imaging. 2003;18:611–622. 77. Mosher TJ, Smith MB. A DANTE tagging sequence for the evaluation of translational sample motion. Magn Reson Med. 1990;15:334–339. 78. Axel L, Dougherty L. Heart wall motion: improved method of spatial modulation of magnetization for MR imaging. Radiology. 1989; 171:349–350. 79. Bolster Jr BD, McVeigh ER, Zerhouni EA. Myocardial tagging in polar coordinates with use of striped tags. 1990;177:349–350. 80. Kuijpers D, Ho KY, van Dijkman PR, Vliegenthart R, Oudkerk M. Dobutamine cardiovascular magnetic resonance for the detection of myocardial ischemia with the use of myocardial tagging. Circulation. 2003;107(12):1592–1597. 81. Paetsch I, Foll D, Kaluza A, et al. Magnetic resonance stress tagging in ischemic heart disease. Am J Physiol Heart Circ Physiol. 2005;288: H2708–H2714. 82. Osman NF, Kerwin WS, McVeigh ER, Prince JL. Cardiac motion tracking using CINE harmonic phase (HARP) magnetic resonance imaging. Mag Reson Med. 1999;(42):1048–1069. 83. Osman NF, McVeigh, Prince JL. Imaging heart motion using harmonic phase MRI. IEEE Tran on Med Imag. 2000;19(3):186–202. 84. Pan L, Prince JL, Lima JA, Osman NF. Fast tracking of cardiac motion using 3D-HARP. IEEE Trans Biomed Eng. 2005;52(8):1425–1435. 85. Wilson RF, Laughlin DE, Ackell PH, et al. Transluminal, subselective measurement of coronary artery blood flow velocity and vasodilator reserve in man. Circulation. 1985;72(1):82–92. 86. Lette J, Tatum JL, Fraser S, et al. Safety of dipyridamole testing in 73,806 patients: the Multicenter Dipyridamole Safety Study. J Nucl Cardiol. 1995;2(1):3–17. 87. Pennell DJ, Underwood SR, Longmore DB. The detection of coronary artery disease by magnetic resonance imaging using intravenous dipyridamole. J Comput Assist Tomogr. 1990;14:167–170. 88. Pennell DJ, Underwood SR, Ell PJ, et al. Dipyridamole magnetic resonance imaging: a comparison with thallium-201 emission tomography. Br Heart J. 1990;64:362–369. 89. Pingitore A, Lombardi M, Scattini B, et al. Head to head comparison between perfusion and function during accelerated high-dose dipyridamole magnetic resonance stress for the detection of coronary artery disease. Am J Cardiol. 2008;101:8–14. 89a. Gebker R, Jahnke C, Manka R, et al. Additional value of myocardial perfusion imaging during dobutamine stress magnetic resonance for the assessment of coronary artery disease. Circ Cardiovasc Imaging. 2008;1:122–130. 90. Roest AA, Kunz P, Lamb HJ, Helbing WA, van der Wall EE, de Roos A. Biventricular response to supine physical exercise in young adults assessed with ultrafast magnetic resonance imaging. Am J Cardiol. 2001;87:601–605. 91. Rerkpattanapipat P, Darty SN, Hundley WG, et al. Feasibility to detect severe coronary artery stenoses with upright treadmill exercise magnetic resonance imaging. Am J Cardiol. 2003;92 (5):603–606. 91a. Jekic M, Foster EL, Ballinger MR, Raman SV, Simonetti OP. Cardiac function and myocardial perfusion immediately following maximal treadmill exercise inside the MRI room. J Cardiovasc Magn Reson. 2008;10(1):3.

Stress Cardiovascular Magnetic Resonance: Myocardial Perfusion Juerg Schwitter

Considerable progress in myocardial perfusion cardiovascular magnetic resonance (CMR) was achieved in the past few years and could be documented in recent multicenter trials.1–3 From single-center trials and particularly from multicenter trials, the knowledge of how to perform and interpret perfusion CMR studies improved considerably. Nevertheless, many aspects of perfusion CMR are still under debate. Some of them were recognized only through the analysis of multicenter data, for example, the variability in data quality and the robustness of the technique when performed at different sites or the importance of the reader experience and data analysis algorithms on the test results. Still, it is clear that large efforts are needed to solve these issues in the near future, which is considering the epidemiologic and economic consequences of introducing new techniques for diagnosing coronary artery disease (CAD). The CMR technique offers an almost unbeatable versatility for exploring any aspect of cardiac disease. Therefore, the chapter starts by illustrating how perfusion CMR is embedded in the concepts of atherosclerosis development and the vulnerable plaque theory. It then addresses established and not yet established aspects of the perfusion CMR protocol, including the different approaches proposed for data interpretation and analysis. In this chapter, the current performance of clinical perfusion CMR is presented. A perspective for the future of perfusion CMR is provided at the end of the chapter.

THE RATIONALE FOR PERFUSION IMAGING Endothelial dysfunction is generally recognized as a crucial initial step in the development of atherosclerosis.4–6 Among many alterations in dysfunctional endothelium, the production of nitric oxide is reduced, which promotes the recruitment of leukocytes to the vessel intima7,8 through enhanced expression of endothelial adhesion molecules.7 In the vessel wall, mononuclear leukocytes become foam cells by lipid accumulation and form reversible fatty streaks. In addition recruitment of smooth muscle cells and their production of extracellular matrix induce the formation of fibrous lesions. During the course of atherogenesis, inflammatory triggers and other factors can affect the balance between the production of matrix-degrading enzymes9 and collagen production by macrophages,10 which can cause fibrous cap destabilization. Subsequent repetitive plaque ruptures (and healing),

even when clinically silent, may cause a stepwise progression of atherosclerotic lesions.11,12 A final rupture then of larger plaques can cause the ultimate thrombus formation and acute vessel occlusion.13 In line with this concept, invasive studies conducted mainly in the preinterventional era demonstrated a positive correlation between stenosis degree in coronary angiography and risk for plaque rupture and acute myocardial infarction.12,14–17 Figure 16-1 demonstrates in approximately 17,000 patients an exponential increase in the likelihood of plaque rupture and coronary occlusions with an increasing degree of coronary artery stenoses. Accordingly, vulnerable plaques were defined by four major criteria:18 (1) thin cap with large lipid core, (2) stenosis degree greater than 90%, (3) endothelial denudation with superficial platelet aggregation, and (4) fissured plaque. Given these events in the development of CAD, a major goal of noninvasive imaging is to identify vulnerable plaques in the coronary circulation. Ideally, identification of these plaques would include determination of both plaque mass (to some degree related to stenosis severity) and plaque composition. While direct visualization of plaques in the coronary circulation and assessment of plaque components is still in an early clinical phase, assessment of stenosis degree by means of perfusion CMR has been evaluated in large clinical single-center and even multicenter trials.

THE PERFUSION CARDIOVASCULAR MAGNETIC RESONANCE PROTOCOL Stress-Only Versus Stress-Rest Examination for the Detection of Hemodynamically Significant Lesions In fundamental canine experiments, a reduction of coronary artery cross-sectional area was related to resting and hyperemic flow in the vessel.19 To assess the hemodynamic consequences of a coronary artery stenosis, the maximum hyperemic coronary blood flow during intracoronary adenosine infusion was measured. To compare maximum hyperemic flow in different coronary vessels, hyperemic flow was normalized with resting flow to correct for differences Cardiovascular Magnetic Resonance 213

16 STRESS CARDIOVASCULAR MAGNETIC RESONANCE: MYOCARDIAL PERFUSION

CHAPTER 16

Pooled data n = 16,884 y = 0.17 e0.55x

25

Occlusion rate (%)

ISCHEMIC HEART DISEASE

30

20

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5

0 0

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40

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Diameter reduction (%) Figure 16-1 The relationship between the degree of diameter stenosis in invasive coronary angiography and occlusion rate per year in 4554 patients (light15,17 and dark 12 brown bars) as well as the relationship between pathologic/normal SPECT scans (brown bars) and complication rate per year (deaths or nonfatal myocardial infarctions) in 12,360 patients.112 For this bar graph, the finding of ischemia in SPECT scans was assigned to the stenosis class of 50% diameter reduction, the normal scans to <50% diameter reduction. For the coronary angiography studies, occlusions at sites of stenoses were considered. For the SPECT studies, the occlusions cannot be related to the location of stenoses; therefore, the SPECT data are indirect support for the idea that low- and highgrade stenoses are associated with low and high risk for plaque rupture, respectively. Overall, these compiled data from almost 17,000 patients12,15,17,112 demonstrate a considerable (i.e. exponential) increase in risk of complications with increasing stenosis severity. This graph also illustrates that the risk of a plaque rupture is not zero in mild stenoses (in line with smaller retrospective studies). Nevertheless, diagnostics in this low-risk subpopulation would require a very high “number to test” to identify patients with significant stenoses, that is, patients at increased risk.

in mass of the vessel-dependent myocardium. Thus, the assessment of hemodynamic significance of a coronary artery stenosis was achieved by dividing hyperemic blood flow by resting blood flow, that is, by calculating the coronary flow reserve (CFR). However, the influence of resting hemodynamics on resting flow and thus on CFR was recognized. By altering resting heart rate and blood pressure, CFR ranged from 2 to 12 in their experiments.19 To solve this problem, the relative CFR was proposed as hyperemic blood flow with stenosis divided by hyperemic blood flow without stenosis. This relative CFR eliminates the influence of resting hemodynamics on CFR, since hyperemic stenosis flow is normalized by maximum adenosine-induced hyperemic flow that is uncoupled from myocardial oxygen demand. Consequently, this relative CFR remained stable during varying resting hemodynamics (standard deviation of 17% for relative CFR versus 45% for the absolute CFR). These findings are in line with a recent CMR study in humans, demonstrating the important effect of resting hemodynamics such as heart rate, contractility, and loading conditions on resting myocardial blood flow and hence, on absolute 214 Cardiovascular Magnetic Resonance

CFR.20 Further evidence that absolute CFR is affected by resting hemodynamics is provided by positron emission tomography (PET) studies on myocardial perfusion performed in patients with CAD. With this technique, a correlation of CFR with percent area stenosis was demonstrated; however, there was a tendency toward even better correlations for hyperemic myocardial flow alone.21–24 These data suggest that relating hyperemic flow supplied by stenosed coronary arteries to hyperemic flow supplied by nonstenosed coronary arteries (e.g., represented in a normal database) is advantageous, since it eliminates the influence of resting hemodynamics as encountered with the CFR approach. This stress-only approach, which compares hyperemic stenosis flow with hyperemic flow of a normal database, has been successfully applied in several clinical CMR studies.1,25 While the CFR measurements are affected by resting hemodynamics, two additional problems should be kept in mind: (1) Matching myocardial regions, for example, the subendocardial layer, for both rest and hyperemic conditions may be difficult, since geometry of the heart is changing with changing heart rate and loading, and (2) to obtain accurate results for the CFR calculation, the technique must guarantee a linear relationship between CMRderived perfusion parameters and true flow over a wide range of flow conditions covering resting and hyperemic flow. A synopsis on advantages and disadvantages on the stress-only and stress-rest protocols is shown in Table 16-1.

Options for Inducing Stress in Cardiac Perfusion Studies From the considerations mentioned above, it appears mandatory to apply some kind of stress to the myocardium to assess stenoses that reduce maximum blood flow but not resting flow. If oxygen demand of the myocardium is increased either by positive inotropic agents such as dobutamine or by direct physical stress, the supplydemand imbalance induces a vasodilation and a hyperemic reaction in the myocardium to meet demand (Fig. 16-2). However, in myocardial regions supplied by a stenosed coronary artery, increased oxygen demand is not met by supply, causing ischemia and consequently regional hypokinesia or akinesia. In this latter situation, assessments of both myocardial (contractile) dysfunction and reduced hyperemic reaction allow detection of a hemodynamically significant stenosis. Alternatively, hyperemia can be induced directly by administration of vasodilators such as dipyridamole or adenosine. This strategy detects coronary artery stenoses by assessment of a compromised hyperemic flow, while ischemia (O2 supply-demand imbalance) is very rarely induced by vasodilators. To cause a steal effect by vasodilators, the stenosis has to be severe, that is, more than 90% as demonstrated in experimental studies.26 Consequently, when contractile function was assessed during vasodilator-induced stress, the sensitivity for detection of CAD in clinical studies was very low27 (Fig. 16-3). Moreover, in comparison with positive inotropic approaches, the vasodilator approach as used for perfusion tests appears advantageous by reducing ischemia-associated side effects such as angina pectoris and arrhythmias.

Stress-Only Perfusion Protocol

Stress-Rest Perfusion Protocol

Acquisitions

Hyperemic data only

Hyperemic and rest data waiting time between acquisitions

Oxygen demand–perfusion relation For resting perfusion

Not applicable

For hyperemic perfusion Analysis

Oxygen demand-supply uncoupled Only hyperemic data

“Ideal” MR parameter–perfusion relation

Linear for low perfusion range (no perfusion increase during hyperemia in myocardium supplied by signficantly stenosed vessel) With normal data base for regional hyperemic perfusion With late enhancement acquisition advantage: acquisition window: minutes

Determined by “confounding factors” Heart rate Contractility Loading condition Oxygen demand-supply uncoupled Two analyses (stress and rest) matching anatomy for rest/stress condition “late enhancement effect” for second acquisition Linear for both resting and hyperemic perfusion, since CFR is calculated as hyperemic data/rest data With normal data base for regional CFR

Normal/abnormal discrimination Viability assessment Assessment of perfusion and viability CMR: Perfusion Viability SPECT: Perfusion Viability PET: Perfusion Viability

Hyperemic perfusion Late enhancement Hyperemic perfusion (stress tracer injection) Rest-tracer redistribution (rest tracer injection) Perfusion (stress flow tracer injection) Rest FDG uptake (rest metabolic tracer injection)

Low resting perfusion indicates scar disadvantage: acquisition window: seconds (first pass) CFR Resting perfusion — — CFR —

CFR, coronary flow reserve; stress indicates hyperemic condition induced by adenosine or dipyridamole, SPECT, single-photon emission computed tomography, FDG, fluoro-deoxyglucose; PET, positron emission tomography.

Physical stress, dobutamine O2 demand ↑

Vasodilator

Vasodilation

Vasodilation Stenosis

Hyperemia ↓

Hypoperfusion

O2 supply ≠ O2 demand “Ischemia”

Dysfunction Figure 16-2 The two major options to provoke perfusion deficits in the presence of a stenosis. Ischemia can be detected by wall motion analysis as well as by perfusion analysis. A compromised hyperemic response is detected by perfusion analysis but not by wall motion analysis unless the vasodilator induces ischemia by a steal effect.

ENDOGENOUS VERSUS EXOGENOUS CONTRAST MEDIA Endogenous Contrast Media Arterial spin labeling exploits the fact that unsaturated protons entering saturated tissue shorten the tissue T1.28–31

Using appropriate electrocardiogram (ECG)-triggered pulse sequences, T1 measurements are performed after global and slice-selective spin preparation. Absolute tissue perfusion is then calculated by assuming a two-compartment model. It should be kept in mind that this approach assumes that the direction of flowing blood in intramural vessels is orthogonal to the slice orientation, which is not the case for all myocardial layers. Furthermore, fiber orientations change during contraction and relaxation, making this approach problematic in the beating heart. In blood oxygen level–dependent (BOLD) imaging, deoxygenated and thus paramagnetic hemoglobin shortens T2 relaxation time and therefore can be used as an endogenous contrast medium (while oxygenated hemoglobin is slightly diamagnetic, causing less T2 shortening). A T2- or T2*-weighted pulse sequence allows for an estimation of an increased content of oxygenated blood. During pharmacologically induced hyperemia, oxygen content will increase in well-perfused myocardium but not in myocardium supplied by a stenosed vessel (which is associated with a higher O2 extraction). However, signal differences in normally perfused regions versus hyperemic regions (with a fourfold increase in flow) were reported to be as low as 32% in an experimental study.32 These BOLD data correlated closely with microsphere data, but the slope of the correlation was as low as 0.08. This limitation in signal difference might be problematic, since contrast media first pass studies suggested several hundred percent of signal change being required for a reliable stenosis detection.33 In another study applying BOLD, △r2 measurements yielded an adequate slope of 0.94.34 But again, the sensitivity to absolute flow changes was low, Cardiovascular Magnetic Resonance 215

16 STRESS CARDIOVASCULAR MAGNETIC RESONANCE: MYOCARDIAL PERFUSION

Table 16-1 Assessment of Coronary Artery Disease: Characteristics of Stress-Only Protocol and Stress-Rest Protocol

100

100

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Specificity (%)

Sensitivity (%)

ISCHEMIC HEART DISEASE

Detection of coronary artery disease: >50% diameter stenosis

*

40 20

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0 Adoecho

Adomibi

DipCMR

Dobecho

DobCMR

Dobmibi

Adoecho

Adomibi

DipCMR

Dobecho

DobCMR

Dobmibi

Figure 16-3 On the basis of the concept given in Figure 16-2, various modalities such as echocardiography, scintigraphy, and CMR can be applied to detect ischemia27,113 and/or hypoperfusion,25,27 provoked by either a physical stress/adrenergic drug or a vasodilator, respectively. The various modalities performed similarly in these single-center studies. Only detection of wall motion abnormalities in the case of compromised hyperemic response is limited by a low sensitivity in line with experimental data.26 Ado, adenosine; Dip, dipyramide; Dob, dobutamine; echo, echocardiography.

since a 100% increase in flow yielded a signal increase of only 5% (that is, a change in myocardial r2 of 0.94/sec) in that study.34 The BOLD approach is also sensitive to magnetic field inhomogeneities, typically occurring in the posterior wall close the coronary sinus draining deoxygenated blood.35 Considering these aspects, the robustness of the method has not yet been fully explored.36,37

in perfusion CMR preserves its high spatial resolution of data acquisition: on the order of 1 to 3 mm 1 to 3 mm. This is not the case for scintigraphic techniques with acquisition windows of several minutes, which preclude breath holding for elimination of respiratory motion, and ECG triggering during SPECT studies requires higher tracer amounts to improve counts statistics.

Extravascular Contrast Media

Exogenous Contrast Media for Magnetic Resonance Perfusion Imaging These contrast media are typically injected into a peripheral vein, and the signal change in the myocardium occurring during the first pass of the contrast media is measured by a fast MR acquisition. All contrast media first pass techniques either under development or in clinical routine are designed to meet the following requirements: (1) provide high spatial resolution to permit detection of small subendocardial perfusion deficits, (2) provide adequate cardiac coverage to allow for assessment of the extent of perfusion deficits, (3) feature high contrast media sensitivity to generate optimum contrast between normally and abnormally perfused myocardium during contrast media first pass, and (4) allow acquisition of perfusion data every one to two heartbeats to yield signal intensity–time curves of adequate temporal resolution that allow for extraction of various perfusion parameters (see below). To reach these goals, high-speed data acquisition and time-efficient magnetization schemes are most important. Since the first pass of contrast media during hyperemia lasts only 5 to 10 seconds, breathing motion is minimized by a breath hold maneuver, while cardiac motion is eliminated by ECG triggering. This control of cardiac and breathing motion 216 Cardiovascular Magnetic Resonance

For perfusion CMR techniques, the relationship between myocardial contrast media concentration and myocardial signal depends on a variety of factors. Normal perfusion can cause a signal increase during first pass of a gadolinium (Gd) chelate when combined with a T1-weighted pulse sequence, while a T2-weighted sequence with a higher dose of a Gd chelate can even cause a signal drop during first pass.38,39 This is fundamentally different from ischemia detection based on the assessment of wall motion, where new onset of dysfunction unambiguously indicates the presence of ischemia.40 Today, extravascular Gd chelates are most commonly used for MR first pass studies in combination with heavily T1-weighted pulse sequences. These contrast media are excluded from the intracellular compartment, that is, from viable cells with intact cell membranes; therefore, a perfusion deficit during the first pass reflects either hypoperfused viable myocardium (which would become ischemic during inotropic stress) and/or scar tissue (with severe reduction of perfusion even at rest). To differentiate hypoperfused tissue further, it is recommended to inject another dose of contrast media and to wait for the establishment of a dynamic equilibrium of contrast media concentrations in the various compartments (blood, viable myocardium, scar tissue), where contrast media concentration is governed by distribution volumes (and no longer by perfusion).41 During this condition, which typically occurs within 10 to

Intravascular Contrast Media For albumin-targeted MS-32544 (Ablavar gadofosurset trisodium, Lantheus Medical, Billerica, MA) and poly-lysineGd compounds,45,46 differences between normally perfused and stenosis-dependent hypoperfused myocardium were reported in animals. However, to our knowledge, these intravascular Gd-based contrast media have not yet been tested in humans. As an alternative, intravascular superparamagnetic iron oxide nanoparticles with a starch

coating were used for perfusion studies in humans.47 In combination with a T2-weighted turbo spin echo sequence, a signal drop in normal myocardium of 59% was observed (this contrast medium is no longer under investigation, owing to iron accumulation in the liver). For T2*enhancing contrast media, not only the concentration of contrast media in the voxel but also its intravoxel distribution (homogeneous versus inhomogeneous) determines its T2*-shortening effect, rendering such a T2* approach susceptible to vessel architecture (vessel orientation, intervessel distance). Since T2* approaches are affected by these geometry factors (which are not known a priori), measurements of tissue contrast media concentrations by T1 techniques are considered superior.

Hyperpolarized Contrast Media Conventional Gd chelates modulate the MR signal by accelerating the relaxation rates of surrounding water protons. However, at a field strength of 1.5 T, the polarization level of the spin population of the MR-active nuclei at the thermal equilibrium is low (several parts per million), and only the vast abundance of water molecules in the human body enables generation of a measurable signal. Since the polarization level of water protons increases with magnetic flux density, a higher signal is achieved at a higher field strength. Alternatively, however, the polarization level of the spin population of specific nuclei (such as liquid 13C in various compounds) could be increased by a factor of up to 100,000 (compared with polarization of water protons at the thermal equilibrium) by dynamic nuclear polarization48 or parahydrogen-induced polarization.49 With these hyperpolarized contrast media, signal is

Subendocardial stress data

ROC analysis

400

1

0.75

200 <0.002 <0.05

Sensitivity

SI change (%)

300

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D1: 0.56 + 0.12 D3: 0.86 + 0.08

0.25 0.15 mmol/kg 0.10 mmol/kg 0.05 mmol/kg

P <0.05 D1 vs. D3

0

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20

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Figure 16-4 A, Increasing doses of the extravascular contrast medium Gd-DTPA from 0.05 to 0.10 and 0.15 mmol/kg administered intravenously induce an increasing myocardial signal in the subendocardium (inner half of myocardial wall) during first pass.1 B, Even a dose as high as 0.15 mmol/kg did not cause measurable susceptibility-induced signal loss in the subendocardial layer using a hybrid echo planar pulse sequence (echo time: 1.3 to 2.2 msec). Accordingly, the highest dose detected stenoses with 50% diameter reduction significantly better than the 0.05 mmol/kg dose. SI, signal intensity. Cardiovascular Magnetic Resonance 217

16 STRESS CARDIOVASCULAR MAGNETIC RESONANCE: MYOCARDIAL PERFUSION

20 minutes after contrast media injection in humans, late enhancement imaging (with the inversion time set to null normal myocardium) is ideal to discriminate hypoperfused but viable myocardium as dark tissue from scar, which appears bright.42,43 For viability assessment, scintigraphic techniques also exploit the equilibrium distribution of tracers, which is observed after rest injection or rest reinjections. However, the radioactive tracers are not taken up by scar tissue, which consequently appears as a cold spot, while viable tissue appears as a hot spot (see Table 16-1). Several multicenter studies were performed for the assessment of the optimal contrast media dose for perfusion imaging. This is important, since higher contrast media doses can cause susceptibility artifacts at the subendocardium, where differences in contrast media concentrations between blood and myocardium are high during first pass. Strongly T1weighted sequences showed absence of a susceptibilityinduced signal drop in the subendocardial layer up to doses of 0.15 mmol/kg of an extravascular Gd chelate, which resulted in superior diagnostic performance of doses of 0.1 and 0.15 mmol/kg versus 0.05 mmol/kg for semiautomatic analysis of the upslope parameter1 (Fig. 16-4).

ISCHEMIC HEART DISEASE

received only from the 13C-nuclei; therefore, no signal from background tissue is obtained. This feature appears ideal for absolute quantification of perfusion, since signal is directly proportional to the amount of 13C molecules similar to radioactive tracers, where radioactivity is directly proportional to tracer concentration. Similarly to radioactive tracers, the signal from hyperpolarized 13C-nuclei also decays (with specific time constants depending on the type of the 13C-compound), since the spin population of the 13 C-molecules is far away from thermal equilibrium. In addition, repetitive radiofrequency pulsing further destroys longitudinal magnetization depending on the pulse sequence type and the imaging parameters. Johansson and colleagues showed that depolarization can be approximated by a monoexponential function with a time constant TD and successfully applied this concept for cerebral perfusion quantification.50 The extraordinary amount of signal available for a short time period also allows to study for example, the Krebs cycle in myocytes. It was hypothesized that a short ischemic insult would alter the energy metabolism in the cells for a time lasting considerably longer than the insult in order to restore normal function as early as possible.51 This concept of a “metabolic memory” was confirmed in a pig model of ischemia applying these hyperpolarized contrast media.51

THE PERFUSION CARDIOVASCULAR MAGNETIC RESONANCE EXAMINATION: WHAT IS ESTABLISHED AND WHAT IS NOT CMR Data Readout Echo planar readout strategies are well suited to speed up data acquisition,52,53 particularly when segmented to reduce the echo time and consequently to render the pulse sequence less prone to susceptibility artifacts (hybrid echo planar pulse sequences).25,33,44,54,55 With these accelerated pulse sequences, several k-lines are acquired following a single radiofrequency excitation, reducing the repetition time (TR) per k-line down to less than 2 msec and consequently reducing the total acquisition window per slice substantially. This enables the acquisition of a stack of slices every one to two heartbeats, allowing for true multislice acquisitions. In contrast, conventional fast gradient recalled echo (GRE) pulse sequences were often used in the past for perfusion imaging. With this technique, each k-line is preceded by a separate radiofrequency excitation, which resulted in a readout duration on the order of 350 to 450 msec depending on the number of phase-encoding steps and the duration of TR.40,56–62 Fast imaging inherently conflicts with signal-to-noise ratio (SNR). Readout strategies during steady-state conditions of magnetization [steady-state free precession (SSFP) sequences] appear promising, since they preserve magnetization and thus a high SNR. This SSFP technique has been employed successfully in animal models for monitoring 218 Cardiovascular Magnetic Resonance

contrast medium first pass through the myocardium.63 In a human volunteer study, however, ECG triggering was problematic as well as the presence of banding artifacts (probably due to longer readout periods or off resonance). Furthermore, the myocardial signal increase was rather low (approximately 40% of baseline signal).64 In a recent volunteer study,65 the contrast-to-noise ratio (CNR) was higher for a saturation recovery SSFP acquisition than for a saturation recovery hybrid echo planar acquisition. However, the delay time (between magnetization preparation and readout) for the hybrid echo planar pulse sequence was 30 msec versus 90 msec for the SSFP acquisition, which is known to reduce CNR (see below). The maximum achievable CNR is crucial for reliable CAD detection and was more important than cardiac coverage.33 Nevertheless, improving cardiac coverage without the need to reduce spatial and/or temporal resolution would be beneficial, since the extent of CAD correlates with outcome. To this end, parallel imaging approaches were tested to increase coverage without a compromise in spatial and temporal resolution. In a recent study in human volunteers, the loss in SNR given by g * √R (g being the so-called geometry factor and R being the accelerating factor) was compensated for by a longer TR (due to implementation of time-adaptive sensitivity encoding [TSENSE], a modification of sensitivity encoding [SENSE]66) and increasing the readout flip angle from 20 to 30 . This accelerated hybrid echo planar saturation recovery technique yielded twice as many slices as were obtained with the nonparallel approach, while SNR improved by approximately 20%.67 In a recent study, we investigated the additional benefit of spatiotemporal correlations of perfusion data over parallel imaging alone.68 By a 5 kt-SENSE approach, relative peak signal enhancement (150% of precontrast) and data quality were similar to those in a parallel (SENSE) approach only, while spatial resolution was increased from 2.6 2.6 mm2 to 1.5 1.5 mm2 (3.3-fold), and SNR further increased 1.4-fold in comparison to parallel imaging (with identical cardiac coverage).

Magnetization Preparation A magnetization preparation by a 90 pulse is now generally accepted as the most efficient way to achieve T1 weighting. It shortens the delay time to 100 to 150 msec25,33 and renders the sequence heart rate independent. The relationship between contrast media concentration and MR signal is also dependent on the flip angle of the readout and preparation pulses as shown in Figure 16-5. A delay time (TD) of 120 msec also places the data collection into phases of minimal cardiac motion, that is, into late systole and middiastole. Since the length of systole varies relatively little with changing heart rate (as occurs during pharmacologically induced hyperemia), these delay times are robust to place data collection into stable heart phases irrespective of heart rate. To further accelerate data collection, it was proposed to play out one single non-slice-selective 90o saturation pulse and to acquire the entire stack of slices thereafter.69 With this scheme, however, the delay time varies from slice to slice; consequently, contrast media sensitivity becomes dependent on slice position and acquisition order, which is

Field Strength: 1.5 T Versus 3.0 T

16

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SNR

10

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1 0

0.2

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GdDTPA concentration (mmol/L) 90º/50º/200 msec 90º/50º/150 msec – 4 slices 90º/50º/100 msec 90º/50º/50 msec 90º/30º/150 msec 90º/10º/150 msec 90º/10º/10 msec 60º/10º/10 msec – 6 slices Figure 16-5 These phantom measurements demonstrate that optimization of a hybrid echo planar pulse sequence54 with respect to the delay time between magnetization preparation and readout, readout flip angle, and preparation flip angle can improve signal response of Gd-DTPA-doped phantoms from 80% (of nondoped phantoms) to 250%, which ultimately increased diagnostic performance in patients.33

likely to affect data analysis. In a modified version of a saturation recovery preparation, the entire myocardium is prepared by a 90o saturation pulse except the slice that is imaged immediately after preparation.70 With this scheme, the time for readout of slicen equals the preparation time for slicenþ1 and so on. This approach allows acquisitions of up to seven slices every two heartbeats (at a heart rate up to 115 beats per minute). Since a high contrast between abnormally and normally perfused myocardium during the peak effect of the bolus is crucial for reliable detection of perfusion differences, preparation of the myocardium by a 180o inversion pulse (IR technique) was performed frequently in the past. Thereby, maximum contrast is obtained by nulling the signal of precontrast myocardium by applying inversion recovery times on the order of 300 to 400 msec.40,56–62,71 However, this IR technique results in a total data acquisition window of 650 to 750 msec per slice when combined with a fast GRE readout and precludes the acquisition of multiple slices during one to two heartbeats. To circumvent this problem, partial preparation flip angles, for example, of 45o to 60o, were suggested, which allowed for shorter delay times.44,54 However, this limited the dynamic range of signal response33 and is no longer recommended (see Fig. 16-5).

Systematic data comparing the diagnostic performance of perfusion CMR at 1.5 T and 3.0 T are still sparse. Increasing the field strength from 1.5 T to 3.0 T increases the available magnetization, as was demonstrated in a study in volunteers, in which myocardial signal during the first pass (at 0.1 mmol/kg Gd-DTPA IV) was 2.6 times higher at 3.0 T than at 1.5 T (relative to baseline signal).72 However, this increase was observed in the anterior wall, while signal increased by only 1.7 times in the posterior wall. This inhomogeneous signal increase might be disadvantageous, and larger clinical trials would be required to demonstrate a potential advantage of perfusion imaging at 3.0 T. In a recent study, we implemented a novel spatiotemporal correlation approach for perfusion data (5 ktSENSE) on a 1.5-T and a 3.0-T machine.73 In patients with suspected CAD, the 1.5-T and 3.0-T machines yielded spatial resolutions of 1.5 1.5 mm2 and 1.3 1.3 mm2 (with otherwise identical temporal resolution and coverage). The areas under the receiver operator curve (ROC) curve for CAD detection (defined as >50% diameter reduction on coronary angiography) were not different with 0.80 and 0.89, respectively (p ¼ 0.21), despite slightly higher SNR on 3.0 T, indicating that the high diagnostic performance of perfusion CMR at 1.5 T is hard to improve.73

ANALYSIS OF PERFUSION DATA With increasing experience, an observer will be able to visually differentiate hypoperfused regions from artifacts, both typically causing lower signal during first pass conditions.2,62,71,74 However, the advantage of experience is traded for subjectivity, that is, reduced interreader reproducibility. Therefore, a computer-assisted or automatic analysis of perfusion data would be highly desirable and could potentially render CMR perfusion analyses fully reproducible. Since many different analysis procedures are currently in use, some common definitions for analysis characteristics are proposed in Figure 16-6.

Visual Assessment While visual reading can provide accurate diagnoses, adequate reproducibility should rely not on individual experience but on standardized criteria that in particular would differentiate signal loss due to hypoperfusion from artifacts. This indicates that reading criteria should also assess image quality, which in turn depends on the pulse sequence, the contrast media type and dose, and, importantly, the experience and care in data acquisition. In a recent multicenter trial performed with a single MR machine type and an identical pulse sequence, experienced readers applying predefined diagnostic criteria yielded k-values as low as 0.302. Signal loss in the subendocardium, which is most sensitive for perfusion abnormalities, can be caused by susceptibility artifacts occurring during the time window when contrast media concentration differs most between the blood pool Cardiovascular Magnetic Resonance 219

16 STRESS CARDIOVASCULAR MAGNETIC RESONANCE: MYOCARDIAL PERFUSION

Effect of imaging parameters: αprep, delay time, αread-out

ISCHEMIC HEART DISEASE

Perfusion data

Visual analysis

Quantitative analysis

Perfusion-related parameters e.g., upslope, peak SI, etc.

Manual

Semiautomatic

Automatic

Absolute perfusion unit: mL/min/g

Manual

Semiautomatic

Automatic

Figure 16-6 A scheme for better definition of possible analysis strategies. In this scheme, “quantitative” results are given in numbers, which allow for an objective comparison of studies both longitudinally (e.g., monitoring disease activity) as well as from patient to patient (e.g., patient data versus normal database). Such quantitative results are obtained either manually (and thus are associated with some observer dependence), in a semiautomatic fashion, that is, with some observer interaction, or automatic, thus eliminating any observer interaction with the data and thereby completely eliminating analysis variability.

and the myocardial tissue. Thus, long persistence of signal loss during first pass is suggestive of the presence of hypoperfusion in the subendocardium. However, these timing criteria also depend on the dose of contrast media injected, the injection rate and site, and the dispersion of the bolus within the pulmonary circulation. A signal reduction consistent with a territory supplied by an individual coronary artery is also supporting the diagnosis of CAD, while artifacts may distribute along phase or frequency directions or may occur in regions with inhomogeneous magnetic field characteristics as observed in the inferior wall (where air is extending between diaphragm and inferior wall of the left ventricle).35

Quantitative Approach Dedicated algorithms for perfusion data analysis would allow comparison of signal responses in an individual patient with a normal database, rendering the technique less observer dependent. Since myocardial signal response is strongly dependent on imaging parameters, a normal database should be updated in case the pulse sequence and/or the imaging parameters would be modified (e.g., by a hardware or software upgrade). On the basis of such a normal database, perfusion studies can be compared between patients, and even more important, the perfusion situation, that is, the activity or progression of CAD, can be monitored by serial perfusion CMR studies in the same patient. To allow for better comparisons between various analysis strategies, it appears reasonable to clarify some confusion that may exist with expressions such as semiquantitative and semiautomatic. In Figure 16-6 quantitative relates to the analysis output obtained as numbers (which may be related to perfusion or may represent perfusion itself). These aspects should be separated from the issue of how the data are extracted from the imaging set, which may occur manually (with considerable observer interaction with the data and corresponding time requirements), 220 Cardiovascular Magnetic Resonance

semiautomatically (with some observer interaction), or fully automatically. A fully automatic approach is time-saving (no labor needed for analysis) and eliminates any observer variability, while it is anticipated that high-quality data would be required for an automatic approach. A quantitative approach is also ideal to generate ROC, which are the adequate means for assessment of test performance. Once a ROC curve has been determined (for a given acquisition and analysis protocol), the optimum cost-effectiveness of the protocol can be calculated (optimum point on the ROC), since cost-effectiveness changes with changing portions of false and true negatives and positives.

Quantitative Approach: Perfusion-Related Parameters First pass perfusion studies are typically performed during breath holding, and therefore any type of quantitative analysis should start with a registration of the first pass data over time, that is, motion in the data caused by breathing and/or diaphragmatic drift should be eliminated either by a manual procedure or by (semi)-automatic algorithms.75–77 To improve SNR cardiac perfusion studies are performed with phased-array coils. Therefore, analyses of signal intensity–time curves have to correct for inhomogeneous coil sensitivities. For this purpose, it has been suggested to subtract precontrast signal from the first pass signal intensities.69 However, signal reception by a phased array coil does not cause a constant offset of signal over the field of view; therefore, a division of first pass signal by precontrast signal for correction is required.1,25 From the resulting signal intensity–time curves calculated for various transmural or subendocardial segments covering ideally the entire left ventricular myocardium, an abundant number of parameters can then be extracted. The applicability was demonstrated for the peak signal intensity,40,44,56,78 signal change over time (upslope),25,33,59–61,78–81 arrival time, time to peak signal, mean transit time,78,79,82,83 area under the signal intensity–time curve,63 and others. For the upslope of the signal intensity–time curve, a relatively close correlation with perfusion data is reported in both animal79,84 and human studies,25 at least for the lowflow range. Its robustness could also be demonstrated in a multicenter trial.1 This upslope parameter is relatively insensitive to contrast media recirculation, since it uses the initial portion of the signal intensity–time curve only, which also reduces its sensitivity for motion (most patients are able hold their breath for this short time period). The upslope parameter is also proposed to correct for differences in arterial input by dividing myocardial upslope data by the blood pool signal upslope.25,33,60,61,80,81 This approach is suboptimal for input correction, since the upslope-contrast media concentration relationship is flat at higher contrast media concentrations, which occur in the blood pool during first pass and which are even higher during hyperemic condition. An experimental study demonstrated linearity between the upslope parameter and perfusion measured by microspheres for low perfusion values only (below approximately 1.5 mL/ min/g) (Fig. 16-7A).84 A similar limitation for upslope estimates was described for humans using PET perfusion measurements as the standard of reference (see Fig. 16-7B).25

2.25 Slopemeasured/slopethreshold

Myocardial/LV upslope ratio (normalized)

6 5 4 3 2 1

2 1.75 1.5 1.25 1 0.75

y = 1.122 log (x) + 1.124 r = 0.69 p <0.0001

0.5

0

0.25 0

1

2

3

4

5

6

0

0.5

1

A

1.5

2

2.5

3

3.5

4

4.5

MBFPET (mL/min/g)

Microsphere blood flow (mL/min/g)

B

Stenosis: 50–69% Stenosis: 70–99% Occlusion Volunteers (mean + SD) Patients without CAD Remote region of CAD patients (supplied by nonstenosed vessel)

Figure 16-7 A, The nonlinear relationship between the myocardial upslope parameter (“corrected” by the signal upslope in the left ventricular [LV] blood pool) and the microsphere blood flow measurements in canine experiments.84 Linearity exists for low-flow situations only (up to approximately 1.5 mL/min/g). B, Similar nonlinearity between the myocardial upslope parameter (normalized by the threshold upslope value) and the myocardial blood flow measured by positron emission tomography in patients with and without coronary artery stenoses. Normalization of the myocardial upslope by the threshold upslope value (which discriminates between stenotic and nonstenotic flow) results in a ratio of 1 for flow in nonstenosed vessels.25 MBF, mean blood flow. Source: Reproduced with kind permission from the Radiological Society of North America84 and from Lippincott Williams & Wilkins.25

Quantitative Approach: Absolute Tissue Perfusion For blood pool contrast media, which do not mix homogeneously in the tissue compartment but are restricted to the intravascular compartment, the so-called bolus tracking approach can be applied, which is based on the central volume principle given by VB ¼ F MTT

(1)

where MTT is the mean transit time (expected distribution of transit times for the blood through the tissue volume) and VB is blood volume. Since the conventional extravascular Gd chelates do not cross the intact blood-brain barrier, this model is generally used for cerebral perfusion measurements. Also, hyperpolarized 13C-contrast media (see below) can act as intravascular tracers;50 hence, this model could be applied for cardiac studies using these contrast media. In the myocardial tissue, however, the conventional Gd chelates behave as extravascular contrast media. Therefore, the contrast media first pass techniques with residue detection will be discussed next. It is assumed that contrast medium diluted in blood enters the tissue via the artery, transits the capillaries, and leaves the tissue through the venous system. Then the Fick principle relates the contrast media concentration in the tissue to the arterial input and venous

outflow by a convolution integral. This model takes both contrast media inflow and outflow into account and assumes a freely diffusible contrast medium that is homogeneously mixed within a single tissue compartment.85 It is given by Ðt CT ðtÞ ¼ F CA ðtÞ eF=lðttÞdt ¼ CA ðtÞ F eF=l : t (2) 0

where CT is the tissue contrast media concentration at time t, CA is the contrast media concentration in blood, F is tissue perfusion (in milliters per minute pre gram), l is the tissue-blood partition coefficient, and denotes convolution. If CT and CA are known, a two-parameter fit yields F and l (by application of a three-parameter fit, blood volume VB can be incorporated into the model as well). Several deconvolution procedures have been used to obtain F.84,86–89 For non-freely diffusible contrast media, the extraction fraction E must also be taken into account,90–92 where E is related to F by Ki ¼ E F ¼ Fð1 ePS=F Þ

(3)

where PS is the permeability-surface-area product (in milliters per gram per second) and Ki is the unidirectional influx constant (in milliters per gram per second). For all these models, knowledge about the arterial input function, that is, the measurement of the arterial contrast media concentration over time, is mandatory. Consequently, the signal intensity –contrast media concentration relationship over the full range of contrast media concentrations occurring in the blood pool during first pass conditions must be known.45,87,93,94 Cardiovascular Magnetic Resonance 221

16 STRESS CARDIOVASCULAR MAGNETIC RESONANCE: MYOCARDIAL PERFUSION

2.5

ISCHEMIC HEART DISEASE

While higher contrast media doses yield an appropriate CNR level for the signal response in the myocardium, such high doses are likely to cause a clipping of the signal intensity–time curve in blood from which a conversion of signal intensities into contrast media concentration is problematic. To solve this problem, a dual-bolus approach was presented,84 in which a small contrast media bolus is injected first for determination of an arterial input function, followed by a larger contrast media bolus to achieve an adequate signal response in the myocardium. In dogs, this approach yielded absolute values of tissue perfusion in close agreement with microsphere measurements. Alternatively, Gatehouse and colleagues proposed a dual-T1-sensitivity method.95 This approach utilizes a single high-dose contrast media bolus injection, which provides a high signal response in the myocardium while preventing clipping of the blood pool signal at peak effect of the bolus. To measure blood signal (with low T1-sensitivity for very short T1), a short saturation recovery time is applied, while it is longer for measurement of myocardial signal (high T1-sensitivity for longer myocardial T1). Furthermore, the low-T1-sensitive blood pool measurements are performed at a low spatial resolution to accelerate acquisition. With this dual-T1-sensitivity single-bolus approach, CFR estimates in volunteers closely matched those from dual-bolus measurements. Once a reliable arterial input function has been obtained, diffusion of water molecules between the intravascular and extravascular compartments must be taken into account, since conventional Gd chelates exert their effect indirectly through water proton relaxation, which modifies the MR signal during first pass.96–99 An extravascular contrast medium mixed homogeneously in the extravascular space by diffusion would generate a maximum signal during first pass of 20% of fully relaxed magnetization (assuming an extracellular compartment of 20% and the absence of water exchange between compartments). For an intravascular contrast medium, the maximum achievable signal during first pass would further decline to approximately 10% of fully relaxed magnetization (assuming a blood volume in tissue of 10% and no water exchange between compartments). In rat experiments, extravascular and intravascular contrast media yielded approximately 70% and 50%, respectively, of fully relaxed magnetization during first pass, indicating that water exchange across both capillary vessel wall and cell membranes strongly affects signal.96 Water exchange conditions can be categorized into fast, intermediate, and slow. If tissue 1/T1 (¼ r1) increases linearly with intravascular r1 and hence myocardial signal response in the presence of a contrast media would be unaffected by water exchange, the water exchange regimen is called fast. Fast exchange conditions are met if the rate of water exchange between compartments is considerably higher than the difference in r1 between the compartments in the presence of contrast media. Larsson and colleagues100 performed simulations based on the tracer kinetic model to calculate the unidirectional influx constant Ki for diffusion of contrast media over the capillary membrane (in milliliters per minute per gram) using Equations 2 and 3. For low doses of extravascular contrast media (0.1 mmol/kg body weight to limit differences in r1 between compartments and assuming E greater 0.3), water exchange only minimally affected Ki, while signal modification was substantial for intravascular contrast media.100 These and other studies97–99 demonstrate that intravascular-extravascular water exchange is the rate222 Cardiovascular Magnetic Resonance

limiting step for signal behavior in tissue (slow exchange regime). Similar to radioactive tracers, E of extravascular contrast media varied with increasing myocardial perfusion.92 By assuming rather small changes in E for resting and hyperemic conditions and therefore using E as a constant, Ki at rest and during hyperemia was calculated, from which a perfusion index or CFR was derived.101,102 Besides precision, the variability of parameter estimates is another important aspect of perfusion measurements. On the basis of simulations for the determination of regional blood volume, intravascular contrast media103,104 were less sensitive for noise than were extravascular contrast media.90 Despite a large body of computer simulations90,103 and experimental data,45,87 the clinical value of absolute quantitative measures of perfusion for the detection of CAD in patients remains unproven. A 95% confidence interval of 45% to 82% to þ45% to þ82% for absolute perfusion quantification45,86 suggests that the sensitivity of these techniques to detect changes in perfusion might be limited. Also short periods of ischemia can cause endothelial leakage,105,106 which would complicate absolute perfusion assessment by intravascular contrast media in patients. For intravascular contrast media based on hyperpolarization (hyperpolarized 13 C-contrast media), water exchange does not affect the MR signal; consequently, perfusion quantification models could become easier. However, for these compounds, the time course of depolarization has to enter the formula for perfusion models (see below).50

CLINICAL PERFORMANCE OF PERFUSION CARDIOVASCULAR MAGNETIC RESONANCE Single-Center Studies: Visual Interpretation Visual discrimination of normal from severely abnormal findings is certainly feasible; however, it should be kept in mind that development of CAD is a progressive process that passes from mild to moderate to severe lesions (most likely by repetitive plaque ruptures). It might be in this intermediate range of lesions where visual assessment becomes difficult, and relating the patient data to a normal database might be particularly advantageous for this population advancing from nonobstructive to mildly obstructive CAD. Results of several studies applying a visual reading is given in Table 16-2. In a recent study, first pass perfusion CMR was compared with SPECT.107 A visual assessment of MR stress and rest dynamic data of 69 patients yielded a sensitivity and specificity of 90% and 85%, respectively, for the detection of stenoses with 70% or more diameter reduction on quantitative coronary angiography (QCA) at 0.075 mmol/kg of Gd-DTPA (injected at 4 mL/sec into a peripheral vein). CMR performed significantly better with areas under the ROC curves (AUC) of 0.89 to 0.91 for two observers versus AUC of 0.71 to 0.75 for SPECT.107

Table 16-2 Performance of Perfusion CMR: Detection of CAD Patients (N 7, 1.5-T Systems) Authors

N

Prep Slices*

CM

Dose

Rest

Stress

Analysis

Reference Standard

Sens

Spec

AUC

Single-Center Studies Hartnell et al.74 Eichenberger et al.60 Walsh et al.71 Bertschinger et al.33 Al-Saadi et al.61 Schwitter et al.25 Schwitter et al.25 Nagel et al.69 Ishida et al.107 Doyle et al.81 Plein et al.73

14 8 46 24 34 57 43 84 104 138 37

IR-single Multi IR-multi SR-multi IR-single SR-multi SR-multi SR-multi SR-multi SR-multi SR-multi

Gd-DTPA Gd-DOTA Gadoteridol Gd-DTPA-BMA Gd-DTPA Gd-DTPA-BMA Gd-DTPA-BMA Gd-DTPA Gd-DTPA Gd-BOPTA Gd-Butrol

0.04 0.05 0.10 0.10 0.025 0.10 0.10 0.025 0.075 0.04 0.10

+ + + + + + +

Dip Dip Dip Dip Dip Dip Dip Ado Ado Dip Ado

Visual Upslope Visual Upslopetrans Upslope CFR Upslopesubendo Upslopesubendo Upslope CFR Visual Upslope CFR Visual

QCA (70%) QCA (80%) 201 Th/99m Tc scinti QCA(50%) QCA (75%)} QCA(50%) 13 NH3-PET (CFR) QCA (75%)} QCA (70%) QCA (70%) QCA (50%)

83%{ 65%{ 89% 82% 90% 87% 91% 88% 90% 57% 90%

100%{ 76%{ 44% 73% 83% 85% 94% 90% 85% 85% 83%

0.76 0.91 0.93 0.93 0.90} 0.80

Multicenter Studies# Wolff et al.2 Giang et al.1 Schwitter et al.3

99 99 241

SR-multi SR-multi SR-multi

Gd-DTPA Gd-DTPA Gd-DTPA-BMA

0.05 0.10 0.10

+ +

Ado Ado Ado

Visual Upslopesubendo Visual

QCA (50%) QCA (50%) QCA (50%)

93% 91% 90%

75% 78% 73%

0.90 0.91 0.87

16 STRESS CARDIOVASCULAR MAGNETIC RESONANCE: MYOCARDIAL PERFUSION

Cardiovascular Magnetic Resonance 223

Ishida et al. Patients with history of myocardial infarction and/or resting wall motion abnormalities excluded. *Prep slices: magnetization preparation by saturation recovery (SR) and inversion recovery (IR), respectively. Single and multi, single-slice, and multislice acquisition, respectively. { Comparison on a regional basis (five segments/heart evaluated). { Comparison on a regional basis (48 segments/heart evaluated). } Vessel area reduction (in all other studies: vessel diameter reduction). } AUC (mean of two readers) for subset of 69 patients who also had a single-photon emission computed tomography (SPECT) study (AUC for SPECT: 0.73; p < 0.001 versus magnetic resonance). # For multicenter studies, diagnostic performance is given for best CM dose (for dose ranges tested, see text). AUC, area under the receiver-operator curve; CM, contrast media (all gadolinium-chelates); QCA, quantitative coronary angiography (diameter reduction); PET-CFR, coronary flow reserve calculated from positron emission tomography data (threshold: CFR < 1.7); scinti, scintigraphy; Upslopetrans, Upslopesubendo: Upslope data derived from full wall thickness and subendocardial layer, respectively; Upslope CFR: coronary flow reserve calculated from upslope data (ratio of slope: rest/hyperemia).

The rate of myocardial signal change (upslope) during contrast medium first pass is most widely used to quantitatively assess perfusion data.25,33,58–61,69,78,108 From a reststress protocol, a CFR index (slopehyperemia/sloperest) was derived and yielded a sensitivity and specificity of 90% and 83%, respectively, for the detection of stenoses with 75% or more area reduction on quantitative coronary angiography (QCA) (at a CFR threshold of 1.5).61 For an arterial input correction, the upslope of signal in the left ventricular blood pool was used; furthermore, the dose of the extravascular contrast media Gd-DTPA was kept at 0.025 mmol/kg body weight to minimize clipping of the blood pool signal intensity–time curve. The same CFR index derived from multislice perfusion data (with peripheral administration of the same contrast media dose) yielded a similar performance with a sensitivity and specificity of 88% and 90%, respectively, for the detection of stenoses with 75% or more area reduction on QCA.69 When CFR was determined by perfusion CMR in women, a specificity of 85% for the detection of CAD (defined as 70% or more diameter reduction on QCA) was achieved, while sensitivity was only 57%.81 This test performance at 0.04 mmol/kg of Gd-BOPTA compares well with the results of a recent dose-finding study for the dose of 0.05 mmol/kg of Gd-DTPA (see Fig. 16-4B).1 A stress-only protocol provided a comparable performance of the hyperemic upslope parameter when thresholds were derived from a normal database, resulting in a sensitivity and specificity of 87% and 85%, respectively, for the detection of stenoses of 50% or greater diameter reduction on QCA.25 This stress-only protocol avoids the need for matching myocardial regions for rest and stress condition and therefore offers easy analysis of the subendocardial layer, where perfusion abnormalities are most severe (Fig. 16-8).25,108 In a comparison of the subendocardial CMR upslope data with PET perfusion data, a sensitivity and specificity of 91% and 94%, respectively, were reported.25

Multicenter Studies Perfusion CMR is potentially the most accurate method currently available for the noninvasive assessment of myocardial perfusion in humans. However, if perfusion CMR should become the first-line method in clinical routine, its robustness, reliability, and safety have to be demonstrated across a large number of sites, which can be done in multicenter trials only. In a single-vendor multicenter trial, a stress-only protocol combined with a semiautomatic data analysis (upslope) performed best with contrast media doses as high as 0.10 to 0.15 mmol/kg Gd-DTPA, yielding an AUC of 0.91 0.07 and 0.86 0.08, respectively, with corresponding sensitivity/specificity of 91/78% and 94/71%, respectively.1 Thus, this multicenter trial confirmed the high diagnostic performance of the stress-only approach proposed in an earlier single-center study.25 However, in this multicenter study, 224 Cardiovascular Magnetic Resonance

1

0.8

Sensitivity

ISCHEMIC HEART DISEASE

CMR VS. QCA (Stenosis >50%)

Single Center Studies: Quantitative Semiautomatic Analysis

0.6

0.4 Upslope - subendocardial Upslope - transmural Upslope - transmural

0.2

0 0

0.2

0.4

0.6

0.8

1

1-specificity Figure 16-8 These receiver operating characteristic curves demonstrate superiority of subendocardial first pass perfusion data over transmural perfusion data for detection of stenoses with 50% diameter reduction on quantitative coronary angiography25,33 supporting the general knowledge of the subendocardium being most sensitive for hypoperfusion. In these two studies, perfusion data of patients were compared with a normal database. If susceptibility artifacts were to reduce signal response in the subendocardium, this approach would correct to a certain extent, since the normal database would also contain lower threshold values (provided that both patient data and normal data are acquired and analyzed with the identical protocol and software).

the high diagnostic performance was achieved in data with adequate quality only. For this purpose, a blinded quality reading was performed first, which eliminated 14% of all studies (Fig. 16-9A). An example from a study at the dose of 0.15 mmol/kg is shown in Figure 16-10. With adequate data quality, a k-value of 0.73 was obtained for interobserver agreement,1 which compares favorably with k-values of 0.80 for multicenter SPECT data.109 As expected with an increasing number of participating sites, homogeneity of data quality, and hence test performance, is likely to deteriorate as demonstrated in a comparison between single-center and multicenter perfusion CMR studies (see Fig. 16-9B). As was discussed above, perfusion CMR exhibits advantages over SPECT with respect to spatial resolution, motion suppression, and lack of attenuation artifacts. This suggests superiority of perfusion CMR over SPECT.107 In the MRIMPACT multicenter program (MR-Imaging for Myocardial Perfusion Assessment in Coronary Artery Disease Trial),3 241 patients underwent both invasive coronary angiography and SPECT within 4 weeks of a perfusion CMR study.3 The trial was conducted in 18 centers in the United States and Europe and is the largest multicenter SPECT trial utilizing 99mTc-tracers available so far. In this trial, which was monitored by official authorities (FDA, EMEA), no serious adverse events occurred in the 234 patients who were dosed with Gd-DTPA-BMA. MR-IMPACT confirmed the high diagnostic performance of the first pass perfusion CMR approach at a dose of 0.1 mmol/kg Gd-DTPA-BMA with an AUC of 0.86 for the detection of CAD (defined as

0.75

0.75

Sensitivity

Sensitivity

1

0.5

0.25

0.5

0.25

Adequate quality Intermediate quality All quality

0

Single center 3 centers - single vendor 18 centers - multi-vendor

0 0

0.25

0.5

A

0.75

1

0

0.25

B

1-specificity

0.5

0.75

1

1-specificity

Figure 16-9 A, The important influence of data quality on test performance given by the area under the receiver operating characteristic curve (AUC)1. If all categories of data quality are entered in the analysis, AUC is rather low. Applying a blinded read to eliminate data with low quality (most of them caused by breathing motion) resulted in 86% of data (intermediate category), and AUC improves. In this study, a hybrid echo planar pulse sequence was applied and yielded five to seven slices every two RR intervals (with triggering on every other R-wave only). Thus, only some slices are acquired during cardiac phases with minimal motion (end-systolic and mid-diastolic phase), while other slices fall within rapid motion phases. Eliminating these slices degraded by motion further increased test performance yielding an AUC of 0.91 (adequate quality). B, A similar trend is demonstrated for quality reduction in studies with increasing numbers of participating sites,1,25 which is most evident for multivendor data.3

Anterior

Hypoperfusion Normal Septal

Lateral

Low Normal Hypoperfusion

B

A

Inferior

E

G

Anterior

Scar Thrombus

Viable myocardium Septal

Lateral

Scar

C

D

F

Inferior

H

Figure 16-10 In this example of a 59-year-old patient, a first pass perfusion deficit is shown in panel A (at the peak effect of the bolus) encompassing approximately two thirds of the left ventricular circumference. From the time series of first pass data, an upslope map is generated and compared with a normal database, allowing the color-coding of pixels with reduced and normal wash-in kinetics in shades of blue and red, respectively (B). Late enhancement imaging for detection of scar tissue is shown in C. Combining perfusion (A) and viability (C) data into a single map (D) allows discrimination of hypoperfused viable (jeopardized) tissue (blue area in panel D) from hypoperfused scarred tissue (white area overlaid in panel D). Perfusion and viability are also demonstrated in polar map format of the subendocardial layer (inner half of the myocardial wall) in panels E and F, respectively. G, Coronary angiography reveals an occlusion of the left anterior descending coronary artery in this patient (arrow) with collateral vessels that protected the anterior and septal wall from transmural infarction. The large circumflex artery is not stenosed, explaining the normal perfusion pattern in the lateral wall in the polar map (E), whereas a stenosis at the crux of the right coronary artery (arrow in H) causes hypoperfusion of the inferior wall. With this stress-only perfusion approach combined with a late enhancement approach, the various tissue components (normal viable, ischemic, and scar) can be defined within various myocardial layers with a single MR one-stop-shop examination. Owing to the extensive areas of hypoperfusion in viable myocardium, the patient was revascularized.

Cardiovascular Magnetic Resonance 225

16 STRESS CARDIOVASCULAR MAGNETIC RESONANCE: MYOCARDIAL PERFUSION

1

ISCHEMIC HEART DISEASE

50% diameter reduction on QCA). Most important, it also showed the superiority of perfusion CMR over SPECT imaging for the detection of CAD.3 From this trial, detailed reading and quality criteria were derived and tested in the MR-IMPACT II trial performed with 33 centers in Europe and the United States.110 This MR-IMPACT II demonstrated in 465 patients a strong superiority of perfusion CMR over SPECT (area under the ROC curve: 0.75 versus 0.65, p ¼ 0.0004, respectively), which was also superior in comparison to gated-SPECT and gated-SPECT in multivessel disease.110

PERSPECTIVES A main feature of CAD is its chronic course over several decades. CAD may become clinically overt by symptoms in an early stage of disease, but it may also remain silent for many years, becoming symptomatic suddenly by an acute myocardial infarction. Since symptoms are not always present heralding an imminent infarction or symptoms may be atypical and therefore may be misinterpreted, up to 60%

of cardiac deaths occurred in the United States outside the hospital or before the patients reached the catheterization laboratory for rescue intervention in the year 2004.111 It can be concluded that such a reactive strategy, that is, evaluating a patient primarily when the patient is exhibiting symptoms, is often inadequate. A more successful strategy would focus on an earlier detection of CAD, that is, a noninvasive detection of severe and, thus, dangerous coronary lesions by CMR and to treat these lesions before they rupture, thereby preventing acute myocardial infarctions. Such an active strategy would monitor the disease process over many years and would indicate the need for action in case of a significant lesion. Since the factors that trigger progression of CAD are not yet known, a repetitive evaluation of the coronary circulation would be required. Perfusion CMR is an ideal tool to monitor CAD over time, since it can be repeated when needed, owing to a lack of radiation. It yields a high sensitivity and specificity for detection of obstructing lesions,1–3 and as a short examination, it is well tolerated by patients. An active strategy could reduce the high number of “prehospital” cardiac deaths significantly, which would represent a major success in cardiac care.

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Comparison of Perfusion and Wall Motion Cardiovascular Magnetic Resonance Imaging Ingo Paetsch and Eike Nagel

A crucial problem in clinical decision making is the accurate noninvasive diagnosis and functional evaluation of patients with suspected or known coronary artery disease to enable appropriate selection of management strategies. With cardiovascular magnetic resonance (CMR), several methods for the detection of myocardial ischemia are available. Owing to space limitations within the scanner bore and impairment of image quality resulting from patient movement, pharmacologic stress agents such as dobutamine or adenosine have mainly been used to define the clinical usefulness of stress CMR examinations for the detection of inducible ischemia. Most commonly, dobutamine stress CMR (DSMR) cine imaging is performed for the visualization of ischemic wall motion abnormalities, or adenosine stress perfusion imaging is performed during the first pass of a contrast agent bolus for the visualization of blood flow alterations. Even though both tests aim at the detection of hemodynamically significant coronary artery stenoses, the underlying concepts show differences that may lead to seemingly discrepant test results. The most important differences are the capability of the pharmacologic stress agent to induce myocardial ischemic reactions, the imaging strategy for visualizing ischemia, and the assessment and interpretation of the images.

they branch into smaller, penetrating arteries which traverse the myocardial wall from the epicardium to the endocardium. The anatomic structure of this coronary microcirculation is purpose-built for rapid oxygen delivery: a high-density network of small-diameter capillaries ensures rapid delivery of oxygen and exchange of substrates, the relative capillary surface area being 15 times larger than in skeletal muscle. Thus, this system is ideally suited to meet the metabolic demands of the ever-working myocardial contractile cells. In the presence of chronic obstructive or occlusive coronary artery disease, the human heart can form anastomoses between different vascular regions that may serve as natural bypasses for blood to reach myocardial territories distal to occlusions or severe stenoses, thereby preventing or alleviating myocardial infarction and demand ischemia, respectively. Such collateral vessels are located mainly in the subendocardial layers. The presence of collateral vessels adds additional complexity to the diagnosis of patients with coronary artery disease.

CORONARY ARTERY PHYSIOLOGY AND PATHOPHYSIOLOGY

Coronary circulation is tightly controlled by a combination of hydrodynamic influences (arterial perfusion pressure, extravascular compressive forces, flow characteristics of the blood), myogenic, neuronal, humoral, local metabolic and endothelium derived factors. These mechanisms regulate total coronary resistance and, thus, coronary blood flow. In healthy humans, resting myocardial oxygen consumption ranges from 8 to 10 mL/min/100 g, while maximal myocardial oxygen consumption ranges from 30 to 40 mL/min/100 g. Theoretically, though, a maximal myocardial oxygen supply of up to 60 mL/min/100 g may be achieved (on the basis of a maximal coronary blood flow of 400 mL/min/100 g and an arteriovenous difference in oxygen pressure of 15 volume percent). This strongly implies that in subjects with healthy coronary arteries, myocardial blood flow is unlikely to be the limiting factor of myocardial function.

To achieve a better understanding of the effects of pharmacologic stimulation, a short description of the basic physiologic and pathophysiologic principles of coronary arterial blood flow regulation is given. The coronary circulation is designed to supply the heart with oxygen and nutrients for a wide variety of metabolic needs encountered in an active human being. Alterations of cardiac function occur rapidly, requiring strong and fast compensatory mechanisms for oxygen supply, the latter being a direct function of coronary blood flow. The epicardial coronary arteries serve mainly as transport vessels with minimal resistance. Further distal,

CORONARY AUTOREGULATION

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CHAPTER 17

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CORONARY FLOW RESERVE With such a high oxygen demand of myocardial contractile cells, one of the unique findings in the coronary circulation is its maximum degree of oxygen extraction even under basal conditions. Therefore, adjustment to changing metabolic needs of the heart is virtually impossible by further increasing oxygen extraction from incoming arterial blood. Accordingly, changes in myocardial demand mandate concomitant changes in coronary blood supply. In humans with normal coronary arteries, blood supply of the heart is downregulated at rest. This downregulation is done by the small vessels intersecting the myocardium (known as resistance vessels or microvessels) which constrict at rest, thereby increasing resistance and decreasing blood flow. During physical exercise, though, these microvessels dilate, and the resultant regulatory response is an increase in blood flow. This recruitment of blood flow (known as flow reserve) represents the most important regulatory mechanism and can amount to a factor of 2 to 5 with a wide range of interindividual variability. Per definition, coronary flow reserve (CFR) is the ratio between the maximally achievable coronary flow during hyperemia and the resting coronary flow. Since coronary blood flow velocity is proportional to coronary flow, CFR can be estimated by measuring the coronary blood flow velocity at rest and during maximal vasodilation (e.g., during infusion of a coronary vasodilator such as adenosine).1 Such measurements can be performed during invasive catheterization by using a standard angioplasty guidewire with a Doppler transducer mounted at its tip. CFR in a patient with normal coronary structure and function should be greater than 2 (normal range: 2 to 5). Conversely, a CFR of less than 2 indicates a functionally significant coronary stenosis. In humans with epicardial coronary artery stenoses (<80% diameter reduction), blood supply of the heart is often nearly normal at rest, owing to a compensatory dilation of the microvascular bed distal to the stenosis.2 As a result, in most patients with chronic obstructive coronary disease, ischemia cannot be detected at rest. However, under stress conditions (e.g., induction of a hyperemic state), no further dilation of the distal resistance vessels is possible; the result is a CFR of less than 2 or a maldistribution of blood flow within the respective myocardial territories, which can be measured with perfusion imaging techniques. Unfortunately, CFR measurements are limited by the dependence on heart rate and blood pressure, and this has mainly been quoted to explain the observed limited reproducibility. As a result, the concept of fractional flow reserve (FFR) has been developed. FFR is derived by comparing the mean coronary pressure distal to a stenosis (as measured by a coronary pressure wire) with the mean proximal coronary pressure (as measured by a guide catheter) at maximal hyperemia.3–5 FFR has a clear normal value of 1.0 and a wellestablished abnormal value of less than 0.75. FFR is advantageous not only because it is independent of blood pressure and heart rate but also because it takes into account the contribution of a collateral coronary circulation. FFR measurements have been well validated in 230 Cardiovascular Magnetic Resonance

comparison with other, noninvasive imaging modalities, such as stress echocardiography and single photon emission computed tomography (SPECT), for the detection of myocardial ischemia. and a high concordance of results (93%) has been described.6

DETECTION OF MYOCARDIAL ISCHEMIA WITH NONINVASIVE IMAGING The following pathophysiologic concepts are pivotal to understanding the imaging abnormalities that can be detected with noninvasive imaging methods that visualize either myocardial blood flow alterations or myocardial contractile function.

Concept of the Ischemic Cascade During myocardial ischemia, the temporal sequence of reduced perfusion, a decline in myocardial contractile function that then is followed by an abnormal electrocardiogram (ECG), and anginal symptoms has been termed the ischemic cascade.7 Since its first description in an animal model of coronary artery occlusion, various investigators have shown that the concept is valid under the conditions of demand ischemia in humans as well. Interestingly, it was found that perfusion abnormalities not only precede an impaired contractile response, but frequently also show a varying spatial extent in comparison to the observed wall motion abnormality (segmental mismatch). Both phenomena are therefore referred to as the spatiotemporal disparity between myocardial perfusion and wall motion abnormalities and are deemed to be responsible for the observed higher sensitivity of perfusion abnormalities over dysfunctional wall motion and the higher prevalence of both over a pathologic ECG or anginal symptoms during stress testing.8–11

Concept of Absolute Versus Relative Ischemia Generally, the onset of ischemia is defined as a reduction in myocardial blood flow to a specific myocardial territory. Thus, by definition, any absolute reduction in blood flow can be considered an ischemic reaction, and this is usually referred to as absolute ischemia. From the myocardial dysfunction point of view, however, an imbalance between myocardial oxygen demand and supply represents the foundation for the development of wall motion abnormalities. Since the heart is an aerobic organ, it relies almost exclusively on the continuous delivery of substrates and oxygen to properly maintain its contractile function. In contrast to skeletal muscle, the heart’s ability to develop an oxygen debt is negligible. Thus,

STRESS AGENTS Pharmacologic Effects Both adenosine and dobutamine are routinely used to induce ischemic reactions of the myocardium, and their relative value for the induction of perfusion abnormalities and dysfunctional wall motion has been examined by using different imaging modalities, including echocardiography, radionuclide scintigraphy, and CMR. Dipyridamole is an inactive prodrug and is activated to adenosine during liver metabolism, which is then responsible for the observed coronary vasodilatory effects. Thus, the vasodilation induced by dipyridamole depends on individual metabolism rate, usually resulting in a longer half-life, prolonged side effects, and varying vasodilatory capacity in comparison with a sole adenosine infusion. Hence, adenosine is used in most institutions and will mainly be described here (see Table 17-1).

Adenosine Adenosine, an endogenous nucleotide, is a potent vasodilator of most vascular beds (except for hepatic and renal arterioles). It exerts its pharmacologic effects through the activation of purine A1 and A2 cell-surface adenosine receptors. The essence of the pharmacologic mechanism lies in the inhibition of the slow inward Ca2þ current, thereby reducing calcium uptake, and in the activation of adenylate cyclase through A2 receptors in smooth muscle cells. The major pharmacologic effect of adenosine on the coronary circulation is vasodilation of resistance vessels. This vasodilation is deemed responsible for the observed alteration of blood flow distribution within the myocardium. Adenosine exhibits no direct effects on contractile function; the increase in myocardial wall thickening that is seen during adenosine infusion represents an indirect effect: In most patients, adenosine causes a reflex increase in heart rate, which then leads to some hypercontractile

response of the myocardium, which is most likely due to the frequency-inotropy relationship.12,13 As described above, coronary autoregulation keeps blood flow constant even in the presence of high-grade coronary artery stenoses. However, in the presence of a highgrade coronary stenosis, the resistance vessels are already fully dilated even under resting conditions in an effort to normalize myocardial blood flow distal to the epicardial stenosis (autoregulation).2 During adenosine stimulation, all resistance vessels become maximally dilated; in such a state of maximal hyperemic response, autoregulation is completely abolished, and coronary blood flow is directly related to the driving pressure. Consequently, myocardial resistance in areas supplied by normal coronary arteries is further downregulated while in areas supplied by stenotic epicardial coronary arteries no additional decrease is possible. Even though these effects must not cause relative ischemia in the sense of oxygen demand exceeding supply, blood flow inhomogeneities are found between normal and stenotic coronary arterial territories (the so-called steal effect). Adenosine is usually administered via the peripheral intravenous route at a dosage of 140 mg/kg/min over 6 minutes total infusion duration. As with all tests that are aimed at detection of absolute reductions of myocardial blood flow, it is crucial to achieve the condition of true steady-state hyperemia. The accuracy of measurements of myocardial blood flow distribution profoundly depends on persistent maximal hyperemia during the time frame needed for imaging. ADENOSINE: STRESS-INDUCIBLE WALL MOTION ABNORMALITIES In the past, the value of adenosine stress echocardiography was examined for the detection of ischemic wall motion abnormalities. Those investigators found low sensitivities, ranging from 40% to 58%, and a high specificity (from 87% to 100%).14,15 In a recent CMR study directly comparing adenosine and DCMR for detection of inducible wall motion abnormalities related to the presence of epicardial stenoses (>50% diameter stenosis), we reported a similarly low sensitivity and a high specificity for adenosine stress CMR (40% and 96%, respectively; diagnostic accuracy: 58%), while DCMR proved to be superior (89% and 80%, respectively; diagnostic accuracy: 86%).16 In this study, the perfusion status of segments with or without inducible wall motion abnormalities was determined as well. Inducible wall motion abnormalities under adenosine stress occurred exclusively in segments that showed

Table 17-1 Most Common Side Effects of Dobutamine and Adenosine During Stress Testing DOBUTAMINE

ADENOSINE

Noncardiac

Cardiovascular

Noncardiac

Cardiovascular

Nausea Urinary urgency Chills

Chest pain Dyspnea Severe increase in blood pressure (> 240/120 mm Hg) Decrease in systolic blood pressure (> 40 mm Hg) Paroxysmal atrial fibrillation Ventricular extrasystoles Ventricular fibrillation (sustained/nonsustained)

Nausea Headache Dizziness/lightheadedness

Chest pain Dyspnea Atrioventricular block Tachyarrhythmias Flushing Hypotension Hypertension

Cardiovascular Magnetic Resonance 231

17 COMPARISON OF PERFUSION AND WALL MOTION CARDIOVASCULAR MAGNETIC RESONANCE IMAGING

adequate oxygen and substrate delivery directly depends on the maintenance of adequate blood supply under all conditions. The three major determinants of myocardial oxygen demand are heart rate, myocardial contractility, and ventricular wall tension. Thus, all stressors capable of sufficiently increasing these parameters can be useful for the induction of relative ischemia (i.e., myocardial oxygen demand exceeding supply).

ISCHEMIC HEART DISEASE

concomitant high-grade segmental perfusion deficits (>75% transmurality). Thus, adenosine-inducible wall motion abnormalities are strictly linked to extensive blood flow maldistribution, the transmurality of the perfusion deficit being a strong predictor for the development of concomitantly occurring segmental myocardial dysfunction. Such extensive absolute reductions in myocardial blood flow were related to the presence of high-grade epicardial stenosis only (>75% diameter stenosis). Thus, adenosine stress cannot be recommended for the detection of wall motion abnormalities related to the presence of epicardial coronary stenoses with 50% to 75% luminal narrowing. ADENOSINE: STRESS-INDUCIBLE PERFUSION ABNORMALITIES Several studies dealing with the diagnostic performance of adenosine stress CMR perfusion imaging have reported high diagnostic accuracies for the detection of significant coronary artery disease, ranging from 81% to 89%.16–20

Dobutamine Dobutamine is a potent synthetic catecholamine that acts mainly via dominant b1-receptor stimulation, with b2- and a1-receptor stimulation occurring to a much lesser degree. The major pharmacologic effects of dobutamine are dosedependent: In the low-dose range (5 to 15 mg/kg/min), marked positive inotropy is observed, whereas higher doses (up to 40 mg/kg/min) lead to a progressive increase of heart rate. In addition, peripheral b2 stimulation may reduce systemic vascular resistance, which is most likely the reason for the lack of an increase (or sometimes even decrease) in systolic blood pressure seen in some patients. Supply-demand mismatch resulting from increased myocardial oxygen demand has been recognized as the mechanism that is responsible for inducible regional systolic dysfunction during intravenous infusion of dobutamine.21 At doses of up to 20 mg/kg/min, dobutamine has been reported to exert a direct vasodilatory effect on the coronary artery circulation as well.22 As has already been mentioned, some patients experience a drop in systolic blood pressure at higher dobutamine doses, owing to a decrease of peripheral vascular resistance. However, this is well tolerated in almost all cases, and the dobutamine infusion should be discontinued only if hypotension is severe (systolic blood pressure drop of >40 mm Hg) or symptomatic. To provide a marker of goodness of pharmacologic stress testing, the peak rate pressure product (heart rate systolic blood pressure) can be calculated, as this provides a reasonable estimate of myocardial workload (a heart rate pressure product greater than 20,000 mm Hg/min usually indicating an adequate hemodynamic response). DOBUTAMINE: STRESS-INDUCIBLE WALL MOTION ABNORMALITIES Over the years, several investigators have consistently reported on the high diagnostic value of dobutamine stress-inducible wall motion abnormalities for the detection of significant epicardial coronary disease (>50% luminal narrowing) with an overall diagnostic accuracy of around 86% (range: 83% to 90%). In general, in patients who are amenable to DCMR wall motion imaging, the method has 232 Cardiovascular Magnetic Resonance

the potential to become the preferred imaging technique for the detection of inducible wall motion abnormalities and can be regarded as the method of first choice in patients with limited echocardiographic image quality.23,24 DOBUTAMINE: STRESS-INDUCIBLE PERFUSION ABNORMALITIES Dobutamine in high doses (20 to 40 mg/kg/min) increases the three main determinants of myocardial oxygen demand (i.e., heart rate, systolic blood pressure, and myocardial contractility), thereby eliciting a secondary increase in myocardial blood flow and provoking myocardial perfusion abnormalities. The flow increase (twofold to threefold baseline values) is reportedly less than that elicited by adenosine. However, dobutamine stress myocardial perfusion imaging using radioactive isotopes including 201Tl, 99mTC-sestamibi, and 99mTC-tetrofosmin has been successfully performed. The overall sensitivity, specificity, and accuracy of dobutamine stress nuclear studies were 85%, 72%, and 83%, respectively.25 Data on DCMR perfusion imaging is relatively limited; in a feasibility study, Al-Saadi and colleagues26 showed that a quantitatively derived myocardial perfusion reserve index differentiated ischemic and nonischemic myocardial segments reaching a sensitivity, specificity, and diagnostic accuracy of 81%, 73%, and 77%, respectively. In this study, however, stress imaging has been done at low-dose dobutamine (20 mg/kg/min level) only to avoid the marked increase of heart rate that occurs at higher doses. The main obstacle for performing myocardial perfusion imaging at high-dose dobutamine levels has been the inability of previous perfusion CMR techniques to acquire multiple slices at high heart rates (i.e., >120 beats per minute) while preserving a high spatial resolution. With the more recent development of parallel imaging techniques (e.g., sensitivity encoding, SENSE),27–29 dynamic, multislice CMR perfusion imaging has become much faster and therefore can be used to assess myocardial perfusion abnormalities occurring under high-dose dobutamine stress (Fig.17-1C). Though no large-scale study on the diagnostic value of dobutamine stress CMR perfusion imaging has been published yet, we do perform CMR perfusion imaging as an integral part of dobutamine stress testing for wall motion assessment (as exemplified in Fig. 17-2C). In our experience, the diagnostic accuracy of high-dose dobutamine myocardial perfusion CMR imaging is comparable to the accuracy of high-dose dobutamine stress nuclear studies. Hence, DCMR might be used as an alternative pharmacologic stress test in patients with contraindications to vasodilator stress.

Safety Aspects Adenosine The vasodilatory effect of adenosine may result in a mild to moderate reduction in systolic, diastolic, and mean arterial blood pressure (<10 mm Hg) with a reflex increase in heart rate. Most patients complain about diffuse chest pain and shortness of breath. These effects are transient, however, and usually do not require medical intervention.

A 0 min

Cine

Cine Perfusion 6 min

10 min

Break (>15 min)

Adenosine (140 μg/kg/min)

≈31 min

≈45 min

Stop

Cine

B 0 min

Cine

Rest

Start Survey

Perfusion ≈30 min

12 min

Stop

Start

Stop

Start Survey

Dobutamine (up to 40 μg/kg/min, +/− atropine)

Rest

Perfusion 6 min

10 min

Break (10 min) 12 min

Perfusion

≈22 min

Break (10 min)

Scar imaging

≈23 min

≈40 min

Survey

C 0 min

Cine

Perfusion

Stop

Start

Dobutamine (up to 40 μg/kg/min, +/− atropine )

Cine (repetitive) 6 min

9 min

12 min

Perfusion 15 min

18 min

Break (10 min) Scar imaging 19 min

≈30 min

≈35 min

Figure 17-1 Time course of stress testing (dobutamine and adenosine administration) and corresponding CMR (cine, dynamic perfusion, and late gadolinium enhancement imaging). A, Combined adenosine stress perfusion and high-dose dobutamine/atropine stress CMR. B, Assessment of rest and adenosine stress perfusion CMR. Optionally, stress testing may be followed by late gadolinium enhancement CMR for myocardial scar detection; if done so, rest perfusion might be spared, and the observer needs to rely on judging the fixed portion of a perfusion deficit from scar extent only. C, Assessment of high-dose dobutamine/atropine stress perfusion CMR at rest and during maximal stress in combination with repetitive cine imaging during graded dobutamine infusion. Though dobutamine stress testing is already capable of accurately diagnosing viable myocardium, stress testing may be followed by scar imaging.

Since adenosine exerts a direct depressant effect on the sinoatrial and atrioventricular (AV) nodes, the transient occurrence of first-, second-, and third-degree AV block and sinus bradycardia has been reported in 2.9%, 2.6%, and 0.8% of patients, respectively.30 Also, adenosine can cause significant hypotension. Patients with an intact baroreceptor reflex are able to maintain blood pressure in response to adenosine by increasing cardiac output and heart rate. Adenosine can also cause a paradoxical increase in systolic and diastolic blood pressure, which develops mostly in individuals with significant left ventricular hypertrophy, but these effects are usually transient and resolve spontaneously. Because adenosine is a respiratory stimulant primarily through activation of carotid body chemoreceptors, intravenous administration showed increases in minute ventilation and a reduction in arterial pCO2, resulting in respiratory alkalosis.12 Approximately 20% to

30% of patients complain of dyspnea and an urge to breathe deeply during adenosine infusion. Because of the above reported adverse effects, a number of studies have been carried out investigating the safety of intravenous adenosine infusions in different diagnostic modalities of cardiac imaging.30,31 So far, there is evidence that was accumulated in over 10,500 patients who were studied with thallium radionuclide imaging, echocardiography, SPECT, and CMR that shows that pharmacologic stress with adenosine presents a safe method of acquiring stress imaging data. Safety of an adenosine infusion at 140 mg/kg/min was evaluated during radionuclide imaging of 9256 consecutive patients.30 The infusion protocol was completed in 80% of patients, required dose reduction in 13%, and was terminated early in 7%. Interpretable imaging studies were obtained in 98.7% of patients, and 0.8% of patients Cardiovascular Magnetic Resonance 233

17 COMPARISON OF PERFUSION AND WALL MOTION CARDIOVASCULAR MAGNETIC RESONANCE IMAGING

Adenosine (140 μg/kg/min)

Rest

508) without alteration in the adenosine infusion. There were no sustained episodes of AV block.

Dobutamine The safety of high-dose DCMR has been investigated. In a routine clinical setting, high-dose DCMR was attempted on 1075

DOBUTAMINE WALL MOTION Low dose

Max

End diastole

Rest

End systole

ISCHEMIC HEART DISEASE

received aminophylline. Minor and well-tolerated side effects were reported in 81.1% of patients. There were no deaths, one myocardial infarction, seven episodes of severe bronchospasm, and one episode of pulmonary edema. Transient AV node block occurred in 706 patients (firstdegree in 256, second-degree in 378, and third-degree in 72) and resolved spontaneously in most patients (n ¼

ADENOSINE PERFUSION Rest

SCAR Stress

A Figure 17-2 Stress CMR examples corresponding to the proposed imaging protocols in Figure 17-1A–C. A, Dobutamine stress CMR showed an extensive inducible wall motion abnormality at maximum stress level (anterior and septal segments, white arrows). Adenosine stress CMR demonstrated an extensive inducible perfusion deficit; note that the inducible perfusion deficit covered the inferolateral segment as well (white arrowheads), but no visually assessable wall motion abnormality is demonstrated (segmental mismatch). Since myocardial scar was not present, the absence of resting perfusion abnormalities could be confirmed. 234 Cardiovascular Magnetic Resonance

Stress

B DOBUTAMINE WALL MOTION Low dose

Max

End systole

End diastole

Rest

ADENOSINE PERFUSION Rest

SCAR Stress

C Figure 17-2—cont’d B, Perfusion CMR at rest was found to be normal; under adenosine stress, though, a subendocardial perfusion deficit of varying transmurality is demonstrated (inferior and inferolateral segment, white arrows). Myocardial scarring was absent. C, Dobutamine stress CMR detected an inducible wall motion abnormality of the inferolateral segment (white arrows). Dobutamine stress perfusion CMR demonstrated an inducible perfusion deficit in identical location. Myocardial scarring was absent. Cardiovascular Magnetic Resonance 235

17 COMPARISON OF PERFUSION AND WALL MOTION CARDIOVASCULAR MAGNETIC RESONANCE IMAGING

SCAR

ADENOSINE PERFUSION Rest

ISCHEMIC HEART DISEASE

occasions in 1035 consecutive patients using the graded standard dobutamine protocol of 10, 20, 30, and 40 mg/kg/min dobutamine for 3 minutes each plus atropine if needed to reach target heart rate. A high number of diagnostic examinations was found (89.5%). Reasons for nondiagnostic tests included failure of ECG triggering (n ¼ 4), suboptimal image quality (n ¼ 6), submaximal stress (n ¼ 21), and limiting side effects (n ¼ 74), including chest pain, dyspnea, and nausea. Atrial fibrillation was provoked in five patients, and nonsustained atrial fibrillation was provoked in 16. Only one patient suffered sustained ventricular tachycardia and underwent successful emergent defibrillation, while no cases of death or myocardial infarction occurred over the observed 5-year period. The patients who were being examined were typical of those evaluated for coronary disease, and over half had ischemia induced at the time of dobutamine stress, thereby closely reflecting the clinical reality. This study demonstrated that the technique can be undertaken with a safety profile virtually identical to that previously reported for dobutamine stress echocardiography.32–35

Drug Interactions

Adenosine infusion should be exercised with caution in patients with chronic lung diseases not associated with bronchoconstriction (emphysema, bronchitis, etc.) and should be avoided in patients with bronchoconstriction or bronchospasm (e.g., allergic asthma). If a patient develops severe respiratory difficulties, the adenosine infusion is ceased immediately, and theophylline can be applied intravenously to rapidly dissolve bronchoconstriction (Tables 17-2 and 17-3).

Dobutamine In general, the dobutamine infusion is to be stopped if target heart rate—defined as age-predicted submaximal heart rate ([220 age] 0.85)—is reached. Additional termination criteria are given in Table 17-3. Contraindications include acute coronary syndromes, severe aortic stenosis, hypertrophic cardiomyopathy, uncontrolled hypertension, and uncontrolled heart failure. If atropine is needed to reach the target heart rate, its contraindications must be taken into consideration (i.e., narrow-angle glaucoma, myasthenia gravis, obstructive uropathy, or obstructive gastrointestinal diseases; see also Tables 17-2 and 17-3).

Dobutamine

Practicability

It is very important that the patient be instructed that betablockers be withheld for at least 24 hours prior to the examination, since beta-blockers reduce the inotropic and chronotropic effects of dobutamine and thus reduce the sensitivity of the stress test. Preferably, other antianginal medication should be withdrawn as well.

A detailed description of the cognitive and training skills necessary for competent performance of pharmacologic stress testing has been released by the AHA/ACC36; in general, these skills do apply to CMR stress testing as well. However, CMR has some unique complexities resulting from the static magnetic field and the examination of a patient in an enclosed magnet.

Adenosine The patient should refrain from caffeine-containing beverages or food (tea, coffee, chocolate, etc.), smoking, and any antianginal medication 24 hours prior to the examination. Caffeine and other methylxanthine derivatives act like theophylline, which is the clinically used antidote of adenosine.

Contraindications and Termination Criteria Adenosine Adenosine should be used with caution in patients with preexisting AV block or bundle branch block and should be avoided in patients with high-grade AV block or sinus node dysfunction. Adenosine should be used with caution if a patient is receiving any medication that already depresses the sinus node and/or AV conduction (e.g., beta-blockers, calcium channel blockers, cardiac glycosides). Adenosine needs to be discontinued in patients who develop persistent or symptomatic high-grade block or a significant drop in systolic blood pressure (>20 mm Hg). The drug should be discontinued in case of persistent or symptomatic hypotension. Also, adenosine should be used with caution in patients with autonomic dysfunction, stenotic valvular heart disease, pericarditis and pericardial effusion, high-grade carotid artery stenoses and cerebrovascular insufficiency or uncorrected hypovolemia. 236 Cardiovascular Magnetic Resonance

Monitoring Monitoring needs are similar for the different stressors. Heart rate and rhythm need to be registered throughout

Table 17-2 Contraindications for Dobutamine/ Atropine and Adenosine Stress DOBUTAMINE

ADENOSINE

General Contraindications

General Contraindications

Unstable angina Severe arterial hypertension Significant aortic stenosis (gradient > 50 mm Hg or area < 1 cm2) Significant obstructive hypertrophic cardiomyopathy Complex cardiac arrhythmias

Unstable angina Arterial hypotension Myocardial infarction < 3 days Asthma

Myocardial inflammation (perimyocarditis) Caution Comedication with diuretics (hypokalemia) Atropine Contraindications Narrow-angle glaucoma, myasthenia gravis, obstructive uropathy

Severe obstructive pulmonary disease Preexisting AV block Caution Stenotic valvular disease Cerebrovascular insufficiency Comedication with b-blockers/Caantagonists/ digitalis Sinoatrial disease

Dobutamine

Adenosine

Target heart rate ([220 age] 0.85) reached Hypertension (blood pressure > 240/120 mm Hg) Systolic blood pressure drop > 40 mm Hg Intolerable symptoms (chest pain, dyspnea, etc.) Significant supraventricular/ ventricular arrhythmias

Persistent or symptomatic AV block or other arrhythmia Persistent or symptomatic hypotension Significant blood pressure drop > 20 mm Hg Severe respiratory difficulties

the stress examination. Basically, changes of the ST segment are nondiagnostic as a result of the ECG wave distortion in the static magnetic field. However, since wall motion abnormalities precede ST segment changes and the former can readily be detected with CMR imaging, monitoring with rapid cine sequences is effective without a diagnostic ECG (e.g., cine CMR real-time scans can be run repetitively to detect new or worsening wall motion abnormalities at the very first occurrence).37 Blood pressure monitoring can easily be done with a conventional monitoring system outside the scanner room with an extension line placed through a waveguide in the radiofrequency cage or special CMR compatible equipment may be used.

Patient Evacuation and Emergency Equipment CMR stress testing can be done safely and successfully if certain caveats are closely followed. In general, monitoring during a stress CMR examination requires the same precautions and emergency equipment as any other stress test.36 A physician who is appropriately trained in advanced cardiac life support must be present throughout the stress examination and during the recovery phase (personal supervision). Precautions for rapid patient evacuation must be taken; that is, a trolley should be permanently placed under the table, and a button for manual table release must be present. The staff must regularly practice the maneuver for rapid patient evacuation consisting of the immediate cessation of the dobutamine infusion, disconnection of the cardiac receiver coil and the ECG cable, and evacuation of the patient using the table-trolley unit, all of which are usually achievable in less than 30 sec by two staff members. Outside the scan room, cardiac resuscitation can be performed according to emergency guidelines.

Image Display and Analysis During the pharmacologic stress procedure, the examiner continuously evaluates the CMR cine loops in an automatic view window as displayed on the console of the scanner. The reconstruction speed of the cine images has become very fast and allows on-line visual assessment of wall motion, with the cine loops being displayed almost instantaneously. Alternatively, with a workstation next to the scanner console, the cine loops can be automatically transferred and displayed in a synchronized quadscreen. For

assessment of perfusion abnormalities with CMR, the interpretation of both stress and rest images is required. Thus, on-line monitoring of the very first occurrence of perfusion abnormalities is limited.

Duration of Stress CMR Examinations A comparison of the examination times required for the different imaging strategies is shown in Figure 17-1. Whereas the adenosine stress CMR perfusion protocol per se is quite rapid (maximal hyperemic stress is achieved after 4 minutes, and imaging takes approximately 2 minutes), most centers add either a perfusion study at rest (required for the calculation of myocardial perfusion reserve) or a delayed enhancement study for the detection of viable and scarred myocardium. In contrast, DCMR takes longer (12 to 15 minutes of graded dobutamine infusion with repetitive cine imaging). However, a dobutamine stress test provides all information (i.e., presence of infarction, viability, ischemia) without the need for additional, contrast-enhanced studies. Consequently, for both pharmacologic stress CMR imaging approaches, the duration of time when the patient is within the scanner is similar.

Pitfalls and Advanced Issues Coverage CMR perfusion imaging usually covers 16 out of 17 myocardial segments according to the standardized myocardial segmentation of the heart using three short axis views (apical, equatorial, and basal).38 Although segment 17 (i.e., the apical cap) is not covered, studies incorporating the apex with imaging of an additional long axis view during CMR perfusion imaging failed to show ischemia in this region.39 Furthermore, the analysis of only the inner three out of five to eight short axis slices resulted in a significantly improved diagnostic accuracy for the detection of coronary artery disease.18 These findings illustrate that diagnostic quality may differ with the order of slice acquisition and/or slice location. Thus, the question of optimal coverage and segmentation cannot be answered definitely at the present time.

Functional Assessment of Viable Myocardium With CMR, myocardial viability can be assessed from a morphologic and a functional perspective. The morphologic approach aims at detecting regions of scarred myocardium after intravenous injection of an extracellular, gadolinium contrast agent. This late gadolinium enhancement (LGE) technique distinguishes scarred myocardium (appearing bright) from nonscarred myocardium (appearing dark). The transmural extent of LGE can then be used to predict the likelihood of functional recovery after restoration of blood flow to the respective myocardial territory. As an alternative, the functional, contractile response to low-dose dobutamine stimulation can be determined. The low-dose dobutamine challenge was found to be superior in predicting recovery of function, particularly in segments with a scar transmurality of 1% to 74%.40 As a possible explanation for this finding, it was suggested that even Cardiovascular Magnetic Resonance 237

17 COMPARISON OF PERFUSION AND WALL MOTION CARDIOVASCULAR MAGNETIC RESONANCE IMAGING

Table 17-3 Termination Criteria for Dobutamine and Adenosine Stress

ISCHEMIC HEART DISEASE

though scar imaging depicts the area of myocardial fibrosis, it does not assess the functional state of the surrounding (potentially viable) myocardial tissue. As a consequence, its capability for the prediction of functional recovery of nontransmurally scarred myocardium was found to be limited. As can be seen in Figure 17-1, LGE can easily be integrated when combined stress CMR perfusion and/or stress CMR wall motion imaging is performed.

myocardial blood flow reserve.43 Therefore, judging the induction of a maximal hyperemic response from these parameters alone might be misleading.

Route and Duration of Administration

Combined Adenosine Perfusion and Dobutamine Wall Motion CMR

In general, for noninvasive stress testing, both dobutamine and adenosine are given intravenously as continuous infusions via a peripheral line. DOBUTAMINE For dobutamine, the likelihood of adverse effects increases with infusion duration rather than dosage. Hence, the examiner is advised to abide by the 3-minute intervals for each stress level. In addition, in a minority of patients (15%), no relevant increase in heart rate might be observed up to the 30 mg/kg/min level (premaximal level); in this case, an early, low-dose application of atropine (e.g., 0.25 mg) should be considered, since with alleviating a dominant vagotonic state, dobutamine rapidly exerts its full effects. Side effects of dobutamine should preferably be reverted by using a short-acting beta-blocker (e.g., esmolol), since then the nearly identical half-life of stressor and antagonist (both 2 to 3 minutes) ensures that side effects of the betablocker do not occur long after the dobutamine effect is gone. Nitroglycerin might be used to treat chest pain symptomatically, but it does not specifically antagonize the effects of dobutamine and should be reserved for patients in whom beta-blockers are contraindicated. ADENOSINE The standard intravenous dosage of adenosine is 140 mg/ kg/min applied during a total infusion duration of 6 minutes at maximum. In a detailed invasive dose-response study, this dose regimen was compared with intracoronary papaverine, which is considered to be the strongest coronary vasodilator.41 Of note, with the standard adenosine regimen, 16% of patients did not reach a maximal hyperemic response as compared with papaverine. In addition, the reagibility of the coronary system to adenosine may vary considerably between patients not only with regard to the maximally achievable hyperemic response, but also with regard to the onset of maximal hyperemia (after 112 48 sec).42 At our department, care is taken to start perfusion CMR not earlier than after 4 minutes of adenosine infusion. In addition, the patient is asked to abstain from substances that antagonize the adenosine effects (i.e., methylxanthine derivatives such as caffeine, theophylline, and theobromine) for at least 24 hours prior to the examination. Recently, the common practice of using heart rate and blood pressure to assess adenosine response during noninvasive stress testing has been questioned. In a study examining the relationship between myocardial blood flow as assessed by PET and changes in heart rate and blood pressure during adenosine infusion, the investigators found that both parameters were extremely poor predictors of 238 Cardiovascular Magnetic Resonance

COSTS It should be recognized that in most countries, the cost of direct vasodilators (in particular adenosine) is higher than that of dobutamine.

In 100 patients, Wahl and colleagues used a combined singlesession protocol for adenosine perfusion and high-dose DCMR.44 The combined protocol was reported to be safe and feasible. Reasons for nondiagnostic tests were poor image quality on perfusion CMR (n ¼ 3 patients) and inability to reach the target heart rate during dobutamine/atropine stress CMR (n ¼ 1 patient). Invasive X-ray angiography was the standard of reference for determination of diagnostic accuracy, with significant epicardial stenoses defined as 50% diameter reduction. Figure 17-3 shows the diagnostic performance of DCMR alone, perfusion CMR alone, and their respective combination (dobutamine and adenosine stress test pathologic, dobutamine or adenosine stress test pathologic). Perfusion CMR and dobutamine stress CMR reached similarly high diagnostic accuracies (86% and 85%, respectively). Compared with the combined analysis, only a marginal improvement in diagnostic accuracy (89%) could be achieved. Yet the number of diagnostic examinations

100 80

99 97 96100

97 87 88

90

84 77

77

86 85

89 81

71

60 40 20 0 Sensitivity, % Specificity, %

Accuracy, %

Diagnostic examinations, %

Dobutamine stress CMR Adenosine stress perfusion CMR Dobutamine stress and adenosine perfusion pathologic Dobutamine stress or adenosine perfusion pathologic

Figure 17-3 Preliminary data on the diagnostic performance of a combined stress examination (high-dose dobutamine stress wall motion and adenosine stress perfusion CMR) as assessed during a single-session examination in 100 patients (see also Fig. 17-1A). Dobutamine stress wall motion and adenosine stress perfusion CMR showed a similarly high diagnostic accuracy (86% and 85%, respectively) for the detection of significant epicardial coronary stenosis of 50% or more. If significant coronary artery disease was assumed in case of at least one pathologic stress test (dobutamine stress wall motion or adenosine stress perfusion CMR), diagnostic accuracy increased only marginally (89%).

CONCLUSION CMR offers the unique possibility of assessing perfusion and wall motion during pharmacologic stress in a singlesession examination. For this purpose, one can rely on either vasodilator (adenosine) or adrenergic-stimulating agent (dobutamine) stress testing. Vasodilator agents cause heterogeneous myocardial perfusion, which under specific conditions (transmurality of perfusion deficit > 75%) is sufficient to cause dysfunctional myocardial motion. However, vasodilator stress

should be used mainly in conjunction with perfusion CMR for which high diagnostic accuracies have been consistently reported. The adrenergic-stimulating agent dobutamine causes an increase in myocardial oxygen demand by increasing contractility, blood pressure, and heart rate, which are the main determinants of myocardial workload. As such, the detection of inducible wall motion abnormalities can best be accomplished by using adrenergic stimulants and a consistently high diagnostic performance for the detection of inducible, dysfunctional wall motion related to the presence of epicardial coronary stenoses; more than 50% has been demonstrated by dobutamine stress wall motion CMR. Preliminary data suggest that a combined single-session evaluation of vasodilator and adrenergic-stimulating stress testing compared to stress testing with either of the agents alone does not lead to a clear improvement in the detection of significant coronary artery disease. A single-session CMR examination of wall motion and perfusion during high-dose dobutamine stress is feasible and represents an alternative for patients with contraindications to vasodilator stress. Although initial results are encouraging, the value of high-dose dobutamine stress perfusion CMR still requires investigation in large, unselected patient populations.

References 1. Fearon WF, Yeung AC. Evaluating intermediate coronary lesions in the cardiac catheterization laboratory. Rev Cardiovasc Med. 2003;4(1):1–7. 2. Wilson RF. Assessing the severity of coronary-artery stenoses. N Engl J Med. 1996;334(26):1735–1737. 3. de Bruyne B, Bartunek J, Sys SU, Pijls NH, Heyndrickx GR, Wijns W. Simultaneous coronary pressure and flow velocity measurements in humans: feasibility, reproducibility, and hemodynamic dependence of coronary flow velocity reserve, hyperemic flow versus pressure slope index, and fractional flow reserve. Circulation. 1996;94(8): 1842–1849. 4. De Bruyne B, Pijls NH, Heyndrickx GR, Hodeige D, Kirkeeide R, Gould KL. Pressure-derived fractional flow reserve to assess serial epicardial stenoses: theoretical basis and animal validation. Circulation. 2000;101(15):1840–1847. 5. Pijls NH, De Bruyne B, Peels K, et al. Measurement of fractional flow reserve to assess the functional severity of coronary-artery stenoses. N Engl J Med. 1996;334(26):1703–1708. 6. Pijls NH, De Bruyne B, Bech GJ, et al. Coronary pressure measurement to assess the hemodynamic significance of serial stenoses within one coronary artery: validation in humans. Circulation. 2000;102(19): 2371–2377. 7. Nesto RW, Kowalchuk GJ. The ischemic cascade: temporal sequence of hemodynamic, electrocardiographic and symptomatic expressions of ischemia. Am J Cardiol. 1987;59(7):23C–30C. 8. Kaul S, Kiess M, Liu P, et al. Comparison of exercise electrocardiography and quantitative thallium imaging for one-vessel coronary artery disease. Am J Cardiol. 1985;56(4):257–261. 9. Carlson RE, Kavanaugh KM, Buda AJ. The effect of different mechanisms of myocardial ischemia on left ventricular function. Am Heart J. 1988;116(2 Pt 1):536–545. 10. Elhendy A, Geleijnse ML, Roelandt JR, et al. Dobutamine-induced hypoperfusion without transient wall motion abnormalities: less severe ischemia or less severe stress? J Am Coll Cardiol. 1996;27(2): 323–329. 11. Forster T, McNeill AJ, Salustri A, et al. Simultaneous dobutamine stress echocardiography and technetium-99m isonitrile single-photon emission computed tomography in patients with suspected coronary artery disease. J Am Coll Cardiol. 1993;21(7):1591–1596.

12. Biaggioni I, Olafsson B, Robertson RM, Hollister AS, Robertson D. Cardiovascular and respiratory effects of adenosine in conscious man: evidence for chemoreceptor activation. Circ Res. 1987;61(6):779–786. 13. Belardinelli L, Linden J, Berne RM. The cardiac effects of adenosine. Prog Cardiovasc Dis. 1989;32(1):73–97. 14. Nguyen T, Heo J, Ogilby JD, Iskandrian AS. Single photon emission computed tomography with thallium-201 during adenosine-induced coronary hyperemia: correlation with coronary arteriography, exercise thallium imaging and two-dimensional echocardiography. J Am Coll Cardiol. 1990;16(6):1375–1383. 15. Marwick T, Willemart B, D’Hondt AM, et al. Selection of the optimal nonexercise stress for the evaluation of ischemic regional myocardial dysfunction and malperfusion: comparison of dobutamine and adenosine using echocardiography and 99mtc-MIBI single photon emission computed tomography. Circulation. 1993;87(2):345–354. 16. Paetsch I, Jahnke C, Wahl A, et al. Comparison of dobutamine stress magnetic resonance, adenosine stress magnetic resonance, and adenosine stress magnetic resonance perfusion. Circulation. 2004;110 (7):835–842. 17. Al-Saadi N, Nagel E, Gross M, et al. Noninvasive detection of myocardial ischemia from perfusion reserve based on cardiovascular magnetic resonance. Circulation. 2000;101(12):1379–1383. 18. Nagel E, Klein C, Paetsch I, et al. Magnetic resonance perfusion measurements for the noninvasive detection of coronary artery disease. Circulation. 2003;108(4):432–437. 19. Schwitter J, Nanz D, Kneifel S, et al. Assessment of myocardial perfusion in coronary artery disease by magnetic resonance: a comparison with positron emission tomography and coronary angiography. Circulation. 2001;103(18):2230–2235. 20. Wolff SD, Schwitter J, Coulden R, et al. Myocardial first-pass perfusion magnetic resonance imaging: a multicenter dose-ranging study. Circulation. 2004;110(6):732–737. 21. Fung AY, Gallagher KP, Buda AJ. The physiologic basis of dobutamine as compared with dipyridamole stress interventions in the assessment of critical coronary stenosis. Circulation. 1987;76(4):943–951. 22. Bartunek J, Wijns W, Heyndrickx GR, de Bruyne B. Effects of dobutamine on coronary stenosis physiology and morphology: comparison with intracoronary adenosine. Circulation. 1999;100(3):243–249.

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17 COMPARISON OF PERFUSION AND WALL MOTION CARDIOVASCULAR MAGNETIC RESONANCE IMAGING

(i.e., at least one of the two stress tests was performed successfully with adequate image quality) for the combined approach increased to 100%. Thus, although this study represents preliminary data only, it may serve as an indicator that adenosine stress perfusion and dobutamine stress wall motion CMR might not necessarily detect the same “ischemic” abnormality, which may at least in part be attributed to the different pathophysiologies being detected (maldistribution of myocardial blood flow during vasodilator-induced hyperemia versus myocardial oxygen supply-demand mismatch during adrenergic stimulation).

ISCHEMIC HEART DISEASE

23. Hundley WG, Hamilton CA, Thomas MS, et al. Utility of fast cine magnetic resonance imaging and display for the detection of myocardial ischemia in patients not well suited for second harmonic stress echocardiography. Circulation. 1999;100(16):1697–1702. 24. Nagel E, Lehmkuhl HB, Klein C, et al. [Influence of image quality on the diagnostic accuracy of dobutamine stress magnetic resonance imaging in comparison with dobutamine stress echocardiography for the noninvasive detection of myocardial ischemia]. Z Kardiol. 1999;88(9):622–630. 25. Elhendy A, Bax JJ, Poldermans D. Dobutamine stress myocardial perfusion imaging in coronary artery disease. J Nucl Med. 2002;43 (12):1634–1646. 26. Al-Saadi N, Gross M, Paetsch I, et al. Dobutamine induced myocardial perfusion reserve index with cardiovascular MR in patients with coronary artery disease. J Cardiovasc Magn Reson. 2002;4(4):471–480. 27. Pruessmann KP, Weiger M, Scheidegger MB, Boesiger P. SENSE: sensitivity encoding for fast MRI. Magn Reson Med. 1999;42(5):952–962. 28. Plein S, Radjenovic A, Ridgway JP, et al. Coronary artery disease: myocardial perfusion MR imaging with sensitivity encoding versus conventional angiography. Radiology. 2005;235(2):423–430. 29. Weiger M, Pruessmann KP, Boesiger P. Cardiac real-time imaging using SENSE. Sensitivity Encoding scheme. Magn Reson Med. 2000;43 (2):177–184. 30. Cerqueira MD, Verani MS, Schwaiger M, Heo J, Iskandrian AS. Safety profile of adenosine stress perfusion imaging: results from the Adenoscan Multicenter Trial Registry. J Am Coll Cardiol. 1994;23(2): 384–389. 31. Cortigiani L, Picano E, Coletta C, et al. Safety, feasibility, and prognostic implications of pharmacologic stress echocardiography in 1482 patients evaluated in an ambulatory setting. Am Heart J. 2001;141(4): 621–629. 32. Mathias Jr W, Arruda A, Santos FC, et al. Safety of dobutamineatropine stress echocardiography: a prospective experience of 4,033 consecutive studies. J Am Soc Echocardiogr. 1999;12(10):785–791. 33. Secknus MA, Marwick TH. Evolution of dobutamine echocardiography protocols and indications: safety and side effects in 3,011 studies over 5 years. J Am Coll Cardiol. 1997;29(6):1234–1240. 34. Geleijnse ML, Fioretti PM, Roelandt JR. Methodology, feasibility, safety and diagnostic accuracy of dobutamine stress echocardiography. J Am Coll Cardiol. 1997;30(3):595–606. 35. Picano E, Mathias Jr W, Pingitore A, Bigi R, Previtali M. Safety and tolerability of dobutamine-atropine stress echocardiography:

240 Cardiovascular Magnetic Resonance

36.

37.

38.

39.

40. 41.

42.

43.

44.

a prospective, multicentre study. Echo Dobutamine International Cooperative Study Group. Lancet. 1994;344(8931):1190–1192. Rodgers GP, Ayanian JZ, Balady G, et al. American College of Cardiology/American Heart Association Clinical Competence Statement on Stress Testing. A Report of the American College of Cardiology/ American Heart Association/American College of Physicians-American Society of Internal Medicine Task Force on Clinical Competence. Circulation. 2000;102(14):1726–1738. Nagel E, Lorenz C, Baer F, et al. Stress cardiovascular magnetic resonance: consensus panel report: detecting left ventricular myocardial ischemia during intravenous dobutamine with cardiovascular magnetic resonance imaging (MRI). J Cardiovasc Magn Reson. 2001;3 (3):267–281. Cerqueira MD, Weissman NJ, Dilsizian V, et al. Standardized myocardial segmentation and nomenclature for tomographic imaging of the heart: a statement for healthcare professionals from the Cardiac Imaging Committee of the Council on Clinical Cardiology of the American Heart Association. Circulation. 2002;105(4):539–542. Elkington AG, Gatehouse PD, Prasad SK, Moon JC, Firmin DN, Pennell DJ. Combined long- and short-axis myocardial perfusion cardiovascular magnetic resonance. J Cardiovasc Magn Reson. 2004; 6(4):811–816. Wellnhofer E, Olariu A, Klein C, et al. Magnetic resonance low-dose dobutamine test is superior to SCAR quantification for the prediction of functional recovery. Circulation. 2004;109(18):2172–2174. Wilson RF, Marcus ML, White CW. Prediction of the physiologic significance of coronary arterial lesions by quantitative lesion geometry in patients with limited coronary artery disease. Circulation. 1987;75 (4):723–732. De Bruyne B, Pijls NH, Barbato E, et al. Intracoronary and intravenous adenosine 5’-triphosphate, adenosine, papaverine, and contrast medium to assess fractional flow reserve in humans. Circulation. 2003;107(14):1877–1883. Mishra RK, Dorbala S, Logsetty G, et al. Quantitative relation between hemodynamic changes during intravenous adenosine infusion and the magnitude of coronary hyperemia: implications for myocardial perfusion imaging. J Am Coll Cardiol. 2005;45(4):553–558. Wahl A, Paetsch I, Roethemeyer S, Gebker R, Klein C, Nagel E. A combined single session analysis of adenosine perfusion and of high-dose dobutamine stress cardiovascular magnetic resonance improves diagnosis of ischemia. Circulation. 2003;108(17):403. [Supplement IV].

Acute Myocardial Infarction: Cardiovascular Magnetic Resonance Detection and Characterization Andrew E. Arai

Noninvasive imaging in patients presenting with acute myocardial infarction (AMI) can serve several useful roles. Cardiovascular magnetic resonance (CMR) in particular is useful in assessing a wide range of clinically relevant issues. Measurement of left ventricular (LV) ejection fraction (LVEF) provides prognostic value and is a critical element in determining which patients should be treated with implanted cardiodefibrillators. CMR is well accepted as a reference standard modality of determining myocardial viability and the methods work well in patients with recent AMI. In most patients, the culprit vessel can be predicted on the basis of the spatial distribution of abnormalities on the late gadolinium enhancement (LGE) scan or other findings. CMR stress testing is capable of detecting residual myocardial ischemia. Various important complications of myocardial infarction can be imaged, such as LV dysfunction, LV thrombus, ventricular septal defect, aneurysm, pseudoaneurysm, and valvular complications. Thus, information that has prognostic value or is useful in risk stratification can be acquired in a single CMR examination in patients who survive acute myocardial infarction. This chapter will focus on the role of CMR in the detection and characterization of AMI. This goal overlaps to some degree with the goals of other chapters, such as those on stress testing and viability assessment, so the material presented will focus on validations, particularly on the perspective of detecting or characterizing AMI.

GLOBAL AND REGIONAL LEFT VENTRICULAR FUNCTION: CINE CARDIOVASCULAR MAGNETIC RESONANCE At 1.5 T, global and regional LV systolic function is relatively easily imaged with steady-state free precession (SSFP) cine CMR.1,2 SSFP cine CMR has been accepted as a reference standard for assessing LV mass, volumes, and LVEF.3–6 Cine CMR is a powerful method for detecting regional wall motion abnormalities, as is evidenced by the diagnostic accuracy of dobutamine stress tests. Recent comparisons indicate that cine CMR performs significantly better than noncontrast-enhanced echocardiography and invasive X-ray left

ventriculography in assessing regional LV systolic function.7 Only contrast-enhanced echocardiography performed slightly better than cine CMR. Cine CMR can assess regional wall motion with high image quality, allowing diagnosis of very small wall motion abnormalities with confidence that might be missed by other methods. For example, Figure 18-1 illustrates a case in which a presumptive diagnosis of clinically unrecognized myocardial infarction could be made on the basis of a focal anteroseptal wall motion abnormality that was missed on a good-quality non-contrast-enhanced transthoracic echocardiogram. Dobutamine stress testing provides the strongest evidence supporting the accuracy of cine CMR in diagnosing regional wall motion abnormalities. In publications summarizing experience using dobutamine stress testing to detect or manage coronary artery disease (CAD) in a combined experience in 877 patients, cine CMR performed with a sensitivity ranging from 82% to 87% and a specificity between 80% and 90%.8–15 This is an important set of validations, because dobutamine stress testing poses some of the most extreme challenges for assessing regional wall motion. The diagnostic information must be acquired quickly—usually within about 2 minutes—at a time when the heart rate is 85% of predicted maximum and patients may be experiencing angina. Cine CMR also performed well in detecting acute coronary syndrome in patients presenting to an emergency department with at least 30 minutes of chest pain. Both qualitatively and quantitatively, regional wall motion abnormalities by cine CMR had an accuracy of 82% for acute coronary syndrome, 89% for non-ST elevation myocardial infarction, and 98% for ischemic heart disease.16 One important reason cine CMR performed so well relates to the acuity of the CMR scan in this study. Specifically, 87 of subjects were scanned before the 4-hour troponin was available. By studying patients within 6 hours of presentation to the emergency department, stunned myocardium offers an additional mechanism that is capable of detecting infarction or unstable angina. The statistics summarizing use of cine CMR to detect acute coronary syndrome illustrate an important concept related to the diagnostic utility of using regional wall motion abnormalities. Most importantly, a single study of regional wall motion does not differentiate acute from chronic conditions. In patients with no history of Cardiovascular Magnetic Resonance 241

18 ACUTE MYOCARDIAL INFARCTION: CARDIOVASCULAR MAGNETIC RESONANCE DETECTION AND CHARACTERIZATION

CHAPTER 18

ISCHEMIC HEART DISEASE

End diastole

infarction, a definite regional wall motion abnormality is reasonably specific for CAD even though there are other conditions that can cause regional wall motion abnormalities such as myocarditis. On the other hand, in a patient with a prior infarction, the presence or absence of a regional wall motion abnormality frequently does not alter the clinical approach to a patient. In the latter situation, evidences of inducible ischemia or residual viability become additional pieces of information that are needed to understand how to deal with the regional wall motion abnormality. Fortunately, CMR is well suited to imaging these other important aspects of ischemic heart disease. In summary, cine CMR is a reference standard method of assessing global and regional LV systolic function. Cine CMR is both sensitive and specific for acute coronary syndrome in patients without prior infarction but cannot unequivocally differentiate AMI from chronic infarction. Focal myocarditis may mimic wall motion abnormalities associated with myocardial infarction.

LATE GADOLINIUM ENHANCEMENT TO DETECT ACUTE MYOCARDIAL INFARCTION LGE CMR has become accepted as a reference standard method for imaging myocardial scar. Before discussing applications of LGE, it is important to understand the mechanisms leading to hyperenhancement of AMI and chronic myocardial infarction, as these are somewhat different mechanisms. Finally, it is important to differentiate perfusion abnormalities from viability abnormalities because CMR is capable of differentiating these different clinically relevant scenarios. Current generation gadolinium-based contrast agents are generally a combination of gadolinium with a chelating agent that facilitates excretion and minimizes likelihood of toxic effects from free gadolinium in the body17 (see Chapter 6). These contrast agents are designated as extracellular contrast agents. After intravenous injection, the 242 Cardiovascular Magnetic Resonance

End systole

Figure 18-1 Small dyskinetic wall motion abnormality associated with myocardial infarction that was missed by transthoracic echocardiography. The cine CMR is shown at end diastole and end systole. The arrow points to the wall motion abnormality.

extracellular gadolinium-based contrast agent rapidly diffuses into the interstitial space but is excluded from the intracellular space in healthy tissues. Thus, viable myocardium enhances to a degree based on the amount of gadolinium that is in the intravascular volume and the interstitial space. In contrast, AMI and chronic myocardial infarctions enhance to a much greater degree and thus can be distinguished as regions that are generally brighter than normal myocardium on T1-weighted images obtained typically 10 to 20 minutes after contrast administration.18 A gadolinium dose of 0.1 and 0.2 mmol/kg and a time window of 5 to 30 minutes after contrast administration provide similar results, provided that the TI is adjusted.18a The vascular, interstitial, and intracellular distribution or exclusion is summarized schematically in Figure 18-2. In AMI, nonviable or infarcted cardiomyocytes lose cell membrane integrity, leading to a larger percentage of the myocardium that enhances with gadolinium (see Fig. 18-2) as gadolinium enters what had previously been intracellular space. Thus, there is enhancement in the area of AMI as gadolinium is found in the intravascular space, in the interstitial space, and inside nonviable cells. With chronic myocardial infarction, LGE enhancement is due to the characteristics of collagen scar (see Chapter 1). The collagen scar that forms after myocardial infarction is relatively cellular in comparison with normal myocardium. There is an extensive increase in interstitial collagen deposition but a relatively small intracellular space within fibroblasts. Thus, both AMI and chronic myocardial infarctions enhance to a greater extent than normal myocardium does. Inversion recovery methods have become the de facto standard in obtaining LGE images of myocardial infarction.19 While gadolinium deposition and T1 shortening in the area of infarction had been recognized for many years, the inversion recovery approach provides much higher contrast than previous T1-weighted images did. Based on the physics of inversion recovery, there is an optimal inversion time that results in uniform nulling of normal myocardium (see Chapter 1). Nulling normal myocardium refers to a situation whereby the signal intensity of normal myocardium becomes uniformly very dark. Thus, any abnormalities of excess gadolinium accumulation appear as a bright patch on a dark background, a situation that leads to high sensitivity in detecting myocardial infarction.

LGE: Nulled to define normal T2: Dark

Viable myocardium within area at risk

LGE: Essentially normal T2: Bright

Acutely infarcted myocardium

LGE: Bright T2: Bright

A

B

C

D

Infarcted myocardium with microvascular obstruction

LGE: Dark or paradoxically gray on IR T2: Bright (sometimes dark)

Figure 18-2 In the setting of acute myocardial infarction, different types of myocardium have distinctive appearances on late gadolinium enhancement (LGE) images and T2-weighted images. A, Normal myocardium outside the ischemic area at risk is characterized by healthy cardiomyocytes (striped cylinders) with intact cell membranes and a characteristic intracellular to extracellular ratio and blood volume. The extracellular gadolinium contrast agents (white crosses) arrive via the bloodstream and rapidly enter the interstitial space but are excluded from the intracellular space. Thus, normal myocardium has a shorter T1 after contrast administration, but the amount of enhancement is modest, and the inversion recovery time can be adjusted to make normal myocardium appear uniformly dark (nulled). B, Viable myocardium within the area at risk has intact cell membranes and therefore excludes gadolinium contrast agents from the intracellular space. However, some aspect of the ischemic period reversibly damaged this part of the heart enough to raise the water content and thus lengthen the T2 of the tissue. Therefore, the area at risk is brighter than normal myocardium on T2-weighted images. Because LGE images are essentially normal, there must be relatively balanced increases in the intracellular and extracellular volumes. Thus, the myocardium appears essentially normal on LGE CMR. C, Acutely infarcted myocardium is characterized by loss of cell membrane integrity. Hence, gadolinium contrast agents rapidly enter not only the extracellular space but also what used to be the intracellular space as long as there has been adequate reperfusion. As indicated by the large number of white crosses, acutely infarcted myocardium enhances significantly in comparison to normal myocardium and appears bright on LGE CMR. Acutely infarcted myocardium also appears bright on T2-weighted images, owing to the tissue swelling. D, Infarcted myocardium with microvascular obstruction has a very different appearance. Despite opening of the epicardial vessels, sometimes the microvessels remain occluded (black noentry of white crosses symbols). With severe enough occlusions, the gadolinium contrast agents cannot enter the myocardium, and the T1 remains long. Sometimes it remains long enough that the signal intensity is slightly positive and is gray on conventional magnitude inversion recovery images. In our experience thus far, non reperfused infarcts are bright on T2-weighted images, but there are times where dark patches are present within the otherwise bright edematous zone due to microvascular obstruction or intramyocardial hemorrhage.

VALIDATION OF LATE GADOLINIUM ENHANCEMENT CARDIOVASCULAR MAGNETIC RESONANCE FOR ACUTE MYOCARDIAL INFARCTION LGE CMR is currently the highest-resolution reliable method for detecting myocardial infarction and provides an exquisitely sensitive diagnostic test. Although it was known for many years that myocardial infarctions enhance with gadolinium, there were important advances in methodology,19 basic validations, and clinical testing (Table 181).20–30 These provide important clues as to why LGE CMR is such a powerful clinical tool. As was previously discussed, the spatial localization of gadolinium corresponds closely to areas of both AMI and chronic myocardial infarction and the LGE mass appears highest in patients with Q-wave AMI, intermediate in patients with ST

elevation but non–Q wave AMI, and lowest in patients with non–ST elevation AMI.31 In animal studies in which histopathology can be used as a reference standard for determining what is normal or infarcted myocardium, the bright zones on LGE CMR correlate closely in size and location to both AMI and chronic myocardial infarction.18,32 In patients, AMI size correlates well with the size of myocardial infarction determined by single photon emission tomography (SPECT) imaging.25 Although acute infarcts have not been studied by both positron emission tomography (PET) scans and CMR, there are good correlations between these two modalities in measuring chronic infarct size.33,34 However, investigators have noted that small subendocardial infarcts are routinely missed on SPECT and PET scans, owing to issues related to spatial resolution.33–35 One must recognize that it is possible for a patient to have an AMI despite a “normal” SPECT scan. Very small myocardial infarctions (<1 g) associated with side branch occlusions at the time of percutaneous interventions are easily detectable by LGE CMR21 but may be missed by SPECT.36 Among patients with inferior AMI, LGE CMR detects right ventricular (RV) infarction more frequently than transthoracic Cardiovascular Magnetic Resonance 243

18 ACUTE MYOCARDIAL INFARCTION: CARDIOVASCULAR MAGNETIC RESONANCE DETECTION AND CHARACTERIZATION

Normal myocardium

ISCHEMIC HEART DISEASE

Table 18-1 Validations of Late Gadolinium Enhancement Cardiovascular Magnetic Resonance in Acute Myocardial Infarction

Year

Authors

N

Acute Versus Chronic

2001

Ricciardi MJ et al.

14 6

Acute Chronic

Microinfarcts were detected in patients who had PCI-related elevations in CKMB. Two patterns of MI were observed: small side branch occlusion and distal embolization.

2001

Choi KM et al.

24

Acute Chronic

Within 7 days of AMI, MI size correlates with peak CKMB. There is an inverse relationship between the transmural extent of MI and the likelihood of regional recovery of function.

2002

Gerber BL et al.

20

Acute Chronic

In patients 4 days after AMI, LGE predicted recovery of regional function better than did perfusion images, which significantly underestimated the amount of irreversible injury.

2003

Beek AM et al.

30

Acute Chronic

In patients 1 week after AMI, there was an inverse relationship between the transmural extent of infarction and the likelihood of regional recovery of function.

2004

Ingkanisorn WP et al.

33 20

Acute Chronic

In patients 2 days after acute PCI, MI size correlated with peak troponin. AMI size predicted chronic LVEF and regional function. Infarct size decreased from 16% of the LV to 11% of the LV.

2004

Lund GK et al.

60

Acute

SPECT and CMR infarct size correlated, but SPECT had 80% sensitivity and CMR had 100% sensitivity. SPECT missed 6 of 30 inferior AMI.

2005

Ibrahim T et al.

33

Acute

The extent of AMI was reasonably stable from 7 to 42 minutes post-contrast and correlated well with SPECT infarct size.

2005

Baks T et al.

22

Acute Chronic

LGE predicted recovery of function after AMI better than perfusion imaging.

2005

Selvanayagam JB et al.

50 24

Acute Chronic

All patients with a troponin elevation (average of 3.7 mg/L) associated with PCI exhibited new abnormal LGE. Troponin level correlated with AMI size.

2005

Ichikawa et al.

18

Acute Chronic

Thickness of nonenhanced myocardium compared with percent transmural enhancement had better diagnostic accuracy for predicting improvement in systolic wall thickening.

2006

Baks T et al.

27

Acute Chronic

LGE could predict dysfunctional segments distal to chronic total occlusions that improved after revascularization.

2006

Gerber BL

16 21

Acute Chronic

MI size by CT and CMR correlated well (r ¼ 0.89), although the 2 standard deviation limits of agreement were about 35 g. Interobserver agreement and intrastudy agreement were good.

2006

Kumar

37

Acute

Among patients with inferior AMI, LGE CMR detects right ventricular infarction more frequently than does echocardiography or electrocardiography for right ventricular infarction.

2007

Ibrahim et al.

78

Acute

AMI was detected more often by CMR than by SPECT, specifically for non–Q-wave AMI and for the left circumflex territory.

Plein

25

Acute

Total

477

Acute

LGE mass is highest in Q-wave AMI, intermediate in non–Q ST elevation AMI, and lowest in non-ST elevation AMI.

2008

Major Findings

AMI, acute myocardial infarction; CKMB, creatine kinase MB; CMR, cardiovascular magnetic resonance; CT, computed tomography; LGE, late gadolinium enhancement; LV, left ventricle; LVEF, left ventricular ejection fraction; MI, myocardial infarction; PCI, percutaneous coronary intervention; SPECT, single photon emission computed tomography.

echocardiography and electrocardiography do.37 From studies like these, we can conclude that LGE CMR is very sensitive to infarction and capable of detecting small myocardial infarctions, even those that are not identified by other imaging methods. Measurements of thickness of nonenhanced myocardium compared with measurements of percent transmural enhancement also provide prognostic accuracy for predicting improved systolic wall thickening following AMI.38 Physiologic correlations to the spatial distribution of gadolinium have been studied at high resolution by electron probe X-ray microanalysis and high-resolution nuclear methods. In both AMI and chronic myocardial infarction, gadolinium concentration in the heart tracks the spatial distribution of sodium and follows a distribution that is inversely related to potassium.39 This distribution is exactly what one would expect for an extracellular contrast agent and forms the basis of gadolinium distribution outlined in Figure 18-2. At similar 244 Cardiovascular Magnetic Resonance

resolution, it has been possible to map the spatial distribution of the chelate used in one of the gadolinium contrast agents and compare the spatial distribution of gadolinium-DTPA with that of sestamibi and FDG.40 The subendocardial vulnerability to ischemia provides an important pathophysiologic basis for distinguishing myocardial infarction from other causes of LGE.41 When the resolution and quality of LGE CMR methods became good enough to routinely determine the transmural extent of infarction in patients,42 the subendocardial vulnerability to ischemia and infarction became an important diagnostic characteristic relative to other causes of gadolinium accumulation in the heart.43 Thus, LGE CMR is a valuable clinical tool for detecting AMI. This method has been extensively validated in both animal and human studies. In general, LGE CMR performs as well as or better than alternative clinical imaging methods for AMI (see Table 18-1).

The degree of LGE enhancement is modulated by the reperfusion status of the myocardium and the degree of residual coronary artery stenosis. Postinfarct perfusion status modulates the wash-in and wash-out of contrast in the infarcted myocardium44 and thus affects the appearance of LGE images as a function of time. In the simplest case in which a coronary artery is occluded for a period of time followed by successful reperfusion with no residual perfusion defect, the simple descriptions in the previous paragraphs are adequate to understand the relative signal intensity on LGE CMR. However, a complete understanding of contrast enhancement after AMI needs to consider whether the coronary artery has been reperfused, whether there is a residual coronary artery stenosis, the amount of collateral blood flow to the ischemic bed, and the status of myocardial perfusion at a microvascular level. In a nonreperfused myocardial infarction, T1-weighted images obtained during the arrival of contrast into the heart show a dense perfusion defect or dark zone relative to normal myocardium. In fact, it is possible to detect the perfusion defect even in patients presenting to an emergency department before serum biomarkers of infarction become abnormal (Fig. 18-3). In our experience, the perfusion defects associated with nonreperfused AMI are more severe and last longer during the first pass perfusion study than do stress-induced perfusion defects, which are very transient phenomena.

Collateral blood flow can bring gadolinium contrast into the infarcted region as long as circulating levels of gadolinium remain high. During the first pass after bolus gadolinium administration, vascular pool contrast agent concentration is highest. Contrast rapidly leaves the intravascular space and enters the interstitial space.17 Because gadolinium clears from the blood pool with a half-life of about 60 to 90 minutes, collateral blood flow can cause a perfusion defect to gradually fill in over a several-minute time span after contrast administration. Despite successfully opening the epicardial coronary artery during percutaneous coronary interventions for AMI, a perfusion defect due to microvascular obstruction is frequently observed (Fig. 18-4). Microvascular obstruction is defined as a residual perfusion defect despite an open epicardial coronary artery. This condition is also widely known as the “no-reflow phenomenon.”45 Several mechanisms can lead to microvascular obstruction, including (1) thrombosis of microvessels, (2) embolization of microvessels, (3) white blood cell plugging of microvessels, (4) compression of microvessels due to tissue swelling, and (5) decreased numbers of functional capillaries. In some patients, the perfusion defect associated with microvascular obstruction will be so severe that the ischemic myocardium shows virtually no enhancement even for many minutes after injection of contrast. In severe microvascular obstruction, the contrast arrival is so slow that regions of the AMI remain hypoenhanced or dark for a prolonged period of time.44 The kinetics of enhancement of AMI with microvascular obstruction are thus very different from the kinetics associated with wellreperfused infarcts. In microvascular obstruction, the signal intensity after contrast delivery is dominated by delayed wash-in or poor residual perfusion. This is quite different from well-reperfused AMI, in which delivery of contrast may be close to normal but wash-out of contrast is delayed relative to normal myocardium and the volume of distribution is higher than normal. Doppler flow wire measurements suggests that early systolic retrograde flow is more common in patients with AMI and CMR complicated by microvascular obstruction.46

Cine

Rest perfusion

Figure 18-3 Perfusion defect (red arrowheads) associated with acute myocardial infarction in a patient with no prior coronary disease. This patient presented to the emergency department within the first hour after onset of chest pain. CMR was performed within 1 hour after arrival in the hospital, and the initial electrocardiogram and troponin-I were normal. The diagnosis of acute myocardial infarction was subsequently confirmed by troponin-I assays 8 hours after presentation (6 hours after CMR) and by coronary angiography. Multiple parallel short axis images allow determination of the amount of myocardium at risk. Cardiovascular Magnetic Resonance 245

18 ACUTE MYOCARDIAL INFARCTION: CARDIOVASCULAR MAGNETIC RESONANCE DETECTION AND CHARACTERIZATION

PERFUSION STATUS MODULATES THE SIGNAL INTENSITY OF ACUTE MYOCARDIAL INFARCTION: RESIDUAL CORONARY OCCLUSION VERSUS MICROVASCULAR OBSTRUCTION

ISCHEMIC HEART DISEASE

Cine

Perfusion

Delayed enhancement

Figure 18-4 Appearance of acute myocardial infarction complicated by severe microvascular obstruction. The signal intensity of myocardium appears relatively uniform on the cine image in this acute anteroseptal myocardial infarction. A first pass perfusion image reveals a dark zone consistent with low perfusion (red arrow). Severe microvascular obstruction is present, because the perfusion defect is still present on LGE CMR 10 minutes after the injection of contrast. As is often seen, the dark patch of microvascular obstruction is surrounded by a rim of bright infarcted myocardium. One can also see that the resolution of the LGE image and overall image quality is better than the perfusion image.

In summary, there is more than one pattern of contrast enhancement in acute myocardial infarction. The temporal course of contrast enhancement is dominated by contrast wash-in, wash-out, and the volume of distribution within the myocardium. In general, AMI will enhance to a greater extent than normal myocardium on LGE CMR obtained about 10 to 20 minutes after contrast administration. Even in cases of microvascular obstruction in which some portions of the infarct may appear hypoenhanced, there will usually be a rim of hyperenhanced myocardium around the dark central core when imaging is performed at this late period.

PROGNOSTIC SIGNIFICANCE OF MICROVASCULAR OBSTRUCTION AFTER MYOCARDIAL INFARCTION Microvascular obstruction portends poor prognosis and adverse postinfarction remodeling. Wu and colleagues were among the first to report that LGE evidence of microvascular obstruction was associated with adverse clinical outcome.47 In a study of 44 patients with AMI, they found that microvascular obstruction was associated with an increased rate of combined cardiovascular end points during an average 16-month follow-up. Although microvascular obstruction was associated with larger infarcts, the presence of microvascular obstruction was independently associated with adverse outcomes in multivariable analysis. In a study of 110 patients with AMI, Hombach and colleagues found that the presence of microvascular obstruction had prognostic significance while infarct size did not.48 One surprising aspect of this study relates to the small amount of myocardium affected by microvascular obstruction. While LGE infarct size averaged 12% of the LV myocardium, microvascular obstruction averaged less than 3% of the myocardium. LV ejection fraction shortly after acute intervention did not predict adverse prognosis well. LV systolic function measured early after AMI is 246 Cardiovascular Magnetic Resonance

abnormal, owing to a combination of stunned and infarcted myocardium. In an era in which many patients receive acute percutaneous or thrombolytic interventions, the amount of myocardium that is permanently injured seems to predict future events. So although it is not clear why microvascular obstruction affecting such a small amount of myocardium should have such powerful prognostic value, it is an important clinical tool, and the imaging is relatively simple to perform. The CMR relationships between postinfarct microvascular obstruction and prognosis mirror similar observations in larger clinical trials. In a thrombolytic-treated population with AMI, Gibson and colleagues found that TIMI frame count evidence of microvascular obstruction was associated with a worse 2-year outcome.49 Myocardial blush grade, an angiographic measure of perfusion, was able to predict worse prognosis after acute percutaneous interventions for myocardial infarction even in patients with TIMI 3 flow.50 Thus, a broad range of clinical perspectives indicate that microvascular obstruction is clinically important. Furthermore, there is evidence that subtler degrees of microvascular obstruction are relevant and that LGE CMR is a sensitive way of identifying and quantifying microvascular obstruction.

MICROVASCULAR OBSTRUCTION ALSO PREDICTS REGIONAL RECOVERY OF FUNCTION AFTER ACUTE MYOCARDIAL INFARCTION Although the transmural extent of LGE predicts the likelihood that regional myocardial function will improve after AMI and successful revascularization or intervention, the presence of microvascular obstruction portends more severe adverse remodeling and worse recovery of function. Considering that many studies have equated LGE with infarcted or scarred myocardium, it seems counterintuitive

Figure 18-5 Adenosine stress perfusion defect (three red arrows) clearly extends beyond the edges of the very small inferior myocardial infarction on the late gadolinium enhancement image.

Cine

ROLE OF STRESS PERFUSION IMAGING AFTER ACUTE MYOCARDIAL INFARCTION After a patient with AMI is stabilized, it is sometimes necessary to determine whether parts of the heart have inducible ischemia. CMR allows acquiring perfusion and viability imaging in the same imaging planes and image orientation (see Chapters 15 and 16). This facilitates comparisons of complementary clinical data. As Figure 18-5 illustrates, the stress-induced perfusion defect is more extensive than the infarct in a patient who presented to the emergency department with chest pain. This case illustrates that managing patients with CAD requires more information than simply detecting disease. This patient had known CAD and presented with prolonged chest pain. Serial troponin levels were normal. The adenosine stress perfusion images are consistent with peri-infarct-inducible ischemia. At cardiac catheterization, she had a new severe stenosis in the left circumflex coronary artery that represented progression of disease from known prior problems in the right coronary artery. Thus, stress perfusion imaging takes the postinfarct CMR evaluation one step further than a rest study. The vasodilator stress test allows detecting perfusion defects beyond the infarcted myocardium.

USE OF T2-WEIGHTED CARDIOVASCULAR MAGNETIC RESONANCE OF ACUTE MYOCARDIAL INFARCTION Many current clinical decisions after AMI can be guided by information gathered during cine, perfusion, and LGE CMR. However, T2-weighted imaging provides additional characterization of the myocardium that may be useful in specific patients. Most imaging tests have difficulty differentiating AMI from chronic myocardial infarction. This is generally true for echocardiography, SPECT, and PET. CMR using cine CMR, perfusion, and LGE also cannot reliably differentiate

Late gadolinium enhancement

Adenosine perfusion

Cardiovascular Magnetic Resonance 247

18 ACUTE MYOCARDIAL INFARCTION: CARDIOVASCULAR MAGNETIC RESONANCE DETECTION AND CHARACTERIZATION

that microvascular obstruction should have further predictive value over an assessment of viability. This question is not fully resolved, but the explanation may lie in a few possibilities. First, any imaging method is limited by partial volume errors. Because each pixel in the image represents a small volume of tissue that is thousands of times larger than a cardiomyocyte, each pixel could represent a mixture of viable and nonviable cells. Thus, pixels that are statistically brighter than normal myocardium may include more viable cells than nonviable cells. This idea is supported by analyses that indicate that a threshold of 50% between bright and dark pixels corresponds most closely to infarcted myocardium.51–53 This problem could be amplified at the infarct edge if the borders are convoluted rather than smooth and could result in partial volume problems affecting more than a single pixel at the edge of some infarcts but not all infarcts. Second, microvascular obstruction tends to occur in more severe infarcts with later reperfusion.54 Thus, microvascular obstruction tends to be associated with more transmural infarcts. Third, the severity of a residual perfusion defect associated with microvascular obstruction may alter postinfarct healing or remodeling. Severe microvascular obstruction may be associated with more thinning of the myocardium in the infarct zone29 and thus worse peak systolic strain, further impairing regional wall thickening.55–57 Fourth, severe microvascular obstruction may compromise the survival of residual viable cardiomyocytes due to apoptosis, whereas in a well-reperfused infarct, any residual cardiomyocytes may hypertrophy and compensate for the acute loss in regional wall thickening. Finally, changes in regional function may be affected by the border zone around the infarct. In a well-reperfused infarct, the ischemic zone and acutely stunned myocardium may be much larger in circumferential extent than the ultimate infarct size. As a result of improved ventricular geometry, the regional afterload in the infarcted segments may improve sufficiently to allow some recovery of wall thickening. In summary, residual microvascular obstruction has an additional clinical value that makes it superior to simple LGE alone. Microvascular obstruction means that a segment is less likely to recover regional wall thickening, and a patient with residual microvascular obstruction has a worse prognosis than does someone with a similar-sized (LGE) infarct but no residual perfusion defect.

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AMI from chronic infarction. In the case of a patient with both an AMI and a prior myocardial infarction, clinicians often have to resort to the electrocardiogram and the appearance of the coronary angiograms to determine which vessel was acutely occluded. Although this combined clinical approach works in most patients, it does not work well in many patients, particularly in non-ST elevation myocardial infarction and patients with left bundle branch blocks. T2-weighted CMR can uniquely differentiate AMI from chronic myocardial infarction and from normal myocardium on the basis of differences in signal intensity. In general, CMR detects protons. Because about 70% of the body is water, water protons dominate the CMR signal, and fat protons are the second most important source of CMR signal. Just as soft tissue such as skin and muscle swell after mechanical injuries, myocardium swells after acute myocardial ischemia and infarction. The extra water in the myocardium changes the mobility of water molecules enough that the T2 of acutely ischemic myocardium increases. We can image such changes in myocardial water content as a bright region on T2-weighted cardiac CMR (Fig. 18-6). Although many early studies showed that T2-weighted abnormalities were present in AMI, image quality, image artifacts, and subtlety of the findings precluded significant clinical application. Abdel-Aty and colleagues published the first large series of patients in which T2-weighted images were used to differentiate AMI from chronic myocardial infarction.58 With very high sensitivity and specificity, they were able to determine the infarct-related artery and differentiate AMI from chronic myocardial infarction. Subsequently, Cury and colleagues demonstrated the combination of T2-weighted CMR and LV wall thickness to have increased specificity, positive predictive value, and overall accuracy (in comparison to cine, perfusion, and LGE CMR) for detection of AMI.59 The observation that the T2-weighted abnormality associated with AMI was almost always nearly transmural, while the LGE abnormality was usually subendocardial,60 indicated that the T2-weighted images were detecting some injury more extensive than the nonviable myocardium. Studies have correlated the size of the T2-weighted abnormality with the ischemic area at risk and showed that the T2-weighted abnormality was significantly larger than the amount of infarcted myocardium.61,62 The fact that T2-weighted abnormalities persist for many days after an AMI means that it should be possible to assess both infarct size and the ischemic area at risk with CMR a few days after

Cine

Late gadolinium enhancement

AAo LA

LV

248 Cardiovascular Magnetic Resonance

the acute event. This may prove to be much more convenient than radionuclide studies performed during AMI. In summary, T2-weighted CMR can differentiate AMI from chronic myocardial infarction. T2-weighted CMR shows the ischemic area at risk as brighter than normal myocardium. T2-weighted imaging is complementary to LGE CMR, as each method assesses different aspects of the response to AMI: recent ischemia versus cell death.

CORONARY ARTERY CARDIOVASCULAR MAGNETIC RESONANCE In a time when multislice computed tomography provides remarkably high-resolution three-dimensional images of the coronary arteries in a single breath hold, coronary artery CMR is more lengthy. In a series of patients with probable acute coronary syndrome studied by CMR and by coronary angiography, adenosine stress CMR perfusion had the highest sensitivity and specificity of any single component of the exam. Coronary artery CMR was a close second.63 Coronary artery CMR had higher diagnostic accuracy than did both cine CMR and LGE. Thus, coronary artery CMR may have a role at some time in the future, particularly as methods continue to improve.

COMPLICATIONS OF ACUTE MYOCARDIAL INFARCTION Echocardiography is the first-line test for assessing most complications of AMI. In experienced hands, CMR can provide high-quality images of mechanical complications of AMI and may be more sensitive than echocardiography for certain problems. LV thrombus is important to detect, owing to the risk of systemic embolism and the need for more aggressive anticoagulation. Although transthoracic echocardiography is more sensitive than invasive left ventriculography, clinical experience suggests that LGE CMR is more sensitive than is transthoracic echocardiography with and without echocardiographic contrast.64 In general, LV thrombus does not enhance early or late after administration of gadolinium-based contrast agents (Fig. 18-7).

T2-weighted black blood

Figure 18-6 Appearance of acute anteroseptal myocardial infarction on cine CMR (left), LGE (middle), and T2-weighted images (right). Note that the LGE CMR shows a thin rim of viable dark myocardium on the right ventricular side of the interventricular septum (red arrows). The bright zone on the T2-weighted black-blood images is nearly transmural in this entire distribution. AAo, aortic arch; LA, left atrium; LV, left ventricle.

Late gadolinium enhancement

Cine

LV

LV

LA

CMR can also detect mechanical complications after AMI. Ventricular aneurysm and pseudoaneurysm can readily be identified and characterized by CMR (Fig. 18-8). Infarctrelated ventricular septal defect can be identified and characterized with velocity-encoded phase contrast cine CMR Figure 18-8 Pseudoaneurysm of the basal lateral wall after acute myocardial infarction shown in the four-chamber view. The pseudoaneurysm is partially filled with a thrombus (arrow on LGE) but not seen on the cine image, owing to a slightly different image plane. LA, left atrium; LV, left ventricle; RA, right atrium; RV, right ventricle.

LA

(Fig. 18-9). Rupture of papillary muscles can be detected with a combination of cine CMR, phase contrast CMR, and LGE CMR. Detailed CMR images of these complications may help to guide surgical repair of these uncommon but serious complications of acute myocardial infarction.

Cine

Late gadolinium enhancement

RV LV RA

LA

Cine

Late gadolinium enhancement

Magnitude image (PC)

Velocity PC (y)

*

Figure 18-9 Ventricular septal defect 1 day after percutaneous intervention for large anterior and septal acute myocardial infarction. The mid to apical septum already shows an aneurysmal contour with severe thinning of some portions of the septum on cine CMR (left). The septum and apex are transmurally infarcted on LGE (green arrows) with dark patches consistent with microvascular obstruction (second panel ). The magnitude images from the velocity-encoded phase contrast (PC) images show a dark jet consistent with turbulent flow through the interventricular septum (red arrow). On the velocity-encoded images, the red arrow points to a white patch that appears to indicate that the left-to-right flow angles toward the apex in this image, where white represents high velocities toward the apex, gray is zero velocity, and black means high velocity toward the base of the heart. The asterisk indicates systolic blood flow through the left ventricular outflow tract. Cardiovascular Magnetic Resonance 249

18 ACUTE MYOCARDIAL INFARCTION: CARDIOVASCULAR MAGNETIC RESONANCE DETECTION AND CHARACTERIZATION

Figure 18-7 Left ventricular apical thrombus associated with acute anterior myocardial infarction shown in the twochamber view (left: end-systolic frame from cine CMR; right: LGE images). The arrow points to the apical thrombus. LA, left atrium; LV, left ventricle.

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Table 18-2 Appropriateness of Indications for CMR Appropriateness Score*

Description of Indication of Cardiac MRI 1. Evaluation of LV function following myocardial infarction OR in heart failure patients with technically limited images from echocardiogram. 2. Quantification of LV function when there is discordant information that is clinically significant from prior tests. 3. Evaluation of LV function following myocardial infarction OR in heart failure patients, since they should be evaluable by echocardiography. 4. Intermediate pre-test probability of CAD in a patient presenting to an emergency department with no ECG changes and serial cardiac enzymes negative 4. Evaluation of myocarditis or myocardial infarction with normal coronary arteries in the setting of positive cardiac enzymes without obstructive atherosclerosis on angiography. 5. Evaluation of cardiac mass (suspected tumor or thrombus) with the use of contrast for perfusion and enhancement. 6. To determine the location and extent of myocardial necrosis including “no reflow” regions post-acute myocardial infarction. 7. To detect post PCI myocardial necrosis. 8. To determine viability prior to revascularization to establish likelihood of recovery of function with revascularization (PCI or CABG) or medical therapy. 9. To determine viability prior to revascularization when viability assessment by SPECT or dobutamine echo has provided “equivocal or indeterminate“ results. 10. Evaluation of post-infarct complications including aneurysm, ventricular septal defect, ruptured/infarcted papillary muscle, myocardial rupture, pericardial effusion in cases where diagnosis remains uncertain after echocardiography or additional information is needed to plan surgical repair.

8 8 6 6 8 9 7 6 9 9 Not assessed

CABG, coronary artery bypass graft surgery; CAD, coronary artery disease; ECG, electrocardiography; LV, left ventricular; PCI, percutaneous coronary intervention; SPECT, single photon emission computed tomography. *The appropriateness score was adopted from Hendel RC, Patel MR, Kramer CM, Poon M, Carr JC, Gerstad NA, Gillam LD, Hodgson JM, Kim RJ, Lesser JR, et al. ACCF/ACR/SCCT/ SCMR/ASNC/NASCI/SCAI/SIR 2006 appropriateness criteria for cardiac computed tomography and cardiac magnetic resonance imaging: a report of the American College of Cardiology Foundation Quality Strategic Directions Committee Appropriateness Criteria Working Group, American College of Radiology, Society of Cardiovascular Computed Tomography, Society for Cardiovascular Magnetic Resonance, American Society of Nuclear Cardiology, North American Society for Cardiac Imaging, Society for Cardiovascular Angiography and Interventions, and Society of Interventional Radiology. J Am Coll Cardiol. 2006;48(7):1475–1497. The numbering of each indication was revised, since the original paper included many indications not applicable to acute myocardial infarction or complications of acute myocardial infarction. Indications were also reworded slightly to improve readability. Appropriateness was graded on a 10 point scale by a board of experts. Average scores of 1–3 were considered inappropriate; scores of 4–6 were considered of uncertain appropriateness; scores of 7–10 were considered appropriate. “Not assessed” means that the committee did not consider the indication.

CONCLUSION CMR offers a wide range of imaging methods that are suitable for detecting AMI and assessing many clinically important questions. The combination of cine, perfusion, and LGE CMR is a powerful and useful test. Some patients may benefit from more detailed imaging with velocity-

encoded CMR or T2-weighted images. Table 18-2 provides a summary of reasonable clinical indications for CMR after AMI, which are a subset of recently published appropriateness criteria.65 It is important to note that postinfarct patients must be imaged efficiently and quickly, so significant CMR experience is important for high-quality clinical results.

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11. Nagel E, Lehmkuhl HB, Bocksch W, Klein C, Vogel U, Frantz E, Ellmer A, Dreysse S, Fleck E. Noninvasive diagnosis of ischemiainduced wall motion abnormalities with the use of high-dose dobutamine stress MRI: comparison with dobutamine stress echocardiography. Circulation. 1999;99:763–770. 12. Paetsch I, Jahnke C, Ferrari VA, Rademakers FE, Pellikka PA, Hundley WG, Poldermans D, Bax JJ, Wegscheider K, Fleck E. Determination of interobserver variability for identifying inducible left ventricular wall motion abnormalities during dobutamine stress magnetic resonance imaging. Eur Heart J. 2006;27:1459–1464. 13. Paetsch I, Jahnke C, Wahl A, Gebker R, Neuss M, Fleck E, Nagel E. Comparison of dobutamine stress magnetic resonance, adenosine stress magnetic resonance, and adenosine stress magnetic resonance perfusion. Circulation. 2004;110:835–842. 14. Wahl A, Paetsch I, Gollesch A, Roethemeyer S, Foell D, Gebker R, Langreck H, Klein C, Fleck E, Nagel E. Safety and feasibility of high-dose dobutamine-atropine stress cardiovascular magnetic resonance for diagnosis of myocardial ischaemia: experience in 1000 consecutive cases. Eur Heart J. 2004;25:1230–1236. 15. Wahl A, Paetsch I, Roethemeyer S, Klein C, Fleck E, Nagel E. Highdose dobutamine-atropine stress cardiovascular MR imaging after coronary revascularization in patients with wall motion abnormalities at rest. Radiology. 2004;233:210–216. 16. Kwong RY, Schussheim AE, Rekhraj S, Aletras AH, Geller N, Davis J, Christian TF, Balaban RS, Arai AE. Detecting acute coronary syndrome in the emergency department with cardiac magnetic resonance imaging. Circulation. 2003;107:531–537. 17. Bellin MF. MR contrast agents, the old and the new. Eur J Radiol. 2006;60:314–323. 18. Kim RJ, Fieno DS, Parrish TB, Harris K, Chen EL, Simonetti O, Bundy J, Finn JP, Klocke FJ, Judd RM. Relationship of MRI delayed contrast enhancement to irreversible injury, infarct age, and contractile function. Circulation. 1999;100:1992–2002. 18a. Wagner A, Mahrholdt H, Thomson L, Hager S, Meinhardt G, Rehwald W, Parker M, Shah D, Sechtem U, Kim RJ. Effects of time, dose, and inversion time for acute myocardial infarct size measurements based on magnetic resonance imaging-delayed contrast enhancement. J Am Coll Cardiol. 2006;47:2027–2033. 19. Simonetti OP, Kim RJ, Fieno DS, Hillenbrand HB, Wu E, Bundy JM, Finn JP, Judd RM. An improved MR imaging technique for the visualization of myocardial infarction. Radiology. 2001;218:215–223. 20. Choi KM, Kim RJ, Gubernikoff G, Vargas JD, Parker M, Judd RM. Transmural extent of acute myocardial infarction predicts longterm improvement in contractile function. Circulation. 2001;104: 1101–1107. 21. Ricciardi MJ, Wu E, Davidson CJ, Choi KM, Klocke FJ, Bonow RO, Judd RM, Kim RJ. Visualization of discrete microinfarction after percutaneous coronary intervention associated with mild creatine kinase-MB elevation. Circulation. 2001;103:2780–2783. 22. Gerber BL, Garot J, Bluemke DA, Wu KC, Lima JA. Accuracy of contrast-enhanced magnetic resonance imaging in predicting improvement of regional myocardial function in patients after acute myocardial infarction. Circulation. 2002;106:1083–1089. 23. Beek AM, Kuhl HP, Bondarenko O, Twisk JW, Hofman MB, van Dockum WG, Visser CA, van Rossum AC. Delayed contrast-enhanced magnetic resonance imaging for the prediction of regional functional improvement after acute myocardial infarction. J Am Coll Cardiol. 2003;42:895–901. 24. Ingkanisorn WP, Rhoads KL, Aletras AH, Kellman P, Arai AE. Gadolinium delayed enhancement cardiovascular magnetic resonance correlates with clinical measures of myocardial infarction. J Am Coll Cardiol. 2004;43:2253–2259. 25. Lund GK, Stork A, Saeed M, Bansmann MP, Gerken JH, Muller V, Mester J, Higgins CB, Adam G, Meinertz T. Acute myocardial infarction: evaluation with first-pass enhancement and delayed enhancement MR imaging compared with 201Tl SPECT imaging. Radiology. 2004;232:49–57. 26. Baks T, van Geuns RJ, Biagini E, Wielopolski P, Mollet NR, Cademartiri F, Boersma E, van der Giessen WJ, Krestin GP, Duncker DJ. Recovery of left ventricular function after primary angioplasty for acute myocardial infarction. Eur Heart J. 2005;26:1070–1077. 27. Ibrahim T, Nekolla SG, Hornke M, Bulow HP, Dirschinger J, Schomig A, Schwaiger M. Quantitative measurement of infarct size by contrastenhanced magnetic resonance imaging early after acute myocardial infarction: comparison with single-photon emission tomography using Tc99m-sestamibi. J Am Coll Cardiol. 2005;45:544–552.

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46. Hirsch A, Nijveldt R, Haeck JDE, Beek AM, Koch KT, Henriques JP, van der Schaaf RJ, Vis MM, Baan J, Jr., de Winter RJ. Relation between the assessment of microvascular injury by cardiovascular magnetic resonance and coronary Doppler flow velocity measurements in patients with acute anterior wall myocardial infarction. J Am Coll Cardiol. 2008;51:2230–2238. 47. Wu KC, Zerhouni EA, Judd RM, Lugo-Olivieri CH, Barouch LA, Schulman SP, Blumenthal RS, Lima JA. Prognostic significance of microvascular obstruction by magnetic resonance imaging in patients with acute myocardial infarction. Circulation. 1998;97:765–772. 48. Hombach V, Grebe O, Merkle N, Waldenmaier S, Hoher M, Kochs M, Wohrle J, Kestler HA. Sequelae of acute myocardial infarction regarding cardiac structure and function and their prognostic significance as assessed by magnetic resonance imaging. Eur Heart J. 2005;26: 549–557. 49. Gibson CM, Cannon CP, Murphy SA, Marble SJ, Barron HV, Braunwald E. Relationship of the TIMI myocardial perfusion grades, flow grades, frame count, and percutaneous coronary intervention to long-term outcomes after thrombolytic administration in acute myocardial infarction. Circulation. 2002;105:1909–1913. 50. Henriques JP, Zijlstra F, van’t Hof AW, de Boer MJ, Dambrink JH, Gosselink M, Hoorntje JC, Suryapranata H. Angiographic assessment of reperfusion in acute myocardial infarction by myocardial blush grade. Circulation. 2003;107:2115–2119. 51. Amado LC, Gerber BL, Gupta SN, Rettmann DW, Szarf G, Schock R, Nasir K, Kraitchman DL, Lima JA. Accurate and objective infarct sizing by contrast-enhanced magnetic resonance imaging in a canine myocardial infarction model. J Am Coll Cardiol. 2004;44:2383–2389. 52. Hsu LY, Ingkanisorn WP, Kellman P, Aletras AH, Arai AE. Quantitative myocardial infarction on delayed enhancement MRI. Part II: Clinical application of an automated feature analysis and combined thresholding infarct sizing algorithm. J Magn Reson Imaging. 2006;23:309–314. 53. Hsu LY, Natanzon A, Kellman P, Hirsch GA, Aletras AH, Arai AE. Quantitative myocardial infarction on delayed enhancement MRI. Part I: Animal validation of an automated feature analysis and combined thresholding infarct sizing algorithm. J Magn Reson Imaging. 2006; 23:298–308. 54. Tarantini G, Cacciavillani L, Corbetti F, Ramondo A, Marra MP, Bacchiega E, Napodano M, Bilato C, Razzolini R, Iliceto S. Duration of ischemia is a major determinant of transmurality and severe microvascular obstruction after primary angioplasty: a study performed with contrast-enhanced magnetic resonance. J Am Coll Cardiol. 2005;46:1229–1235. 55. Choi CJ, Haji-Momenian S, Dimaria JM, Epstein FH, Bove CM, Rogers WJ, Kramer CM. Infarct involution and improved function during healing of acute myocardial infarction: the role of microvascular obstruction. J Cardiovasc Magn Reson. 2004;6: 917–925. 56. Kramer CM. When two tests are better than one: adding late gadolinium enhancement to first-pass perfusion cardiovascular magnetic resonance. J Am Coll Cardiol. 2006;47:1639–1640.

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57. Rogers WJ, Jr., Kramer CM, Geskin G, Hu YL, Theobald TM, Vido DA, Petruolo S, Reichek N. Early contrast-enhanced MRI predicts late functional recovery after reperfused myocardial infarction. Circulation. 1999;99:744–750. 58. Abdel-Aty H, Zagrosek A, Schulz-Menger J, Taylor AJ, Messroghli D, Kumar A, Gross M, Dietz R, Friedrich MG. Delayed enhancement and T2-weighted cardiovascular magnetic resonance imaging differentiate acute from chronic myocardial infarction. Circulation. 2004;109:2411–2416. 59. Cury RC, Shash K, Nagurney JT, Rosito G, Shapiro MD, Nomura CH, Abbara S, Bamberg F, Ferencik M, Schmidt EJ. Cardiac magnetic resonance with T2-weighted imaging improves detection of patients with acute coronary syndrome in the emergency department. Circulation. 2008;118:837–844. 60. Friedrich MG, Abdel-Aty H, Taylor A, Schulz-Menger J, Messroghli D, Dietz R. The salvaged area at risk in reperfused acute myocardial infarction as visualized by cardiovascular magnetic resonance. J Am Coll Cardiol. 2008;51:1581–1587. 61. Aletras AH, Tilak GS, Natanzon A, Hsu LY, Gonzalez FM, Hoyt RF, Jr., Arai AE. Retrospective determination of the area at risk for reperfused acute myocardial infarction with T2-weighted cardiac magnetic resonance imaging: histopathological and displacement encoding with stimulated echoes (DENSE) functional validations. Circulation. 2006;113:1865–1870. 62. Garcia-Dorado D, Oliveras J, Gili J, Sanz E, Perez-Villa F, Barrabes J, Carreras MJ, Solares J, Soler-Soler J. Analysis of myocardial oedema by magnetic resonance imaging early after coronary artery occlusion with or without reperfusion. Cardiovasc Res. 1993; 27:1462–1469. 63. Plein S, Greenwood JP, Ridgway JP, Cranny G, Ball SG, Sivananthan MU. Assessment of non-ST-segment elevation acute coronary syndromes with cardiac magnetic resonance imaging. J Am Coll Cardiol. 2004;44:2173–2181. 64. Weinsaft JW, Kim HW, Shal DJ, Klem I, Crowley AL, Brosnan R, James OG, Patel MR, Heitner J, Parker M. Detection of left ventricular thrombus by delayed enhancement cardiovascular magnetic resonance prevalence and markers of systolic dysfunction. J Am Coll Cardiol. 2008;52:148–157. 65. Hendel RC, Patel MR, Kramer CM, Poon M, Carr JC, Gerstad NA, Gillam LD, Hodgson JM, Kim RJ, Lesser JR, et al. ACCF/ACR/SCCT/ SCMR/ASNC/NASCI/SCAI/SIR 2006 appropriateness criteria for cardiac computed tomography and cardiac magnetic resonance imaging: a report of the American College of Cardiology Foundation Quality Strategic Directions Committee Appropriateness Criteria Working Group, American College of Radiology, Society of Cardiovascular Computed Tomography, Society for Cardiovascular Magnetic Resonance, American Society of Nuclear Cardiology, North American Society for Cardiac Imaging, Society for Cardiovascular Angiography and Interventions, and Society of Interventional Radiology. J Am Coll Cardiol. 2006;48(7):1475–1497.

Acute Myocardial Infarction: Ventricular Remodeling Rajan A. G. Patel and Christopher M. Kramer

Left ventricular (LV) remodeling after myocardial infarction occurs in response to the initial injury and myocyte loss. It is a process in which LV cavity anatomy, myocyte contractile function, and myocardial electrical features are modified as a consequence of local mechanical aberrations, systemic hemodynamic changes, alterations in the neurohormonal milieu, paracrine and intercellular factors, and intracellular factors including modifications of gene expression.1–3 Remodeling after infarction consists of two phases. The early phase occurs during the first 72 hours, and the late phase occurs thereafter.3 Both the early and late phases of LV remodeling are the consequence of a complex series of mechanical and molecular triggers that result in alterations of cellular signaling and gene expression in the infarct, peri-infarct, and remote zones. During the last 30 years, investigators have begun to unravel the complex series of cues for postinfarction remodeling. Only recently has the literature begun to paint a coherent picture of the nuclear, cytoplasmic, and cell surface phenomena associated with this process.4,5 The early phase of LV remodeling occurs at the time of acute myocardial infarction (MI). The initial extracellular response to MI involves the activation of an inflammatory cascade. This may be followed by microvascular obstruction (MO).6 Monocytes, macrophages, and neutrophils infiltrate the infarct zone as a consequence of local cytokine expression from endothelial cells and myocytes. Within hours of myocyte injury, infarct expansion occurs.7 Neutrophils release serine proteases and activate matrix metalloproteases, resulting in degradation of the extracellular matrix (ECM) collagen struts. Histology of the infarct zone reveals a reduction in the number of myocytes spanning the infarct zone. This local wall thinning is thought to occur as a result of slippage between muscle bundles once the ECM has been degraded.8 The extent of infarct expansion has been shown to correlate more closely with the extent of MO than with the initial infarct size.9 Infarct expansion results in abnormal wall stress on peri-infarct and remote-infarct myocytes. Abnormal wall stress disrupts the Frank-Starling relationship within these areas. Locally, angiotensin 2 is released, which stimulates synthesis of contractile units. Systemically, the sympathetic nervous system is activated.3 The subsequent increased sympathetic tone will augment heart rate and increase ionotropy in addition to increasing afterload. The late phase of ventricular remodeling is characterized by scar formation at the site of myocyte injury, cavity dilation,

and hypertrophy in the peri-infarct and remote-infarct zones. From a teleologic perspective, cavity dilation may represent a compensatory mechanism for a depressed LV ejection fraction (LVEF) post-infarction. With cavity dilation stroke volume can increase, thus maintaining cardiac output.10 However, under these conditions, cardiac output is maintained at the cost of increased wall stress throughout the cardiac cycle which, in turn, stimulates further cavity dilation.2 The response of myocytes in the noninfarct zones to increased wall stress is hypertrophy,11 specifically an increase in cell length.12 The result is an increase in overall myocardial muscle mass without a corresponding increase in wall thickness as a consequence of greater cavity size. Hypertrophy causes a further increase in wall stress and, subsequently, additional cavity dilation. The result is a cycle of adverse ventricular remodeling (Fig. 19-1).

CARDIOVASCULAR MAGNETIC RESONANCE EVALUATION OF VENTRICULAR ANATOMY DURING REMODELING Cardiovascular magnetic resonance (CMR) allows the assessment of LV volumes and mass, the serial comparison of these measurements in both animals and humans, and the study of myocardial fiber structure during remodeling. These metrics not only are important for prognosis post-infarction but also lend insight into the remodeling process. Many studies have documented the importance of LV volumes after infarction for predicting prognosis.13–16 White and colleagues17 established LV end-systolic volume as the most powerful predictor of long-term outcome after infarction. The tomographic capability of CMR allows for comprehensive/volumetric coverage of the LV from the base to apex (Fig. 19-2), thus obviating the need for geometric assumptions in determining LV volume. Grothues and colleagues18 demonstrated that fast breath hold CMR techniques have a better interstudy reproducibility than did echocardiography in serially assessing LV end-systolic volume, LVEF, and mass. This point is particularly important during the assessment of therapeutic agents for MI or heart failure, as it results in a considerable reduction in the sample

Cardiovascular Magnetic Resonance 253

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CHAPTER 19

ISCHEMIC HEART DISEASE

Baseline

48 hours

1 week

8 weeks

Figure 19-1 Serial phase sensitive inversion recovery gradient recalled echo images of a canine that underwent left anterior descending occlusion/reperfusion. Images were obtained at baseline, 48 hours, 1 week, and 8 weeks post-infarction. Note the progressive LV cavity dilation, expansion of myocardial thinning, and regression of late gadolinium enhancement defined scar during the observation period.

Figure 19-2 Diagram demonstrating the tomographic approach to CMR of the LV. Short axis slices are imaged, spanning the entire LV from base (slice 1) to apex (slice 10). A segmental approach to analysis can be performed. In this example, each short axis slice is divided into eight segments. The inferolateral segments are shaded. Source: Dubach et al, Circulation 1997;95:2060–2067 with permission.

Anteroseptal 1

8 1 8 2 1 7

2

3

4

2 5

6

3

7

7 8

9 10

6 5

3 6

4 5

size necessary to demonstrate changes in LV dimensions and function that are clinically important. Matheijssen and colleagues19 have demonstrated that the intraobserver and interobserver variability of LV volumes and mass is in the 2% to 4% range. CMR has also been assessed for studying the right ventricle (RV). The interstudy reproducibility of RV parameters has been reported from a group of 60 patients undergoing CMR. The percent of variance is listed after each parameter: RV end-diastolic volume, 6.2%; RV 254 Cardiovascular Magnetic Resonance

4 Inferolateral

end-systolic volume, 14.1%; RV ejection fraction (RVEF), 8.3%; and RV mass, 8.7%. In this study, the authors noted higher variability for RV measurements relative to LV metrics in the same patients. However, this difference was statistically significant only for ejection fraction (EF).20 Various investigators have validated the CMR measurement of LV mass after infarction in comparison with LV mass at autopsy and have found very high r values of 0.93 to 0.99.21–23 In an ovine model of anteroapical MI

analysis demonstrated that LV end-diastolic volume (p ¼ 0.038), LVEF (p ¼ 0.0057), and MO (p ¼ 0.044) at the initial post-MI study predicted major adverse cardiac events. RV changes after LV infarction have also been studied by using serial CMR in human patients and in animal models. The study by Hombach and colleagues28 cited above included 6 of the 18 patients with RV involvement were found to have reduced RVEF. Wiesmann and coworkers studied a group of mice after LAD ligation,29 finding a greater than 50% decrease in LVEF with a greater than twofold increase in LV end-diastolic volume and an almost sixfold increase in LV end-systolic volume. While there was no significant change in RV end-diastolic volume or end-systolic volume, RV stroke volume decreased by almost 56%, and RV cardiac output decreased by over 50%. Nahrendorf and colleagues30 used CMR at 2 weeks and 8 weeks to assess RV changes in response to LAD infarction in a rat model. Between 2 and 8 weeks post-infarction, the increase in RV end-diastolic volume was more than 2.5 times that experienced by controls. At 2 weeks, RVEF was significantly greater in controls than in mice that experienced an infarction, but there was no difference at 8 weeks. Diffusion tensor magnetic resonance imaging (DTMRI) allows quantitative assessment of myocardial fiber orientation with high spatial resolution within the excised intact myocardium. This method of imaging may provide the means to quantify and model ventricular remodeling at the myocardial fiber level.31 It relies on the assumption that water diffusion is greatest along the length of a myocardial fiber. If this assumption is true, then the primary eigenvector of the diffusion tensor coincides with the local fiber orientation. Chen and colleagues32 performed DTMRI on Fischer 344 formalin-fixed rat hearts post-MI. They reported greater water diffusivity and decreased diffusion anisotropy in infarct zones. Post-MI, myocytes in the infarct zone remodeled such that the angular shift per millimeter of transmural depth was significantly greater than in myocardium of control rats. However, the overall fiber orientation remained relatively unchanged.

CMR EVALUATION OF REGIONAL LEFT VENTRICULAR FUNCTION DURING REMODELING CMR methods have also been used to study the natural history of regional LV function during remodeling after infarction. These include cine CMR,25 three-dimensional (3D) wall thickening,33 and CMR tagging.21,26,34 Konermann and colleagues25 assessed endocardial motion toward the centroid of the LV cavity. Within the infarction zone (defined as a region of systolic thickening less than 2 mm), endocardial motion fell during the 6-month followup from 2.0 1.6 mm to 0.5 2.9 mm, the extent being dependent on infarct size, but independent of infarct location. Within noninfarcted regions, endocardial motion decreased in the group as a whole, falling from 7.1 2.4 mm to 6.3 2.7 mm from week 1 to week 26. The decline was most dramatic in anterior infarcts. Overall, wall thickening did not change in noninfarcted regions. Cardiovascular Magnetic Resonance 255

19 ACUTE MYOCARDIAL INFARCTION: VENTRICULAR REMODELING

induced by coronary ligation, a tomographic set of shortaxis CMR images analyzed by summation of discs was used to follow LV mass, end-diastolic, and end-systolic volumes in the first 6 months after the initial event. A stepwise increase in LV mass, LV end-diastolic volume, and LV endsystolic volume was found. The increase in LV end-diastolic volume was disproportionate such that the LV volume-tomass ratio increased over the 6-month period. Infarct wall thickness fell during the 6-month follow-up, whereas no change was observed in wall thickness of noninfarcted regions.21 These noninfarcted regions demonstrated an increase in segment length during the 6-month follow-up, suggesting that the hypertrophy that occurs during LV remodeling is of the eccentric type.24 In patients, cine CMR has been used to measure LV volumes and mass after infarction. Konermann and colleagues25 reported the serial imaging of 61 patients at 1, 4, and 26 weeks after nonreperfused infarction. Between weeks 1 and 26, the average LV end-diastolic volume index increased from 74 23 mL/m2 to 85 28 mL/m2, and the LV end-systolic volume index increased from 40 19 m2 to 51 29 mL/m2. Most of the increase was found in the subgroup of 32 patients who suffered an anterior infarction. No change in LV stroke volume was found over the study period. Average LV mass increased from 246 66 g to 276 80 g. The change in the volume-to-mass ratio was directly related to the enzymatic infarct size. The ratio decreased in smaller infarcts, remained unchanged in moderate-sized infarcts, and increased in larger infarcts. CMR was used in a study of 26 patients who suffered anterior infarctions and received reperfusion therapy. Patients were imaged on day 5 2 and week 8 1 after infarction. All patients had single-vessel left anterior descending (LAD) disease and regional LV dysfunction with an initial LVEF of 50% or less. The LV mass index trended downward during the 8-week period, falling from 109 19 g/m2 to 102 18 g/m2. The average LV end-diastolic index increased from 83 24 mL/m2 to 96 27 mL/m2, and the LVEF increased from 39 12% to 45 14%, whereas the LV end-systolic volume index did not change. Therefore, despite an increase in LV end-diastolic volume, global LV function improved in the 8-week postinfarction period. By multivariate analysis, the only significant predictor of an increase in LV end-diastolic volume index was peak creatinine kinase.26 Schroeder and colleagues27 performed serial CMR studies at 1 week, 13 weeks, 26 weeks, and 52 weeks postinfarction in 51 patients. They found a 5% increase in LV mass index among those treated with percutaneous transluminal coronary angioplasty (PTCA) compared with 10% for those treated with thrombolytics. Furthermore, the increase in LV mass index occurred without a change in thickness of the noninfarcted myocardium. Subsequently, Hombach and colleagues28 reported on 110 patients who underwent CMR 6.1 2.2 days after acute infarction, including 89 who had a follow-up study 225 92 days later. These included 81 patients who were treated with percutaneous coronary intervention (PCI) (94% stents) 6.5 4.7 hours after symptom onset. Over 90% of the population was discharged on aspirin, beta-blockers, and angiotensin-converting enzyme inhibitors (ACE-I). Contrastenhanced techniques were used to assess MO and infarct size (discussed below). Upon follow-up, multivariate

ISCHEMIC HEART DISEASE

Holman and colleagues33 performed cine CMR imaging in 25 patients 3 weeks after anterior infarction and analyzed functional data with a 3D wall thickening using a centerline method. Wall thickening was reduced in myocardial territories perfused by the LAD and left circumflex (LCx) arteries compared with a normal database and in the LAD compared with the other territories within this patient group. The quantity of dysfunctional myocardium correlated well with an enzymatic estimate of infarct size. In an ovine model of LV infarction/remodeling,21 CMR tagging35,36 was used to evaluate regional function in circumferential and longitudinal planes at baseline, 1 week, 8 weeks, and 6 months after infarction. Shortening within infarcted regions was reduced throughout the study period. A persistent difference in intramyocardial shortening was found between noninfarcted regions adjacent to and remote from the infarct border. Function in adjacent noninfarcted regions fell markedly at 1 week after infarction and partially improved at 8 weeks after infarction but remained depressed relative to baseline and to remote regions.21 Moulton and colleagues used the same model and CMR tagging to demonstrate that adjacent or border zone fibers were stretched during isovolumic systole and that this contributed to reduced fiber shortening during systolic ejection.37 Epstein and colleagues38 studied regional systolic function in a murine model of LAD occlusion and reperfusion imaged at baseline and 1 day post-infarction. A CMR tagging sequence demonstrated that a gradient of contractile dysfunction occurs early after infarction between the infarcted and remote regions. These investigators calculated the percentage of circumferential shortening (%CS) at baseline to be 14.5 3.4%. After infarction, the infarct region %CS dropped to 0.7 4.4% (p < 0.01), while the peri-infarct and remote regions had a %CS of 7.4 4.4% ( p < 0.01) and 11.8 4.2% (p < 0.01), respectively. The spectrum of functional changes that occurs up to 12 weeks postinfarction has been described by Thomas and colleagues in a rat model of LAD ligation with CMR performed at 1 to 2, 3 to 4, 6 to 8, and 9 to 12 weeks post-infarction.39 In addition to changes in LV volumes and mass, CMR identified maximum (E1) and minimum (E2) principal strains and changes in the orientation angle of the maximum principal stretch. Both principal strains were found to be decreased in the infarct, the peri-infarct, and the remote-infarct regions throughout the study period. Furthermore the previously described gradient of abnormal strains extending from the infarct region to the remote regions early after infarction was documented throughout this longitudinal study. A CMR tagging study demonstrated mild dysfunction within remote noninfarcted regions in 28 patients at 5 2 days after their first reperfused anterior infarction.34 Function within the apex, anterior wall, and septum was depressed relative to a normal database. Basal lateral percentage of intramyocardial %CS was only 17 %, significantly less than the 22 7% in normal subjects. In addition, midinferior %CS (12 10%) was lower than normal subjects (19 5%). When this patient group was followed for 8 weeks after anterior infarction,26 improvement in regional function was found in both infarcted and noninfarcted regions, with apical %CS improving (9 6%% to 13 5%) as well as midanterior (6 6% to 10 7%) and midseptal regions (8 7% to 12 6%). The dysfunction demonstrated on day 5 in remote 256 Cardiovascular Magnetic Resonance

midinferior and basal lateral regions resolved by 8 weeks, while LVEF increased from 39 12% to 45 14%. However, LV end-diastolic volume increased during this time. Therefore, as in the ovine model, the improvement in regional function from week 1 to week 8 after infarction was uncoupled from changes in global LV volume. CMR myocardial tagging was used to assess the correlation between regional function and regional loading conditions in 16 patients after their first anterior infarction40 with imaging at 1 week and 3 months post-infarction. The LV myocardium was divided into a 32-segment model. Regional LVEF was calculated by using a pie-shaped volume defined by endocardium and the center of the LV. The regional load was defined as the product of the systolic blood pressure and mean radius of curvature of a given segment divided by the wall thickness of that segment. In remote regions, deformation decreased as loading conditions increased between the initial and follow-up scans, while in peri-infarct regions, no significant change in deformation was appreciated as loading conditions increased. However, in infarct regions, deformation increased, while loading conditions increased.40 In the future, it is likely that CMR sequences that have higher spatial resolution than tagged imaging does, such as cine displacement encoding with stimulated echoes, will allow better assessment of regional displacement and strain.41 Beyond describing the phenomenon of postinfarction ventricular remodeling, the utility of such sequences will be the quantitative assessment of the effects of pharmacologic, mechanical, or electrical interventions on regional ventricular function during remodeling.

CMR SPECTROSCOPY EVALUATION OF ENERGETICS DURING LEFT VENTRICULAR REMODELING CMR spectroscopy (CMRS) has been used in animal models to define the metabolic alterations that occur within remodeled myocardium after myocardial injury. McDonald and colleagues42 studied a canine model of LV remodeling after transmyocardial direct current (DC) shock, which causes localized myocardial necrosis to approximately 20% of the anteroapical left ventricle and leads to LV remodeling over the ensuing 12 months. These investigators used a surface coil over the lateral wall, remote from the direct region of myocardial damage, and performed spatially localized 31 P spectroscopy. The ratio of creatine phosphate (CP) to adenosine triphosphate (ATP), a measure of high-energy phosphate stores, was reduced in the subendocardium and subepicardium of the remodeled LV. In the same model,43 pacing reduced the subendocardial/subepicardial blood flow ratio, which was associated with a fall in the subendocardial CP-to-ATP ratio and an increase in inorganic phosphate–to-CP ratio (Fig. 19-3). This finding suggests that redistribution of blood flow may play a role in the alteration of myocardial high-energy phosphate levels in remodeled myocardium. The extent of remodeling, as indexed by the increase in end-diastolic volume and mass over time, correlated with the subendocardial CP-to-ATP ratio.

Pi

*

ENDO

*

*

MID

* CP

ATPγ ATPα

ATPβ

*

EPI

A

In a study using 31P spectroscopy in isolated, perfused remodeled rat myocardium after MI induced by coronary ligation, Friedrich and colleagues demonstrated similar changes in high-energy phosphates.44 CP concentration was reduced in noninfarcted myocardium in comparison to control heart. The extent of reduction correlated with the size of the infarction. However, the ATP content was unchanged within noninfarcted myocardium. The CP-to-ATP ratios were reduced in comparison to those in controls. These investigators postulated that assessment of high-energy phosphates may be an excellent marker of the magnitude of dysfunction within noninfarcted myocardium. In support of this concept, Zhang and colleagues performed a study in a porcine model of infarction.45 Among 18 animals with infarction, 6 developed clinical congestive heart failure, while the other 12 developed asymptomatic LV dysfunction. The CP-to-ATP ratio in remodeled noninfarcted myocardium was decreased, especially within the subendocardium. The decrease in the CP-to-ATP ratio in the animals with congestive heart failure was transmural and greater than that in the animals with asymptomatic LV dysfunction and remodeling. Therefore, the extent of bioenergetic abnormalities reflects the severity of LV dysfunction in the remodeled LV. Straeter-Knowlen and colleagues46 used 1H CMRS to assess ex vivo myocardial triglyceride concentration after infarction in a canine model of LAD ligation. Twenty-four hours after infarction, Evans blue was infused to assess the risk area, followed by euthanasia. 1H CMRS detected the highest triglyceride concentrations in the peri-infarct regions within the risk area. The lowest triglyceride concentrations were within the remote-infarct regions. An intermediate triglyceride level was detected within the infarct zone. These findings were confirmed with histologic staining techniques. Other research of myocardial lipid metabolism post-infarction has suggested similar changes

B

in triglyceride concentrations.47,48 Straeter-Knowlen and colleagues46 hypothesized that 1H CMRS may be useful in assessing myocardial viability and reversible injury. Subsequently, Hillenbrand and coworkers49 performed serial studies on a canine model of infarction comparing infarct size using 23Na CMRS and 1H CMRS. The 23Na signal ratio of infarcted to noninfarcted regions was shown to correlate with 1H CMRS data. In similar experiments on rats, the 23Na signal intensity was also shown to decrease as scar formation increased after infarction. The extent of scar formation was confirmed with histology after euthanasia at 3 days and 6 weeks after infarction. MRS has been used to serially document established molecular changes in the infarct, peri-infarct, and remote zones. However, at present, it is not used in routine clinical practice.

CONTRAST-ENHANCED CMR AND PREDICTORS OF LEFT VENTRICULAR REMODELING Contrast-enhanced CMR with agents that shorten T1 relaxation have been particularly useful in the assessment of ventricular remodeling. Gadolinium-based agents have been used to study MO, scar formation, and viability. Gadoliniumdiethylenetriamine pentaacetic acid (Gd-DPTA) is an extracellular agent that is retained within infarcted tissue. This is hypothesized to occur as a result of delayed wash-out due to a decrease in the functional capillary density within the infarct zone and also due to an increase of ECM within collagenous scar.50 The approach is highly reproducible in both the acute and chronic infarct settings.51 Hypoenhancement observed on early imaging after Gd-DPTA contrast infusion is a marker of MO due to the “no-reflow” phenomenon.52 In dogs that were reperfused Cardiovascular Magnetic Resonance 257

19 ACUTE MYOCARDIAL INFARCTION: VENTRICULAR REMODELING

Figure 19-3 A transmurally localized 31 P-MRS spectrum that shows highenergy phosphate levels in a canine heart with LV remodeling after localized myocardial damage after direct current shock. A, Levels during sinus rhythm. B, Levels during pacing at 240 beats per minute. Pacing caused a significant decrease of the creatine phosphate (CP) to adenosine triphosphate (ATP) ratio and an increase of △Pi/CP in these hearts. At baseline, the CP-to-ATP ratio is lower in remodeled hearts than in controls. ENDO, subendocardial voxel; EPI, subepicardial voxel; MID, midwall voxel; Pi, inorganic phosphate. Source: Zhang J, McDonald KM. Bioenergetic consequences of left ventricular remodeling. Circulation. 1995;92:1011– 1019 with permission.

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after infarction, Gerber and colleagues demonstrated that the extent of hypoenhancement or MO correlated with an increase in LV end-diastolic volume at 10 days.9 Wu and colleagues53 performed CMR on 44 patients 10 6 days after infarction. MO was defined as areas of hypoenhancement 1 or 2 minutes after gadolinium injection. Eleven patients had MO. These patients experienced a greater incidence of cardiovascular events, including cardiac death, reinfarction, congestive heart failure, and stroke, in comparison to patients without MO (45% versus 9%, p ¼ 0.10). During multivariate analysis controlling for infarct size, MO remained an independent predictor of postinfarction complications. Additionally, evidence of MO by CMR was a better predictor of cardiovascular complications than was angiographic flow score (5 of 11 patients with MO by CMR had TIMI grade 3 flow). Choi and colleagues54 performed CMR on 25 patients at 1 week and 8 weeks postinfarction. LVEF did not improve in patients with MO. Partial functional improvement was noted in segments with nontransmural late gadolinium enhancement (LGE), but no functional improvement was observed in segments that demonstrated LGE and MO. Imaging at least 10 to 15 minutes after gadolinium infusion has been used to assess scar size. Areas of scar appear bright on inversion recovery sequences and are referred to as LGE. In a mouse model, Yang and colleagues55 established a correlation between myocardial infarction size and area of late hyperenhancement based on 2,3,5-triphenyl tetrazolium chloride (TTC) staining in a murine model of coronary ligation/infarction and reperfusion, Bland-Altman

3 Days

8 Weeks

Dog 1

Dog 1

Dog 2

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258 Cardiovascular Magnetic Resonance

analysis revealed a mean difference of 0.6% between TTC assessment of infarct size and the area of LGE. Rochitte and colleagues56 performed serial imaging of infarct size in dogs after 90 minutes of LAD occlusion with an angioplasty balloon with CMR at 2 hours, 6 hours, and 48 hours postinfarction. They demonstrated an increase in scar size and the MO zone during the first 48 hours. Fieno and colleagues57 used gadoteridol (Gd-HP-D03A) to serially assess infarct size in 17 dogs that underwent mid-LAD occlusion (after the first diagonal) for 45 minutes, 90 minutes, or permanently. LGE CMR was performed at 3 days, 10 days, 4 weeks, and 8 weeks to assess the circumferential and longitudinal extents of infarction. Infarct resorption during the 8week follow-up period correlated with initial infarct size (Fig. 19-4). At the 4-week scan, the extent of infarct resorption was related to the presence of reperfusion. These authors also demonstrated an increase in the LV expansion index with serial CMR scans. In the previously cited study from Hombach and colleagues,28 in which patients received stenting with contemporary medical therapy, LGE CMR was used to assess scar size and MO. Patients with evidence of MO were found to have higher enzyme levels and larger infarctions in comparison with patients without MO. The increased infarct size at baseline among patients with MO (16 7% versus 8 5%, p < 0.001) persisted at the late (9-month) follow-up scan (10 5% versus 6 5%, p < 0.001). Furthermore, at 100 days post-infarction, there was a 12.6% (95% confidence interval: 1.4 to 23.8) increase in the incidence of a major adverse cardiovascular event among those with MO (Fig. 19-5).

Figure 19-4 Late gadolinium enhancement gradient recalled echo images in a canine infarct model demonstrating resorption of scar between 3 days and 8 weeks post-infarction. Source: Fieno DS, et al. Infarct resorption, compensatory hypertrophy, and differing patterns of ventricular remodeling following myocardial infarctions of varying size. J Am Coll Cardiol. 2004;43(11):2124–2131 with permission.

Additionally, LGE CMR provided evidence of papillary muscle infarction in 28 patients and RV infarction in 18 patients. Lund and colleagues58 performed CMR at 5 3 days and at 8 3 months in 55 patients after acute infarction treated with primary angioplasty or thrombolytics. At 5 days, the LGE mass was directly proportional to the LVEF. An infarct size representing at least 24% of LV mass predicted adverse remodeling with a sensitivity and specificity of 92% and 93%, respectively. In addition, the extent of LGE predicted adverse ventricular remodeling better than did the peak creatinine kinase or the extent of MO. At the 8-month scan, patients with myocardial remodeling experienced a larger relative reduction in LGE infarction size in comparison to patients who did not experience remodeling (34 26% versus 12 35%, p < 0.05). Kaandorp and coworkers59 studied 29 patients at 1 week and 9 months after infarction, including 94% on beta-blockers and 83% on ACE-I. Receiver operating characteristic curve analysis demonstrated that a LGE scar size exceeding 23% of LV mass had a 95% sensitivity and a 95% specificity for predicting a greater than 10% increase in LV end-diastolic volume. An LGE mass of at least 36 g had a 100% sensitivity and 95% specificity to predicting a greater than 10% increase in LV end-diastolic volume. Baks and colleagues60 performed CMR at 5 days and 5 months after first infarction in 22 patients who underwent primary angioplasty. LGE-defined infarct shrinkage occurred to

EVALUATION OF THERAPY FOR LEFT VENTRICULAR REMODELING: ANIMAL STUDIES Owing to the precision with which CMR can be used to measure LV mass and volumes, it has been increasingly used to evaluate pharmacologic, mechanical, and stem cell therapy for the prevention of adverse LV remodeling after infarction. The majority of the CMR studies with Cardiovascular Magnetic Resonance 259

19 ACUTE MYOCARDIAL INFARCTION: VENTRICULAR REMODELING

Figure 19-5 Short axis LGE CMR in a patient who presented to the emergency department several hours after the onset of chest pain with a large anterior ST-segment elevation infarction. Note the subendocardial hypoenhancement with surrounding hyperenhanced myocardium. These dark areas represent regions of microvascular obstruction and portend a poor prognosis.

the same extent in small and large infarctions, with a mean decrease of 41%. Dysfunctional segments without MO had increased wall thickness at 5 days compared with remote or noninfarcted myocardium and thickness similar to remote myocardium at 5 months, with improved systolic thickening. Segments with MO demonstrated wall thinning at 5 months with no significant recovery in systolic wall thickening. Finally, Bodi and colleagues61 performed an LV function, LGE, and low-dose dobutamine wall motion assessment in 214 consecutive patients with ST elevation MI treated with thrombolytic therapy or primary angioplasty. During a median follow-up of 553 days, there were 21 major adverse cardiac events (cardiac death, non-fatal MI, heart failure admissions). Multivariate analysis of clinical, electrocardiographic, biomarkers, angiography and CMR indexes demonstrated that the extent of systolic dysfunction and transmural necrosis provided independent prognostic information. While Gd-DPTA can reveal areas of scar in both acute and chronic infarcts, other agents may be useful in delineating acute infarctions. Watzinger and colleagues62 performed 1 hour of coronary occlusion followed by reperfusion in 27 rats with CMR at 2 days and 8 weeks. Twelve hours before the first CMR, 0.05 mmol/kg of Gdmesoporphyron was infused. Acutely infarcted regions enhanced with Gd-mesoporphyron, while chronic infarcts did not. Gd-mesoporphyron is not approved for use in humans; however, agents such as this may prove useful in studying remodeling in patients with multiple infarctions. Noncontrast CMR methods to localize infarction are also available. Friedrich and colleagues63 performed T2-weighted imaging and LGE CMR in 92 patients with acute infarction and successful reperfusion. All patients had high T2 signal in the infarct region. In the slice with maximum infarct extension, the proportion of the area with high T2 signal was 38 15% of the myocardial slice compared with 22 10% in LGE. The implications for remodeling remain unexplored. Finally, Ganame and coworkers employed T2-weighted CMR to differentiate between hemorrhagic (hypointense core) and non-hemorrhagic (hyperintense core) infarcts with microvascular obstruction and infarct size determined by LGE in 98 patients with infarction reperfused.64 One quarter of the patients had hemorrhagic infarcts, which were related to larger infarct size and infarct transmurality and lower LVEF. At 4 months, myocardial hemorrhage was an independent predictor of adverse LV remodeling, defined as an increase in LV end-systolic volume.

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pharmacologic agents have focused on the reninangiotensin-aldosterone axis. CMR evaluation of interventional techniques is not yet widespread, but this area has the potential for significant growth as investigators study novel therapies or seek to identify the mechanisms underlying favorable therapeutic outcomes in population-based studies. Stem cell transplantation as a means of preventing or treating adverse ventricular remodeling is also gaining increasing popularity. A primary objective of myocardial stem cell research is to demonstrate gain of myocyte function with transplanted stem cells. CMR is an ideal modality to study the effects of myocardial stem cell transplantation, as it allows assessment of scar size and regional function. One of the first studies to apply CMR in the assessment of pharmacologic agents was that of Saeed and colleagues.65 Using an infarction model, they studied the impact of cilazapril, an ACE-I on LV wall thickness and chamber diameter. Cilazaipril preserved wall thickness in the infarct-containing segment of the anterior and septal walls and limited the increase in chamber surface area, a surrogate for LV volume. CMR techniques were also employed in an ovine model of LV remodeling, demonstrating that use of an ACE-I during the first 8 weeks after anteroapical infarction limited the increase in LV end-diastolic volume that was noted in untreated controls.66 The limitation in LV dilation was associated with maintenance of regional function in noninfarcted myocardium adjacent to the infarction, suggesting that dysfunction in this region may be an important determinant of the remodeling process. CMR has been utilized to assess LV remodeling in a canine model of DC-induced myocardial necrosis. In one CMR study performed at 16 weeks after DC shock, highdose ACE-I limited the increase in LV mass and enddiastolic volume seen in controls, whereas neither low-dose ACE-I, alpha-1 receptor blockade, nor angiotensin II type 1 receptor blockade was effective.67 To further characterize the mechanisms underlying the reduction in remodeling with ACE inhibition, McDonald and colleagues68 performed baseline and 4-week CMR in a canine model of DC shock in the setting of a bradykinin antagonist added to ACE-I. Bradykinin antagonism counteracted the reduction in LV mass that is seen with ACE-I alone, suggesting that bradykinin may be largely responsible for the prevention of hypertrophy seen with ACE-I. No change in LV volumes was noted at the 4-week time point in any of the groups. These investigators used similar methods to evaluate therapy with captopril and metoprolol administered late after injury (11 months after DC shock) and continued for 3 months.69 Both therapies were associated with a reduction in LV mass and LV end-diastolic volume compared to untreated controls. Using CMR and the ovine model of anteroapical infarction, Kramer and colleagues70 demonstrated that the addition of beta-blockade to ACE-I during LV remodeling did not change LV end-diastolic volume. However, LVEF did improve over the first 8 weeks in comparison to ACE-I alone. The most powerful pharmacologic regimen for attenuating LV remodeling and preserving regional and global function was the combination of ACE-I and losartan71 (Fig. 19-6). 260 Cardiovascular Magnetic Resonance

The role of the angiotensin 2 receptor (AT2R) in LV remodeling has been investigated in a series of studies involving transgenic mice with cardiac overexpression of the AT2R. CMR was performed at 1, 7, and 28 days postinfarction. At baseline, end-systolic volume index was lower and LVEF was higher in the transgenic mice. The initial infarct size was similar in both groups. After controlling for baseline differences, the AT2R-overexpressing mice had a significantly lower end-systolic volume index and a higher LVEF at 1 week and 4 weeks.72 The prevention of adverse remodeling was found to be, at least in part, due to preservation of contractile function in the peri-infarct regions.73 Nitric oxide has been hypothesized to mediate the benefit of AT2R overexpression. This was demonstrated in a study of AT2R-overexpressing mice treated with a nitric oxide synthase (NOS) inhibitor in which these mice remodeled similarly to wild type mice.74 In a study of the relative contribution of AT2R and angiotensin 1 receptor (AT1R) in postinfarction remodeling, these investigators demonstrated equivalent remodeling in AT2R-overexpressing mice and wild type mice treated with losartan, an AT1R antagonist. These data support the hypothesis that much of the benefit of AT1R antagonism is mediated through the AT2R. Transgenic mice with AT2R overexpression and AT1R knockout remodeled even less, likely owing to further blood pressure lowering in the latter group.75 Isbell and colleagues76 utilized AT2R-overexpressing mice with and without knockout of the bradykinin 2 receptor to ascertain whether it was involved in the attenuation of remodeling ascribed to AT2R after infarction. Serial CMR performed at baseline, 1, 7, and 28 days after infarction demonstrated the lower end-systolic volume index and LVEF previously observed in AT2R-overexpressing mice was mediated by the bradykinin 2 receptor. However, the reduction in adverse remodeling attributed to AT2R appeared to be independent of the bradykinin receptor. As was alluded to above, inducible nitric oxide synthase (iNOS) may play an important role in limiting ventricular remodeling after infarction. Using CMR, Gilson and colleagues77 reported on the effect of iNOS in a murine LAD occlusion/reperfusion model. LV volumes were lower and LVEF was higher in iNOS knockout mice than in wild type mice. Furthermore, the circumferential extent of wall thinning was reduced in iNOS knockout mice. This effect was evident by day 7 post-infarction and persisted at 28 days. The canine model of DC-induced myocardial injury has also been used to study the role of nitrates in reducing LV remodeling.22 Dogs were randomized to nitrate therapy (N ¼ 10) or control (N ¼ 17) and studied at baseline and at 1 and 16 weeks after injury. Cine CMR demonstrated that nitrate therapy prevented the increase in LV mass seen in control animals at 1 week post-damage and the increase in LV mass and LV end-diastolic volume at 16 weeks. Pharmacologic agents that may limit reperfusion injury and MO might also be useful in limiting ventricular remodeling post-infarction. Nicorandil is an ATP-sensitive potassium channel agonist used as an antianginal agent in Asia and Europe because of its nitrate-like effect.78 Krombach and colleagues79 demonstrated that nicorandil treatment in rats after coronary artery ligation prior to reperfusion reduced infarct size assessed by LGE. In another study, they fed nicrorandil to rats prior to coronary ligation and reperfusion and for 8 weeks post-infarction. At 8 weeks, the

19 ACUTE MYOCARDIAL INFARCTION: VENTRICULAR REMODELING

RV

LA

LV

A

C

B

D

E

Figure 19-6 End-diastolic four-chamber long-axis cine gradient recalled echo image at 8 weeks post-infarction in an ovine model with an example from each of five therapeutic groups. A, No therapy. B, Angiotensin type 1 receptor blocker (ARB). C, Standard-dose angiotensin-converting enzyme inhibitor (ACE-I). D, High-dose ACE-I. E, Combination therapy with low-dose ACE-I and ARB. Note that there is visually less infarct expansion with combination therapy. LA, left atrium; LV, left ventricle. Source: Mankad S, et al. Combined angiotensin II receptor antagonism and angiotensin-converting enzyme inhibition further attenuates postinfarction left ventricular remodeling. Circulation. 2001;103(23):2845–2850 with permission.

nicorandil-treated groups had improved LVEF, LV enddiastolic volume, LV end-systolic volume, and LV wall thickening in remote-infarct regions and peri-infarct regions in comparison to controls that experienced initial infarcts of the same size.80 Nahrendorf and colleagues81 studied the role of HMGCoA-reductase inhibition in LV remodeling to determine whether the mechanism of any positive effect on remodeling was NOS dependent. Adult rats underwent left coronary ligation followed by treatment with placebo, cerivastatin, or cerivastatin plus N-methyl-L-arginine methyl ester (L-NAME), a NOS inhibitor. Rats that were fed cerivastatin had a significantly lower LV mass at 4 and 12 weeks post-infarction. However, LV volume was similar among all three groups. L-NAME abolished the effects of cerivastatin. The authors concluded that HMG-CoAreductase inhibitors diminished LV hypertrophy and helped to preserve function but not LV dilation in a mechanism that appears to involve NOS. The same group also tested the effect of testosterone on LV remodeling based on data demonstrating a higher 2-year mortality post-infarction in women. Rats were treated with placebo, testosterone, or orchiectomy with CMR at 2 weeks and 8 weeks after LAD ligation. At 8 weeks. LVEF and cardiac output decreased in all three groups.82 The effect of transmyocardial laser revascularization (TMLR) post-infarction is an example of a percutaneous intervention affecting remodeling that has been evaluated

by CMR. First, infarction was produced in a rat model of coronary ligation followed by baseline CMR at 8 weeks. TMLR was then performed on remote myocardium in the treatment group. Twelve weeks post-infarction, CMR was repeated with dobutamine stress perfusion imaging. LVEF was similar in both groups. However, the TMLR group experienced greater LV dilation relative to the control group (82 16 mL versus 25 17 mL, p < 0.03) and a greater increase in LV mass (124 31 mg versus 55 19 mg, p < 0.03). During dobutamine stress, the TMLR group expressed a greater augmentation of LVEF and superior perfusion. Therefore, CMR demonstrated that in this rat model, TMLR resulted in adverse ventricular remodeling but better stress function and perfusion.83 CMR has also been used to study the role of mechanical restraint devices, such as the Acorn cardiac support device, to limit LV remodeling after acute infarction. Blom and colleagues84 utilized an ovine model of LAD occlusion to create an anterior myocardial infarction in 10 animals. Five sheep received the Acorn device 1 week later. At 12 weeks after infarction, animals treated with the Acorn device had a smaller LV epicardial surface area in comparison to controls. Furthermore, border zone systolic regional radial strain was improved at 12 weeks in sheep that were treated with the Acorn device. Stem cell therapy after myocardial infarction is in its early stages. However, Limbourg and colleagues85 used CMR to demonstrate an improvement in myocardial Cardiovascular Magnetic Resonance 261

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function and prevention of LV dilation in a mouse infarction model that received stem cell transplantation. However, scar size did not decrease. Amado and colleagues86 used CMR with a porcine model of infarction to demonstrate an improvement in function and reduction in scar. This group also transplanted stem cells labeled with an iron-containing compound that allowed localization of the transplanted cells within the myocardium. Van den Bos and colleagues87 used CMR to demonstrate gain of regional function and prevention of cavity dilation after myoblast transplantation in a cryoinfarction rabbit model. Zeng and colleagues88 recently published data from a porcine permanent LAD occlusion model in which allogenic adherent stem cells were transplanted 60 minutes after LAD ligation. Volumetrics were assessed with CMR at 25 days after LAD ligation. Interestingly, myocardial energetics were also studied with 31P CMRS. LVEF was higher in animals treated with stem cells (p < 0.05). The ratio of endocardial phosphocreatine to ATP was also higher among animals treated with stem cells in comparison to untreated controls (1.31 0.29 versus 1.06 0.30, p < 0.05). The question of whether effects such as these are due to stem cell adherence and differentiation, paracrine effects, or other processes remains to be answered.

EVALUATION OF THERAPY FOR LEFT VENTRICULAR REMODELING: HUMAN STUDIES CMR has made significant contributions to clinical trials of LV remodeling. The very high accuracy and reproducibility of CMR have made it the reference standard for the noninvasive evaluation of LV function. These factors have also allowed clinical trials evaluating LV remodeling to be adequately powered to detect end points with fewer patients than would be necessary if other noninvasive imaging modalities were used.18,23 Clinical trials using CMR in the assessment of therapeutics to modify LV remodeling have tested pharmacologic agents, percutaneous interventions, stem cell transplantation, and even exercise. ACE-I and beta-blockers have been the most extensively studied pharmacologic agents with regard to LV remodeling post-infarction. Schulman and colleagues89 randomized 43 patients with a Q-wave acute MI within 24 hours of symptom onset to intravenous enalaprilat and then oral ACE-I therapy or placebo for 1 month post-infarction. Twentythree of the patients underwent CMR at 1 month for evaluation of infarct expansion, defined as the ratio of the infarct to noninfarct endocardial segment length. Other parameters that were measured included infarct segment length and wall thickness. ACE-I therapy was associated with a reduced infarct segment length (7.9 1.0 cm versus 10.6 0.9 cm in controls) and a lower infarct expansion index (1.1 0.3 versus 1.8 0.3 in controls). The greatest impact of ACE-1 therapy on infarct expansion was found in the subgroup with anterior infarcts. Johnson and colleagues90 utilized CMR to assess ACE-I therapy postinfarction in 35 patients with an LVEF greater than 40% and 262 Cardiovascular Magnetic Resonance

their first acute Q-wave infarction. Studies were performed at 1 week and 3 months post-MI. Stacked short-axis cine CMR slices from apex to base were summed to measure LV end-diastolic and end-systolic volume index and mass. Therapy with ramipril contributed to a fall in LV mass index (from 82 18 g/m2 to 79 23 g/m2), whereas there was no significant change in control patients (from 77 15 g/m2 to 79 23 g/m2). No significant change in LV end-diastolic volume index was noted in either the treatment group or controls. The same group91 used CMR to study 29 patients with an LVEF less than 40% treated with ACE-I. Between day 5 and 3 months post-infarction, LV mass increased, LV end-diastolic volume increased, and LVEF improved despite an increase in calculated endsystolic wall stress. Groenning and colleagues92 performed CMR three times over 6 months on 41 patients enrolled in the MERIT-HF study. These were patients with chronic stable heart failure who had been randomized to metoprolol succinate or placebo. At 6 months, the metoprolol group experienced a significant decrease in LVend-diastolic volume index (150 mL/m2 at baseline decreasing to 126 mL/m2 at 6 months, p ¼ 0.007) and LV end-systolic volume index (107 mL/m2 at baseline decreasing to 81 mL/m2 at 6 months, p ¼ 0.001), with a concomitant increase in LVEF (29% at baseline increasing to 37% at 6 months, p ¼ 0.005). There were no changes among these variables in the placebo group. Dubach and coworkers93 performed CMR on 26 patients with a mean LVEF ¼ 26 6% randomized to bisoprolol fumarate or placebo. After 1 year, the treatment group experienced an increase in LVEF (25 7% versus 36 9%; p < 0.05) and a trend towards smaller LV end-diastolic volume and LV end-systolic volume. The control group experienced no change. More recently, Bellenger and colleagues94 performed CMR on 34 patients with chronic stable heart failure who were participating in the CHRISTMAS trial, a double-blind study comparing carvedilol and placebo and including CMR at baseline and at 6 months. The carvedilol group experienced a decrease in LV end-systolic volume index (9 versus þ3 mL/m2, p ¼ 0.001) and in LV end-diastolic volume index (8 versus 0 mL/m2, p ¼ 0.05) with a concomitant increase in LVEF (þ3 versus 2%, p ¼ 0.003) relative to control (Fig. 19-7). These studies demonstrate that differences in LV structure due to pharmacologic therapy of acute infarction can be demonstrated with CMR in relatively small-sized patient cohorts. Studies using echocardiography post-infarction have required a much larger group of patients (on the order of several hundred) to demonstrate quantifiable differences between treatment and control groups.95 CMR has been used to assess post-infarction LV remodeling after PCI. Studies have been conducted to assess the effect of both early and late recanalization of infarctrelated arteries on remodeling. One of the first studies using CMR to determine the beneficial effects of early infarct artery patency on LV remodeling documented changes in LV mass and volume over time. Sixteen patients with an occluded LAD 10 days after anterior infarction were randomized to PTCA of the LAD within the following 2 weeks or on a delayed basis (3 months post-infarction). Early mechanical intervention resulted in improved LVEF, improved regional function, and reduced end-systolic volume. These findings were not observed in the delayed therapy group.96 Subsequently, Silva and colleagues97 reported

160 140

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120 p <0.001

100

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80 60 p = 0.008

40 20 0 LVEDVI

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p = ns

120 100

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Figure 19-7 Change in variables over 6 months post-infarction in the carvedilol and placebo groups in the CHRISTMAS trial. LVEDVI, left ventricular end-diastolic volume index; LVESVI, left ventricular end-systolic volume index; LVEF, left ventricular ejection fraction; LVMl, left ventricular mass index. *p < 0.05 and {p < 0.01 for the comparison of the change in variable over time between carvedilol and placebo. Source: Bellenger NG, et al. Effects of carvedilol on left ventricular remodelling in chronic stable heart failure: a cardiovascular magnetic resonance study. Heart. 2004;90(7): 760–764, with permission.

on 36 patients with an occluded infarct-related artery who were randomized to PCI or medical management. All patients in this study received a contemporary medical regimen including aspirin, beta-blocker, ACE-I, and a lipidlowering agent in addition to clopidigrel if a stent was deployed. CMR was performed upon enrollment and at 6 months. To detect a mean change of 5% in end-systolic volume with alpha ¼ 0.05 and power ¼ 80%, only 34 patients were necessary. At 6 months, there was no difference in LV volumes. LVEF improved in the PCI group and became worse in the conservative management group. Both groups had similar mean circumferential shortening fractions at infarct, peri-infarct, and remote-infarct regions. Finally, no difference was noted in LV size and mass at 6 months. Bellenger and colleagues98 used CMR to assess whether the presence of infarct zone viability would effect

LV remodeling after late recanalization of an infarct related artery. Twenty-six patients from The Open Artery Trial (TOAT) were randomized to PCI (14 patients) or medical management (12 patients) after an anterior infarction with PCI performed on average of 26 days post-infarction. The initial CMR study, which included a dobutamine stress protocol, occurred at an average of 7.7 weeks. The follow-up CMR study occurred at an average of 10.8 months post-infarction. In the PCI group, there was a statistically significant correlation between the number of viable segments and the improvement in LVEF and end-systolic volume. No significant relationship existed between end-diastolic volume and viable segments in the PCI group. No correlation between viability and any of the aforementioned metrics of remodeling was noted in the conservatively treated group. Kirschbaum and colleagues99 studied 21 patients who underwent PCI for treatment of chronic total occlusions. In patients with LGE comprising less than 25% of the wall thickness of segments supplied by the chronically occluded artery, the affected segments experienced an improvement in systolic wall thickening at the 5-month and 3-year follow-up scans. In patients with LGE between 25% and 75% of the wall thickness of affected segments, no change in systolic wall thickening was observed at 5 months. However, a significant improvement was observed at 3 years. Among those patients with LGE greater than 75% of the wall thickness of affected segments, no improvement was seen at 5 months or 3 years. Finally, Baks and colleagues100 studied 27 patients with cine and LGE CMR before and at 5 months after successful drug-eluting stent implantation for total coronary occlusion. They found significant declines in LV end-systolic volume index and end-diastolic volume index, whereas the LVEF was unchanged. The extent of the LV that was dysfunctional yet viable on LGE was related to improvements in endsystolic volume index and LVEF. CMR is a particularly useful modality for assessing PCI, not only, as was previously mentioned, because of adequate power to detect end points with small sample sizes, but also because of the capability of localizing scar, assessing scar size, and determining viability in a single study. As was described above, CMR offers all the noninvasive parameters that one would demand when assessing postinfarction LV remodeling after myocardial stem cell transplantation. In the BOOST trial, 60 patients were randomized to autologous bone marrow cell intracoronary injection or medical therapy after successful PCI for an acute STsegment elevation infarction. Baseline CMR was performed on average at 3.5 days after PCI and at 6 months. While LVEF was similar in both groups at baseline, at 6 months, the control EF increased only 0.7%, while the stem cell group EF increased by 6.7% (p < 0.003). LV end-diastolic volume index, LV end-systolic volume index, LV mass index, and scar size did not differ between the control and stem cell groups from baseline to 6 months.101 The TOPCARE-AMI trial consisted of 59 patients randomized to receive an intracoronary injection of circulating progenitor cells or bone marrow derived progenitor cells. Stem cell infusion occurred 4.9 1.5 days after acute infarction. CMR was performed at baseline, 4 months, and 1 year. At 1 year, EF improved by 9.3 8% (p < 0.001), and scar volume decreased by 34 34% (p < 0.001).102 Cardiovascular Magnetic Resonance 263

19 ACUTE MYOCARDIAL INFARCTION: VENTRICULAR REMODELING

CARVEDILOL

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Finally, CMR has also been used in studies assessing nontraditional therapies of LV remodeling post-infarction. Dubach and colleagues103 studied 25 patients with LV dysfunction (EF 32 6%) and randomized them to 2 months of exercise in a rehabilitation program or control. Cine CMR was performed, and LV volumes, mass, and EF were measured. At the end of the training period, no differences were noted between groups in any of the aforementioned parameters. In addition, no differences were noted within groups over time. Exercise capacity increased in the treated group, but no deleterious effects on global parameters of LV remodeling were found, contrary to previously published data.104

CONCLUSION Owing to the acquisition of a volumetric dataset that provides superior reproducibility, CMR is particularly well suited to monitoring changes in LV size, shape, and scar burden in addition to regional and global function after infarction. Investigators are increasingly utilizing CMR in both animal and human/clinical studies to study the effects of pharmacologic and other interventions on the remodeling process. With its inherent precision and accuracy, CMR should allow studies of these interventions to occur with more power and in fewer subjects, increasing their cost-effectiveness.

References 1. Burlew BS, Weber KT. Connective tissue and the heart: functional significance and regulatory mechanisms. Cardiol Clin. 2000;18(3): 435–442. 2. Pfeffer MA, Braunwald E. Ventricular remodeling after myocardial infarction: experimental observations and clinical implications. Circulation. 1990;81(4):1161–1172. 3. Sutton MG, Sharpe N. Left ventricular remodeling after myocardial infarction: pathophysiology and therapy. Circulation. 2000;101(25): 2981–2988. 4. Lindsey ML, Mann DL, Entman ML, Spinale FG. Extracellular matrix remodeling following myocardial injury. Ann Med. 2003;35(5):316–326. 5. Laframboise WA, Bombach KL, Dhir RJ, et al. Molecular dynamics of the compensatory response to myocardial infarct. J Mol Cell Cardiol. 2005;38(1):103–117. 6. Reffelmann T, Kloner RA. The “no-reflow” phenomenon: basic science and clinical correlates. Heart. 2002;87(2):162–168. 7. Warren SE, Royal HD, Markis JE, Grossman W, McKay RG. Time course of left ventricular dilation after myocardial infarction: influence of infarct-related artery and success of coronary thrombolysis. J Am Coll Cardiol. 1988;11(1):12–19. 8. Weisman HF, Bush DE, Mannisi JA, Weisfeldt ML, Healy B. Cellular mechanisms of myocardial infarct expansion. Circulation. 1988;78(1): 186–201. 9. Gerber BL, Rochitte CE, Melin JA, McVeigh ER, Bluemke DA, Wu KC, et al. Microvascular obstruction and left ventricular remodeling early after acute myocardial infarction. Circulation. 2000;101(23): 2734–2741. 10. Gulch RW, Jacob R. Geometric and muscle physiological determinants of cardiac stroke volume as evaluated on the basis of model calculations. Basic Res Cardiol. 1988;83(5):476–485. 11. Anversa P, Loud AV, Levicky V, Guideri G. Left ventricular failure induced by myocardial infarction. I. Myocyte hypertrophy. Am J Physiol. 1985;248(6 Pt 2):H876–H882. 12. Kramer CM, Rogers WJ, Park CS, et al. Regional myocyte hypertrophy parallels regional myocardial dysfunction during post-infarct remodeling. J Mol Cell Cardiol. 1998;30(9):1773–1778. 13. Hammermeister KE, DeRouen TA, Dodge HT. Variables predictive of survival in patients with coronary disease: selection by univariate and multivariate analyses from the clinical, electrocardiographic, exercise, arteriographic, and quantitative angiographic evaluations. Circulation. 1979;59(3):421–430. 14. Migrino RQ, Young JB, Ellis SG, et al. End-systolic volume index at 90 to 180 minutes into reperfusion therapy for acute myocardial infarction is a strong predictor of early and late mortality. The Global Utilization of Streptokinase and t-PA for Occluded Coronary Arteries (GUSTO)-I Angiographic Investigators. Circulation. 1997;96(1): 116–121. 15. Nicolosi GL, Latini R, Marino P, et al. The prognostic value of predischarge quantitative two-dimensional echocardiographic measurements and the effects of early lisinopril treatment on left ventricular structure and function after acute myocardial infarction in the GISSI-3 Trial. Gruppo Italiano per lo Studio della Sopravvivenza nell’Infarto Miocardico. Eur Heart J. 1996;17(11):1646–1656. 16. St John SM, Pfeffer MA, Moye L, et al. Cardiovascular death and left ventricular remodeling two years after myocardial infarction: baseline

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55. Yang Z, Berr SS, Gilson WD, Toufektsian MC, French BA. Simultaneous evaluation of infarct size and cardiac function in intact mice by contrast-enhanced cardiac magnetic resonance imaging reveals contractile dysfunction in noninfarcted regions early after myocardial infarction. Circulation. 2004;109(9):1161–1167. 56. Rochitte CE, Lima JA, Bluemke DA, et al. Magnitude and time course of microvascular obstruction and tissue injury after acute myocardial infarction. Circulation. 1998;98(10):1006–1014. 57. Fieno DS, Hillenbrand HB, Rehwald WG, et al. Infarct resorption, compensatory hypertrophy, and differing patterns of ventricular remodeling following myocardial infarctions of varying size. J Am Coll Cardiol. 2004;43(11):2124–2131. 58. Lund GK, Stork A, Muellerleile K, et al. Prediction of left ventricular remodeling and analysis of infarct resorption in patients with reperfused myocardial infarcts by using contrast-enhanced MR imaging. Radiology. 2007;245(1):95–102. 59. Kaandorp TA, Lamb HJ, Viergever EP, et al. Scar tissue on contrastenhanced MRI predicts left ventricular remodelling after acute infarction. Heart. 2007;93(3):375–376. 60. Baks T, van Geuns RJ, Biagini E, et al. Effects of primary angioplasty for acute myocardial infarction on early and late infarct size and left ventricular wall characteristics. J Am Coll Cardiol. 2006;47:40–44. 61. Bodi V, Sanchis J, Nunez J, et al. Prognostic value of a comprehensive cardiac magnetic resonance assessment soon after a first ST-segment elevation myocardial infarction. J Am Coll Cardiol Img. 2009;2(7):835–842. 62. Watzinger N, Lund GK, Higgins CB, Wendland MF, Weinmann HJ, Saeed M. The potential of contrast-enhanced magnetic resonance imaging for predicting left ventricular remodeling. J Magn Reson Imaging. 2002;16(6):633–640. 63. Friedrich MG, Abdel-Aty H, Taylor A, et al. The salvaged area at risk in reperfused acute myocardial infarction as visualized by cardiovascular magnetic resonance. J Am Coll Cardiol. 2008;51:1581–1587. 64. Ganame J, Messali G, Dymarkowski S, et al. Impact of myocardial haemorrhage on left ventricular function and remoelling in patients with reperfused acute myocardial infarction. Eur Heart J. 2009;30: 1440–1449. 65. Saeed M, Wendland MF, Seelos K, Masui T, Derugin N, Higgins CB. Effect of cilazapril on regional left ventricular wall thickness and chamber dimension following acute myocardial infarction: in vivo assessment using MRI. Am Heart J. 1992;123(6):1472–1480. 66. Kramer CM, Ferrari VA, Rogers WJ, et al. Angiotensin-converting enzyme inhibition limits dysfunction in adjacent noninfarcted regions during left ventricular remodeling. J Am Coll Cardiol. 1996; 27(1):211–217. 67. McDonald KM, Garr M, Carlyle PF, et al. Relative effects of alpha 1adrenoceptor blockade, converting enzyme inhibitor therapy, and angiotensin II subtype 1 receptor blockade on ventricular remodeling in the dog. Circulation. 1994;90(6):3034–3046. 68. McDonald KM, Mock J, D’Aloia A, et al. Bradykinin antagonism inhibits the antigrowth effect of converting enzyme inhibition in the dog myocardium after discrete transmural myocardial necrosis. Circulation. 1995;91(7):2043–2048. 69. McDonald KM, Rector T, Carlyle PF, Francis GS, Cohn JN. Angiotensinconverting enzyme inhibition and beta-adrenoceptor blockade regress established ventricular remodeling in a canine model of discrete myocardial damage. J Am Coll Cardiol. 1994;24(7):1762–1768. 70. Kramer CM, Nicol PD, Rogers WJ, Seibel PS, Park CS, Reichek N. Beta-blockade improves adjacent regional sympathetic innervation during postinfarction remodeling. Am J Physiol. 1999;277(4 Pt 2): H1429–H1434. 71. Mankad S, d’Amato TA, Reichek N, et al. Combined angiotensin II receptor antagonism and angiotensin-converting enzyme inhibition further attenuates postinfarction left ventricular remodeling. Circulation. 2001;103(23):2845–2850. 72. Yang Z, Bove CM, French BA, et al. Angiotensin II type 2 receptor overexpression preserves left ventricular function after myocardial infarction. Circulation. 2002;106(1):106–111. 73. Bove CM, Gilson WD, Scott CD, et al. The angiotensin II type 2 receptor and improved adjacent region function post-MI. J Cardiovasc Magn Reson. 2005;7(2):459–464. 74. Bove CM, Yang Z, Gilson WD, et al. Nitric oxide mediates benefits of angiotensin II type 2 receptor overexpression during post-infarct remodeling. Hypertension. 2004;43(3):680–685. 75. Voros S, Yang Z, Bove C, et al. The interaction between the AT1 and AT2 receptor during post-infarction left ventricular remodeling. Am J Physiol Heart Circ Physiol. 2005;.

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31. Helm P, Beg MF, Miller MI, Winslow RL. Measuring and mapping cardiac fiber and laminar architecture using diffusion tensor MR imaging. Ann N Y Acad Sci. 2005;1047:296–307. 32. Chen J, Song SK, Liu W, et al. Remodeling of cardiac fiber structure after infarction in rats quantified with diffusion tensor MRI. Am J Physiol Heart Circ Physiol. 2003;285(3):H946–H954. 33. Holman ER, Buller VG, de Roos A, et al. Detection and quantification of dysfunctional myocardium by magnetic resonance imaging: a new three-dimensional method for quantitative wall-thickening analysis. Circulation. 1997;95(4):924–931. 34. Kramer CM, Rogers WJ, Theobald TM, Power TP, Petruolo S, Reichek N. Remote noninfarcted region dysfunction soon after first anterior myocardial infarction: a magnetic resonance tagging study. Circulation. 1996;94(4):660–666. 35. Axel L, Dougherty L. Heart wall motion: improved method of spatial modulation of magnetization for MR imaging. Radiology. 1989;172(2): 349–350. 36. Zerhouni EA, Parish DM, Rogers WJ, Yang A, Shapiro EP. Human heart: tagging with MR imaging: a method for noninvasive assessment of myocardial motion. Radiology. 1988;169(1):59–63. 37. Moulton MJ, Downing SW, Creswell LL, et al. Mechanical dysfunction in the border zone of an ovine model of left ventricular aneurysm. Ann Thorac Surg. 1995;60(4):986–997. 38. Epstein FH, Yang Z, Gilson WD, Berr SS, Kramer CM, French BA. MR tagging early after myocardial infarction in mice demonstrates contractile dysfunction in adjacent and remote regions. Magn Reson Med. 2002;48(2):399–403. 39. Thomas D, Ferrari VA, Janik M, et al. Quantitative assessment of regional myocardial function in a rat model of myocardial infarction using tagged MRI. MAGMA. 2004;17(3–6):179–187. 40. Rademakers F, Van de WF, Mortelmans L, Marchal G, Bogaert J. Evolution of regional performance after an acute anterior myocardial infarction in humans using magnetic resonance tagging. J Physiol. 2003;546(Pt 3):777–787. 41. Kim D, Gilson WD, Kramer CM, Epstein FH. Myocardial tissue tracking with two-dimensional cine displacement-encoded MR imaging: development and initial evaluation. Radiology. 2004;230(3):862–871. 42. McDonald KM, Yoshiyama M, Francis GS, Ugurbil K, Cohn JN, Zhang J. Myocardial bioenergetic abnormalities in a canine model of left ventricular dysfunction. J Am Coll Cardiol. 1994;23(3):786–793. 43. Zhang J, McDonald KM. Bioenergetic consequences of left ventricular remodeling. Circulation. 1995;92(4):1011–1019. 44. Friedrich J, Apstein CS, Ingwall JS. 31P nuclear magnetic resonance spectroscopic imaging of regions of remodeled myocardium in the infarcted rat heart. Circulation. 1995;92(12):3527–3538. 45. Zhang J, Wilke N, Wang Y, et al. Functional and bioenergetic consequences of postinfarction left ventricular remodeling in a new porcine model. MRI and 31 P-MRS study. Circulation. 1996;94(5):1089–1100. 46. Straeter-Knowlen IM, Evanochko WT, den Hollander JA, Wolkowicz PE, Balschi JA. Caulfield JB et al. 1H NMR spectroscopic imaging of myocardial triglycerides in excised dog hearts subjected to 24 hours of coronary occlusion. Circulation. 1996;93(7): 1464–1470. 47. Katz AM, Messineo FC. Lipid-membrane interactions and the pathogenesis of ischemic damage in the myocardium. Circ Res. 1981; 48 (1):1–16. 48. van der Vusse GJ, van Bilsen M, Jans SW, Reneman RS. Lipid metabolism in the ischemic and reperfused heart. EXS. 1996;76:175–190. 49. Hillenbrand HB, Becker LC, Kharrazian R, et al. 23Na MRI combined with contrast-enhanced 1H MRI provides in vivo characterization of infarct healing. Magn Reson Med. 2005;53(4):843–850. 50. Wu KC, Lima JA. Noninvasive imaging of myocardial viability: current techniques and future developments. Circ Res. 2003;93(12):1146–1158. 51. Thiel H, Kappl MJE, Conradi S, et al. Reproducibility of chronic and acute infarct size measurement by delayed enhancement-magnetic resonance imaging. J Am Coll Cardiol. 2006;47:1641–1645. 52. Judd RM, Lugo-Olivieri CH, Arai M, et al. Physiological basis of myocardial contrast enhancement in fast magnetic resonance images of 2-day-old reperfused canine infarcts. Circulation. 1995;92(7): 1902–1910. 53. Wu KC, Zerhouni EA, Judd RM, et al. Prognostic significance of microvascular obstruction by magnetic resonance imaging in patients with acute myocardial infarction. Circulation. 1998;97(8):765–772. 54. Choi CJ, Haji-Momenian S, DiMaria JM, et al. Infarct involution and improved function during healing of acute myocardial infarction: the role of microvascular obstruction. J Cardiovasc Magn Reson. 2004; 6(4):917–925.

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76. Isbell DC, Voros S, Yang Z, et al. Interaction between bradykinin subtype 2 and angiotensin II type 2 receptors during post-MI left ventricular remodeling. Am J Physiol Heart Circ Physiol. 2007;293(6): H3372–H3378. 77. Gilson WD, Epstein FH, Yang Z, et al. Borderzone contractile dysfunction is transiently attenuated and left ventricular structural remodeling is markedly reduced following reperfused myocardial infarction in inducible nitric oxide synthase knockout mice. J Am Coll Cardiol. 2007;50(18):1799–1807. 78. Roland E. Safety profile of an anti-anginal agent with potassium channel opening activity: an overview. Eur Heart J. 1993;14(suppl B): 48–52. 79. Krombach GA, Higgins CB, Chujo M, Saeed M. Blood pool contrastenhanced MRI detects suppression of microvascular permeability in early postinfarction reperfusion after nicorandil therapy. Magn Reson Med. 2002;47(5):896–902. 80. Saeed M, Watzinger N, Krombach GA, et al. Left ventricular remodeling after infarction: sequential MR imaging with oral nicorandil therapy in rat model. Radiology. 2002;224(3):830–837. 81. Nahrendorf M, Hu K, Hiller KH, et al. Impact of hydroxymethylglutaryl coenzyme a reductase inhibition on left ventricular remodeling after myocardial infarction: an experimental serial cardiac magnetic resonance imaging study. J Am Coll Cardiol. 2002;40 (9):1695–1700. 82. Nahrendorf M, Frantz S, Hu K, et al. Effect of testosterone on postmyocardial infarction remodeling and function. Cardiovasc Res. 2003;57(2):370–378. 83. Nahrendorf M, Hiller KH, Theisen D, et al. Effect of transmyocardial laser revascularization on myocardial perfusion and left ventricular remodeling after myocardial infarction in rats. Radiology. 2002; 225(2):487–493. 84. Blom AS, Pilla JJ, Arkles J, et al. Ventricular restraint prevents infarct expansion and improves borderzone function after myocardial infarction: a study using magnetic resonance imaging, three-dimensional surface modeling, and myocardial tagging. Ann Thorac Surg. 2007; 84(6):2004–2010. 85. Limbourg FP, Ringes-Lichtenberg S, Schaefer A, et al. Haematopoietic stem cells improve cardiac function after infarction without permanent cardiac engraftment. Eur J Heart Fail. 2005;7 (5):722–729. 86. Amado LC, Saliaris AP, Schuleri KH, et al. Cardiac repair with intramyocardial injection of allogeneic mesenchymal stem cells after myocardial infarction. Proc Natl Acad Sci U S A. 2005;102(32): 11474–11479. 87. van den Bos EJ, Thompson RB, Wagner A, et al. Functional assessment of myoblast transplantation for cardiac repair with magnetic resonance imaging. Eur J Heart Fail. 2005;7(4):435–443. 88. Zeng L, Hu Q, Wang X, et al. Bioenergetic and functional consequences of bone marrow-derived multipotent progenitor cell transplantation in hearts with postinfarction left ventricular remodeling. Circulation. 2007;115(14):1866–1875. 89. Schulman SP, Weiss JL, Becker LC, et al. Effect of early enalapril therapy on left ventricular function and structure in acute myocardial infarction. Am J Cardiol. 1995;76(11):764–770.

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90. Johnson DB, Foster RE, Barilla F, et al. Angiotensin-converting enzyme inhibitor therapy affects left ventricular mass in patients with ejection fraction > 40% after acute myocardial infarction. J Am Coll Cardiol. 1997;29(1):49–54. 91. Foster RE, Johnson DB, Barilla F, et al. Changes in left ventricular mass and volumes in patients receiving angiotensin-converting enzyme inhibitor therapy for left ventricular dysfunction after Q-wave myocardial infarction. Am Heart J. 1998;136(2):269–275. 92. Groenning BA, Nilsson JC, Sondergaard L, Fritz-Hansen T, Larsson HB, Hildebrandt PR. Antiremodeling effects on the left ventricle during beta-blockade with metoprolol in the treatment of chronic heart failure. J Am Coll Cardiol. 2000;36(7):2072–2080. 93. Dubach P, Myers J, Bonetti P, et al. Effects of bisoprolol fumarate on left ventricular size, function, and exercise capacity in patients with heart failure: analysis with magnetic resonance myocardial tagging. Am Heart J. 2002;143(4):676–683. 94. Bellenger NG, Rajappan K, Rahman SL, et al. Effects of carvedilol on left ventricular remodelling in chronic stable heart failure: a cardiovascular magnetic resonance study. Heart. 2004;90(7):760–764. 95. St John SM, Pfeffer MA, Plappert T, et al. Quantitative twodimensional echocardiographic measurements are major predictors of adverse cardiovascular events after acute myocardial infarction. The protective effects of captopril. Circulation. 1994;89(1):68–75. 96. Pfisterer ME, Buser P, Osswald S, Weiss P, Bremerich J, Burkart F. Time dependence of left ventricular recovery after delayed recanalization of an occluded infarct-related coronary artery: findings of a pilot study. J Am Coll Cardiol. 1998;32(1):97–102. 97. Silva JC, Rochitte CE, Junior JS, et al. Late coronary artery recanalization effects on left ventricular remodelling and contractility by magnetic resonance imaging. Eur Heart J. 2005;26(1):36–43. 98. Bellenger NG, Yousef Z, Rajappan K, Marber MS, Pennell DJ. Infarct zone viability influences ventricular remodelling after late recanalisation of an occluded infarct related artery. Heart. 2005;91(4):478–483. 99. Kirschbaum SW, Baks T, van den EM, et al. Evaluation of left ventricular function three years after percutaneous recanalization of chronic total coronary occlusions. Am J Cardiol. 2008;101(2):179–185. 100. Baks T, van Geuns RJ, Duncker DJ, et al. Prediction of left ventricular function after drug-eluting stent implantation for chronic total coronary occlusions. J Am Coll Cardiol. 2006;47:721–725. 101. Wollert KC, Meyer GP, Lotz J, et al. Intracoronary autologous bonemarrow cell transfer after myocardial infarction: the BOOST randomised controlled clinical trial. Lancet. 2004;364(9429):141–148. 102. Schachinger V, Assmus B, Britten MB, et al. Transplantation of progenitor cells and regeneration enhancement in acute myocardial infarction: final one-year results of the TOPCARE-AMI Trial. J Am Coll Cardiol. 2004;44(8):1690–1699. 103. Dubach P, Myers J, Dziekan G, et al. Effect of exercise training on myocardial remodeling in patients with reduced left ventricular function after myocardial infarction: application of magnetic resonance imaging. Circulation. 1997;95(8):2060–2067. 104. Jugdutt BI, Michorowski BL, Kappagoda CT. Exercise training after anterior Q wave myocardial infarction: importance of regional left ventricular function and topography. J Am Coll Cardiol. 1988; 12(2):362–372.

Myocardial Viability Udo P. Sechtem and Heiko Mahrholdt The detection of residual myocardial viability in a patient with regional or global severe left ventricular (LV) dysfunction is of clinical importance to plan the therapeutic strategy because revascularization of dysfunctional but viable myocardium may improve LV function.1 Several imaging techniques have been shown to be successful in detecting myocardial viability; these include LV angiography using appropriate interventions,2 perfusion scintigraphy, positron emission tomography (PET),3 and echocardiography.4 More recently, cardiovascular magnetic resonance (CMR) techniques gained widespread acceptance as a technique to identify viable myocardium and distinguish it from myocardial necrosis and scar.5–8 This chapter will review the current knowledge of how these techniques can be used in humans to guide clinical decision making and predict recovery of function after revascularization of dysfunctional myocardium. All scientific papers on the value of imaging techniques for detecting myocardial viability use the well-known statistical terms of sensitivity, specificity, and the predictive values. The sensitivity of a test is commonly defined as the number of true positives divided by the sum of true positives and false negatives. In other words, the sensitivity of a test is the number of diseased persons with a positive test divided by the total number of diseased persons. Common sense might suggest that scar tissue and hence absence of viability might indicate the presence of disease. The total number of diseased people or segments would thus be the scarred people or segments without recovery of function. The presence of viability would be synonymous with a healthy state (relatively speaking), and the total number of viable, healthy people or segments would appear as the denominator in the formula for calculating specificity. In the viability literature, however, sensitivity indicates the ability of a test to identify viable myocardium, and specificity is an indicator how well the test performs when it comes to detecting scar. Positive and negative predictive values are used accordingly. This needs to be borne in mind when it comes to the interpretation of test results reported later in this chapter.

FEATURES OF VIABLE MYOCARDIUM DETECTABLE BY CARDIOVASCULAR MAGNETIC RESONANCE Scar Formation and Left Ventricular Wall Thickness Myocardium is commonly defined as viable if it shows severe dysfunction at baseline but recovers function with time either spontaneously (myocardial stunning) or following revascularization (hibernating myocardium). Clinically,

stunned myocardium may be found in patients with early reperfusion of an infarct-related artery. If there is no residual high-grade stenosis, blood flow at rest will be normal and the myocardium will recover spontaneously after a few days. Patients with hibernating myocardium often present with severe triple vessel disease, globally depressed LV function, and prominent dyspnea but surprisingly little angina.9 This type of dysfunction is often more chronic, and previous myocardial infarction may or may not be reported in the history. Pathology may reveal regions of transmural scar, regions with predominantly subendocardial scar, and regions with mixtures of scar and viable myocardium. Severe wall thinning is the hallmark of transmural chronic myocardial infarction. However, wall thinning is the end result of infarct healing, and it may take up to 4 months before the remodeling process is completed.10 In contrast to the severe thinning of chronic transmural scar, the best example for which is the thin-walled anterior LV aneurysm, acute and subacute transmural infarcts may not yet have reached the stage of thinning because local infarct remodeling is incomplete.11 In contrast to transmural myocardial infarction, which may or may not appear thinned, depending on infarct age, healed nontransmural infarcts usually do not develop severe thinning. Some thinning may be observed, however, depending on the degree to which the endocardially located infarct extends throughout the wall. Therefore, the finding of preserved myocardial wall thickness in diastole in a patient with a known chronic infarct that is more than 4 months old will likely represent nontransmural infarction with a substantial rim of viable myocardium surrounding the endocardial scar. If the infarct is more recent than approximately 4 months, preserved end-diastolic wall thickness cannot be used to distinguish between viable and nonviable myocardium. Recently, a patient was described with a small subendocardial infarct with a thickness of 1.5 mm covered by an epicardially located rim of viable myocardium measuring 3.5 mm.12 Although diastolic wall thickness was only 5 mm, recovery of the myocardium was demonstrated. Thus, regional wall thinning may be possible under exceptional circumstances in myocardium that has only little subendocardial scarring probably due to ventricular remodeling, and full recovery of function may occur following revascularization. The ratio of viable to total myocardium (viable plus nonviable) irrespective of wall thickness in the dysfunctional region may therefore be more accurate than end-diastolic wall thickness in predicting functional improvement.

Contractile Reserve of Viable Myocardium A well-known feature of viable myocardium is augmented contractility in response to a suitable stimulus.13 Such stimuli Cardiovascular Magnetic Resonance 267

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include sympathomimetic agents14 or postextrasystolic potentiation.2 In contrast, necrotic or scarred tissue will not respond to such stimulation. Today, the most widely used mode of stimulation is to infuse low doses of dobutamine up to 10 mg/kg/min. If a contractile reserve can be elicited, the responsive myocardium will usually recover function after appropriate revascularization.15 However, it appears that there is also some spontaneous improvement in the response to dobutamine over the course of infarct healing after reperfused myocardial infarction, which may affect the accuracy of this viability marker after myocardial infarction.11

Noninvasive Observation of Tissue Edema Irreversible damage to myocardial occurs after approximately 30 to 120 minutes of ischemia. Very early changes can be observed by electron microscopy; these changes include intracellular edema and swelling of the entire cell, including the mitochondria. The sarcolemma ruptures, and there is free exchange between the extracellular and intracellular compartments. In some infarcts, light microscopy reveals changes just a few hours after the onset of ischemia; these changes are most pronounced at the periphery of the infarct. After 8 hours, there is edema of the interstitium, and infiltration of the infarct zone by neutrophils and red blood cells becomes evident. Small blood vessels undergo necrosis and karyolysis of muscle cell nuclei can be observed.16 Plugging of capillaries by erythrocytes is most pronounced in the center of the infarct. If reperfusion can be achieved at an early stage, the resulting infarcts contain a mixture of necrosis and hemorrhage within zones of irreversibly injured myocytes. Myocardial edema is associated with prolonged magnetic resonance (MR) relaxation times, and this leads to characteristically increased signal intensity on MR images, which are sensitive to such changes.17 By using modern T2-weighted pulse sequences, the edema associated and surrounding acute infarcts can be depicted, and this can be used to differentiate between acute and chronic infarcts.18

The No-Reflow Phenomenon and Early Hypoenhancement with Gadolinium A feature of the central necrotic region within a myocardial infarct is intracapillary red blood cell stasis.19 Plugging of the capillaries leads to tissue hypoperfusion. This hypoperfusion is primarily related to a reduced functional capillary density rather than reduced microvascular flow rates.20 A decrease in functional capillary density results in a prolonged washin time constant. This lack of reperfusion after restoration of flow in epicardial vessel is known as the noreflow phenomenon. When the myocardium is imaged by CMR early after injection after gadolinium (early gadolinium enhancement), no-reflow zones appear dark in comparison with the surrounding rim regions of the infarct.21 268 Cardiovascular Magnetic Resonance

Late Gadolinium Enhancement in Infarcted Tissue Gadolinium chelates are commonly used as CMR contrast agents. These metabolically inert molecules are distributed extracellularly, and they shorten both T1 and T2 relaxation times. Rupture of myocyte membranes leads to an increased volume of distribution of CMR contrast agents with a corresponding increase in the effective voxel concentration of such agents.22–24 Thus, a higher concentration of gadolinium contrast agents leads to a more pronounced shortening of relaxation times. As ECG gated CMR images are usually T1 weighted (because the R-R interval is approximately 800 msec, which corresponds to the T1 value of myocardium), this will result in a higher signal intensity of infarcted as compared with normal tissue once the contrast material has fully penetrated the infarct region. The time-concentration curve of MR contrast agents in infarct tissue does not correspond to that in blood or normal tissue kinetics. Thus, while early hypoenhancement of infarcted regions after injection of contrast material is due to delayed contrast penetration,20 late enhancement in infarction is due to both increased volume of distribution and slow contrast washout.19,25 The enhancement pattern that is seen will depend on regional differences in tissue washin/washout kinetics as well as the time after injection of contrast when the image is acquired. Late gadolinium enhancement (LGE) has now been extensively validated in animal and human studies, and with improved imaging sequences (notably the use of inversion recovery to null signal from normal myocardium), the signalto-noise ratio of enhanced to unenhanced tissue is dramatically higher than with previous sequences, at approximately 500%. This has led to greatly improved image quality and a substantial increase in use of the technique. In animal experiments, the area of LGE has been shown to correlate closely with areas of infarction,26 and for the first time in vivo, high-quality imaging of the distribution of scar is possible.

High-Energy Phosphates and Viability The primary energy reserve in living myocardial cells is stored in form of creatine phosphate and ATP. Depletion of total myocardial creatine, creatine phosphate, and ATP follows severe ischemic injury as shown in biopsy samples obtained from patients during cardiac surgery27 or necropsy.28 Using 31 P magnetic resonance spectroscopy (MRS), it is possible to measure the myocardial content of phosphocreatine and ATP.29 1H-MRS has a higher sensitivity than 31P-MRS and has the ability to detect the total pool of phosphorylated plus unphosphorylated creatine in skeletal and cardiac muscle.30

Sodium and Potassium Cardiovascular Magnetic Resonance Imaging Although much of this work has been performed in animals, there are now some early human results; therefore, this brief section is included for completeness. When cell

CARDIOVASCULAR MAGNETIC RESONANCE TO DETECT VIABLE MYOCARDIUM IN ACUTE MYOCARDIAL INFARCTION Signal Intensity Changes on T2-Weighted Images Myocardial edema accompanies acute myocardial necrosis. On T2-weighted spin echo images, the increased water content leads to an increase in signal intensity.36 In animal models, a good correlation between water content and T2 relaxation time or T2-weighted signal intensity has been described.37 The area at risk measured by CMR correlates well with that determined at pathology, but the zone of increased signal intensity overestimates the infarct area determined by triphenyltetrazolium chloride staining.36 T2-weighted spin echo images acquired early after myocardial infarction (within 10 days) demonstrate that the infarct site is a region of high signal intensity in comparison with normal myocardium.38 The advent of modern rapid pulse sequences, which produce a higher contrast between edematous and normal myocardium, has led to some revival of T2-weighted CMR.39 Friedrich and colleagues studied patients with reperfused acute myocardial infarction using a T2-weighted breath hold short inversion time (short TI) inversion recovery pulse sequence and LGE images.40 In comparison to the zone of high signal intensity on late enhancement images corresponding to the area of irreversible ischemic damage, the high signal intensity zone on T2weighted images was larger, corresponding to the area of reversible damage. Hence, the combination of these two techniques allows for quantifying the extent of the salvaged area after revascularization. However, there are several potential pitfalls to the interpretation of T2-weighted images, including the necessity to differentiate signal from slowly flowing blood in the ventricle from increased signal intensity from a region of infarction and to recognize artifactual variation of signal intensity in the myocardium due to respiratory motion or residual cardiac motion.

Contrast-Enhanced Studies Using Spin Echo Cardiovascular Magnetic Resonance Detection of areas of acute myocardial infarction can be improved by using the effects of gadolinium-based contrast agents and image quality on contrast-enhanced T1weighted images is superior to unenhanced T2-weighted images.24 However, the signal intensity difference between enhancing and nonenhancing regions of the left ventricle was rather small, and image quality was still suboptimal in many patients owing to long acquisition times resulting in breathing artifacts,41,42 which prevented major clinical applications of T1-weighted spin echo imaging for the detection of viable myocardium.

Late Enhancement with Gadolinium in Acute Infarction In the mid and late 1990s, Kim and Judd developed late enhancement imaging.19,25 Within a short time, LGE CMR became widely used because of its high image quality, simplicity, and high resolution allowing clear demarcation of the transmural extent of necrosis and scar. The T1weighted segmented inversion recovery pulse sequence that is employed in most centers acquires images in mid-diastole when cardiac motion is minimal (Fig. 20-1).43 Segmentation of k-space makes it possible to acquire images during a breath hold, which reduces motion artifacts as compared with the older T1-weighted spin echo techniques. With the appropriate choice of the inversion time and imaging 10 to 20 minutes after the intravenous application of contrast material, the signal intensity of normal myocardium is nulled but the infarcted tissue becomes very bright (Fig. 20-2). Kim and colleagues26 clearly showed that this technique provided precise images of acute and subacute infarcts irrespective of transmurality and reperfusion status (Fig. 20-3). The time between contrast injection and starting the imaging process is important, as contrast concentration and hence enhancement of a region vary over time. Normal regions with normal contrast washin and washout reach a constant enhancement within 2 to 3 minutes.19 The border zones of an infarct, however, have longer time

R

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TI

Figure 20-1 Segmented inversion recovery fast gradient echo sequence with TI set to null normal myocardium after contrast agent administration. Source: From Simonetti OP et al. An improved MR imaging technique for the visualization of myocardial infarction. Radiology. 2001;218:215–223. ECG, electrocardiogram. Cardiovascular Magnetic Resonance 269

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membranes break down at the time of infarction, there are two biochemical changes with inorganic elements that can be detected using CMR. The first is the increase in sodium concentration, and this occurs because the volume of muscle previously occupied by the intact cell has low sodium concentration. Sodium imaging using CMR therefore shows a bright signal at the site of acute infarction.31–33 By contrast, the potassium concentration, which is high in the intracellular space but low extracellularly, falls markedly with the loss of cell membrane integrity, and potassium imaging therefore shows a dark area with infarction.34,35 Such imaging requires a multifrequency CMR system, which is very limited in availability, but the technique has the potential to distinguish acute from chronic infarction on the basis of these acute cation fluxes, and it is conceivable that clinical application might be found.

ISCHEMIC HEART DISEASE

Figure 20-2 Nontransmural inferior myocardial infarct (left panel). This patient presented with acute onset of ST-elevation myocardial infarction and was revascularized by catheterization approximately 120 minutes after first onset of symptoms. The transmural inferolateral myocardial infarct displayed in the right panel is several months old. Note the thinning of the infarcted area, which is considered to be a hallmark of transmural infarcts.

Figure 20-3 Postmortem images of a dog with a 3-day-old myocardial infarct. Left, Triphenyltetrazolium chloride (TTC) stained slices of the left ventricle showing the infarct as nonstained white areas, whereas normal myocardial cells stain red. Slices are arranged from base to apex, starting in the upper left and advancing from left to right and then from top to bottom. Right, Insert showing a magnification. There is excellent matching of the enhanced areas in the ex vivo magnetic resonance images with the necrotic zones in the necropsy slices. Source: From Kim RJ et al. Relationship of MRI delayed contrast enhancement to irreversible injury, infarct age, and contractile function. Circulation. 1999;100(19):1992–2002. 270 Cardiovascular Magnetic Resonance

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A

B

C

Figure 20-4 Temporal changes in signal intensity after 0.2 mmol/kg gadolinium-DTPA administered intravenously in a patient after primary percutaneous transluminal coronary angioplasty plus stenting of the left anterior descending coronary artery for anterior myocardial infarction. Magnified view of three-chamber view at 2 minutes (A), 15 minutes (B), and 30 minutes (C). Source: From Beek AM et al. Delayed contrast-enhanced magnetic resonance imaging for the prediction of regional functional improvement after acute myocardial infarction. J Am Coll Cardiol. 2003;42(5):895–901.

constants for washin and washout as compared to normal, and these time constants are even longer in the center of the infarct (see below). Consequently, both core and border zone of the infarct appear dark during the early phase of contrast washin (Fig. 20-4). Later, core and normal myocardium have similar signal intensities, but the border zone begins showing some enhancement. When contrast washout begins in normal myocardium at about 10 minutes after injection, core and border zones of the infarct are enhanced, and this persists in the core of the infarct at late stages of contrast washout. At this time, the border zones have already returned to near-normal signal intensities. There has been debate about whether LGE occurs exclusively in regions with myocardial necrosis or also in edematous viable border zones around infarcted areas.44 However, the observation by Saeed and colleagues that the Gd-DTPA-enhanced region overestimates true infarct size by approximately 8% was not supported by other animal studies.45 Fieno and colleagues compared ex vivo MR with triphenyltetrazolium chloride–stained sections and confirmed that the spatial extent of enhancement was the same as the spatial extent of infarction at every stage of healing from day 1 to 8 weeks.46 The ischemic bed at risk was defined by fluorescent microparticles injected into the left atrium with the infarct-related artery occluded. Enhanced regions were smaller than the ischemic bed at risk at every stage of healing (Fig. 20-5). Image intensities of viable myocardium within the risk region were the same as those of remote, normal myocardium. A possible explanation for the discrepant findings of Saeed and colleagues and the group of Kim and Judd is that Saeed and colleagues used a different pulse sequence than that found to be highly and reproducibly accurate for infarct sizing. It is important to note that accurate infarct imaging requires constant adjustment of the TI if imaging cannot be completed within 5 minutes. This is necessary, as the null point of normal myocardium depends on the concentration of the contrast agent at any given point in time. As contrast washout from normal myocardium occurs at a comparatively rapid pace, TI increases continuously, necessitating corresponding increases of TI to achieve perfect nulling of normal myocardium. Typical TI values range

from 310 msec at 10 minutes to 385 msec at 30 minutes (Fig. 20-6). With appropriate adjustment of TI, there are no differences in infarct size measurement by contrastenhanced MR between 10 and 30 minutes after contrast injection (Fig. 20-7).47 If TI is not adjusted, the spatial extent of enhancement decreases by up to 30% over time after contrast application.48

No-Reflow by Cardiovascular Magnetic Resonance MR imaging of acutely necrotic myocardium sometimes reveals a central core of low signal intensity surrounded by the typical bright infarct zone (Fig. 20-8). In the center of the infarct region, myocytes and capillaries may undergo necrosis simultaneously because of profound and sustained ischemia. In that situation, capillaries become occluded by dying blood cells and debris, to the extent that even with restoration of epicardial blood flow, the infarct core will not promptly reperfuse. This area of microvascular obstruction is called the no-reflow region.18 Microvascular obstruction following acute infarction correlates with greater myocardial damage by echocardiography and poorer global LV function in the early postinfarction phase.49 The central low signal intensity region seen on CMR images within the first 2 to 3 minutes after contrast injection corresponds to experimentally produced no-reflow regions.50 CMR enables in vivo monitoring of the time course of microvascular obstruction. It could be shown that this zone increases threefold during the 48 hours after reperfusion, indicating progressive microvascular injury well beyond coronary occlusion and reflow. Moreover, the extent of microvascular obstruction is the strongest predictor of LV volume at 10 days after reperfusion. Large areas of microvascular obstruction within infarcts lead to significant reductions of radial thickening in adjacent noninfarcted regions.51 CMR could also prove in a dog model that intraaortic balloon counterpulsation improves myocardial perfusion at the tissue level and reduces the extent of no-reflow caused by microvascular obstruction.52 CMR may also be used to study the effects of the noreflow phenomenon in patients with acute infarcts. Wu and co-workers studied 44 patients by using perfusion CMR.53 Almost all of these patients had thrombolysis or Cardiovascular Magnetic Resonance 271

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Figure 20-5 Dog with a 1-day-old reperfused infarct. Upper left, The ex vivo magnetic resonance enhancement image. Middle left, The matching TTC-stained myocardial slice. Bottom left, The myocardium at risk as the blue-appearing region without fluorescent microparticles. Right, Light microscopy views of region 1 (not at risk, not infarcted), region 2 (at risk but not infarcted), and region 3 (infarcted). Arrows point to contraction bands. Source: From Fieno DS et al. Contrast-enhanced magnetic resonance imaging of myocardium at risk: distinction between reversible and irreversible injury throughout infarct healing. J Am Coll Cardiol. 2000;36(6):1985–1991.

direct angioplasty. To study LV remodeling, 17 patients underwent repeated CMR 6 months after the initial study. Microvascular obstruction was defined as hypoenhancement seen 1 to 2 minutes after contrast injection. Infarct size was assessed as percent LV mass enhanced 5 to 10 minutes after contrast. Patients with microvascular obstruction (N ¼ 11) had more cardiovascular events than did those without (45% versus 9%, p ¼ 0.02). The risk of adverse events increased with infarct extent (30%, 43%, and 71% for small [N ¼ 10], mid-sized [N ¼ 14], and large [N ¼ 14] infarcts, respectively; p < 0.05). Even after infarct size was controlled for, the presence of microvascular 272 Cardiovascular Magnetic Resonance

obstruction remained a prognostic marker for postinfarction complications (Fig. 20-9). Rogers and colleagues confirmed the adverse outcome of patients with low signal intensity regions within the infarct zone after injection of gadolinium contrast.54 However, Gerber and associates reported that early gadolinium enhancement in the infarct area was inferior to LGE for predicting functional improvement in dysfunctional segments in patients studied 4 days after acute myocardial infarction.55 Receiver operating characteristic analysis showed an accuracy of 74% for the absence of late enhancement but only 49% for the absence of early hypoenhancement.

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Figure 20-6 Solid line, Monoexponential fit to serum contrast agent concentration (left-hand y-axis) as function of time after administration calculated (0.125 mmol/kg gadolinium-DTPA) based on data of Weinmann HJ et al. Physiol Chem Phys Med NMR. 1984;16:167–172. Dashed line, MR inversion time (right-hand y-axis) calculated based on data of solid line. Solid circles indicate data measured in the study of Mahrholdt and colleagues. Source: From Mahrholdt H et al. Reproducibility of chronic infarct size measurement by contrast-enhanced magnetic resonance imaging. Circulation. 2002;106(18):2322–2327.

segment with transmural infarction will usually not recover wall thickening. This concept was confirmed in a dog experiment of acute myocardial infarction. Segments with less than 25% of enhancement on day 3 after induction of the infarct had an 87% probability of improvement by day 28.56 In contrast, no segment with 100% transmurality of enhancement improved. Choi and colleagues57 confirmed these experimental findings in 24 patients who presented with their first myocardial infarction and were successfully revascularized. Improvement in segmental contractile function at 2 to 3 months was inversely related to the transmural extent of infarction as demonstrated by LGE within 7 days of the acute event. Improvement in global contractile function, as assessed by ejection fraction and mean wall thickening score, was not predicted by peak creatine kinase-MB or by total infarct size, as defined by CMR. Instead, the best predictor of global improvement was the extent of dysfunctional myocardium that was not infarcted or had infarction comprising less than 25% of LV wall thickness. Some early reports in humans with acute infarction suggested that spontaneous recovery usually occurs after acute infarction in LGE regions, which is contrary to the above

SCAN 1

SCAN 2 - Patient repositioned, later time, different scanner operator

Figure 20-7 Full set of short-axis views of magnetic resonance (MR) scan 1 performed 9 minutes after contrast injection and MR scan 2 acquired after removing and repositioning of the patient in the scanner with images acquired 32 minutes after contrast injection. TI was increased continuously to achieve perfect nulling of the myocardium. Presence, location, and size of the enhanced region were similar in both MR scans. Source: From Mahrholdt H et al. Reproducibility of chronic infarct size measurement by contrast-enhanced magnetic resonance imaging. Circulation. 2002;106(18):2322–2327.

Late Gadolinium Enhancement and Recovery of Function The transmural extent of myocardial infarction as defined by LGE CMR predicts the likelihood of improvement of contractile function. Spontaneous recovery of function can be expected after ischemic injury in a segment with a limited amount of subendocardial scar. In contrast, a

findings.52,58 However, these studies examined enhancement in a different way, within the first few minutes after gadolinium injection, and did not differentiate the transmural extent of enhancement in their analysis. Moreover, a short inversion time of 100 msec was employed. These confounding factors make direct comparison of the results between the studies of the two groups difficult.59 Overall opinion remains that with high-resolution imaging, Cardiovascular Magnetic Resonance 273

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1.0

ISCHEMIC HEART DISEASE

Figure 20-8 Zone of no reflow represented by a central core of low signal intensity surrounded by the typical bright infarct zone a few days after transmural RCX infarct (left panel, white arrow). The no-reflow phenomenon is detectable up to 2 to 4 weeks after acute infarction and is associated with poor prognosis (see text for details). Subsequently, at follow-up 12 weeks after infarction (right panel), the no-reflow zone has resolved.

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Figure 20-9 Event-free survival (clinical course without cardiovascular death, reinfarction, congestive heart failure, or stroke) for patients with and without CMR microvascular obstruction. Source: From Wu KC, Zerhouni EA, Judd RM, et al. Prognostic significance of microvascular obstruction by magnetic resonance imaging in patients with acute myocardial infarction. Circulation. 1998;97:765.

significant transmural enhancement at 5 to 15 minutes reflects nonviable myocardium, which has also been confirmed by the results from other centers.60–62 The use of LGE CMR for predicting improvement of myocardial function was recently challenged by contrast echocardiography.63 In patients after primary coronary intervention, follow-up at 60 days after the infarct showed improvement of function by echocardiography in 115 of 192 segments (gold standard). The sensitivity, specificity, and accuracy of qualitative scores by contrast echocardiography, first-pass CMR, and LGE CMR were similar for all three techniques. The authors stress the advantage of MCE as a bedside technique. However, contrast echocardiography is still not widely used clinically for this purpose. 274 Cardiovascular Magnetic Resonance

After an acute ischemic event, structural changes occur within the infarct zone, and infarct healing with scar formation is completed by approximately 3 to 4 months.10 Thinning of the infarct region may occur early, especially in large anterior myocardial infarcts, but transient thickening of the infarcted segment due to edema64 has also been observed. The consequence of infarct thinning is an increase in the size of the infarcted segment, known as infarct expansion.65 However, even in transmural infarcts, infarct expansion may not occur in patients with open infarct-related arteries. Such patients are encountered more often today with the widespread use of thrombolysis and angioplasty of the infarct artery.66 Therefore, transmural necrosis and nontransmural necrosis may have the same wall thickness early after myocardial infarction. Both conditions may also be associated with complete absence of resting function early after the acute event. However, it should be remembered that even a small amount of wall thickening in a region of interest indicates the presence of residual contracting cells and hence of viable myocardium.

CARDIOVASCULAR MAGNETIC RESONANCE IN CHRONIC MYOCARDIAL INFARCTION As was mentioned above, chronic myocardial infarcts are structurally different from acute myocardial infarcts. The most obvious difference is that chronic transmural infarcts may be very thin, owing to infarct expansion and remodeling.67,68 Consequently, this feature can be detected by CMR and can be used to distinguish between chronic transmural scar and

p < 0.0001

Rel. FDG-uptake (%)

100

The hypothesis that thinned and akinetic myocardium represents chronic scar has been tested by comparing CMR findings with those obtained by PET and single photon emission computed tomography (SPECT) in identical myocardial regions.70,71 Comparison of CMR with scintigraphic techniques is easily accomplished because identical regions can be matched, owing to the three-dimensional nature of both techniques. To define transmural scar by end-diastolic wall thickness, a cutoff value of 5.5 mm was selected. This value corresponded to the mean end-diastolic wall thickness in normal individuals: 2.5 standard deviations. It also corresponded well to the wall thickness of less than 6 mm found in a histopathologic study of transmural chronic scar.65 Regions with a mean end-diastolic wall thickness of less than 5.5 mm had a significantly reduced fluorodeoxyglucose (FDG) uptake as compared to regions with an end-diastolic wall thickness of 5.5 mm or more (Fig. 20-10).69 In 29 of 35 patients studied, the diagnosis of viability based on FDG uptake was identical to the one based on myocardial wall thickness as assessed by CMR. Importantly, relative FDG uptake did not differ between segments with systolic wall thickening at rest or akinesia at rest as long as wall thickness was preserved. These findings were extended in another population of patients who underwent revascularization and control CMR at 3 months after revascularization.72 Of 125 segments with an end-diastolic wall thickness less than 5.5 mm in 43 patients with chronic infarcts, only 12 segments recovered (corresponding to a negative predictive accuracy of 90% for the finding of end-diastolic wall thinning to predict transmural scar). In contrast, the positive predictive accuracy was only 62% for preserved end-diastolic wall thickness of 5.5 mm or greater for predicting the presence of viable myocardium with the potential for recovery. This is likely explained by the fact that simply looking at wall thickness by CMR will not define the thickness of the remaining viable rim, which is the most important factor in determining whether a segment will recover function or not.11 Definition of this viable rim is possible only by using contrastenhanced CMR, and this will be discussed below. The relationship between end-diastolic wall thickness and viability has been disputed by other researchers,73 who found FDG-uptake on PET images largely independent of regional end-diastolic wall thickness. However, this study included recent and chronic infarcts and used a suboptimal conventional spin echo technique with a short echo time of 20 msec to measure wall thickness.

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Myocardial Wall Thickness as a Feature of Viable Myocardium

n.s.

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47% WTh < 1 mm N = 248

DWT > 5.5 mm DWT < 5.5 mm (viable) (scar)

Figure 20-10 Normalized uptake of 18-F fluorodeoxyglucose (FDG) uptake stratified by end-diastolic wall thickness as measured from gradient echo magnetic resonance images. Regions with preserved wall thickening (WTh) of 1 mm or more and preserved end-diastolic wall thickness (DWT) of > 5.5 mm had a relative FDG uptake similar to that of regions without wall thickening (WTh < 1 mm) but preserved end-diastolic wall thickness. In contrast, akinetic regions (WTh < 1 mm) with reduced end-diastolic wall thickness had significantly reduced FDG uptake (mean: 47%), indicating scar formation. Source: Data from Baer FM, Voth E, Schneider CA, Theissen P, Schicha H, Sechtem U. Comparison of low-dose dobutamine–gradient-echo magnetic resonance imaging and positron emission tomography with [18F] fluorodeoxyglucose in patients with chronic coronary artery disease: a functional and morphological approach to the detection of residual myocardial viability. Circulation. 1995;91:1006–1015.

Contractile Reserve During Low-Dose Dobutamine Infusion Although severely reduced end-diastolic wall thickness is helpful in identifying scarred myocardium, the positive predictive value of a preserved end-diastolic wall thickness for predicting recovery of function following revascularization is disappointingly low. However, CMR offers the possibility of measuring wall thickening not only at rest but also during low-dose dobutamine infusion. Until recently, a protocol with acquisition of cine CMR images in multiple short axes and two long axes sections at rest and at 5 and 10 mg/kg/min dobutamine required an imaging time of more than 60 minutes. The advent of fast CMR sequences now permits completion of the same protocol within approximately 30 to 45 minutes, and image quality is often better with breath hold cine CMR images than with conventional techniques. The sensitivity of dobutamine CMR for detection of viable myocardium as defined by a normalized FDG uptake on PET images is 81% with a specificity of 95%.69 When recovery of wall thickening following revascularization was considered to be the gold standard, the sensitivity of dobutamine CMR in predicting recovery of function after revascularization was 89% with a specificity of 94%. The latter analysis was patient related, which is clinically more meaningful than a segment-by-segment analysis.70 New techniques such as tissue-tagged MR imaging with three-dimensional strain analysis may further enhance the accuracy of dobutamine CMR.74 Using a healthier patient group and a different methodology, Trent and colleagues74a found less satisfactory values for sensitivity and specificity of wall motion (50% and 72%, Cardiovascular Magnetic Resonance 275

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residual viable myocardium in the infarct area. However, caution must be used in observing a small area of pronounced wall thinning in order not to assume that the entire region perfused by an occluded coronary artery is completely scarred. Frequently, myocardial cells in the border zone survive, and ischemia of this border zone alone may cause substantial symptoms in a patient. Therefore, in a patient with single-vessel disease, previous myocardial infarction and anginal symptoms, restoration of blood flow by reestablishing patency of the occluded artery may be justified, despite evidence of complete necrosis in the center of the infarct zone.69

ISCHEMIC HEART DISEASE

respectively) or wall thickening (50% and 68%, respectively). This was possibly related to the fact that they used higher dobutamine doses. They also included segments with worsening wall dynamics, which were considered viable. In contrast, Baer and colleagues70 looked only at akinetic segments, which by definition could not become worse. Baer and coworkers also presented data on the relative value of dobutamine CMR and dobutamine transesophageal echocardiography (TEE).75 Normalized FDG uptake on PET images was used as the standard against which both techniques were compared. The sensitivity and the specificity of dobutamine TEE and dobutamine CMR for FDG PET– defined myocardial viability were 77% versus 81% and 94% versus 100%, respectively. Thus, the two imaging techniques provide similar accuracy. In choosing the appropriate technique, patient acceptance becomes an important consideration. Although claustrophobia may be a problem with CMR, only a small fraction of patients are affected. In contrast, many patients do not like the experience of a transesophageal echocardiographic examination. On the other hand, there is a clear cost advantage for transesophageal echocardiography because the echo probe costs only a fraction of a CMR scanner, and additional investment is not necessary.

Late Gadolinium Enhancement in Chronic Infarction Similar principles apply to the assessment of myocardial infarction in the chronic phase as to assessment shortly after the event. Late enhancement of the infarct zone is also seen in chronic infarcts,76,77 because of the continuing increase in partition coefficient, and delayed contrast agent kinetics. The tissue in chronic scar consists mainly of fibroblasts and collagen surrounded by a large intercellular space. Because gadolinium contrast agents will diffuse into this space, enhancement may simply be related to the larger volume of distribution in comparison to normal myocardium with its tightly packed myocytes, which the contrast agent cannot enter. LGE CMR is able to identify patients with healed myocardial infarction with high accuracy. Moreover, this technique may permit noninvasive differentiation between ischemic and nonischemic cardiomyopathy. In a study of 71 subjects, 40 patients with healed myocardial infarction were prospectively enrolled after enzymatically proven necrosis and were imaged 3 1 months (N ¼ 32) and/or 14 7 months (N ¼ 19) later.74 They were compared with 20 patients with nonischemic cardiomyopathy and 11 normal volunteers. Twenty-nine of 32 patients (91%) with 3-month-old infarcts (13 non-Q-wave) and all 19 with 14-month-old infarcts (8 non-Q-wave) exhibited LGE. In patients in whom the infarct-related artery was determined at angiography, 24 of 25 patients with 3-month-old infarcts (96%) and all 14 with 14-month-old infarcts had LGE in the correct territory. None of the 20 patients with nonischemic cardiomyopathy or the 11 volunteers had LGE. Regardless of the presence or absence of Q-waves, the majority of patients with LGE had only nontransmural involvement. Normal LV contraction was visualized in 7 patients with 3-month-old infarcts (22%) and 3 with 14-month-old infarcts (16%), but in these cases, LGE was limited to the subendocardium. McCrohon and 276 Cardiovascular Magnetic Resonance

colleagues78 demonstrated in a cohort of 90 patients with heart failure that all patients with coronary artery disease had LGE primarily with a subendocardial or transmural pattern, whereas patients with dilated cardiomyopathy (no stenosis by coronary angiography) had either no enhancement (59%), patchy or midwall enhancement not corresponding to a coronary perfusion bed (28%), or enhancement patterns indistinguishable from those of patients with ischemic cardiomyopathy (13%). The patchy and midwall enhancement patterns are often found in patients with biopsy-proven myocarditis.79 This indicates that CMR may give profound new insights into the causes of LV failure. Prediction of the myocardial response to revascularization remains the holy grail of viability testing. Late enhancement CMR performs well in this discipline. In the first study looking at the ability of contrast CMR to predict improvement in LV function after revascularization, Kim and colleagues80 showed that the likelihood of functional recovery decreased progressively as the transmural extent of LGE observed before revascularization increased (Fig. 20-11). Approximately 80% of segments with no LGE improved function after revascularization (Fig. 20-12), whereas if more than 75% of the transmural tissue was enhanced, only a small percentage improved with revascularization (Fig. 20-13).

Thickness of the Viable Epicardial Rim and Recovery of Function The clear depiction of myocardial scar by gadoliniumenhanced CMR also permits clear identification of epicardially located viable tissue. Knuesel and colleagues80a determined the segmental amount of viable tissue by LGE CMR and compared this to 18F-FDG uptake and resting perfusion using 13N-ammonia. FDG uptake of 50% or more corresponded to a viable rim thickness of 4.5 mm by LGE CMR. Segments that had both a thick viable rim and a FDG uptake of 50% or more showed functional recovery in 85% of segments (Fig. 20-14), whereas thin metabolically nonviable segments (FDG uptake <50%) improved function in only 13% of segments (p < 0.0005). Segments with thin viable rims and preserved FDG uptake and those with a thick viable rim but reduced FDG uptake improved function only in 36% and 23% of segments, respectively. Similar findings were reported by Ichikawa and colleagues,80b who studied 18 patients with a first myocardial infarct within 1 week of the event and more than 5 months later. Receiver operating characteristic analysis demonstrated that the measurement of thickness of nonenhanced myocardium had a slightly better diagnostic accuracy for predicting improved systolic wall thickening from scan 1 to 2 in dysfunctional segments. Obviously, further studies are needed in order to define the respective merits of absolute and relative measurements of infarcted and viable myocardium in the infarct zone for the prediction of functional recovery.

Comparison of Late Contrast Enhancement with Other Imaging Modalities The comparison of gadolinium-enhanced CMR with other modalities for the discrimination of viable and nonviable

Segments with severe hypokinesia, akinesia, or dyskinesia

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20 MYOCARDIAL VIABILITY

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0 1– 25 26 –5 51 0 –7 76 5 –1 00

0

0 1– 25 26 –5 51 0 – 76 75 –1 00

(1

of

58

(1

)

3

of

20

of

12

Improved contractility (%)

Figure 20-11 Relationship between the transmural extent of enhancement before revascularization and the likelihood of increased contractility after revascularization. Data are shown for all 804 dysfunctional segments and separately for the 462 segments with at least severe hypokinesia and the 160 segments with akinesia or dyskinesia before revascularization. For all three analyses, there was an inverse relationship between the transmural extent of enhancement and the likelihood of improvement in contractility. Source: From Kim RJ, Wu E, Rafael A, et al. The use of contrast-enhanced magnetic resonance imaging to identify reversible myocardial dysfunction. N Engl J Med. 2000;343:1445, with permission. Copyright # 2000 Massachusetts Medical Society. All rights reserved.

Transmural extent of hyperenhancement (%)

myocardium has been favorable. Segmental LGE CMR correlated well with diminished thallium-201 uptake measured by rest-redistribution thallium-201 SPECT in patients with globally reduced LV function.81 However, the much better spatial resolution that permits discrimination of subendocardial and transmural infarcts82 and the lack of ionizing radiation still favor CMR. Klein and colleagues found that the area of LGE-measured CMR correlated closely with myocardial infarcts defined by PET in patients with ischemic cardiomyopathy, but CMR showed the extent of infarct transmurality more clearly (Fig. 20-15).83 Kuehl and colleagues84 compared 18FFDG PET and contrast-enhanced CMR in 26 patients. The study showed that segmental glucose uptake by PET was inversely correlated with the segmental extent of LGE and that using a cutoff value of 37% segmental enhancement optimally differentiated viable from nonviable segments defined by PET, resulting in a sensitivity and specificity of contrast-enhanced CMR for detection of viable myocardium, using PET as the gold standard, of 96% and 84%, respectively.

Comparison of Different Cardiovascular Magnetic Resonance Techniques for the Diagnosis of Viability Recently, Wellnhofer and colleagues85 compared LGE CMR to dobutamine CMR (5 and 10 mg/kg/min) for assessment of myocardial viability. Both techniques were performed in 29 patients with chronic CAD and resting LV dysfunction (mean left ventricular ejection fraction: 32% 8%). Cine CMR imaging at rest was repeated

3 months after revascularization to determine wall motion improvement. The transmurality of LGE was assessed visually, using a five-grade scale. Similarly, wall motion was assessed visually. Using a cutoff value of 25% transmurality of LGE, the authors found that contrastenhanced CMR correctly identified 73% of hibernating segments but dobutamine CMR was better and identified 85% correctly. The results for sensitivity and specificity of Wellnhofer and colleagues for dobutamine CMR are slightly worse that those reported by Baer and associates,70 whose patient group was rather healthy, with a mean left ventricular ejection fraction of 42% 10%. However, the sensitivity found by Wellnhofer and colleagues for scar transmurality less than 50% is better than the 50% found by Gunning and colleagues,86 who studied a patient group with more severely depressed LV function (mean left ventricular ejection fraction: 24% 8%). In contrast, the sensitivity of dobutamine CMR in the Wellnhofer study dropped sharply to values below 60% in segments with a scar transmurality of 50% or more. Thus, in severely impaired ventricles, low-dose dobutamine CMR has a suboptimal sensitivity but a preserved high specificity at around 90%. This has been known from studies using dobutamine echocardiography in which similar low sensitivity values of 50% and less were reported. The lower sensitivity of dobutamine echocardiography in detecting viable myocardium in regions with reduced function and perfusion may indicate that some regions of hibernating myocardium are so delicately balanced between the reductions in flow and function, with exhausted coronary flow reserve that any catecholamine stimulation to increase oxygen demands will merely result Cardiovascular Magnetic Resonance 277

ISCHEMIC HEART DISEASE

BEFORE REVASCULARIZATION End diastole

End systole

No hyperenhancement

A AFTER REVASCULARIZATION End diastole

End systole

B Figure 20-12 Representative cine images and contrast-enhanced images obtained by magnetic resonance in a patient with reversible ventricular dysfunction. This patient had severe hypokinesia of the anteroseptal wall (arrows), and this area was not enhanced before revascularization. The contractility of the wall improved after revascularization. Source: From Kim RJ, Wu E, Rafael A, et al. The use of contrast-enhanced magnetic resonance imaging to identify reversible myocardial dysfunction. N Engl J Med. 2000;343:1445, with permission. Copyright # 2000 Massachusetts Medical Society. All rights reserved.

in ischemia and inability to elicit enhanced contractile function.87 However, it is just the patients with global severe depression of function in whom viability testing is clinically most meaningful. In these patients, it has been demonstrated that LGE CMR predicts improvement in severely dysfunctional segments, with 86% improving if no scar is seen by contrast-enhanced CMR but none improving for scar transmurality less than 75%.77 Similar findings were reported by Schvartzman and colleagues.88 In an editorial accompanying the paper by Wellnhofer and colleagues, Kim and Manning comment that contractile reserve has a reduced predictive accuracy if more severe dysfunction is present at rest.89 In this most important subgroup of patients, LGE CMR appears to have a greater accuracy. Further studies are needed to compare LGE CMR and dobutamine CMR for different levels of LV dysfunction. From a practical point of view, LGE CMR has the advantage that it does not require pharmacologic stress and thus involves less risk and subsequently less monitoring of the patient. Furthermore, it is likely easier and less observer dependent for interpretation. Dobutamine and LGE CMR can also be used as complementary techniques to make the diagnosis of 278 Cardiovascular Magnetic Resonance

viability. Kaandorp and associates90 studied 48 patients with chronic coronary artery disease and a mean ejection fraction of 37 10%. Regional dysfunction was present in 41% of segments and 61% of those had contractile reserve. The likelihood of a contractile improvement with dobutamine was highest (75%) in segments with small amounts of subendocardial scar, lowest (17%) in those with transmural scar, but intermediate (42%) in segments with intermediate infarct transmurality. The authors suggest an approach where LGE CMR is sufficient at both extreme ends but dobutamine CMR should be added to optimally predict outcome in segments with intermediate amounts of scar transmurality.

Cardiovascular Magnetic Resonance Spectroscopy The hallmark of viable myocardium is the presence of high-energy phosphates within the cell. As phosphorus31 CMRS (31P-CMRS) is the only available technique to

End systole

No hyperenhancement

A AFTER REVASCULARIZATION End diastole

End systole

B Figure 20-13 Representative cine images and contrast-enhanced images obtained by CMR in a patient with irreversible ventricular dysfunction. This patient had akinesia of the anterolateral wall (arrows), and this area was enhanced before revascularization. The contractility of the wall did not improve after revascularization. Source: From Kim RJ, Wu E, Rafael A, et al. The use of contrast-enhanced magnetic resonance imaging to identify reversible myocardial dysfunction. N Engl J Med. 2000;343:1445, with permission. Copyright # 2000 Massachusetts Medical Society. All rights reserved.

observe high-energy phosphates noninvasively in vivo, it can be employed to detect and quantify this sign of life within a myocardial region. By quantifying the amount of high-energy phosphate compounds, it is possible to determine the amount of viable myocardium present in the regent of interest. Yabe and colleagues evaluated patients with reversible and irreversible thallium defects on exercise redistribution studies.29 The volume of interest in this study was on the order of 30 cm3. Phosphocreatine content was significantly lower in the group without thallium redistribution (which may indicate absence of viability) and in the group with reversible defects (indicating residual viability) as compared to a group of 11 normal individuals. The ATP concentration, however, was significantly lower than that in normals only in the group without thallium redistribution. Clinical applications of the technique are not expected because the large volumes of interest usually incorporate mixtures of scar, normal myocardium, and ischemically injured viable myocardium.

Proton spectroscopy (1H-CMRS) has higher sensitivity than 31P-CMRS and is able to detect the total pool of phosphorylated plus unphosphorylated creatine in both skeletal and cardiac muscle. 1H-CMRS offers about a 20fold net theoretical sensitivity improvement in comparison with 31P-CMRS of phosphorylated creatine. Consequently, 1H-CMRS allows for the first time at magnetic field strengths of many clinical CMR systems the metabolic interrogation of small voxels (<10 mL) in all regions of the left ventricle. Because the entire ventricle can be examined by 1H-CMRS, comparison of viable and nonviable tissue is possible within the same patient. Thus, the patient can serve as his or her own control. Bottomley and Weiss were the first to employ proton spectroscopy in patients with remote myocardial infarction (longer than 1 month).30 In a dog model, they established that enzymatic degradation of creatine in heart extracts resulted in the complete disappearance of the 1H N-methyl resonance peak at 3.0 parts per million. Although myocardial creatine is significantly reduced in infarction, there is Cardiovascular Magnetic Resonance 279

20 MYOCARDIAL VIABILITY

BEFORE REVASCULARIZATION End diastole

ISCHEMIC HEART DISEASE

Contrast-enhanced MR

FDG-Uptake (%)

100

Baseline MR: Diastole

Baseline MR: Systole

50

A

B

C

D

0 75

E

F

G

H

50

Thick metabolically viable segments with functional recovery Scar (contrast-enhanced MR) Thrombus

Perfusion (K1, mL/min/100mL)

25

Follow-up MR: Diastole

Follow-up MR: Systole

Figure 20-14 A 59-year-old patient 5 months after an anteroseptal myocardial infarction. Contrast-enhanced magnetic resonance (MR) (A) shows a small subendocardial zone of scar (bright) with adjacent thrombus (black, segment 8). A thick rim of viable tissue is observed (segments 1 and 6 through 8 in panel A) that corresponds to segments with preserved FDG uptake (B), whereas perfusion in these segments is reduced (D). A schematic in panel C illustrates this situation, with segments 1 and 6 through 8 consisting of metabolically viable segments with a thick rim of viable tissue on MR. Contractile function is severely reduced in these segments before revascularization (E and F) but has normalized 9 months after coronary artery graft bypass surgery (G and H). Source: From Knuesel PR et al. Characterization of dysfunctional myocardium by positron emission tomography and magnetic resonance: relation to functional outcome after revascularization. Circulation. 2003;108(9):1095–1100.

some overlap between noninfarcted myocardium and infarcted myocardium. Again, practical applications of spectroscopy are hampered by the fact that it is technically demanding and existing imaging tools are much easier to use clinically for diagnosing myocardial viability.

CONCLUSION CMR provides a variety of novel methods of obtaining information on residual viability after myocardial infarction. Indirect signs of viability that can be observed by CMR are the absence of increased signal intensity on spin echo images in a myocardial region involved in a recent infarct, any sign of wall thickening at rest (which is detectable with high accuracy by cine CMR), wall thickening after stimulation by low-dose dobutamine, preserved wall thickness, and a preserved viable epicardial rim. On the other hand, myocardial necrosis is characterized by high signal intensity on spin echo images, LGE (possibly with a low-intensity core region due to no reflow) of the infarct area after injection of gadolinium, reduced wall thickness (in chronic infarcts) without a substantial viable rim, and the absence of a contractile reserve during dobutamine stimulation. Direct

280 Cardiovascular Magnetic Resonance

observation of the presence of high-energy phosphates and measurement of total creatine are possible using spectroscopic techniques. Evaluation of viable myocardium is most important in patients with severe LV dysfunction because these patients can gain most from revascularization if substantial amounts of myocardium are present. Revascularization in these patients on the other hand carries a high risk but will bring no benefit if the ventricle is composed mainly of scar tissue. Thus, transplantation should be reserved for those patients for whom all other therapeutic options have been excluded.91 Unfortunately, the available information about the value of CMR techniques in identifying patients with global LV dysfunction who have a high likelihood of profiting from a coronary revascularization is still scarce. Lowdose dobutamine CMR may be less accurate than LGE CMR in this patient group. In patients with regional LV dysfunction, however, in whom the need for revascularization needs to be established, both dobutamine CMR and LGE CMR are well validated and can be used clinically.92 Depiction of zones of acute necrosis by observing enhancement after gadolinium carries substantial promise for detailed studies of the effects of different treatment strategies for acute myocardial infarction as well as for treatment in the early and late postinfarction phase.

Mid-ventricular

Basal

NH3

FDG

CMR

Figure 20-15 Three short-axis views (apical, equatorial, and basal) of a PET viability study with assessment of rest perfusion (NH3) (top) and glucose metabolism (FDG) (middle). Bottom, CMR images in corresponding slices showing enhancement. Note that in segments with reduced perfusion and metabolism, there is an increased CMR signal. Because of better spatial resolution in CMR, distinction between transmural, subendocardial, and papillary defects can be made. The border between enhanced and normal areas is distinct. Source: From Klein C et al. Assessment of myocardial viability with contrast-enhanced magnetic resonance imaging: comparison with positron emission tomography. Circulation. 2002;105(2):162–167.

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55. Gerber BL, Garot J, Bluemke DA, Wu KC, Lima JA. Accuracy of contrast-enhanced magnetic resonance imaging in predicting improvement of regional myocardial function in patients after acute myocardial infarction. Circulation. 2002;106:1083–1089. 56. Hillenbrand HB, Kim RJ, Parker MA, Fieno DS, Judd RM. Early assessment of myocardial salvage by contrast enhanced magnetic resonance imaging. Circulation. 2000;102:1678–1683. 57. Choi KM, Kim RJ, Gubernikoff G, Vargas JD, Parker M, Judd RM. Transmural extent of acute myocardial infarction predicts long-term improvement in contractile function. Circulation. 2001;104:1101–1107. 58. Kramer CM, Rogers WJ, Mankad S, Theobald TM, Pakstis DL, Hu YL. Contractile reserve and contrast uptake pattern by magnetic resonance imaging and functional recovery after reperfused myocardial infarction. J Am Coll Cardiol. 2000;36:1835–1840. 59. Bucciarelli-Ducci C, Wu E, Lee DC, Holly TA, Klocke FJ, Bonow RO. Contrast-enhanced cardiac magnetic resonance in the evaluation of myocardial infarction and myocardial viability in patients with ischemic heart disease. Curr Probl Cardiol. 2006;31:128–168. 60. Beek AM, Ku¨hl HP, Bondarenko O, et al. Delayed contrast-enhanced magnetic resonance imaging for the prediction of regional functional improvement after acute myocardial infarction. J Am Coll Cardiol. 2003;42:895–901. 61. Baks T, van Geuns RJ, Biagini E, et al. Recovery of left ventricular function after primary angioplasty for acute myocardial infarction. Eur Heart J. 2005;26:1070–1077. 62. Ingkanisorn WP, Rhoads KL, Aletras AH, Kellman P, Arai AE. Gadolinium delayed enhancement cardiovascular magnetic resonance correlates with clinical measures of myocardial infarction. J Am Coll Cardiol. 2004;43:2253–2259. 63. Biagini E, van Geuns RJ, Baks T, et al. Comparison between contrast echocardiography and magnetic resonance imaging to predict improvement of myocardial function after primary coronary intervention. Am J Cardiol. 2006;97:361–366. 64. Haendchen RV, Corday E, Torres M, Maurer G, Fishbein MC, Meerbaum S. Increased regional end-diastolic wall thickness early after reperfusion: a sign of irreversibly damaged myocardium. J Am Coll Cardiol. 1984;3:1444–1453. 65. Pirolo JS, Hutchins GM, Moore GW. Infarct expansion: pathologic analysis of 204 patients with a single myocardial infarct. J Am Coll Cardiol. 1986;7:349–354. 66. Ito H, Yu H, Tomooka T, et al. Incidence and time course of left ventricular dilation in the early convalescent stage of reperfused anterior wall acute myocardial infarction. Am J Cardiol. 1994;73:539–543. 67. Dubnow MH, Burchell HB, Titus JL. Postinfarction left ventricular aneurysm: a clinicomorphologic and electrocardiographic study of 80 cases. Am Heart J. 1965;70:753–760. 68. Bellenger NG, Swinburn JM, Rajappan K, Lahiri A, Senior R, Pennell DJ. Cardiac remodelling in the era of aggressive medical therapy: does it still exist? Int J Cardiol. 2002;83:217–225. 69. Braunwald E, Kloner RA. Myocardial reperfusion: a double-edged sword? J Clin Invest. 1985;76:1713–1719. 70. Baer FM, Smolarz K, Jungehulsing M, et al. Chronic myocardial infarction: assessment of morphology, function, and perfusion by gradient echo magnetic resonance imaging and 99mtc-methoxyisobutyl-isonitrile SPECT. Am Heart J. 1992;123:636–645. 71. Baer FM, Voth E, Schneider CA, Theissen P, Schicha H, Sechtem U. Comparison of low-dose dobutamine-gradient-echo magnetic resonance imaging and positron emission tomography with [18F]fluorodeoxyglucose in patients with chronic coronary artery disease: a functional and morphological approach to the detection of residual myocardial viability. Circulation. 1995;91:1006–1015. 72. Baer FM, Theissen P, Schneider CA, et al. Dobutamine magnetic resonance imaging predicts contractile recovery of chronically dysfunctional myocardium after successful revascularization. J Am Coll Cardiol. 1998;31:1040–1048. 73. Perrone-Filardi P, Bacharach SL, Dilsizian V, et al. Metabolic evidence of viable myocardium in regions with reduced wall thickness and absent wall thickening in patients with chronic ischemic left ventricular dysfunction. J Am Coll Cardiol. 1992;20:161–168. 74. Bree D, Wollmuth JR, Cupps BP, et al. Low-dose dobutamine tissuetagged magnetic resonance imaging with 3-dimensional strain analysis allows assessment of myocardial viability in patients with ischemic cardiomyopathy. Circulation. 2006;114(1 suppl):I33–I36. 74a. Trent RJ, Waiter GD, Hillis GS, McKiddie FI, Redpath TW, Walton S. Dobutamine magnetic resonance imaging as a predictor of myocardial functional recovery after revascularization. Heart. 2000;83(1):40–46.

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CHAPTER 21

Coronary Artery and Vein Imaging: Methods Reza Nezafat, Rene´ M. Botnar, Kraig V. Kissinger, Peng Hu, and Warren J. Manning

Despite ongoing progress in both prevention and early diagnosis, coronary artery disease (CAD) remains the leading cause of death in both men and women in the United States1 and throughout the Western world. In 2009, an estimated 785,000 Americans will have a new coronary event, whereas an additional 195,000 will have a silent first myocardial infarction.1 Catheter-based, invasive X-ray coronary angiography remains the “gold standard” for the diagnosis of significant (>50% diameter stenosis) CAD. More than 1 million catheter-based X-ray coronary angiograms are performed annually in the United States,1 and the volume is higher in Europe. Although numerous noninvasive tests are available to help discriminate among those with and without significant angiographic disease, 30% or more of patients referred for catheter-based X-ray coronary angiography are found to have no significant stenoses.2,3 Despite the absence of disease, these individuals are exposed to the cost, inconvenience, and potential morbidity of X-ray angiography.4–6 Data also suggest that in selected high-risk populations, such as patients with aortic valve stenosis, the incidence of subclinical stroke associated with retrograde catheter crossing of the stenotic valve may exceed 20%.7 Percutaneous coronary intervention in single-vessel disease is commonly performed to relieve symptoms or decrease pharmaceutical use, but the greatest effect on mortality occurs with mechanical intervention among patients with left main (LM) and multivessel CAD. Thus, it would be desirable to have a noninvasive method that allows direct visualization of the proximal and mid-native coronary vessels for accurate identification or exclusion of LM or multivessel CAD. Over the last decade, coronary artery cardiovascular magnetic resonance (CMR) has evolved as a clinical option for catheter-based X-ray angiography among patients with suspected anomalous coronary artery disease and coronary artery aneurysms. This technique has reached sufficient maturity that it obviates the need for catheter-based X-ray angiography in the discrimination of patients with multivessel disease. Although it does not offer the spatial resolution of current coronary multidetector computed tomography (MDCT) approaches, both approaches must address similar technical obstacles. This chapter reviews the technical challenges and general imaging strategies for coronary artery and vein CMR. The next chapter (Chapter 22) reviews the clinical data regarding coronary artery magnetic resonance imaging for the assessment of anomalous coronary artery disease, coronary artery aneurysms, native vessel integrity, 284 Cardiovascular Magnetic Resonance

and coronary artery bypass graft disease, including comparisons with MDCT.

CORONARY ARTERY CARDIOVASCULAR MAGNETIC RESONANCE: TECHNICAL CHALLENGES AND SOLUTIONS Technical Challenges Despite the routine clinical use of CMR for the evaluation of similar-sized blood vessels, coronary artery CMR is more technically challenging because of unique issues, including: (1) the small caliber of the vessels (2 to 5 mm in diameter); (2) their near-constant motion during both the respiratory and the cardiac cycles; (3) their tortuosity; and (4) surrounding signal from adjacent epicardial fat and myocardium.

Cardiac Motion Bulk epicardial motion is a major impediment to coronary artery and vein CMR and can be separated into motion related to direct cardiac contraction and relaxation during the cardiac cycle and that caused by superimposed diaphragmatic and chest wall motion from respiration. The magnitude of motion from each component may greatly exceed the coronary artery diameter, leading to blurring artifacts in the absence of motion suppressive methods. To compensate for bulk cardiac motion, a regular rhythm and accurate electrocardiographic (ECG) synchronization with QRS detection are absolute requirements. Vector ECG approaches are preferred.8,9 Although potentially adequate for lower-resolution cine CMR ventricular functional acquisitions, cardiac gating strategies using peripheral pulse detection methods lead to inferior coronary artery CMR. Real-time CMR implementations10,11 have also been proposed using spiral or echo planar acquisition for rapid chromonar artery localization, but further improvements in the signal-to-noise ratio (SNR) are required for this approach to enable coronary artery stenosis evaluation. Both catheter-based X-ray angiography12 and CMR13–17 methods have characterized coronary motion during the cardiac cycle. Both the proximal and mid-right coronary artery

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Figure 21-1 Graph of in-plane right coronary artery (RCA) and left anterior descending (LAD) coronary artery motion during the cardiac cycle. The X-axis shows time as a percentage of the R-R interval. Note the improved image quality of the RCA cross-section when acquired during middiastole compared with early diastole.11

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(RCA) and the left anterior descending (LAD) coronary artery display a triphasic pattern (Fig. 21-1), with the magnitude of in-plane motion nearly twice as great for the RCA. During isovolumic relaxation, approximately 350 to 400 msec after the R-wave, and again at mid-diastole (immediately before atrial systole), coronary motion is minimal. LAD diastasis is longer than RCA diastasis and begins earlier in the cardiac cycle.17 The duration of the mid-diastolic diastasis period is inversely related to heart rate and dictates the coronary imaging data acquisition interval. Because of limitations of gantry speed, beta-blockers are commonly used in MDCT to reduce the heart rate and increase the duration of the rest period.17 For coronary artery CMR, the acquisition interval is adapted to the heart rate/diastasis interval. Although the use of a heartrate-dependent formula for identifying the mid-diastolic diastasis period is effective in many subjects, there may be considerable intersubject variation.18,19 Thus, use of a patientspecific diastasis period is recommended. This can be readily identified by the acquisition of a high-temporal-resolution cine dataset orthogonal to the long axis of the proximal/midRCA and the LAD. Semi-automated tools to identify the optimal data acquisition window have been proposed.20,21 For patients with a heart rate of 60 to 70 bpm, a coronary artery CMR acquisition duration of approximately 80 msec during each cardiac cycle results in improved image quality.22 With higher heart rates, the duration must be further abbreviated (e.g., < 50 msec), whereas with bradycardia, the acquisition interval can be expanded to 120 msec or longer. The use of patient-specific acquisition windows reduces overall scan time,18,19 and correction for heart rate variability improves image quality.23 Sinus arrhythmia leading to heart rate variability is common, especially in younger adults, and may lead to image degradation.24 Correction for heart rate variability using an adaptive real-time arrhythmia rejection algorithm improves coronary artery CMR quality.21

Effect of Nitrates on Coronary Artery Blood Flow The intraluminal signal in gradient echo coronary artery CMR is dependent on the inflow of unsaturated protons. As a result, mid-diastole has been identified as the preferred period for image acquisition because it corresponds to a period of minimal coronary motion and rapid (30 cm/sec) coronary artery blood flow. The addition of shortor longer-acting sublingual nitrates before coronary artery CMR results in 20% to 30% vasodilation and improves the contrast-to-noise ratio (CNR) and SNR of both gradient echo and steady-state free precession (SSFP) coronary artery CMR.25–27 The effect of the nitrate is dependent on the dose and the timing of image acquisition (Fig. 21-2). CNR for contrast agent coronary artery CMR sequences are less sensitive to inflow effects and thus have advantages for systolic acquisitions (discussed later). The effect of nitrates on CNR and SNR is unknown in this setting, although nitrates might still be useful to promote maximal vasodilation.

Respiratory Motion The second major challenge for coronary artery CMR is compensation for respiratory motion. With inspiration, the diaphragm may descend up to 30 mm and the chest wall expands, resulting in inferior displacement and anterior rotation of the heart.28,29 Minimizing respiratory motion artifacts can be achieved with several approaches (Table 21-1), including sustained end-expiratory breath holding, chest wall bellows, respiratory navigators, fat navigators, and self-gating methods. Comparison of singlebreath-hold with multiple-breath-hold and navigator gated respiratory methods suggests that free breathing navigator Cardiovascular Magnetic Resonance 285

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Figure 21-2 Right coronary artery (RCA) cardiovascular magnetic resonance acquired pre- and post-isosorbide dinitrate using a threedimensional free-breathing steady-state free precession sequence with doses of 2.5 mg (top row) and 5 mg (bottom row) as a function of time. Improved RCA vasodilation and signal enhancement can be observed in all post-isosorbide dinitrate images (arrows in top and bottom rows). As a result of an enhanced signal-to-noise ratio, distal segments are visualized better with isosorbide dinitrate.

Table 21-1 Respiratory Suppression Methods Breath Holding Sustained end-expiratory breath hold Hyperventilation Supplemental oxygen Free Breathing Multiple averages Chest wall bellows Magnetic resonance navigators Navigator location Right hemi-diaphram Left hemi-diaphragam Basal left ventricle Coronary artery of interest Anterior thorax Prone vs. supine imaging Single vs. multiple navigators Prospective vs. retrospective navigator triggering Navigator triggering vs. navigator gating with real-time motion correction

methods are similar to a single breath hold.30 Direct monitoring of coronary artery motion with cardiac fat navigator echoes has also been shown as an alternative motion compensation method in coronary CMR.31,32 Preliminary data on respiratory self-gating methods in which motion is extracted from the acquired data are encouraging,33–36 but remain to be further explored.

Breath Hold Methods Initial two-dimensional (2D) coronary artery CMR methods used prolonged (15- to 20-second) end-expiratory breath holds to suppress respiratory motion.37 Although breath holding offers the advantage of relative ease of implementation in compliant subjects, it limits sequence design with regard to temporal acquisition window and image spatial resolution. Slice registration errors (caused by variability in the end-expiratory diaphragmatic position) are very 286 Cardiovascular Magnetic Resonance

common, as is diaphragmatic drift during the breath hold,17,29,30,37,38 and may occur in up to half of patients.17 The use of supplemental oxygen and hyperventilation (alone or in combination) can prolong breath hold duration,39,40 but these methods may not be appropriate for all patients, and both diaphragmatic drift and slice registration errors persist.40 The use of higher field strength and parallel imaging may ultimately allow for single-breath-hold three-dimensional (3D) coronary artery CMR similar to MDCT, but such high-resolution acquisitions are not available today.

Free Breathing Methods Early free breathing coronary artery CMR used signal averaging or chest wall bellows to reduce motion artifacts.41–43 These respiratory compensation methods were quickly supplanted by more flexible and elegant CMR respiratory navigators.

Cardiovascular Magnetic Resonance Navigators: Triggering Alone Diaphragmatic CMR navigators, first proposed by Ehman44 for abdominal MR imaging, overcome the time constraints and patient cooperation requirements imposed by multiple breath holds, thereby offering superior spatial resolution opportunities. CMR navigator implementation varies among CMR vendors. In the ideal implementation, the navigator can be positioned at any interface (see Table 21-1) that accurately reflects respiratory motion, including the dome of the right hemi-diaphragm (Fig. 21-3),38,43,45 the left hemi-diaphragm, the anterior chest wall, the anterior free wall of the left ventricle (LV),42 or even through the coronary artery of interest. The navigator should not cause an image artifact and should be temporally located immediately preceding the imaging portion of the sequence (Fig. 21-4), with data accepted (used for image reconstruction) only when the navigator indicates that the “interface” (e.g., diaphragm position) falls within a user-defined window. For simplicity and ease in set-up, the dome of the right hemi-diaphragm

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Figure 21-4 Schematic of the coronary artery cardiovascular magnetic resonance (CMR) pulse sequence for scout scanning (scout scan) and subsequent higher-resolution coronary CMR (HiRes-Scan). The elements of the sequence (T2-prep navigator [NAV], fat saturation prepulse [FSat], and three-dimensional imaging sequence) are shown in temporal relationship to the electrocardiogram and trigger delay. Note that the electrocardiogram timing and respiratory suppression methods for the scout scan and high-resolution coronary artery CMR are consistent.

has become the preferred location.42,45–47, Using a 3-mm navigator gating window, implementation results in data acceptance efficiency (accepted/total R-R interval) of 25% to 30%.42 The use of multiple navigators further erodes navigator efficiency without definite improvement in image quality. Because of increased susceptibility at the lung-liver interface at higher field strengths (3 Tesla [T]), a more central navigator position has been advocated.48

From CMR studies of cardiac border position during the respiratory cycle, Wang observed that the overwhelming effect of respiration on cardiac position is in the superior-inferior direction.28 At end-expiration, the ratio between cardiac and diaphragmatic displacement is approximately 0.6 for the RCA and approximately 0.7 for the left coronary artery,28,38 although there is variability among subjects29,49,50 and position (e.g., supine vs. prone imaging).38 This relatively fixed relationship offers the opportunity for prospective navigator gating with real-time tracking48,49 during which the position of the interface (diaphragm) is determined. The slice position coordinates can then be shifted in real time (before data collection) to allow appropriate adjustment of the spatial coordinates.39,45,50 This approach allows for the use of wider gating windows and shorter scan times (i.e., increased navigator efficiency). With real-time tracking implementation, a 5-mm diaphragmatic gating window is often used, with navigator efficiency approaching 50%.49–51 Coronary artery CMR with real-time navigator tracking has been shown to minimize registration errors (compared with breath holding), with maintained or improved image quality.42,48,49 Although we use a “fixed” superior-inferior correction factor of 0.6 (with no left-right or anterior-posterior correction),51,52 others have observed significant individual variability in this relationship29 and advocate for patient-specific algorithms. More sophisticated affine motion models that account for displacement in all three coordinates may offer advantages.53,54 Using a navigator to monitor epicardial fat has also been proposed.54 Prone positioning38,55 and abdominal Cardiovascular Magnetic Resonance 287

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Figure 21-3 Coronal (A) and transverse (B) thoracic image with identification of the navigator at the dome of the right hemi-diaphragm (RHD Nav). C, Respiratory motion of the lung-diaphragm interface recorded using a two-dimensional selective navigator with the lung (superior) and liver (inferior) interface. The maximum excursion between endinspiration and end-expiration in this example is approximately 11 mm. The broken line in the middle of C indicates the position of the lung-liver interface at each R-R interval. Data are accepted only if the lung-liver interface is within the acceptance window of 5 mm. Data acquired with the navigator outside of the window are rejected. Accepted data are indicated by the broken line at the bottom of C.

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or thoracic banding55,56 may also affect image quality. Finally, coronary artery CMR quality is improved by using consistent ECG timing as well as respiratory suppression methods for both the coronary localizing and motion scout images and coronary artery CMR acquisitions.57

Spatial Resolution If not for the limitations of SNR and acquisition duration, isotropic coronary artery CMR spatial resolution approximating the 500 µ resolution data of MDCT would be used. Spatial resolution requirements for clinical coronary artery CMR depend on whether the goal is to identify the origin and proximal course of the coronary artery (e.g., issues of anomalous coronary disease) or to identify focal stenoses in the proximal and middle segments. Figure 21-5 shows a projection X-ray coronary angiogram at 300 µ, 500 µ, 1000 µ, and 2000 µ spatial resolution. At 500 µ and 1000 µ resolution, focal disease is readily visible, whereas at resolution of greater than 1000 µ, only the course of the artery is apparent. Phantom studies confirm these observations.58

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Suppression of Signal from Surrounding Tissue The intrinsic contrast between the coronary blood pool and the adjacent myocardium and epicardial fat can be manipulated using the inflow effect for gradient echo sequences and by the application of CMR prepulses. Fat has a relatively short T1. Frequency selective prepulses can be applied to saturate signal from fat tissue, thereby allowing visualization of the underlying coronary arteries.37,41,59 The coronary arteries also run in close proximity to the epimyocardium. Myocardium and blood have relatively similar T1 relaxation values, but different T2 relaxation values. Two methods that can enhance the contrast between the coronary lumen and the underlying myocardium are T2 preparation prepulses22,60,61 and magnetization transfer contrast.41,62 The former is often used for coronary artery CMR because it also suppresses deoxygenated venous blood, whereas the latter is used for coronary vein CMR.63 The effect of fat saturation and T2 preparation prepulses is shown in Figure 21-6. Figure 21-5 X-ray coronary angiogram obtained at 300 µ (A), 500 µ; (B), 1000 µ; (C), and 2000 µ (D), in-plane spatial resolution. The focal coronary stenoses of the proximal left circumflex coronary artery (black arrow) and the proximal left anterior descending coronary artery (white arrow) are appreciated with image resolution of less than 1000 µ. (Courtesy of Daniel Sodickson, MD, PhD.)

Figure 21-6 Effect of T2 prepulses (T2 Prep) and fat saturation (Fat Sat) on coronary artery cardiovascular magnetic resonance (CMR). A, Double-oblique electrocardiogram navigator-gated/corrected coronary artery CMR in the absence of fat saturation and T2 preparation. There is little contrast between the coronary artery lumen and the epicardial fat and myocardium. B, The addition of T2 prepulse and fat saturation allows for clear definition of the right coronary artery.

Over the last two decades, coronary artery CMR sequences have continued to evolve, with increasing current interest in 3-T, parallel imaging, and contrast agent implementation. Coronary artery CMR sequences can be conceptualized as being composed of the following components: (1) cardiac (e.g., vector ECG) triggering to suppress bulk cardiac motion; (2) respiratory motion suppression (e.g., breath hold, navigators); (3) prepulses or exogenous contrast to enhance CNR of the coronary arterial blood (e.g., fat saturation, T2 preparation, magnetization transfer contrast, inversion prepulse with exogenous CMR contrast agents); and (4) image acquisition that optimizes coronary arterial SNR (Fig. 21-7). The imaging sequences (Table 21-2) may include black blood (fast spin echo and dual inversion),

bright blood (segmented k-space spoiled gradient echo and SSFP), and all implemented as 2D (typically breath hold) or 3D (prolonged breath hold or free-breathing navigator) acquisitions using Cartesian, spiral, and radial acquisitions.

Free Breathing Spin Echo Coronary Artery Cardiovascular Magnetic Resonance Early attempts to image the coronary arteries with ECGgated spin echo axial CMR were met with limited success. Lieberman64 used ECG-gated spin echo CMR and was able to visualize portions of the native coronary arteries in only 30% of 23 subjects, whereas Paulin65 studied six patients with ECG-gated spin echo imaging. Despite data acquisition during ventricular systole, the absence of respiratory motion suppression, and data acquisition over several minutes, the origin of the LM coronary artery was seen in all (100%) subjects and the ostium of the RCA was seen in more than half. No stenoses were visualized in either report. In our experience, ECG-gated T1-weighted fast spin echo imaging of the thorax at the level of the coronary arteries will often show their origin. This approach is often used to visualize reverse saphenous vein bypass grafts (see Fig. 21-9) (See Chapter 24).

Breath Hold Two-Dimensional Segmented K-Space Gradient Echo Coronary Artery Cardiovascular Magnetic Resonance

Figure 21-7 Breath hold transverse two-dimensional coronary artery CMR in a healthy subject at the level of the take-off of the right coronary artery (white arrow).53

Table 21-2 Coronary Artery Cardiovascular Magnetic Resonance Methods Black Blood Spin echo Dual inversion fast spin echo Bright Blood Segmented k-space gradient echo Two-dimensional breath hold Three-dimensional free breathing (or breath hold) Steady-state free precession Contrast-enhanced coronary magnetic resonance imaging Extracellular and intravascular agents Aortic root tagging methods Parallel Imaging Techniques Cartesian vs. Spiral vs. Radial Acquisitions High Field (3 Tesla)

The first robust approach to coronary artery CMR was ECGtriggered, breath hold, 2D segmented k-space gradient echo acquisition described nearly two decades ago37 and still applicable today on clinical CMR systems or situations in which anomalous coronary artery disease may be the clinical question. Data were acquired within a 112-msec temporal resolution (repetition time of 14 msec) during a sustained 16-heartbeat breath hold with 1.9 0.9 mm in-plane spatial resolution).66,67 A series of 10 to 15 overlapping transverse or oblique images (each requiring a single breath hold) was acquired at the level of the origin of the RCA and left coronary arteries. Because of variability in the diaphragmatic position between breath holds, the acquisition of repetitive images with the same spatial coordinates showed adjacent coronary regions. The number of breath holds can be reduced by the combination of breath holds with navigator correction.42,49,51 The combination of higher-resolution 2D gradient echo imaging with free breathing navigators is highly reproducible and has been used to detect coronary vasodilation in response to nitroglycerin.27,68 Similar breath hold 2D segmented k-space spoiled gradient echo acquisitions may also be used to image the larger-diameter coronary artery bypass grafts (Fig. 21-8). Reverse saphenous vein grafts are larger in diameter, and both saphenous vein grafts and internal mammary bypass grafts are less mobile than the native coronary arteries, with Cardiovascular Magnetic Resonance 289

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CORONARY ARTERY CARDIOVASCULAR MAGNETIC RESONANCE ACQUISITION SEQUENCES

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Figure 21-8 Oblique breath hold two-dimensional coronary magnetic resonance imaging of a patent saphenous vein bypass graft (SVG). Two adjacent images show the SVG (arrows) extending from its aortic origin (Ao) to the distal touchdown on the posterior descending coronary artery. LV, left ventricle; RV, right ventricle.

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predominant flow during ventricular systole. This facilitates data acquisition during a longer period (150 to 200 msec) within each R-R interval and with less rigorous respiratory motion criteria. Susceptibility artifacts from stainless steel bypass graft markers and vascular clips pose an impediment to bypass graft coronary CMR (Fig. 21-9).

Three-Dimensional Coronary Artery Cardiovascular Magnetic Resonance Methods The higher SNR of thin-slab or thick-slab 3D acquisitions and their flexibility on visualization of the imaging data in any arbitrary orientation have made it the predominant approach for the last decade. The associated respiratory

suppression approach has varied from prolonged breath hold with lower spatial resolution59,69 to higher-spatialresolution methods in combination with free breathing and navigator methods22,41,70 or prolonged breath holding with parallel acquisitions.71–73 Because data from a volume of tissue surrounding the coronary arteries are acquired, the setup of 3D coronary artery CMR is less demanding of the patient and CMR technologist than repetitive 2D breath hold acquisitions. After obtaining thoracic scouts (nine transverse, nine coronal, and nine sagittal interleaved acquisitions), the navigator is positioned at the dome of the right hemi-diaphragm (see Fig. 21-3). Free breathing or breath hold cine (consistent with the subsequent coronary artery CMR sequence) is then acquired perpendicular to the mid-RCA to define the optimal delay and acquisition period (period of minimal in-plane motion). A low-resolution whole heart scout, consisting of an ECG-triggered 3D SSFP with both fat

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Figure 21-9 A, Posterior-anterior chest X-ray in a patient with coronary artery bypass grafts. Note the sternal wires (thick arrow) as well as the coronary artery bypass graft markers (thin arrow). B, Transverse coronary artery cardiovascular magnetic resonance (CMR) in the same patient. Note the large local artifacts (signal voids) related to the sternal wires (thick arrow) and bypass graft markers (thin arrow). The size of the artifacts is related to the type of graft marker used. C, Barium and tantalum markers (arrow) result in the smallest artifacts. The size of the artifacts is somewhat reduced with spin echo and black-blood CMR methods. 290 Cardiovascular Magnetic Resonance

Figure 21-10 Targeted three-dimensional reformatted coronary magnetic resonance imaging of the left coronary system acquired using free breathing and real-time navigator gating with motion correction in a healthy adult subject. The transverse acquisition shows the left main artery, left anterior descending artery, and left circumflex coronary artery. The in-plane spatial resolution is 0.7 1.0 mm. Figure 21-11 Coronal scout showing the whole heart coronary artery cardiovascular magnetic resonance setup with the thick- slab threedimensional axial dataset (yellow box) compared with targeted thin-slab three-dimensional acquisitions for the atrioventricular groove (gray box; right coronary artery, left circumflex coronary artery) and axial dataset (green box; left main artery, left anterior descending artery, left circumflex coronary artery, proximal right coronary artery).

For imaging of the RCA, transverse images from the second scout showing the proximal, mid-, and distal RCA are identified. Using a three-point planscan software tool,70 the imaging plane passing through all three coordinates of the RCA is identified and the targeted 3D coronary sequence is repeated in this orientation. For CMR systems that lack a three-point planscan interactive software tool, an imaging plane along the right and left atrioventricular groove is suggested. Each submillimeter in-plane 3D segmented gradient echo acquisition is typically 6 to 8 minutes in duration (assuming a navigator efficiency of 40% to 50%). A similar approach has also been applied for coronary artery bypass grafts.74 For whole heart acquisitions,75 the scout imaging process is similar, but a single, 3D thick axial volume, positioned approximately 1 cm above the LM and extending to the inferior cardiac border (Fig. 21-11), is prescribed from coronal scouts.75 Whole heart coronary CMR provides for visualization of more distal segments of the coronary arteries75,76 and is more conducive to advanced postprocessing methods (Fig. 21-12). Isotropic voxels are also possible.77 Both segmented k-space gradient echo and SSFP coronary artery CMR methods may be used for targeted thinslab 3D acquisitions. Thin-slab 3D targeted acquisitions with a gradient echo imaging sequence result in a more homogenous blood pool signal. This approach is more heavily dependent on the inflow of unsaturated protons.78 If coronary flow is slow or stagnant, saturation effects will cause a local signal loss that is often relatively exaggerated compared with lumen stenosis. SSFP applications offer superior SNR with a 50% improvement in blood SNR and blood-myocardium CNR compared with gradient echo methods,79 with reduced sensitivity to inflow effects. As a result, at most centers, SSFP methods are increasingly used for both breath hold80–82 and free breathing coronary artery CMR75,79,83–85 at 1.5 T. They offer specific advantages for systolic imaging.56 Fat saturation and T2 prepulses are still used. The reproducibility of free breathing targeted

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saturation and T2 preparation prepulses, is then acquired with diaphragmatic navigator gating to determine the course of the coronary arteries. Subsequently, either thinslab targeted 3D coronary artery CMR or thick-slab whole heart coronary artery CMR can be performed. For the targeted 3D approach, two datasets are acquired. Visualization of the LM, LAD, and left circumflex (LCX) coronary arteries, a 3D volume is interactively prescribed in the axial plane centered around the LM coronary artery (identified in the second scout) using the same ECG delay and navigator parameters as the scout. Typically, a 30-mm slab with 20 overlapping slices is acquired using a segmented k-space gradient echo or SSFP acquisition with submillimeter in-plane spatial resolution (0.7 1.0 mm) and a temporal acquisition of 56 to 84 msec/heartbeat (Fig. 21-10).22,70

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Figure 21-13 Free breathing three-dimensional spiral coronary artery cardiovascular magnetic resonance in a healthy subject. (Courtesy of Peter Bo¨ernert, PhD.) Figure 21-12 Three-dimensional reconstruction of whole heart steady-state free precession coronary magnetic resonance imaging after computer-assisted image segmentation enabled major coronary vessels to be visualized. (Courtesy of Hajime Sakuma, MD.)

3D coronary approaches appears to be excellent.86,87 Although both SSFP and gradient echo approaches can be used for targeted acquisition, only SSFP imaging is used for the whole heart approach because of the more favorable in-flow properties of SSFP.78 The higher SNR of 3D acquisition allows for the application of parallel imaging (e.g., simultaneous acquisition of spatial harmonics [SMASH],88 sensitivity encoding [SENSE]),89 and generalized autocalibrating partially parallel acquisitions [GRAPPA])90 to reduce scan time (see Chapter 3).81,82,91,92

CORONARY ARTERY CARDIOVASCULAR MAGNETIC RESONANCE: ADVANCED METHODS Both SNR and the speed of data acquisition remain major limitations for coronary artery CMR. To overcome these hurdles, several CMR centers continue with the development and implementation of novel approaches, including nonCartesian acquisitions, CMR contrast agents, and higher-field (3 T) imaging.

Spiral and Radial Coronary Magnetic Resonance Imaging Alternative k-space acquisitions, including spiral and radial coronary magnetic resonance imaging, have received attention. Meyer and colleagues69 first reported the use of spiral coronary artery CMR. The advantages of spiral (vs. Cartesian) 292 Cardiovascular Magnetic Resonance

acquisitions include more efficient filling of k-space (Fig. 21-13), enhanced SNR,79,93 and favorable flow properties. Spirals have more complex reconstruction algorithms. Although a single-shot k-space trajectory can be used, interleaved spiral imaging is preferred because it is associated with reduced artifacts.69,93–95 Spiral coronary artery CMR may be implemented with both breath hold (2D) and free breathing or navigator gating.79,93–96 Data suggest that single spiral acquisitions (per R-R interval) afford a nearly threefold improvement in SNR compared with conventional Cartesian approaches (Fig. 21-14).79,93 Acquiring two spiral acquisitions during each R-R interval will halve the acquisition time while maintaining superior SNR (vs. Cartesian acquisition) and CNR. Variable density spirals may also be beneficial.97 Radial approaches also offer the benefit of more rapid acquisitions with decreased sensitivity to motion. Data in healthy subjects appear promising,79,95,96,98 and this technique may be particularly beneficial for coronary wall imaging99–101 (see Chapter 26).

Contrast-Enhanced Coronary Artery Cardiovascular Magnetic Resonance With contrast-enhanced coronary artery CMR methods, blood signal contrast is based on the intravascular T1 relaxation rate, potentially allowing for “true” lumen imaging. Contrast-enhanced magnetic resonance angiography is widely employed for carotid, aortic, renal, and peripheral vascular applications, but the previously described unique timing constraints for coronary artery CMR have limited coronary applications for clinically available extracellular agents. Issues related to nephrogenic systemic fibrosis (see Chapter 6) in this susceptible population have led to some hesitancy in pursuing contrast agent coronary artery CMR, but newer agents are now available and are being explored.

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Exogenous CMR contrast agents can be subcategorized into extracellular (interstitial) and intravascular agents. Extracellular paramagnetic contrast agents (gadolinium chelates) have been used for first-pass breath hold studies through the coronary bed.102–104 As with aortic magnetic resonance angiography, a test bolus is often used. These agents have been shown to be of some value for first-pass and singlebreath-hold approaches,102–104 but clinical studies directly comparing contrast with noncontrast coronary artery CMR are lacking. The requirement for placement of an intravenous catheter, the added expense of the contrast agent, and the dependence of the acquisition on a single first-pass

D

acquisition, without an opportunity for repetition, limits the application of targeted 3D high-resolution approaches. A slow infusion of gadobenate dimeglumine (Gd-BOPTA, MultiHance, Bracco, Milan, Italy) appears particularly intriguing for coronary artery CMR because it has a weak and transient interaction with plasma proteins; therefore, it has up to twice the in vivo relaxivity of other extracellular agents105 (Fig. 21-15). One intravascular CMR contrast agent is currently available (Vasovist, Scherring AG, Germany),106 and several blood pool contrast agents are under development. These include gadolinium-based106–110 and ultra-small-particle

Figure 21-15 A, Contrast improvements achieved with exogenous contrast agents, such as gadobenate dimeglumine (GdBOPTA, MultiHance, Bracco, Milan, Italy). Images were acquired with an inversion recovery three-dimensional whole heart gradient echo imaging sequence. B, Reformatted proximal segment of the right coronary artery (RCA) can be visualized easily.

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Figure 21-14 Double oblique view of the left anterior descending artery (LAD; A and B) and right coronary artery (RCA; C and D) acquired with a targeted three-dimensional fat- and muscle-suppressed (T2 preparation) Cartesian (A and C) and spiral (B and D) k-space sampling technique. Because of more efficient k-space sampling, the spiral technique results in a threefold signal-to-noise improvement compared with the conventional Cartesian segmented k-space technique. (Courtesy of Peter Bo¨rnert, PhD.)

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superparamagnetic iron-oxide-based111–114 contrast agents. These intravascular agents afford longer scan times with free breathing or repeated breath hold methods. A 180 inversion prepulse is often used to highlight the marked reduction in T1 with imaging when the longitudinal magnetization of myocardium crosses the null point.106 Improvements in CNR of 60% to 100% have been reported.

3-Tesla Coronary Magnetic Resonance Imaging

at both 1.5 T122–127 and 3.0 T,128 with clinical safety shown for imaging early after implantation at 1.5 T.122–125 In the United States, both the Cypher (Cordis, Miami Lakes, FL) and Taxus Liberte (Boston Scientific, Natick, MA) and Xeience (Abbott Laboratories, Abbott Park, IL) drug-eluting stents are approved for CMR scanning immediately after implantation, but the local susceptibility artifact that leads to signal voids or artifacts at the site of the stent can be substantial (Fig. 21-17). The signal void is dependent on both the stent material129 and the CMR sequence. There appear to be relatively small artifacts with tantalum

An exciting area of interest has been high-field, 3.0-T coronary artery CMR. SNR is directly related to field strength (B0), and this technique offers the opportunity to double SNR (see Chapter 13). Although the vast majority of coronary artery CMR studies have been performed on 1.5-T systems, commercial 3.0-T systems are increasingly available and are becoming the platform of choice for the testing of many advances. Technical problems include artifacts caused by field inhomogeneities,115 increased susceptibility to artifacts, reduced T2*,48,116,117 T1 prolongation, and the amplified magnetohydrodynamic effect.8,9 Free breathing navigator and breath hold 3D coronary artery CMR studies in healthy volunteers showed greater than 50% improvement in SNR, with impressive image quality, using segmented k-space gradient echo (Fig. 21-16), SSFP,118 spiral, and contrast-enhanced methods.48,119–120a Very preliminary coronary artery CMR studies with a 7.0-T system suggest that it may be the “new CMR frontier.”121

Stent

Special Considerations: Intracoronary Stents Improvements in long-term patency rates for percutaneous coronary interventions using conventional and drug-eluting intracoronary stents have resulted in their widespread use, including more than 80% of the growing number of percutaneous revascularizations. Typically made of high-grade stainless steel, these stents pose a particular imaging problem for CMR. The attractive force and local heating are negligible

RCA

Figure 21-17 Transverse two-dimensional breath hold gradient echo coronary magnetic resonance imaging at the level of the left anterior descending artery in a patient with a patent stent. Note the signal void (black marker) corresponding to the site of the stent. (Courtesy of Christopher Kramer, MD.)

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294 Cardiovascular Magnetic Resonance

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Figure 21-16 Comparative three-dimensional gradient echo coronary artery cardiovascular magnetic resonance at 1.5 Tesla (A) and 3.0 Tesla (B). The increased signal-to-noise ratio afforded by the higher field strength allows for enhanced (0.6 0.6 mm) spatial resolution. (Courtesy of Matthias Stuber, PhD.) Ao, aorta; LAD, left anterior descending; LCX, left circumflex; LM, left main; RCA, right coronary artery.

Coronary Vein Cardiovascular Magnetic Resonance With the advent of cardiac resynchronization therapy for patients with congestive heart failure, there has been increased interest in imaging the coronary vein anatomy to guide the placement of lateral pacing leads (to determine the presence of an appropriate lead and as a guide to fluoroscopy) and to determine whether there is an underlying

myocardial scar in the area of the vein. The technical obstacles associated with coronary vein CMR are similar to those for coronary artery CMR, but spatial resolution requirements are less restrictive because information regarding vein anatomy and vessel size is desired, but not information about focal stenoses. CMR is ideally suited for this application because of the lack of ionizing radiation and the lack of iodinated contrast required by its competitor, MDCT. Both noncontrast63 and contrast131–133 types of coronary vein CMR have been implemented. For noncontrast approaches, the T2 preparation pulse is exchanged for a magnetization transfer pulse to avoid attenuation of deoxygenated blood present in the coronary vein (Fig. 21-18). Imaging during end-systole (vs. mid-diastole for coronary artery imaging) also is preferred because it coincides with the maximum coronary vein size.63

Future Technical Developments Current research on coronary artery CMR is focused on improving motion correction algorithms, enhancing SNR and CNR, and reducing image acquisition time. Several

CS CS LatV

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Figure 21-18 The variations in the coronary venous anatomy can be seen in images from four healthy adult subjects acquired with magnetizationprepared spoiled gradient echo coronary vein cardiovascular magnetic resonance (CMR) during the systolic rest period. There are clear variations in the branching point, angle, and diameter of different tributaries of the coronary sinus (CS), highlighting the potential of coronary vein CMR for noninvasive assessment of the coronary venous anatomy. LatV, lateral branch; PostV; posterior branch; RCA, right coronary artery. Cardiovascular Magnetic Resonance 295

21 CORONARY ARTERY AND VEIN IMAGING: METHODS

stents130 and very prominent artifacts with stainless steel stents. Artifacts are also relatively larger with gradient echo methods. This signal void or artifact precludes direct evaluation of intrastent and peri-stent coronary integrity, although assessment of blood flow and direction proximal and distal to the stent using CMR flow methods or spin labeling methods may provide indirect evidence of patency by documentation of antegrade flow. Animal studies using novel “CMR lucent” stent materials have been reported,130 but mechanical properties, biocompatibility, and long-term patency rates for these novel stents are unknown.

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imaging methods, such as high-magnetic-field (3-T) imaging, highly accelerated acquisition and reconstruction with phased array coils, and the use of exogenous contrast agents, could potentially overcome some of the limitations of current coronary artery CMR approaches. The goal is to provide a noninvasive imaging tool that will allow for

early identification or exclusion of proximal and mid-vessel coronary artery disease. As these methods are further developed, comparative multicenter studies will need to be performed to define the role of coronary artery CMR both within the context of the comprehensive CMR examination and within the field of noninvasive cardiac imaging.

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91. Park J, Larson AC, Zhang Q, et al. High-resolution steady-state free precession coronary magnetic resonance angiography within a breath-hold: parallel imaging with extended cardiac data acquisition. Magn Reson Med. 2005;54(5):1100–1106. 92. Niendorf T, Hardy CJ, Giaquinto RO. Toward single breath-hold whole-heart coverage coronary MRA using highly accelerated parallel imaging with a 32-channel MR system. Magn Reson Med. 2006;56 (1):167–176. 93. Bornert P, Stuber M, Botnar RM, et al. Direct comparison of 3D spiral vs. Cartesian gradient-echo coronary magnetic resonance angiography. Magn Reson Med. 2001;46(4):789–794. 94. Thedens DR, Irarrazaval P, Sachs TS, et al. Fast magnetic resonance coronary angiography with a three-dimensional stack of spirals trajectory. Magn Reson Med. 1999;41(6):1170–1179. 95. Bornert P, Aldefeld B, Nehrke K. Improved 3D spiral imaging for coronary MR angiography. Magn Reson Med. 2001;45(1):172–175. 96. Weber OM, Pujadas S, Martin AJ, Higgins CB. Free-breathing, threedimensional coronary artery magnetic resonance angiography: comparison of sequences. J Magn Reson Imaging. 2004;20(3):395–402. 97. Sussman MS, Stainsby JA, Robert N, et al. Variable-density adaptive imaging for high-resolution coronary artery MRI. Magn Reson Med. 2002;48:753–764. 98. Leiner T, Katsimaglis G, Kissinger KV, et al. Comparison of Cartesian and radial balanced GTFE coronary MRA [abstract]. J Cardiovasc Magn Reson. 2004;6(1):75. 99. Park J, Larson AC, Zhang Q, et al. 4D Radial coronary artery imaging within a single breath-hold: cine angiography with phase-sensitive fat suppression (CAPS). Magn Reson Med. 2005;54(5):833–840. 100. Botnar RM, Stuber M, Kissinger KV, et al. Non-invasive coronary vessel wall and plaque imaging with magnetic resonance imaging. Circulation. 2000;102(12):2582–2587. 101. Katoh M, Spuentrup E, Buecker A, et al. MR coronary vessel wall imaging: comparison between radial and spiral k-space sampling. J Magn Reson Imaging. 2006;23:757–762. 102. Zheng J, et al. Three-dimensional gadolinium-enhanced coronary magnetic resonance angiography: initial experience. J Cardiovasc Magn Reson. 1999;1(1):33–41. 103. Goldfarb JW, Edelman RR. Coronary arteries: breath-hold, gadoliniumenhanced, three-dimensional MR angiography. Radiology. 1998;206 (3):830–834. 104. Deshpande V, Li D. Contrast-enhanced coronary artery imaging using 3D TrueFISP. Magn Reson Med. 2003;50(3):570–577. 105. Rohrer M, Bauer H, Mintorovitch J, Requardt M, Weinmann HJ. Comparison of magnetic properties of MRI contrast media solutions at different magnetic field strengths. Invest Radiol. 2005;40:715–724. 106. Stuber M, Botnar RM, Danias PG, et al. Contrast agent-enhanced, free-breathing, three-dimensional coronary magnetic resonance angiography. J Magn Reson Imaging. 1999;10(5):790–799. 107. Paetsch I, Jahnke C, Barkhausen J, et al. Detection of coronary stenoses with contrast enhanced, three-dimensional free breathing coronary MR angiography using the gadolinium-based intravascular contrast agent gadocoletic acid (B-22956). J Cardiovasc Magn Reson. 2006;8(3):509–516. 108. Huber ME, Paetsch I, Schnackenburg B, et al. Performance of a new gadolinium-based intravascular contrast agent in free-breathing inversion-recovery 3D coronary MRA. Magn Reson Med. 2003;49 (1):115–121. 109. Herborn CU, Barkhausen J, Paetsch I, et al. Coronary arteries: contrast-enhanced MR imaging with SH L 643A: experience in 12 volunteers. Radiology. 2003;229(1):217–223. 110. Paetsch I, Jahnke C, Barkhausen J, et al. Detection of coronary stenoses with contrast enhanced, three-dimensional free breathing coronary MR angiography using the gadolinium-based intravascular contrast agent gadocoletic acid (B-22956). J Cardiovasc Magn Reson. 2006;(8):509–516. 111. Knuesel PR, Nanz D, Wolfensberger U, et al. Multislice breath-hold spiral magnetic resonance coronary angiography in patients with coronary artery disease: effect of intravascular contrast medium. J Magn Reson Imaging. 2002;16(6):660–667.

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112. Taylor AM, Panting JR, Keegan J, et al. Safety and preliminary findings with the intravascular contrast agent NC100150 injection for MR coronary angiography. J Magn Reson Imaging. 1999;9 (2):220–227. 113. Sandstede JJ, Pabst T, Wacker C, et al. Breath-hold 3D MR coronary angiography with a new intravascular contrast agent (feruglose): first clinical experiences. Magn Reson Imaging. 2001;19(2):201–205. 114. Ringgaard S, Pedersen M, Rickers J, et al. Spiral coronary angiography using a blood pool agent. J Magn Reson Imaging. 2005;22:213–218. 115. Nezafat R, Stuber M, Ouwerkerk R, et al. B1-insensitive T2 preparation for improved coronary magnetic resonance angiography. Magn Reson Med. 2006;55:858–864. 116. Noeske R, Seifert F, Rhein KH, Rinneberg H. Human cardiac imaging at 3 T using phased array coils. Magn Reson Med. 2000;44 (6):978–982. 117. Atalay MK, Poncelet BP, Kantor HL, et al. Cardiac susceptibility artifacts arising from the heart-lung interface. Magn Reson Med. 2001;45 (2):341–345. 118. Bi X, Deshpande V, Simonetti O, et al. Three-dimensional breathhold SSFP coronary MRA: a comparison between 1.5T and 3.0T. J Magn Reson Imaging. 2005;22:206–212. 119. Santos JM, Cunningham CH, Lustig M, et al. Single breath-hold whole-heart MRA using variable density spirals at 3T. Magn Reson Med. 2006;55(2):371–379. 120. Bi X, Li D. Coronary arteries at 3.0T: contrast-enhanced magnetization-prepared three-dimensional breathhold MR angiography. J Magn Reson Imaging. 2005;21(2):133–139. 120a. Yang Q, Li K, Liu X, et al. Contrast-enhanced whole heart coronary magnetic resonance angiography at 3T: a comparative study with x-ray angiography in a single center. J Am Coll Cardiol. 2009;54: 69–76. 121. van Elderen SGC, Webb AG, Versluis M, et al. In vivo human coronary magnetic resonance angiography at 7 Tesla. J Cardiovasc Magn Reson. 2009;11(suppl 1):36. 122. Strohm O, et al. Safety of implantable coronary stents during 1Hmagnetic resonance imaging at 1.0 and 1.5 T. J Cardiovasc Magn Reson. 1999;1(3):239–245. 123. Kramer CM, Rogers Jr WJ, Pakstis DL. Absence of adverse outcomes after magnetic resonance imaging early after stent placement for acute myocardial infarction: a preliminary study. J Cardiovasc Magn Reson. 2000;2(4):257–261. 124. Gerber TC, Fasseas P, Lennon RJ, et al. Clinical safety of magnetic resonance imaging early after coronary artery stent placement. J Am Coll Cardiol. 2003;42(7):1295–1298. 125. Hug J, Nagel E, Bornstedt, et al. Coronary arterial stents: safety and artifacts during MR imaging. Radiology. 2000;216(3):781–787. 126. Scott NA, Pettigrew RI. Absence of movement of coronary stents after placement in a magnetic resonance imaging field. Am J Cardiol. 1994;73(12):900–901. 127. Shellock FG, Shellock VJ. Metallic stents: evaluation of MR imaging safety. AJR Am J Roentgenol. 1999;173(3):543–547. 128. Shellock FG. MR safety at 3-Tesla: bare metal and drug eluting coronary artery stents. Signals. 2005;53(2):26–27. 129. Maintz DC, Botnar RM, Fischbach R, et al. Coronary magnetic resonance angiography for assessment of the stent lumen: a phantom study. J Cardiovasc Magn Reson. 2002;4(3):359–367. 130. Spuentrup E, Ruebben A, Schaeffter T, et al. Magnetic resonance– guided coronary artery stent placement in a swine model. Circulation. 2002;105(7):874–879. 131. Chiribiri A, Kelle S, Gotze S, et al. Visualization of the cardiac venous system using cardiac magnetic resonance. Am J Cardiol. 2008;101: 407–412. 132. Chiribiri A, Kelle S, Kohler U, et al. Magnetic resonance cardiac vein imaging: relation to mitral valve annulus and left circumflex coronary artery. JACC: Cardiovasc Imaging. 2008;1:729–738. 133. Younger JF, Plein S, Crean A, et al. Visualization of coronary venous anatomy by cardiovascular magnetic resonance. J Cardiovasc Magn Reson. 2009;11:26.

Coronary Artery Imaging: Clinical Results Thomas H. Hauser, Jonathan Chan, and Warren J. Manning

Chapter 21 reviewed the technical issues and solutions related to coronary artery cardiovascular magnetic resonance (CMR). This chapter reviews the clinical data comparing coronary artery CMR with X-ray coronary angiography for identification of anomalous coronary artery disease, characterization of coronary artery aneurysms, and detection of native vessel disease. It also describes studies comparing CMR with coronary artery multidetector computed tomography (MDCT), the other principal noninvasive modality for imaging the coronary arteries. The majority of data represent single-center experience, with quantitative X-ray coronary angiography used as the reference standard for most of the larger and multicenter studies.

IDENTIFICATION OF ANOMALOUS CORONARY ARTERIES Although unusual (<1% of the general population1,2) and most often benign, congenital coronary anomalies in which the anomalous segment courses anterior to the aorta and posterior to the pulmonary artery are well-recognized causes of myocardial ischemia and sudden cardiac death in children and young adults.3 These adverse events commonly occur during or immediately after intense exercise and are believed to be related to compression of the anomalous segment, vessel kinking, or coexistent eccentric ostial stenoses.3 The ability of coronary artery CMR to reliably identify the major coronary arteries and their relationship to the ascending aorta and pulmonary artery immediately provides for its application for the identification and characterization of anomalous coronary artery disease. The spatial resolution requirements for identifying anomalous coronary vessels are less stringent than for defining native vessel stenoses, allowing for lower resolution and faster CMR imaging. Projection X-ray angiography had traditionally been the imaging test of choice for the diagnosis and characterization of these anomalies. However, the presence of an anomalous vessel is sometimes suspected only after the procedure, particularly in a situation in which there was unsuccessful engagement of a coronary artery. In addition, the declining use of a pulmonary artery catheter has made characterization of the anterior or posterior trajectory of the anomalous vessels more difficult to appreciate on projection X-ray angiography.

Coronary artery CMR has several advantages compared with both coronary MDCT and X-ray angiography in the diagnosis of these coronary anomalies. In addition to being noninvasive and not requiring ionizing radiation or iodinated contrast agents, coronary artery CMR provides a definitive three-dimensional (3D) “road map” of the mediastinum (Fig. 22-1). Early studies applied two-dimensional (2D) breath hold segmented k-space gradient echo coronary artery CMR,4–6 although most centers now use targeted 3D7–9 or whole heart10–12 free breathing navigator coronary artery CMR because of the superior reconstruction capabilities afforded by 3D datasets, with similar results. The ability to acquire these data using CMR without the use of ionizing radiation is likely to be of particular benefit in the generally younger population.12a There have been at least six published series4–9 of patients who underwent a blinded comparison of coronary artery CMR data with X-ray angiography. These studies have uniformly reported excellent accuracy, including several instances in which coronary artery CMR was determined to be superior to X-ray angiography (Table 22-1). These data have been extended to imaging performed at 3 T.12 As a result, CMR is considered a class I indication for suspected anomalous coronary artery disease.13 At experienced CMR centers, clinical coronary artery CMR is the preferred test for patients in whom anomalous disease is suspected, those with known anomalous disease that must be further clarified, and those with another cardiac anomaly associated with coronary anomalies (e.g., tetralogy of Fallot). Although MDCT has also been shown to be efficacious for this indication,14–16 coronary artery CMR is often preferred because there is no need for ionizing radiation or intravenous contrast.

CORONARY ARTERY ANEURYSMS AND KAWASAKI DISEASE Coronary artery aneurysms are relatively uncommon, but have received increasing attention because of their common occurrence in pediatric and young adult patients. In the absence of a percutaneous intervention, the vast majority of acquired coronary aneurysms are caused by mucocutaneous lymph node syndrome (Kawasaki disease), a generalized vasculitis of unknown etiology usually occurring in children younger than 5 years old. Infants and children Cardiovascular Magnetic Resonance 299

22 CORONARY ARTERY IMAGING: CLINICAL RESULTS

CHAPTER 22

ISCHEMIC HEART DISEASE

RCA

DPI/DC PA AA

AO

LAD RCA

LM

LAD LA

A

B

TABLE 22-1 Coronary Artery Cardiovascular Magnetic Resonance for Anomalous Coronary Artery Disease Investigator

No. of Patients

Correctly Classified Vessels

McConnell et al.[4] Post et al.[5] Vliegen et al.[6] Taylor et al.[7] Bunce et al.[8] Razmi et al.[9]

15 19 12 25 26 12

14 19 11 24 26 12

(93%) (100%)* (92%){ (96%) (100%){ (100%)

*Including 3 patients originally misclassified by X-ray angiography. { Including 5 patients unable to be classified by X-ray angiography. { Including 11 patients unable to be classified by X-ray angiography.

Figure 22-1 Free breathing targeted three-dimensional coronary artery cardiovascular magnetic resonance using T2 prepulse navigator gating with real-time motion correction. A, Transverse orientation showing a malignant-type anomalous left anterior descending coronary artery (LAD) originating from the right coronary artery (RCA). B, Transverse image in another patient with a malignant-type anomalous origin of the RCA from the left coronary cusp. AA, AO, ascending aorta; LA, left atrium; LM, left main coronary artery; PA, pulmonary artery; RA, right atrium. (From Simonetti OP et al. An improved MR imaging technique for the visualization of myocardial infarction. Radiology. 2001; 218: 215–223.)

with this syndrome may show evidence of myocarditis or pericarditis, with nearly 20% having coronary artery aneurysms. These aneurysms are the source of both short- and long-term morbidity and mortality.17 Fortunately, half of the children with coronary aneurysms during the acute phase of the disease have a normal-appearing coronary lumen on catheter-based X-ray angiography 2 years later.17,18 Transthoracic echocardiography is often adequate for diagnosing and following aneurysms in very young children, but this modality is deficient after adolescence and in obese children. These young adults are therefore often referred for serial catheter-based X-ray coronary angiography, with the accumulation of significant radiation exposure over time. Coronary artery CMR data from two series of adolescents and young adults with coronary artery aneurysms (Fig. 22-2) have confirmed the high accuracy of coronary artery CMR for both the identification and the

Figure 22-2 A, Transverse targeted three-dimensional T2 prepulse coronary artery cardiovascular magnetic resonance of a subject with a left coronary artery aneurysm. B, corresponding X-ray angiogram, showing good correlation with the cardiovascular magnetic resonance findings.

A

300 Cardiovascular Magnetic Resonance

B

Data support a broad clinical role for coronary artery CMR in the assessment of suspected anomalous coronary artery disease (and coronary artery bypass graft patency, which is addressed later), but data are not yet sufficient to support the use of clinical coronary artery CMR for routine identification of coronary artery stenoses among patients presenting with chest pain or for screening, even in high-risk patients. However, increasing data suggest a role for coronary artery CMR among patients referred for catheter-based X-ray angiography and especially for the discrimination of ischemic versus nonischemic cardiomyopathy.23,24 As previously discussed, gradient echo sequences show rapidly moving laminar blood flow as “bright,” whereas areas of stagnant flow or focal turbulence appear “dark” because of local saturation (stagnant flow) or dephasing (turbulence; Fig. 22-3). Areas of focal stenoses appear as varying severity of “signal voids” in the coronary artery

X-ray diameter stenosis

NATIVE VESSEL CORONARY ARTERY STENOSES

CMR, with the severity of signal loss related to the angiographic stenosis (Fig. 22-4).25 Because of time constraints of breath hold, 2D breath hold coronary artery CMR has relatively limited in-plane spatial resolution, but the technique has successfully shown proximal coronary stenoses in several clinical studies (Table 22-2).25–28 When

p < 0.001

100 90 80 70

p < 0.001

60 50 40 30 20 10 0 Irregular wall

Partial

Severe

Severity of cardiovascular magnetic resonance signal loss Figure 22-4 The severity of signal loss correlates with the X-ray angiographic diameter stenosis. CMR, cardiovascular magnetic resonance imaging. Adapted from Pennell DJ, Bogren HG, Keegan J, et al. Assessment of coronary artery stenosis by magnetic resonance imaging. Heart. 1996;75(2):127–133.

Figure 22-3 A, Targeted threedimensional steady-state free precession coronary artery cardiovascular magnetic resonance showing focal stenoses in the midleft anterior descending artery (LAD) (solid arrow) and proximal left circumflex artery (dashed arrow). B, Corresponding X-ray angiogram with analogous arrows. C, Targeted three-dimensional gradient echo coronary artery cardiovascular magnetic resonance showing a focal stenosis (arrow) in the mid-LAD. D, Corresponding X-ray angiogram with LAD stenosis (arrow).

A

C

B

D

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22 CORONARY ARTERY IMAGING: CLINICAL RESULTS

characterization (diameter and length) of these aneurysms.19–21 Although not specifically examined in long-term follow-up studies, these data suggest that coronary aneurysms can now be effectively followed with serial CMR studies. Similar data have been reported for ectatic coronary vessels.22

Investigator [26]

Manning et al. Duerinckx and Urman[27] Pennell et al.[25] Post et al.[28]

No. of Subjects

No. of (%) Vessels

Sensitivity

Specificity

39 20 39 35

52 27 55 35

90% 63% 85% 63%

92% (78%–100%) 56% (37%–82%) — 89% (73%–96%)

(35%) (34%) (35%) (28%)

reported, the distance from the vessel origin to the focal stenosis on coronary artery CMR correlates closely with X-ray angiography findings (Fig. 22-5).25 However, there have been wide variations in the reported sensitivity and specificity of 2D coronary artery CMR, much of which

60

50 Distance to arterial origin by (X-ray) (mm)

ISCHEMIC HEART DISEASE

TABLE 22-2 Two-Dimensional Breath Hold Coronary Artery Cardiovascular Magnetic Resonance for Identification of Focal Coronary Stenoses 50% or More in Diameter

40

30

20

10

0 0

10

20

30

40

50

60

Distance to arterial origin by MRI (mm)

Figure 22-5 Scatterplot comparing the distance from the coronary origin to the stenosis as measured by X-ray and coronary artery cardiovascular magnetic resonance.

(71%–100%) (0%–73%) (75%–100%) (0%–100%)

may be attributable to technical and methodologic issues, including the wide variation in patient selection, presence of arrhythmias, and prevalence of disease, and the wide range of variation in technical implementation (CMR vendor, echo time, receiver coils, timing of acquisition, acquisition duration, and breath hold maneuvers) and the need for 20 to 40 breath holds to complete a study. To date, no multicenter 2D coronary artery CMR data are available. With the increasing availability of CMR navigators, most CMR centers have migrated to a free breathing 3D gradient echo or steady-state free precession (SSFP) coronary artery CMR with targeted 3D or whole heart approaches. This method has higher patient acceptance (free breathing with no breath holds required), improved signal-to-noise ratio, and datasets that facilitate multiplanar reconstructions for improved visualization of the coronary arteries. As with 2D gradient echo methods, a focal stenosis or turbulent flow appears as a signal void or marked narrowing of the lumen along the course of the vessel, with less dependence on blood flow characteristics (laminar vs. turbulent) with SSFP imaging (Fig. 22-6). Data from numerous single-center sites have been published using modern prospective navigator gating with real-time correction (Table 22-3),29–40 Although assessment of the sensitivity of coronary artery stenosis was found to be similar for both source and projection images,41 our strong preference is to make diagnoses through review of the source images. We then use the projection images to convey our findings visually to the referring physician. An international, multicenter, free breathing, 3D targeted coronary artery CMR study of 109 patients without previous X-ray angiography using common hardware and software showed high sensitivity (although only modest specificity) and negative predictive Figure 22-6 Whole heart threedimensional steady-state free precession coronary artery cardiovascular magnetic resonance. A, Mild proximal left anterior descending coronary artery stenosis (proximal white arrow). B, X-ray coronary angiogram with mild stenosis (red arrow). C, Three-dimensional volume rendering view. (Courtesy of Dr. Hajime Sakuma.)

A

B

302 Cardiovascular Magnetic Resonance

C

Investigator

Cardiovascular Magnetic Resonance Method For Diameter Stenosis 50%

No. of Subjects

Sensitivity

Specificity

Prospective Navigators with Real-time Correction Targeted Three-Dimensional Coronary Artery CMR Bunce et al.[29] Moustapha et al.[30]

34 25

TFE TFE

Sommer et al.[31]

112

TFE

Bogaert et al.[32] Plein et al.[33] Ozgun et al.[34]

19 10 14

Dewey et al.[35] Maintz et al.[36]

15* 12

Ozgun et al.[37] Jahnke et al.[38] Paetsch et al.[39]

20 21 18 18 18

Sommer et al.[40]

88% 92% 90% (proximal) 74% 88% (good quality) 85%–92% 75% 91% 76% 86% 92% 81% 82% 79% 83% 86% 82%

TFE TFE TFE SSFP SSFP TFE SSFP SSFP SSFP SSFP Contrast TFE

72% 55% 92% (proximal) 63% 91% (good quality) 50%-83% 85% 57% 85% 98% 67% 82% 82% 91% 77% 95% 88%

Prospective Navigator with Real-Time Correction Whole Heart Steady-State Free Precession Coronary Artery (CMR) Sakuma et al.[43] Jahnke et al.[44] Sakuma et al.[42] Kato et al.[45] Nagata et al.[46] Klein et al.[47] Pouleur et al.[48]

39 55 113 138 62 46 77

3 Tesla Whole Heart Coronary Artery CMR 18 Sommer et al.[41] 69 Yang et al.[49]

SSFP SSFP SSFP Multicenter SSFP 32-channel SSFP SSFP SSFP

82% 78% 82% 87% 83% 91% 100%

91% 91% 90% 71% 93% 54% 72%

TFE Contrast IR-GRE

82% 94%

89% 82%

IR-GRE, inversion recovery gradient echo; SSFP, steady-state free precession; TFE, turbo field echo. *Based on 60% of patients with good free breathing coronary magnetic resonance imaging images.

value of coronary artery CMR for the identification of coronary disease (>50% diameter stenosis by quantitative coronary angiography; Table 22-4).23 The sensitivity and negative predictive value were particularly high for the identification of left main or multivessel disease. Accordingly, coronary CMR was especially valuable for patients with dilated cardiomyopathy in the absence of clinical myocardial infarction. In our experience, coronary artery CMR is highly accurate and superior to late gadolinium enhancement methods for determining the etiology (ischemic vs. nonischemic) of cardiomyopathy.24 Data are increasingly available on whole heart SSFP coronary artery CMR methods. With an inferior in-plane spatial resolution, the clinical results appear to be at least as accurate as those obtained with targeted acquisition free breathing methods (see Table 22-3).42–48 A comparative study of gradient echo and SSFP coronary artery CMR by Ozgun and colleagues suggested that gradient echo approaches had higher sensitivity, whereas SSFP approaches had better specificity.37 A large number of these studies have been led by Sakuma and colleagues, including a preliminary report using a 32-channel receiver coil46 and a multicenter trial in Japan using this method.45 As expected, multicenter coronary artery CMR data are not as

favorable as those reported in single centers, but multicenter coronary artery CMR and multicenter coronary MDCT data appear similar (discussed later). Available data suggest that 3D gradient echo coronary artery CMR findings at 1.5 T and 3.0 T are similar.40

TABLE 22-4 Free Breathing Three-Dimensional Navigator Coronary Artery Cardiovascular Magnetic Resonance: Multicenter Trial Results Any Coronary Artery Disease Sensitivity Specificity Prevalence Positive predictive value Negative predictive value

Left Main/ MVD

93% 42% 59% 70%

100% 85% 15% 54%

81%

100%

MVD, multi-vessel disease. (Adapted from Kim WY, Danias PG, Stuber M, et al. Coronary magnetic resonance angiography for the detection of coronary stenoses. N Engl J Med. 2001;345:1863–1869.)

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TABLE 22-3 Free Breathing Three-Dimensional Gradient Echo Coronary Magnetic Resonance (CMR) Imaging Using Prospective Navigators for Identification of Focal Coronary Stenoses 50% or More in Diameter

ISCHEMIC HEART DISEASE

A

B

D

C

E

Figure 22-7 Tesla coronary artery cardiovascular magnetic resonance (CMR) image in a 50-year-old man. A and B are reformatted CMR images; C is a volume rendered CMR; and D and E are the corresponding angiograms. Courtesy of Dr. Yang.

Few studies have examined the role of contrast-enhanced coronary artery CMR for the detection of disease. Paetsch and associates39 compared noncontrast 3D targeted coronary artery CMR with an investigational intravascular agent using an inversion recovery prepulse. Superior specificity was found with the addition of contrast. Preliminary data from Yang and coworkers using gadobenate dimeglumine (GdBOPTA, Multihance, Bracco, Milan, Italy) with a 32-channel receiver coil at 3 T appear very promising for contrast coronary artery CMR at 3 T (Fig. 22-7).49

COMPARISON OF CORONARY ARTERY CARDIOVASCULAR MAGNETIC RESONANCE WITH MULTIDETECTOR COMPUTED TOMOGRAPHY Several studies have directly compared MDCT with coronary artery CMR for the detection of coronary artery disease (Table 22-5). Overall, these studies have suggested near equivalence of the current methods when interpretable segments were considered and superiority of 64-slice MDCT for patients with lower calcium scores with “intention to diagnose” analyses. The first direct study by Gerber and colleagues50 compared free breathing targeted 3D coronary artery CMR with four-slice MDCT and showed slight superiority of coronary artery CMR for overall accuracy. A follow-up study of 52 patients performed by the same group with 16-slice MDCT showed equivalence with visual analysis.51 Another 16-slice MDCT comparative study by Dewey and colleagues 304 Cardiovascular Magnetic Resonance

showed superiority of MDCT52 for sensitivity, but the coronary artery CMR technique used free breathing or an inferior multiple-breath-hold approach.36 Interestingly, coronary artery CMR sensitivity (74%) was improved compared with an earlier report from this group,36 suggesting a learning curve for coronary artery CMR. Patients expressed a preference for MDCT,52 with the advantage of very rapid and simplified protocols at the expense of substantial radiation exposure and need for iodinated contrast. A third comparative coronary artery CMR and a 16-slice MDCT comparative study by Maintz and associates53 showed superiority of coronary artery MDCT image quality and disease analysis on a coronary segment basis, but showed equivalence when only evaluable segments were included. Finally, a comparison of whole heart coronary artery CMR with 40- or 64-slice MDCT by Pouleur and coworkers48 showed very high sensitivity of coronary artery CMR, but overall superiority of coronary artery MDCT. This is likely related to their “intention to diagnose” analysis and the classification of nonevaluable segments as “diseased” in a population with a low (22%) prevalence of coronary artery disease. Nonevaluable segments were more common on CMR. Epicardial calcium is associated with aging and with the development of coronary atherosclerosis. It is a marker for increased risk of adverse events associated with coronary artery disease.54 The presence of epicardial calcium is a well-recognized limitation of coronary artery MDCT, with an exaggerated appearance on imaging (frequently referred to as blooming) that interferes with the accurate determination of coronary artery stenosis. Several trials using MDCT have excluded patients with substantial epicardial calcium55 or have shown marked reduction in specificity for patients with Agatston calcium scores of greater than 400.56,57 A comparative study of coronary artery CMR and

Investigator Gerber et al.[50] Kefer et al.[51] Dewey et al.[52] Maintz et al.[53] Pouleur et al.[48] Liu et al.[58]

Cardiovascular Magnetic Resonance/CCT Method

Sensitivity

Specificity

PPV

NPV

Accuracy

CMR:FB-SSFP 4-slice MDCT CMR:FB-SSFP 16-slice MDCT CMR-BH/FB-SSFP SSFP 16-slice MDCT CMR-FB-SSFP 16-slice MDCT CMR-FB-SSFP 40/64 MDCT CMR-FB-SSFP

62% 79%{ 75% 82% 74%{ 92% 82% 84% 100% 94% 81% 75%

84%* 71% 77% 79% 75% 79% 88% 95% 72%{ 88% 75%{ 48%

49% 40% 42% 46% 95% 95% 68% 77% 50%{ 70% — —

90% 93% 93% 95% 84% 90% 94% 95% 100% 98% — —

80%{ 73% 77% 80% — — 87% 93% 78%{ 90% — —

BH, breath hold; CMR, cardiovascular magnetic resonance. FB, free breathing; MDCT, multidetector computed tomography; NPV, negative predictive value; PPV, positive predictive value; SSFP, steady-state free precession. *p < 0.001. { p < 0.05. **Reported values are from patient-based analyses for the detection of coronary artery disease.

LAD LAD LAD AO AO

A

B

C

Figure 22-8 A, Extensive epicardial coronary calcium of the left anterior descending (LAD) coronary artery as visualized by multidetector computed tomography (MDCT). B, Coronary artery cardiovascular magnetic resonance. C, X-ray coronary angiography. The proximal calcium deposit (arrowhead) obscures the lumen on MDCT, but is seen on cardiovascular magnetic resonance as patent. The more distal calcium deposit (arrow) is identified as a stenosis on cardiovascular magnetic resonance and the corresponding X-ray angiogram. AO, aorta. Adapted from Liu X, Zhao X, Huang J, et al. Comparison of 3D free-breathing coronary MR angiography and 64-MDCT angiography for detection of coronary stenosis in patients with high calcium scores. Am J Roentgenol. 2007;189:1326–1332.

64-slice MDCT in patients with high calcium scores showed superiority of coronary artery CMR57 (Fig. 22-8).

CORONARY ARTERY CARDIOVASCULAR MAGNETIC RESONANCE FOR CORONARY ARTERY BYPASS GRAFT ASSESSMENT Assessment of coronary artery bypass graft patency is a common clinical issue.59,60 Compared with the native coronary arteries, reverse saphenous vein and internal mammary artery grafts are relatively easy to image because of

their relatively minimal motion during the cardiac and respiratory cycle and the larger lumen of reverse saphenous vein grafts. Furthermore, their predictable and less convoluted course has allowed imaging of bypass grafts, even with less sophisticated CMR methods. With schematic knowledge of the origin and touchdown site of each graft, conventional free breathing electrocardigraphic (ECG)-triggered 2D spin echo and 2D gradient echo coronary artery CMR in the transverse plane have both been used to reliably assess bypass graft patency (Fig. 22-9 and Table 22-6).61–70 Patency is determined by visualizing a patent graft lumen in at least two contiguous transverse images along its expected course (presenting as a signal void for spin echo techniques and bright signal for gradient echo approaches). If signal consistent with flow is identified in the area of the graft lumen, it is very likely to be patent. If a patent lumen is Cardiovascular Magnetic Resonance 305

22 CORONARY ARTERY IMAGING: CLINICAL RESULTS

TABLE 22-5 Comparative Studies of Coronary Artery Cardiovascular Magnetic Resonance and MDCT**

ISCHEMIC HEART DISEASE

Figure 22-9 A and B, Transverse fast spin echo images in a patient with previous coronary artery bypass grafting. A, The proximal graft is identified (arrow), with subsequent inferior sections (B and C) showing a patent graft (arrows).

TABLE 22-6 Sensitivity, Specificity, and Accuracy of Coronary Artery Cardiovascular Magnetic Resonance for Assessment of Coronary Artery Bypass Graft Patency Investigator

Technique

No. of Grafts

White et al.[61] Rubinstein et al.[66] Jenkins et al.[67] Galjee et al.[63] White et al.[62] Aurigemma et al.[64] Galjee et al.[63]

2D 2D 2D 2D 2D 2D 2D

72 47 41 98 28 45 98

Molinari et al.[68] Engelmann et al.[65]

3D GRE CE-3D GRE

Wintersperger et al.[69] Vrachliotis et al.[70]

CE-3D GRE CE-3D GRE

Spin Spin Spin Spin GRE GRE GRE

echo echo echo echo

51 96 SVG 37 IMA 39 45

Patency 69% 62% 63% 74% 50% 73% 74% 66% (SVG) 76.5% 66% 100% 87% 67%

Sensitivity 86% 90% 89% 98% 93% 88% 98% 92% 91% 92% 100% 97% 93%

Specificity 59% 72% 73% 85% 86% 100% 88% 85% 97% 85% 100% 100% 97%

Accuracy 78% 83% 83% 89% 89% 91% 96% 89% 96% 89% 97% 95%

2D, two-dimensional; 3D, three-dimensional; CE, contrast-enhanced; GRE, gradient recalled echo; IMA, internal mammary artery graft; SVG, saphenous vein graft.

seen at only one level (e.g., for spin echo techniques, a signal void is seen at only one level), a graft is considered indeterminate. If a patent graft lumen is not seen at any level, the graft is very likely occluded. Combining spin echo and gradient echo imaging in the same patient does not appear to improve accuracy.63 3D noncontrast68 and contrast-enhanced coronary artery CMR has also been described for the assessment of graft patency,69,70 with slightly improved results (see Table 22-6). The accuracy of ECG-triggered SSFP sequences appears to be similar to that of spin echo and gradient echo approaches.71 Limitations of coronary artery CMR bypass graft assessment include difficulties related to local signal loss and artifact as a result of implanted metallic objects (hemostatic clips, ostial stainless steel graft markers, sternal wires, coexistent prosthetic valves and supporting struts or rings, and graft stents; Fig. 22-10). The inability to identify severely diseased yet patent grafts is also a hindrance to clinical utility and acceptance. Langerak and colleagues72 found free breathing navigator 3D gradient echo coronary artery CMR to be accurate for the assessment of saphenous vein graft stenoses (Fig. 22-11), with very good agreement with quantitative X-ray angiography for the assessment of both graft occlusion 306 Cardiovascular Magnetic Resonance

and graft stenoses (Table 22-7). This group has also advocated assessment of rest and adenosine stress coronary artery flow assessment using the phase velocity CMR technique,73,74 suggesting superior results for flow assessment.

CONCLUSION Over the last 15 years, coronary artery CMR has been transformed from a scientific curiosity to a clinically useful imaging tool in selected populations. Uses include the identification and characterization of anomalous coronary arteries, characterization of aneurysms, and assessment of coronary artery bypass graft patency. Coronary artery CMR also appears to be of clinical value for the assessment of native vessel integrity in selected patients, especially those with suspected left main or multivessel disease. Normal findings on coronary artery CMR strongly suggest the absence of left main or multivessel disease. Overall, data suggest that successful coronary artery imaging is accomplished more often with MDCT, but when considering well-visualized arteries, accuracy is similar for CMR and MDCT. CMR is superior in

Ao

A

B

C

Figure 22-10 A, Posterior-anterior chest radiograph in a patient with coronary artery bypass grafts. Note the sternal wires (thick arrow) and the coronary artery bypass markers (thin arrow). B, Transverse coronary cardiovascular magnetic resonance in the same patient. Note the large local artifacts (signal voids) related to sternal wires (thick arrow) and bypass graft markers (thin arrow). C, Barium and tantalum markers (arrow) result in the smallest artifacts. The size of the artifacts is reduced somewhat with spin echo and black-blood cardiovascular magnetic resonance.

Figure 22-11 Oblique breath hold two-dimensional ECGtriggered coronary artery cardiovascular magnetic resonance of a patent saphenous vein bypass graft (SVG). Two adjacent images show the SVG (arrows) extending from (A) the aortic origin (Ao) to the (B) distal touchdown on the posterior descending coronary artery. LV, left ventricle; RV, right ventricle.

SVG Ao

LV RV

SVG

A

B

TABLE 22-7 Diagnostic Accuracy of Submillimeter Coronary Artery Cardiovascular Magnetic Resonance for Saphenous Vein Graft Disease Graft occlusion Graft stenosis 50% Graft stenosis 70%

Sensitivity

Specificity

83% (36%–100%) 82% (57%–96%) 73% (39%–94%)

100% (92%–100%) 88% (72%–97%) 80% (64%–91%)

Adapted from Langerak SE, Vliegan HW, de Roos A, et al. Defection of vein graft disease using high resolution magnetic resonance angiography. Circulation. 2002;105:328-333.

patients with prominent epicardial calcium and has the advantage of avoiding ionizing radiation and iodinated contrast. Technical and methodologic advances in motion suppression, along with increasing experience and

development of 3 T platforms, multiple coil arrays, and intravascular contrast agents, will no doubt facilitate improved accuracy and increased clinical use. As with MDCT, data from multicenter trials will be needed to define the clinical role of coronary artery CMR. Cardiovascular Magnetic Resonance 307

22 CORONARY ARTERY IMAGING: CLINICAL RESULTS

SVG marker

ISCHEMIC HEART DISEASE

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59. Shroyer AL, Grover FL, Hattler B, et al. On-pump versus off-pump coronary-artery bypass surgery. N Engl J Med. 2009;361:1827–1837. 60. Lopez RD, Hafley GE, Allen KB, et al. Endoscopic versus open veingraft harvesting in coronary-artery bypass surgery. N Engl J Med. 2009;361:235–244. 61. White RD, Caputo GR, Mark AS, et al. Coronary artery bypass graft patency: noninvasive evaluation with MR imaging. Radiology. 1987;164(3):681–686. 62. White RD, Pflugfelder PW, Lipton MJ, Higgins CB. Coronary artery bypass grafts: evaluation of patency with cine MR imaging. AJR Am J Roentgenol. 1988;150(6):1271–1274. 63. Galjee MA, van Rossum AC, Doesburg T, et al. Value of magnetic resonance imaging in assessing patency and function of coronary artery bypass grafts: an angiographically controlled study. Circulation. 1996;93(4):660–666. 64. Aurigemma GP, Reichek N, Axel L, et al. Noninvasive determination of coronary artery bypass graft patency by cine magnetic resonance imaging. Circulation. 1989;80(6):1595–1602. 65. Engelmann MG, Knez A, von Smekal A, et al. Non-invasive coronary bypass graft imaging after multivessel revascularisation. Int J Cardiol. 2000;76(1):65–74. 66. Rubinstein RI, Askenase AD, Thickman D, et al. Magnetic resonance imaging to evaluate patency of aortocoronary bypass grafts. Circulation. 1987;76(4):786–791. 67. Jenkins JP, Love HG, Foster CJ, et al. Detection of coronary artery bypass graft patency as assessed by magnetic resonance imaging. Br J Radiol. 1988;61(721):2–4. 68. Molinari G, Sardanelli F, Zandrino F, et al. Value of navigator echo magnetic resonance angiography in detecting occlusion/patency of arterial and venous, single and sequential coronary bypass grafts. Int J Card Imaging. 2000;16(3):149–160. 69. Wintersperger BJ, Engelmann MG, Von Smekal A, et al. Patency of coronary bypass grafts: assessment with breath-hold contrastenhanced MR angiography–value of a non-electrocardiographically triggered technique. Radiology. 1998;208(2):345–351. 70. Vrachliotis TG, Bis KG, Aliabadi D, et al. Contrast-enhanced breathhold MR angiography for evaluating patency of coronary artery bypass grafts. AJR Am J Roentgenol. 1997;168(4):1073–1080. 71. Bunce NH, Lorenz CH, John AS, et al. Coronary artery bypass graft patency: assessment with true ast imaging with steady-state precession versus gadolinium-enhanced MR angiography. Radiology. 2003;227 (2):440–446. 72. Langerak SE, Vliegan HW, de Roos A, et al. Detection of vein graft disease using high resolution magnetic resonance angiography. Circulation. 2002;105(3):328–333. 73. Langerak SE, Kunz P, Vliegen HW, et al. MR flow mapping in coronary artery bypass grafts: a validation study with Doppler flow measurements. Radiology. 2002;222(1):127–135. 74. Langerak SE, Vliegen HW, Jukema JW, et al. Value of magnetic resonance imaging for the noninvasive detection of stenosis in coronary artery bypass grafts and recipient coronary arteries. Circulation. 2003;107(11):1502–1508.

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CHAPTER 23

Coronary Artery and Sinus Velocity and Flow Jennifer Keegan and Dudley J. Pennell

A coronary stenosis may be observed during cardiovascular magnetic resonance (CMR) imaging as an area of signal loss caused by turbulent flow, and although both the degree and the extent of signal loss are indicative of the severity of the stenosis, accurate quantification is not possible.1 However, both phasic coronary artery blood flow and flow velocity may be affected by the presence of stenosis, and the ratio of coronary artery flow under maximal vasodilation to coronary artery flow at rest (coronary flow reserve) is a good indicator of the physiologic significance to the myocardium. This makes the measurement of coronary flow and velocity valuable. Measurements of instantaneous and mean coronary flow parameters are most commonly made using an intracoronary Doppler flow wire, positioned in the arterial lumen during X-ray contrast angiography. However, this is an invasive procedure, with a small but definite risk of complications, and the radiation dose to the patient is relatively high for a diagnostic test. The repeated use of such a technique to monitor disease progression or regression in response to drug therapy or lifestyle changes is therefore not acceptable. In addition, the presence of the Doppler wire may itself affect flow,2 as may the stress to the patient resulting from the invasive nature of the procedure. Other techniques capable of assessing mean flow parameters include continuous thermodilution and positron emission tomography (PET), the former also requiring placement under X-ray contrast angiography, with the concomitant disadvantages, and the latter being relatively unavailable and expensive. CMR has the ability to quantify blood flow noninvasively and has the potential to be a useful noninvasive alternative to the intracoronary technique. Since the 1980s, it has been used extensively for the assessment of phasic blood flow velocity and flow in a wide range of cardiovascular applications,3,4 and it has the potential to be implemented in coronary flow studies. One of the main problems relating specifically to applying the technique to the coronary arteries is their small size (typically < 5 mm in diameter), which, for current levels of in-plane resolution, results in only a few pixels across the vessel. This has implications for the accurate measurement of both vessel cross-sectional area and blood flow velocity. Vessel tortuosity is a further problem, giving rise to difficulties in accurately aligning the vessel so that flow is truly through-plane or in-plane, as required. This is exacerbated by the movement of the arteries with the cardiac and respiratory cycles. In addition, the temporal resolution of the velocity encoding sequence must be sufficiently good to 310 Cardiovascular Magnetic Resonance

minimize the blurring of the vessel as a result of motion within the period of acquisition and to resolve the phasic velocity profile. The low peak flow velocities in normal arteries at rest (typically < 25 cm/sec) present a further challenge and require highly sensitive velocity windows, whereas in the presence of stenoses, high velocities are present, together with complex flow, which may lead to signal loss. The combination of these problems is formidable, and it was only in the last 10 to 15 years that CMR techniques started to generate results. This chapter reviews the progress made to date and discusses potential future improvements.

INDIRECT ASSESSMENT OF TOTAL CORONARY FLOW AND FLOW RESERVE Two indirect approaches for assessing total coronary flow and flow reserve by CMR have been reported, the first from velocity mapping of cardiac venous outflow and the second from velocity mapping in the aortic root.

Coronary Sinus Flow Velocity mapping of coronary venous outflow is less problematic than velocity mapping in the coronary arteries because the coronary sinus has a much larger diameter (typically, 7 to 10 mm) and also because effects of signal loss are unlikely because flow is less susceptible to turbulence. In the human heart, coronary sinus flow almost equals total coronary flow because approximately 96% of left ventricular (LV) venous blood flow drains into the right atrium via the coronary sinus.5 van Rossum and colleagues were the first to show that blood flow in the coronary sinus could be measured using CMR.6 This feasibility study assessed the ability of cine CMR velocity mapping to measure phasic and mean coronary venous outflow in the distal coronary sinus of 24 healthy subjects. The flow profiles were generally biphasic and primarily diastolic, with 37% of subjects showing some reverse flow immediately after the R-wave. The mean volumetric flow over the cardiac cycle was 144 62 mL/min, similar to values reported in normal subjects using continuous thermodilution (122 mL/min).7 Although phasic blood flow in the sinus may have been expected to be predominantly systolic, as

One issue with the breath hold approach to measuring coronary sinus flow is that breath holding changes intrathoracic pressure and affects cardiac output and venous blood flow.16 Schwitter and associates17 performed a validation of non-breath-hold velocity mapping against PET. Although taking longer to acquire (typically, 4 minutes), the nonbreath-hold technique has the advantages of higher spatial resolution (0.8 0.8 mm) and improved signal-to-noise ratio (SNR) through the acquisition of multiple averages. Furthermore, the high temporal resolution that is achievable is better suited for resolving the highly pulsatile flow profile and for minimizing the blurring effects of the extensive inplane motion of the coronary sinus during the cardiac cycle. Unlike the initial study of van Rossum and coworkers,6 this free breathing study used retrospective electrocardiographic ECG gating,18 which enables data acquisition throughout the entire cardiac cycle, and respiratory ordered phase encoding,19 which reduces the effects of respiratory motion. From a basal short axis localizer image showing the length of the coronary sinus, a plane was defined that transected the center of both the left atrium and the aortic valve, thereby cutting the coronary sinus perpendicularly, approximately 2 cm proximal to its entrance to the right atrium. Scans were performed both before and after the administration of 0.56 mg/kg dipyridamole in 16 healthy subjects and 10 orthotopic heart transplant recipients, and the results were compared with PET data. Representative pre-dipyridamole magnitude and phase images in a healthy subject are shown in Figure 23-1A, with the flow curves for this subject (preand post-dipyridamole) shown in Figure 23-1B. The mean difference between coronary flow reserve measured by PET and CMR was 2.2%, with limits of agreement of 27.2% and 31.6%. Correlation between CMR-measured myocardial blood flow (coronary sinus flow divided by myocardial mass) and PET was good (r ¼ 0.93), although the slope was considerably less than unity (0.73). This underestimation of blood flow measured by CMR results from the fact that a variable part of the inferior and inferior-septal myocardium is drained by the middle cardiac vein, which either enters the coronary sinus just before its orifice or empties directly into the right atrium.20 If coronary sinus flow is divided instead by the mass of drained myocardium (as measured from a stack of short axis images), the correlation remains good (r ¼ 0.95), whereas the slope approaches unity (1.05). Figure 23-2 shows a Bland-Altman plot of myocardial blood flow (sinus blood flow per unit mass of drained myocardium) measured by CMR and PET, with baseline flows normalized for rate-pressure product. The mean differences between PET and CMR were 3.4 12.8% at resting baseline and 3.9 20.2% during hyperemia and were not significant. The authors have subsequently used this technique to show that administration of 17b-estradiol over 3 months without progestin coadministration does not improve coronary flow reserve in postmenopausal women.21 A validation of non-breath-hold CMR-measured coronary sinus flow has also been performed against flow probes in dogs by Lund and colleagues22 In this study, the correlation between coronary blood flow (measured as the sum of left anterior descending [LAD] and circumflex [LCX] coronary flow) with flow probes against coronary sinus flow measured with CMR was 0.98, with a nonsignificant mean difference of 3.1 8.5 mL/min. Coronary sinus blood flow per unit Cardiovascular Magnetic Resonance 311

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in other venous structures, the authors suggested that the thin, compliant walls of the sinus, together with its drainage into the right atrium, where pressure varies considerably, may be responsible for the predominantly diastolic flow profile. These findings have been supported by other independent studies, both directly, using an ultrasonic transit time technique to measure phasic volumetric coronary sinus flow in conscious dogs,8 and indirectly, by observing areas of signal void caused by accelerating and turbulent flow near the entrance of the coronary sinus in the right atrium in early diastole in conventional cine CMR in healthy subjects.9 The imaging time for the feasibility study was typically 4 minutes, and this may be inappropriate when assessing sinus flow under pharmacologically induced maximal vasodilation. Kawada and associates10 assessed the possibility of using a segmented k-space11 gradient echo CMR approach to the acquisition so that all data could be acquired in less than 25 seconds. This had the further advantage that the full acquisition could be performed in a single breath hold, thereby eliminating the effects of respiratory motion. In a study of 9 healthy subjects and 29 patients with hypertrophic cardiomyopathy (HCM), data were acquired from an oblique coronal plane. Four reference and four velocity encoded views were acquired per data segment, giving a segment duration of 120 msec, but the temporal resolution was effectively improved by view sharing,12 a technique whereby data are generated at intermediate time points from the preceding and following data segments. Hence, depending on the R-R interval, velocity maps were obtained at up to 14 phases in the cardiac cycle. The authors observed the same biphasic velocity and flow profiles6 in both healthy subjects and patients with HCM, with no significant difference noted between the baseline myocardial blood flow (coronary blood flow per unit mass of myocardium) in the two groups (0.74 mL/min/g vs. 0.62 mL/min/g). However, after intravenous administration of 0.56 mg/kg dipyridamole, the increase in myocardial blood flow in patients with HCM (0.62 mL/min/g to 1.03 mL/min/g) was less than that in the healthy subjects (0.74 to 2.14 mL/min/g), resulting in significantly different coronary flow reserves for the two groups (1.72 0.49 vs. 3.01 0.75, respectively, p < 0.01). The authors concluded that, in healthy subjects, myocardial blood flow and coronary flow reserve, as measured by breath hold phase velocity mapping, were similar to those found by other techniques; in addition, the CMR technique was able to distinguish between healthy subjects and patients with HCM. The breath hold approach has also been used by Kennedy and coworkers13 to show significant differences in the coronary flow reserve of healthy subjects (4.59 0.58) and heart transplant patients with mild (2.15 0.44, p < 0.05) and severe (2.21 0.59, p < 0.05) coronary disease, as determined by post-transplant coronary angiography. A validation of the breath hold approach was reported by Koskenvuo and colleagues, who compared myocardial blood flow measured by CMR with that measured by PET in both healthy subjects14 and patients with coronary artery disease (CAD).15 Good correlations were reported for myocardial blood flow in both subject groups (0.82 and 0.80, respectively), whereas for coronary flow reserve, the correlations were 0.76 and 0.50, respectively.

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Figure 23-1 A, Representative example of four magnitude images (a through d) and corresponding phase CMR images (e through h) in a healthy subject at rest. Arrows show the coronary sinus; the time after the R-wave is shown at the top left corners. At the end of left ventricular ejection (b and f), the cross-sectional area of the coronary sinus is largest (b) and the flow is retrograde (f). During atrial contraction, reflux into the coronary sinus occurs (d and h). B, Biphasic flow data with high flow in early systole and early diastole at baseline (gray line), whereas during hyperemia, the highest flow occurs during systole (black line). (From Schwitter J, DeMarco T, Kneifel S, et al. Magnetic resonance-based assessment of global coronary flow and flow reserve and its relation to left ventricular functional parameters. Circulation. 2000;101:2696–2709.)

mass of myocardium was 0.40 0.09 mL/min/g (CMR) compared with 0.44 0.08 mL/min/g (flow probes) and also was not significant. The authors went on to study patients with chronic heart failure and showed that coronary flow reserve was significantly reduced compared with that in healthy subjects (2.3 0.9 vs. 4.2 1.5, p ¼ 0.01).23 This has been recently confirmed by Aras and associates24 (coronary flow reserve 1.45 vs. 3.00 in patients with chronic heart failure and healthy subjects, respectively, p < .001). Similarly, Watzinger and coworkers25 have shown significantly reduced flow reserve in patients with idiopathic 312 Cardiovascular Magnetic Resonance

cardiomyopathy compared with healthy subjects (2.19 0.77 vs. 3.51 1.29, p < .05). In all three studies, there were no significant differences in baseline flow between healthy subjects and those with disease (0.52 mL/min/g vs. 0.46 mL/min/, p ¼ not significant [NS]; 0.83 0.26 mL/ min/g vs. 0.85 0.30 mL/min/g, p ¼ NS; and 0.55 0.19 mL/min/g vs. 0.48 0.07 mL/min/g, p ¼ NS). Data for subjects with idiopathic cardiomyopathy are shown in Figure 23-3. These studies provide useful results, but have a number of important limitations. Cardiac and respiratory motion

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Figure 23-3 Box plot of myocardial blood flow (MBF) reserves in healthy subjects and patients with idiopathic cardiomyopathy (IDC). Each box encloses 50% of the data, with the median value shown as a line. The error bars extending from the top to the bottom of each box mark the minimum and maximum values. MBF reserve is reduced in patients with IDC compared with healthy subjects (p < 0.05). (From Watzinger N, Lund GK, Saeed M et al. Myocardial blood flow in patients with dilated cardiomyopathy: quantitative assessment with velocity-encoded cine magnetic resonance imaging of the coronary sinus. J Magn Reson Imaging. 2005;21:347– 353.)

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Figure 23-2 A, Correlation between estimate of myocardial blood flow (MBF; mL/min/g) derived from positron emission tomography (PET; X-axis) and CMR (Y-axis) measurements. For coronary sinus flow divided by drained myocardium, correlation with PET data is high and the slope of the regression line approximates unity (dashed line). B, Comparison of myocardial blood flow as measured by PET and cardiovascular magnetic resonance (CMR; coronary sinus flow divided by drained myocardium mass) with baseline flow normalized for ratepressure product (Bland-Altman analysis). Mean differences between PET and CMR measurements during baseline (3.4 12.8%) and hyperemia (3.9 20.2%) were not significantly different. CS, coronary sinus; LVM, left ventricular mass; SD, standard deviation. (From Schwitter J, DeMarco T, Kneifel S et al. Magnetic resonance-based assessment of global coronary flow and flow reserve and its relation to left ventricular functional parameters. Circulation. 2000;101:2696– 2709.)

results in blurring of the sinus, and the small number of pixels covering the sinus (typically five across the diameter in diastole) results in considerable partial volume averaging in edge pixels, which is problematic for the accurate assessment of sinus cross-sectional area and for the determination of mean sinus flow velocity. As discussed, drainage of a variable part of the inferior and inferior-septal myocardium by the middle cardiac vein20 results in underestimates of myocardial blood flow measured, and although this estimate is an indicator of total coronary blood flow, no regional assessment of either flow or flow reserve is possible with this technique.

Total Coronary Flow Reserve from Measurements in the Aortic Root In 1993, it was proposed that coronary flow reserve might be derived from flow measurements made in the ascending aorta, which is less susceptible to cardiac and respiratory motion and partial volume effects than the coronary arteries.26 Coronary diastolic flow, which represents the bulk of coronary flow, can be estimated as the retrograde flow in the ascending aorta during systole and diastole minus the antegrade flow during diastole. A variable velocity encoding window was implemented to maintain the accuracy of velocity measurements during periods of both high flow in systole (window ¼ 200 cm/sec) and low flow in diastole (window ¼ 30 cm/sec).27 Although it was suggested that the assessment of absolute diastolic coronary flow with this technique is inaccurate because of known errors associated with the velocity mapping technique, it was argued that these errors should be the same both pre- and post-vasodilation and therefore should be eliminated from the assessment of diastolic coronary flow reserve, defined in this instance as the difference (rather than the ratio) between the two measurements. In seven patients with abnormal findings on myocardial perfusion scintigraphy, the diastolic coronary flow reserve was -50 76 mL/min compared with 260 66 mL/min in eight healthy subjects. The principle of this technique was later refined by taking into account the motion of the coronary arteries during the cardiac cycle.28 A model was developed describing the flow through five transverse parallel aortic slices extending from the base of the aortic valve to above the level of the coronary ostia. This was then solved mathematically, and in five healthy subjects, it was shown Cardiovascular Magnetic Resonance 313

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that the standard error in the measurement of total coronary artery flow was approximately 90 mL/min, or 30% of total coronary artery flow. The error in coronary flow reserve is higher because the errors in baseline and maximal vasodilation flow are additive. In addition, the assumption that flow is predominantly diastolic, although true in healthy subjects, may not hold in the presence of disease, introducing further unknown errors into the technique. A more recent attempt at measuring coronary artery blood flow in this way likewise found that the technique was compromised by poor reproducibility, although significant changes in coronary blood flow with hormone replacement therapy were observed.29

Direct Assessment of Coronary Artery Velocity The problems associated with the assessment of flow velocity and coronary artery flow were discussed earlier and preclude the application of standard CMR techniques. As for coronary imaging, the major breakthrough for coronary flow techniques was the shortening of sequence duration to the extent that data could be acquired over a single breath hold, effectively freezing respiratory motion and eliminating the resulting artifact. Further advances with navigator echo monitoring of the diaphragm position have allowed data to be acquired over multiple reproducible breath holds or during free breathing, enabling increased spatial and temporal resolution and data averaging (resulting in higher SNRs). The following sections describe the approaches used to assess coronary flow velocity and flow.

Bolus Tagging The first report detailing the imaging of coronary artery flow was made in 1991 in rat and mice hearts with bolus tracking.30 With this technique, based on one previously implemented in vivo in the aorta and the carotid arteries,31 a section of blood above the coronary ostia on the aortic root is tagged by the application of a slice-selective presaturation radiofrequency pulse, and imaging is performed after a delay. During this delay period, the tagged volume of blood washes into the coronary artery tree, where it is seen as a signal void. The mean velocity of the tagged blood can be calculated from the degree of tag movement and the wash-in delay time. The technique has also been developed for multi-bolus tracking using stimulated echoes,32 and its application was demonstrated in a 3-mm-diameter tube with laminar flow and in an isovolumetric perfused rat heart. Using this approach, an image of multiple boluses (typically three), each with a different wash-in time, can be obtained simultaneously to show the coronary artery tree. Again, the extent of the arterial pathway seen depends on the flow velocities in the tagged volumes and on the wash-in delay times, with short delays required for visualization of the proximal portions and longer delays for the mid- and distal portions. By tracking multiple boluses simultaneously in this way, this technique essentially results in images of blood flow. 314 Cardiovascular Magnetic Resonance

Echo Planar Time-of-Flight Technique Echo planar techniques are an attractive option for coronary artery investigations because of their fast imaging times that reduce the effects of both cardiac and respiratory motion. In 1993, the first report detailing the use of an echo planar single-shot time-of-flight technique for assessing coronary artery flow velocity in 11 healthy subjects was reported.33 In this study, performed at 1.5 Tesla (T), fat-suppressed, 5to 10-mm-thick, short axis cardiac slices were acquired with an in-plane pixel size of 1.5 3 mm, each taking approximately 95 msec. Before the 90 slice selection radio frequency pulse, another 90 pulse was applied to saturate a thick band centered on the imaging slice. If increasing time delays are programmed between the saturation and slice select pulses, the signal intensity in the vessel changes as a result of blood wash-in through the slice. The rate at which it changes can be used to calculate the blood velocity at the time of imaging. An example of the images obtained at a single time point after the R-wave is shown in Figure 23-4. For this timing in the cardiac cycle, images (b) to (d) show that blood flow is too slow to wash in to the image slice when the saturation delays are less than 80 msec. However, as the delay is increased from 80 msec (e) to 170 msec (g), increasing amounts of blood wash-in occur, and at 200 msec (h), flow is fast enough for it to be complete. The images for each time point in the cardiac cycle were acquired within a single breath hold, with a number of breath holds required to generate the data for a velocity profile throughout the cardiac cycle. The average velocity profile for all 11 subjects is shown in Figure 23-5 and shows the expected peak flow in early diastole. As can be seen, it was not possible to measure velocity at less than 200 msec from the R-wave with this technique because this time period is required for the application of saturation delays. This was not considered a problem for normal subjects because left coronary flow is predominantly diastolic. A further restriction of the technique is that the long echo time of the sequence (echo time [TE] ¼ 28 msec) may lead to signal loss at sites of turbulent flow. At such sites, there may also be a breakdown in the assumption of laminar flow required for the velocity calculations from the wash-in data. The authors reported that in nine subjects who were imaged during continuous hand and lower extremity exercise, eight showed an increase in diastolic velocity (increase over exercise period ¼ 52 24%). Isometric exercise does not produce a maximal physiologic stress response, and more recently, echo planar timeof-flight coronary flow velocity reserve has been measured using this technique in healthy subjects (N ¼ 10) after the infusion of 0.56 mg/kg dipyridamole.34 In these subjects, peak diastolic velocity was observed to increase from 22 7 cm/sec to 90 40 cm/sec, resulting in a coronary flow velocity reserve of 3.9 1.5, with velocity returning to baseline (23 5 cm/sec) after the administration of aminophylline. In this study, the authors acquired low-resolution, singleshot coronary images rather than using a segmented approach to build up higher-resolution images over a number of cardiac cycles. This was prompted by their observation of considerable variability in the beat-to-beat position of the

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Figure 23-4 Echo planar time-of-flight images of a short axis slice acquired 350 msec after the R-wave with the left anterior descending artery perpendicular to the imaging plane (arrow), acquired in a single breath hold. The initial image (A) was obtained with no saturation pulse and acts as a reference. Images B to H were acquired at the same time after the R-wave, but with a thick-slab saturation prepulse at increasingly large saturation delays (SDs) before image slice selection pulse. The intensity change in the artery as a function of SD is used to calculate blood flow velocity, on the assumption that the velocity profile is laminar. Images A to H were all acquired within the same single breath hold. (From Poncelet BP, Weisskoff RM, Weeden VJ, Brady TJ, Kantor H. Time of flight quantification of coronary flow with echo-planar MRI. Magn Reson Med. 1993;30:447–457.)

LAD during both long and short breath hold periods. This usually corresponded to a downward drift in vessel position as the breath hold continued and could be as much as 6 mm (i.e., greater than the vessel diameter). It was unclear, however, how much of this movement was caused by poor breath holding and how much was the result of beat-to-beat variations in cardiac contraction. Regardless of the cause, it would introduce blurring, which is largely avoided in a single-shot image. Beat-to-beat variations in flow, as opposed to spatial position, still have an effect on measured velocity because individual images must be acquired with different saturation delays in consecutive cardiac cycles to generate the wash-in curve. Therefore, measured velocity is affected by changes over the breath hold period. The importance of this study lies in its pioneering and largely successful approach to a previously unsolved problem. It has largely been superseded by phase velocity mapping approaches, as discussed later, which are more robust and have higher availability.

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R-R interval Figure 23-5 Time-of-flight coronary flow velocity profile in 11 healthy subjects. For each subject and for each gating delay, a series of images with different saturation delays was acquired, as shown in Figure 23-4. (From Poncelet BP, Weisskoff RM, Weeden VJ, Brady TJ, Kantor H. Time of flight quantification of coronary flow with echo-planar MRI. Magn Reson Med. 1993;30:447–457.)

Gradient Echo Phase Velocity Mapping Breath Hold Techniques Using a velocity encoded segmented k-space gradient echo technique, velocity maps may be acquired in a single breath hold, thereby eliminating respiratory motion artifact. Cardiovascular Magnetic Resonance 315

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To accomplish this, the segment duration is typically on the order of 100 msec, and to minimize blurring as a result of cardiac motion, the acquisition is best performed in mid-diastole, when the heart is relatively stationary. In addition to limiting the temporal resolution of the sequence, the need to perform data acquisition within a single breath hold also limits the number of phase encoding steps that can be acquired, which in turn limits the spatial resolution and the SNR in the resulting images. The first diastolic through-plane coronary artery phase velocity maps were reported by Edelman and colleagues in 1993.35 These were acquired at 1.5 T using a fatsuppressed sequence consisting of a reference velocity compensated gradient waveform followed by a sensitized (velocity window ¼ 150 cm/sec) gradient waveform, each repeated four times in a cardiac cycle. TE was 8 msec and repetition time (TR) was 15 msec, resulting in a segment duration of 120 msec. The sequence was validated against a Doppler flow meter in vitro using both constant and mildly pulsatile flow in an 8-mm-diameter tube and in vivo against a standard nonsegmented velocity mapping sequence in the descending aorta of three healthy volunteers. Velocity maps were acquired over 24 cardiac cycles with in-plane pixel dimensions of 1.4 0.8 mm. In healthy subjects, mean flow velocities at rest in the mid-portion of the right and left anterior descending coronary arteries were 9.9 cm/sec and 20.5 cm/sec, respectively. These values are lower than those found in Doppler flow studies, a finding that is to be expected because the small number of pixels across the vessel results in partial volume averaging of the velocity profile. In four subjects who received intravenous administration of adenosine, velocities increased by at least a factor of four, suggesting that CMR has the potential to assess the hemodynamic significance of a stenosis by measuring the flow response to vasodilation. Another approach to quantifying the severity of stenosis is to perform in-plane coronary artery velocity mapping with a view to measuring increased velocity at the site of lumen narrowing. In-plane coronary artery velocity mapping was first performed in healthy volunteers by Keegan and associates,36 using a sequence with TE of 10 msec, TR of 20 msec, and segment duration of 160 msec, which effectively limited acquisitions to the period of relative cardiac diastasis in early diastole. The in-plane resolution was 1.6 0.8 mm, and data were acquired over breath holds of 24 to 32 cardiac cycles. The velocity sensitivity used was 50 cm/sec and was achieved by phase map subtraction of two images, one sensitized so that flow velocities of 100 cm/sec gave a phase shift of þ2p radians and the other sensitized so that flow velocities of 100 cm/sec gave a phase shift of 2p radians. This was shown to result in fewer blood flow artifacts than subtracting an image sensitized to flow velocities of 50 cm/ sec from a reference nonsensitized image. Through-plane and in-plane velocities measured with this technique were validated in vitro against a standard nonsegmented velocity mapping sequence in a 5.6-mm-diameter tube with pulsatile flow having a maximum velocity of 30 cm/sec and a maximum rate of change of velocity of 126 cm/sec2, comparable to the values expected in normal human coronary arteries at rest. Phantom work was also carried out to show the ability of the technique to measure a velocity increase at the sites of mild, moderate, and severe area-reducing 316 Cardiovascular Magnetic Resonance

stenoses and hence to quantify severity. However, although velocity increases at the sites of area-reducing stenoses have been observed with this approach,3 the tortuous pathways and small caliber of the coronary vessels result in partial-volume-type effects being more problematic for in-plane than for through-plane velocity mapping. Furthermore, in-plane studies require a high degree of reproducibility in the breath holding position, which is difficult to achieve without techniques such as navigator echo monitoring.

Navigator Techniques The studies discussed earlier used breath holding as a means of respiratory motion control, limiting the sequence parameters to allow the entire dataset to be acquired within the duration of a single breath hold and requiring a high degree of patient cooperation, which is not always forthcoming. In addition, inter- and intra- variations in the breath hold position may be problematic, and hemodynamic changes that occur secondary to breath holding, such as increases in intrathoracic pressure and heart rate, may themselves alter the blood flow being measured. A navigator echo approach under either prospective or retrospective control would enable data to be acquired during free breathing and would avoid these problems. Furthermore, temporal resolution of the sequence could be improved by reducing the segment duration, albeit at the expense of prolonged scan time. The influence of temporal resolution was studied by Hofman and colleagues,37 who compared the use of a segmented breath hold technique (segment duration of 126 msec) with a retrospective respiratory gated technique38 (reference and velocity sensitized view pair duration of 32 msec) for the assessment of flow velocity, vessel cross-sectional area, and volume flow in the right coronary arteries of six healthy subjects. Eight data averages were acquired, and the data were reconstructed with retrospective gating. The residual displacement of diaphragm positions in the reconstructed data was 3.9 mm. In-plane spatial resolution was 0.8 1.6 mm, and velocity sensitivity was 25 cm/sec. Vessel regions of interest were obtained semiautomatically from the magnitude images of the velocity compensated dataset by means of a seed growing algorithm and a 50% threshold. Velocity profiles generated from the retrospectively gated data for all six subjects (uncorrected for through-plane velocity of the vessel itself) are shown in Figure 23-6A and show a sharp peak in systole, a minimum at end-systole, and a second peak in early diastole. The in-plane displacements of the vessels are shown in Figure 23-6B and typically show peaks in systole and early diastole, with minimal displacement in mid- to end-diastole. The authors have since reported a similar pattern of movement for the LAD.39 During times of peak vessel displacement, motion blurring of the artery occurs and the breath hold segmented images are consequently of poorer quality than those acquired with the retrospective respiratory gated short segment duration sequence (Fig. 23-7B). In diastole, however, when in-plane displacement of the vessel is low, the two techniques generate images of comparable quality (see Fig. 23-7A). This results in the breath hold segmented k-space gradient echo sequence overestimating the instantaneous vessel crosssectional area by as much as a factor of four, with an

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coronary artery segments.41 Both breath holding and realtime slice-followed42 navigator echo controlled free breathing CMR techniques (temporal resolution of 140 msec and 46 msec, respectively) were performed within 24 hours of the invasive procedure and the maximal coronary flow velocities in a 2 2 mm pixel region of interest determined. Both CMR techniques were found to significantly underestimate flow velocity, although the correlations with invasive measurements were strong (r ¼ 0.70 and r ¼ 0.86, respectively). The navigator echo controlled technique was significantly more accurate than the breath hold technique (p < 0.02), although this was at the expense of prolonged acquisition time. The underestimation of blood flow velocity by both techniques is largely caused by partial volume averaging of the spatial flow profile in the CMR studies, together with insufficient sampling of the temporal flow profile. The decreased accuracy of the breath hold technique compared with the navigator echo free breathing technique is likely to be caused by a combination of the longer acquisition window of the former, together with hemodynamic changes resulting from the breath holding procedure itself. Figure 23-8 shows an example of the results obtained. Although still underestimating flow velocities, the significantly improved accuracy of the navigator echo free breathing technique compared with the breath hold technique provides a further step toward CMR assessment of coronary flow parameters, albeit at the expense of prolonged acquisition times.

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Figure 23-6 Through-plane cross-sectional averaged velocity (A) and in-plane displacement (B) in the right coronary artery as a function of time in the cardiac cycle and as measured on respiratory gated acquisition (oversampling factor ¼ 8) in six subjects. Different symbols represent different subjects. RA, right atrium; RCA, right coronary artery; RV, right ventricle. (From Hofman MBM, van Rossum AC, Sprenger M, Westerhof N. Assessment of flow in the right human coronary artery by magnetic resonance phase contrast velocity measurement: effects of cardiac and respiratory motion. Magn Reson Med. 1996;35:521–531.)

average increase in the time-averaged cross-sectional area of 90%. Based on this work, the authors suggest that, for acceptable levels of blurring throughout the cardiac cycle, the acquisition window for studies of the left and right coronary arteries (RCA) should be less than 90 msec and less than 25 msec, respectively. This is similar to the values derived by Marcus and associates,40 who suggested an acquisition window of less than 58 msec and less than 23 msec for the left and right arteries, respectively. A comparison between segmented k-space gradient echo phase velocity mapping and Doppler flow wire techniques has been performed in 26 angiographically normal

Interleaved spiral imaging is an alternative technique for generating high-resolution images of the coronary arteries,43 which, compared with those acquired using segmented gradient recalled echo (GRE), have higher SNRs and better temporal resolution.44 This potentially enables velocity mapping in both the left and the more mobile right coronary arteries.45 The results of breath hold spiral, free breathing spiral, and breath hold segmented GRE phase velocity mapping have been compared with those of free breathing GRE in healthy volunteers.46 Eight left and eight right coronary arteries were studied and flow velocity profiles generated with each technique, with the free breathing GRE acquisitions being used as a gold standard. Figure 239A and B shows systolic and diastolic frames from all four techniques for the LAD and RCA, respectively. The lack of contrast in the breath hold segmented GRE acquisition is clearly seen, particularly in systole, when through-plane flow enhancement is low. This is caused by the implementation of view sharing, which although necessary to increase the number of cine phases in the segmented acquisition, precludes the use of fat-suppression techniques. The good contrast apparent with the spiral techniques is a result of the use of a water-only excitation pulse. Examples of right temporal flow profiles (uncorrected for throughplane motion of the vessel) obtained with these techniques are shown in Figure 23-10. Flow profiles generated with the free breathing spiral sequence (100-cardiac-cycle acquisition, assuming 40% respiratory efficiency) agreed closely with those generated with the free breathing GRE Cardiovascular Magnetic Resonance 317

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Figure 23-7 Through-plane magnitude images and velocity maps for retrospective respiratory gated (RRG) acquisitions (top) and breath hold acquisitions (bottom) acquired (A) in mid-diastole (gating delay of 90 msec) and (B) in early systole (gating delay of 760 msec). In mid-diastole, when in-plane displacement is low, good-quality images are obtained with both techniques. In early systole, however, when in-plane displacement is high, the long acquisition window of the breath hold sequence results in considerable blurring of the artery (arrows). (From Hofman MBM, van Rossum AC, Sprenger M, Westerhof N. Assessment of flow in the right human coronary artery by magnetic resonance phase contrast velocity measurement: effects of cardiac and respiratory motion. Magn Reson Med. 1996;35:521–531.)

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Figure 23-8 Original tracing of an invasively determined flow curve of the left coronary artery compared with noninvasively determined flow curves in the same patient. The solid line shows the real-time adaptive navigator correction technique; the dotted line shows the breath-hold technique. (From Nagel E, Bornstedt A, Hug J et al. Noninvasive determination of coronary blood flow velocity with magnetic resonance imaging: comparison of breath-hold and navigator techniques with intravascular ultrasound. Magn Reson Med. 1999;41:544–549.) 318 Cardiovascular Magnetic Resonance

sequence, taking 10 times longer to acquire for the same spatial resolution (see Fig. 23-10A). By comparison, the breath hold segmented GRE sequence (see Fig. 23-10C) did not resolve the sharp peaks in the temporal flow profiles of the RCA, and for the group as a whole, significantly underestimated peak systolic (88 mm/sec vs. 252 mm/sec, p < .001), peak diastolic (114 mm/sec vs. 153 mm/sec, p < 0.01), and mean (56 mm/sec vs. 93 mm/sec, p < 0.001) velocities. For the less mobile left artery, peak systolic, peak diastolic, and mean velocities were also underestimated by the breath hold segmented GRE sequence, but this only reached statistical significance for the peak systolic velocity (80 mm/sec vs. 135 mm/sec, p < 0.01). The breath hold spiral sequence agreed reasonably well with the free breathing GRE sequence (see Fig. 23-10B), although deviations were observed at times of rapid cardiac movement. This was due to misregistration of the velocity encoded and reference datasets that were acquired consecutively in the same cardiac cycle in the

23 CORONARY ARTERY AND SINUS VELOCITY AND FLOW

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Cardiovascular Magnetic Resonance 319

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breath hold acquisition, whereas in the free breathing acquisition, they were acquired at the same time point on alternate cycles. The profiles shown in Figure 23-10 are uncorrected for through-plane motion of the vessel. A region of adjacent myocardium is generally used as a marker of through-plane motion and is used to correct flow profiles. This is feasible for the left coronary artery, but for the RCA, the adjacent myocardium is too thin to provide a reliable correction. For this artery, epicardial fat surrounding the artery has been used for through-plane correction, with the fat signal provided from a fat-excitation image generated from a separate acquisition, or from one interleaved with the water-excitation image used for imaging flow.47 Figure 23-11 shows mean ( SD) flow profiles before and after correction in the left (N ¼ 13) and right (N ¼ 10) coronary arteries of healthy subjects. The corrected flow profiles bear strong resemblances to those found in normal arteries using Doppler flow wire.48 Very recently,49 breath hold interleaved spiral phase velocity mapping was performed in the RCA of nine healthy subjects at 3 T, and it was shown that isometric handgrip exercise induced significant increases in peak diastolic flow velocity (26.8 9.3 cm/sec vs. 37.0 12.8 cm/sec, p < 0.01), which returned to baseline values after 1 minute. The advantages of high SNR, good temporal and spatial resolution, and ability to implement fat suppression give the spiral technique considerable promise, but more research is needed to investigate how the blurring of off-resonance material50 and flow-direction-sensitive “implosion/explosion” artifacts in regions of poor homogeneity51 affect the measurements. Validation studies against Doppler guidewire are also necessary, and the extent of image distortion in the vicinity of stents requires investigation.

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Figure 23-10 Velocity profile for an example right coronary artery as measured by the high temporal resolution free breathing spiral (FB_SP) sequence compared with that measured by the free breathing gradient echo (FB_GRE) sequence (A), the breath hold spiral (BH_SP) sequence (B), and the breath hold GRE (BH_GRE) sequence (C). (From Keegan J, Gatehouse PD, Mohiaddin RH, Yang GZ, Firmin DN. Comparison of spiral and FLASH phase velocity mapping, with and without breath-holding, for the assessment of left and right coronary artery blood flow velocity. J Magn Reson Imaging. 2004;19:40–49.)

320 Cardiovascular Magnetic Resonance

Validation and Feasibility Studies The first directly validated measurements of coronary artery flow and coronary flow reserve using segmented GRE phase velocity mapping were performed by Clarke and colleagues in dogs in 1995.52 Nonmagnetic perivascular ultrasound probes were placed around the isolated LAD and LCX in eight ventilated dogs. A subcritical stenosis was generated in the LAD by placement of a polycarbonate resin constrictor. Breath holding was effectively achieved by turning the ventilator off. Cine phase velocity mapping was performed using a sequence with TE of 11 msec and TR of 19.1 msec, with two to three reference and velocity sensitized view pairs per data

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segment. The segment duration was therefore 76 to 115 msec, allowing the acquisition of four to six image frames over the cardiac cycle. Images were acquired with an in-plane pixel size of 0.70 to 0.94 mm2 over breath holds of up to 40 seconds’ duration. The velocity sensitivity was 138 cm/sec. Data were acquired both before and after the administration of adenosine. A region of interest was drawn around the artery of interest in the magnitude image, and the mean velocity in that region on the corresponding velocity map was calculated. In this study, the size of the region of interest was kept constant from frame to frame and only its position changed because the authors believed that the spatial resolution in the CMR image was insufficient to track phasic changes

in the coronary arterial diameter. The results of a typical experiment are shown in Figure 23-12A, and a plot of CMR flow reserve versus that measured by ultrasound is shown in Figure 23-12B. For the LCX and LAD, mean CMR-measured flow reserves were 2.57 0.92 and 1.38 0.31, respectively (p ¼ 0.011), compared with ultrasound-measured values of 2.55 0.77 and 1.42 0.31, respectively (p ¼ 0.002). The mean difference between the two techniques was 0.01, with limits of agreement of þ0.64 and 0.61. The authors concluded that, although the breath hold period in this study is not feasible for the majority of patients and both temporal and spatial resolution are limited, the results of the phase velocity mapping technique agreed well with Doppler Cardiovascular Magnetic Resonance 321

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Figure 23-12 A, The typical experiment took approximately 90 minutes, with ultrasound measurements of flow in the left circumflex artery (blue line) and left anterior descending artery (red line) being recorded throughout. Two baseline cine velocity CMR sets were acquired in the left anterior descending artery (red circle) and left circumflex artery (blue square). Magnetic resonance velocity cine image sets were recorded in each vessel during adenosine infusion. B, Scatterplot of linear regression of the 16 CMR-estimated and ultrasound-measured coronary flow reserve measurements produced a regression line with a slope of 1.04, an intercept of 0.1, and a correlation coefficient of 0.94. (From Clarke GD, Eckels R, Chaney C, et al., Measurement of absolute epicardial coronary artery blood flow and flow reserve with breath-hold cine phase-contrast magnetic resonance imaging. Circulation. 1995;91:2627–2634.)

ultrasound, suggesting that it could be used for accurate, noninvasive measurement of absolute flow and flow reserve in the major epicardial coronary arteries. The same technique was used to measure flow in the LAD in 12 subjects both before and after the administration of adenosine before assessment with an intracoronary Doppler flow wire during cardiac catheterization.53 In this study, with an in-plane pixel size of 0.8 to 1.0 mm, an inplane presaturation pulse was applied before each cine acquisition, suppressing the signal from tissue in the slice and enabling inflowing blood to be visualized with improved contrast. Depending on the heart rate, four to five cine phases were acquired per cardiac cycle, with the number of phase encoding steps reduced to ensure complete data acquisition in a breath hold of up to 25 seconds. A respiratory gating belt was used to ensure the consistency of the breath hold positions. Excellent agreement was reported between CMR and ultrasound coronary flow and flow reserve measurements, with the limits of agreement 43 and 27 mL/min and 0.5 and 0.3, respectively. Sakuma and associates used a similar technique to validate CMR-measured coronary flow against findings obtained with an open-chest sonographic flowmeter in seven dogs54 under suspended mechanical ventilation. Good correlation was found (r ¼ 0.95), with a slope close to unity and low interobserver variability (8%). The same authors used CMR velocity mapping to study the velocity profiles in the LAD of eight healthy subjects both before and after dipyridamole administration.55 However, in this study, the use of view sharing enabled the effective temporal resolution of the imaging sequence to be improved to 64 msec, and 7 to 13 cine images were acquired in an end-expiratory breath hold of 24 heartbeats. Correction of coronary blood flow velocity for the through-plane velocity of the vessel as a whole was also performed by assuming that the velocity of the vessel was the same as that in an area of adjacent myocardium. This averages to zero over 322 Cardiovascular Magnetic Resonance

the cardiac cycle as a whole and is therefore unimportant for the assessment of mean coronary blood flow parameters, but may vary from as much as þ20 cm/sec in systole to 10 cm/sec in diastole.56 Consequently, it has strong implications for the assessment of instantaneous flow and flow velocity. In this study, volume flow data were not calculated because the authors believed that the small number of pixels across the vessel prevented the accurate assessment of cross-sectional area. The average baseline peak diastolic flow velocity was 14.8 cm/sec, increasing to 46.3 cm/sec after dipyridamole and resulting in an average coronary flow velocity reserve of 3.1 0.6. Interstudy reproducibility (i.e., absolute difference in two measurements divided by the mean of two measurements expressed as a percentage) was 9.5% and 6.8% pre- and post-pharmacologic stress, respectively, with comparable interobserver reproducibility. A similar study performed at the same center in 12 healthy subjects using sustained handgrip exercise to induce vasodilation likewise showed an increase in peak diastolic flow velocity from 21 to 31 cm/sec.57 Further validation against Doppler flow wire was found by Shibata and colleagues58 in a group of 19 patients scheduled for X-ray coronary angiography. Flow velocities measured with CMR were found to be significantly and considerably less than those measured by Doppler flow wire (12.5 4.9 cm/sec vs. 32.4 12.9 cm/sec, p < 0.01), although measures of coronary flow velocity reserve showed good correlation (r ¼ 0.9), with no statistically significant difference. Figure 23-13 shows examples of flow velocity curves at baseline and after dipyridamole injection in a patient with a normal LAD and in a patient with a significant stenosis. The flow curves in the stenosed artery show a less than twofold increase in peak diastolic flow velocity with dipyridamole, resulting in reduced coronary flow velocity reserve. The correlation between coronary flow velocity reserve measured by Doppler guidewire and

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Figure 23-13 A, Flow velocity curves in a patient with a normal left anterior descending artery measured by CMR at baseline and after dipyridamole injection. Diastolic peak velocity showed a threefold increase after the administration of dipyridamole. B, Flow velocity curves in a patient with a significant stenosis in the left anterior descending artery measured by CMR at baseline and after dipyridamole injection. Diastolic peak velocity showed a less than twofold increase after the administration of dipyridamole. ECG, electrocardiogram. (From Shibata M, Sakuma H, Isaka N, et al. Assessment of coronary flow reserve with fast cine phase contrast magnetic resonance imaging: comparison with measurement by Doppler guidewire. J Magn Reson Imaging. 1999;10:563–568.)

that measured by CMR is shown in Figure 23-14. Bedaux and associates59 similarly found that the coronary flow velocity reserve measured by CMR was not significantly different from that measured by Doppler flow wire in both healthy and diseased arteries (2.7 1.0 vs. 3.1 0.6 and 1.5 0.7 vs. 1.9 0.7, respectively). Validation of CMR measurement of left anterior descending coronary artery flow velocity reserve has also been performed against myocardial perfusion reserve in the anterior myocardium using PET and 15O-labeled water.60 In 10 healthy subjects, coronary blood flow velocity reserve measured with CMR was 2.44 1.44 compared with 2.52 0.84 when measured with PET (p ¼ NS), with a correlation between the two measurements of 0.79. In a further feasibility study, coronary volume flow profiles were measured in normal subjects, with and without pharmacologic stress.61 Again, by using view sharing techniques, 7 to 13 images could be obtained per cardiac cycle over breath holds of up to 20 seconds, depending on the R-R interval. In this study, the in-plane resolution was

0.9 1.4 mm. Regions of interest were drawn around the vessel, using the magnitude image as a guide. Within this region, automated edge detection was used to determine the exact vessel area, using a magnitude threshold of 35%, and the region of interest was adjusted (in both position and size) for each frame in the cardiac cycle. Three acquisitions were performed before dipyridamole administration, and one was performed after dipyridamole administration. The mean area covered by the LAD was 15.6 mm2 before dipyridamole administration, increasing to 17.7 mm2 after dipyridamole administration. Peak velocity, mean coronary flow velocity, and mean flow all increased significantly after dipyridamole administration, with coronary flow reserve of 5.0 2.6 compared with coronary flow velocity reserve of 3.5 1.3. There are a number of potential reasons for the discrepancy between the coronary reserves measured with volume flow and peak velocity, the most likely being that the calculations made from peak velocity measurements are based on single-pixel values with higher statistical noise, which, in this case, were uncorrected for the throughplane movement of the vessel itself. Simultaneous flow mapping in the LAD and great cardiac vein has also been performed by careful selection of the imaging plane and has been used for vessel identification.62 These vessels run closely parallel to each other in the interventricular sulcus and can be difficult to distinguish. The results of flow mapping showed the expected diastolic-predominant flow in the LAD, whereas flow in the great cardiac vein is predominantly systolic. These characteristic flow profiles clearly identify the vessels and may be helpful in cases in which early intersection of the great cardiac vein and the LAD make the distinction difficult. For assessment of volume flow rate, accurate delineation of the cross-sectional area of the coronary artery is essential. The majority of studies use phase difference processing of Cardiovascular Magnetic Resonance 323

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the data, with a magnitude threshold of 30%18 or 50%.63 An alternative approach would be to use complex difference processing, which takes into account partial volume averaging of edge pixels. Given the small diameter of the coronary arteries and the relatively large pixel size, this may be expected to perform better than phase difference processing.64,65 A study comparing complex difference and phase difference (30% magnitude threshold) CMR velocity mapping techniques in humans showed that the phase difference technique overestimated volume flow rates by approximately 25%.66 This has been studied in more detail by Wedding and colleagues,67 who compared CMR measurements of coronary flow and flow reserve with those obtained from an ultrasonic transit time probe in dogs when using phase difference processing (both 30%18 and 50% thresholds) and complex difference processing (both with and without correction for in-plane motion).68 They observed that mean CMR flow measurements correlated well with ultrasound data for all techniques (r ¼ 0.92 to 0.94), but both the complex difference technique and the 50% magnitude threshold phase difference technique systematically underestimated flow by 25% to 30%. Against expectations, the phase difference technique with the 30% magnitude threshold provided the best agreement with ultrasound, similar to that reported in other studies,52,53 but it was very sensitive to vessel within the boundary identification and the number of pixels within it. Possible reasons for the unexpectedly poorer performance of the complex difference technique were motion blurring of the vessel and poor SNR.

Patient Studies Despite the plethora of validation and feasibility studies that have been performed, relatively few clinical studies have been reported. The first report of using CMR coronary artery velocity mapping for the functional assessment of proximal and middle LAD stenoses in humans was by Hundley and associates69 Thirty-three patients who were referred for cardiac catheterization because of chest pain were examined with breath hold CMR velocity mapping both before and during intravenous administration of 140 mg/kg/min adenosine. Coronary flow and coronary flow reserve were measured and compared with the results of computerassisted quantitative coronary angiography and, where possible, Doppler flow wire (N ¼ 17). All patients with stenosis severity of greater than 70% by quantitative coronary angiography were identified as having coronary flow reserve of 1.7 or less by CMR. The sensitivity and specificity of CMR coronary flow reserve of 1.7 or less for the identification of stenosis severity of greater than 70% were 100% and 83%, respectively. The correlation between Doppler flow wire and CMR measures of coronary flow reserve was 0.87, with CMR underestimating flow reserve by approximately 35%. The authors concluded that CMRbased measurement of coronary flow reserve may prove useful in identifying patients who are likely to obtain a survival benefit from coronary artery bypass grafting, where benefits are particularly high if there is involvement of the proximal LAD. A further report by the same authors discussed the assessment of coronary arterial restenosis after 324 Cardiovascular Magnetic Resonance

successful percutaneous coronary arterial stenosis.70 In this study, 17 patients with recurrent chest pain more than 3 months after intervention were studied by CMR, and it was found that a coronary flow reserve value of 2.0 or less was 100% and 82% sensitive and 89% and 100% specific for detecting luminal diameter narrowing of 70% or greater and 50% or greater, respectively. Figures 23-15 and 23-16 show examples of data in patients without and with (respectively) a significant stenosis, as determined by Xray coronary angiography. Two further studies have been performed to assess the detection of restenosis after coronary artery stent implantation. In the first, serial changes in CMR coronary flow velocity reserve were evaluated in 10 patients with coronary artery disease who had undergone successful elective stent implantation of a lesion in the LAD.71 Flow velocities were measured distal to the stent every 4 weeks for 6 months, with follow-up angiography performed 6 months after the initial intervention. Patients without restenosis at follow-up angiography (N ¼ 7) had normal coronary flow velocity reserve during the 6 months of follow-up, with values of 2.31 0.30 at 1 month and 2.52 0.25 at 6 months after stent implantation. However, those with restenosis (N ¼ 3) showed a significant decrease, with values of 2.26 0.49 at 1 month and 1.52 0.09 at 6 months (p < .05). In a larger study, 38 patients were studied 24 hours after stent implantation and again 3 months later.72 Results were compared with X-ray coronary angiography, and in 18 cases, results were compared additionally with the results of Doppler flow measurements. CMR measurements could be made in 75% of the patients studied. At follow-up, coronary flow velocity reserve was 1.78 0.16 in vessels without coronary stenosis, 1.46 0.22 in vessels with 50% to 74% stenosis, and 1.10 0.22 in those with 75% or greater stenosis (p < .05 between all groups), as shown in Figure 23-17. A threshold coronary flow velocity reserve of 1.2 resulted in sensitivity of 83% and specificity of 94% for the detection of stenoses of greater than 75%.

CONCLUSION AND FUTURE DEVELOPMENTS A number of validation and feasibility studies have shown that CMR techniques have the potential to assess coronary artery flow velocity, flow, flow reserve, and flow velocity reserve, with the most commonly used technique being segmented GRE phase velocity mapping. Although a number of validation studies have shown good correlation of CMR data with Doppler flow wire and ultrasound transit time techniques, the robustness of the technique must be improved and larger-scale validation studies need to be performed. The use of the interleaved spiral technique, with its short readout window and high temporal resolution, may prove to be a preferable technique and needs to be further investigated. To summarize the requirements, any technique must take into account the following factors: 1. Spatial resolution: For the assessment of volume flow, spatial resolution must be sufficiently good to measure the cross-sectional area throughout the cardiac cycle and to minimize partial volume averaging of flow velocities in edge pixels. These partial volume effects become

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Figure 23-15 Tangential (top left) and cross-sectional (top right) coronary artery CMR of the left anterior descending coronary artery (LAD). On the far right is a selected image from the patient’s contrast coronary angiogram. On tangential cardiovascular magnetic resonance imaging (CMR) and the contrast coronary angiogram, the location of the intracoronary stent is shown. The arrow on the CMR cross-sectional image shows the corresponding cross-sectional view of the LAD imaged 5 to 10 mm distal to the intracoronary stent. From these crosssectional images, the region of interest (ROI) is selected for calculating the cross-sectional area of the vessel. The ROI is then transferred to velocity maps, in which the intensity of the grayscale of each pixel within the ROI represents velocity (darker indicates higher velocity). The bottom row of images shows corresponding velocity maps (cross-sectional images of vessels) from the patient at rest and after pharmacologic stress with adenosine. On the right velocity map, darker pixels within black circles indicate higher velocities after adenosine infusion. This patient had coronary flow reserve of 3.1 and 35% stenosis by quantitative coronary angiography (QCA). LCx, left circumflex; LM, left main. (From Hundley WG, Hillis LD, Hamilton CA, et al. Assessment of coronary arterial restenosis with phase-contrast magnetic resonance imaging measurements of coronary flow reserve. Circulation. 2000;101:2375–2385.)

larger as the vessel becomes smaller and the relative number of pixels in the boundary increases. The degree to which measured velocities are affected depends strongly on the relative magnitude of the stationary material included in the boundary pixels. Several studies have investigated these effects in small vessels. Hofman and colleagues63 compared phase contrast measurements of time-averaged volume flow in the femoral arteries of dogs with measurements made by an ultrasonic transit time meter and found that the proportional difference between the techniques was 0.8% when the number of pixels across the vessel diameter was just three. Tang and associates73 found that, for both in vitro and in vivo studies, volume flow accuracy increases with resolution, as expected, and errors are less than 10% when the ratio of pixel size to vessel radius is less than 0.5. This was also seen by Wolf and coworkers, who showed an error of 9% for five pixels across the vessel diameter.74 Sondergaard and colleagues showed less than 18% error in measured flow

rate for four pixels across the vessel diameter in vitro.75 Arheden and associates76 studied pulsatile flow in tubes with various spatial resolutions (but with a minimum of five pixels across the tube diameter) and found that overestimation of the size of the region of interest resulted in flow errors of less than 20%, provided that the imaging plane was within 10 of being perpendicular to the direction of flow. As has been noted, the effects of partial volume errors could potentially be reduced by using a complex difference, rather than phase difference, approach to the generation of velocity maps.64 2. Temporal resolution: The temporal resolution of the sequence used should be adequate to resolve the phasic velocity profile of coronary blood flow. This ability is determined by the number of views acquired per data segment in a segmented GRE acquisition. For accurate assessment of time-averaged parameters, measurements should be made throughout the entire cardiac cycle, with the accuracy increasing as the number of cine Cardiovascular Magnetic Resonance 325

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Figure 23-16 Tangential (top left) and cross-sectional (top right) coronary artery CMR of the left anterior descending (LAD) coronary artery. On the far right is a selected image from the patient’s X-ray coronary angiogram. On tangential CMR and the X-ray coronary angiogram, the location of the intracoronary stent is shown. On the right velocity map, persistent gray pixels after adenosine administration in the patient indicate impaired coronary flow reserve (value ¼ 0.8) that corresponded to the presence of 57% stenosis by quantitative coronary angiography. LCx, left circumflex; LM, left main. (From Hundley WG, Hillis LD, Hamilton CA, et al. Assessment of coronary arterial restenosis with phase-contrast magnetic resonance imaging measurements of coronary flow reserve. Circulation. 2000;101:2375–2385.)

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frames increases.77 For segmented GRE acquisitions, these points favor the acquisition of small numbers of views per data segment, although this leads to long breath hold periods or the need for navigator free breathing techniques. Alternatively, more rapid and efficient methods of k-space coverage may be used, such as interleaved rectilinear echo planar or spiral techniques. A further reason to reduce the duration of data acquisition is to minimize blurring of the vessel as a result of movement in the acquisition window, which leads to overestimation of the vessel cross-sectional area through partial volume averaging of edge pixels and underestimation of mean flow velocity. 3. Velocity sensitivity: The maximum velocity in normal coronary arteries at rest is typically less than 25 cm/sec. Therefore, the window used for resting studies should be narrow, approximately 50 cm/sec, to allow measurement of these low velocities with maximum accuracy but without aliasing. For maximal vasodilation studies, the window should be increased accordingly. 4. Through-plane movement of the vessel: As has been noted, through-plane movement of the vessel through the cardiac cycle can affect instantaneous measurement of blood flow velocity and volume flow in that vessel considerably, although time-averaged measurements

compared with long segment duration breath holding studies.41 The long scan times could potentially be reduced by using a real-time phase encode reordering technique whereby the most significant central lines of k-space are acquired with the diaphragm position, as measured by a navigator echo, within a narrow range and the outermost lines acquired with the diaphragm position within a larger range. These techniques have already been successfully applied to coronary artery CMR.78,79 Furthermore, realtime slice-following may also be implemented and could potentially result in improved scan efficiency by allowing the use of larger navigator windows without a reduction in image quality.

References 1. Pennell DJ, Bogren H, Keegan J, Firmin DN, Underwood SR. Assessment of coronary artery stenosis by magnetic resonance imaging. Heart. 1996;75:127–133. 2. Torii R, Wood NB, Hughes AD, et al. A computational study on the influence of catheter-delivered intravascular probes on blood flow in a coronary artery model. J Biomech. 2007;40:2501–2509. 3. Mohiaddin RH, Pennell DJ. Magnetic resonance imaging of flow in the cardiovascular system. In: Reichek N, ed. Cardiology Clinics. Philadelphia: WB Saunders; 1998:161–188. 4. Gatehouse PD, Keegan J, Crowe LA, et al. Applications of phasecontrast flow and velocity imaging in cardiovascular MRI. Eur Radiol. 2005;15:2172–2184. 5. Hood WB. Regional drainage of the human heart. Br Heart J. 1968;30:105–109. 6. van Rossum AC, Visser FR, Hofman MBM, et al. Global left ventricular perfusion: noninvasive measurement with cine MR imaging and phase velocity mapping of coronary venous outflow. Radiology. 1992;182:685–691. 7. Ganz W, Tamura K, Marcus HS, et al. Measurement of coronary sinus blood flow by continuous thermodilution in man. Circulation. 1971;54:181–195. 8. Canty JM, Brooks A. Phasic volumetric coronary venous outflow patterns in conscious dogs. Am J Physiol. 1990;258:H1457–H1463. 9. Mirowitz SA, Lee JKT, Guterrez FR, Brown JJ, Eilenberg SS. Normal signal-void patterns in cardiac cine MR images. Radiology. 1990;176: 49–55. 10. Kawada N, Sakuma H, Yamakado T, et al. Hypertrophic cardiomyopathy: MR measurement of blood flow and vasodilator flow reserve in patients and healthy subjects. Radiology. 1999;211:129–135. 11. Edelman RR, Wallner B, Singer A, Atkinson DJ, Saini S. Segmented TurboFLASH: method for breath-holding MR imaging of the liver with flexible contrast. Radiology. 1990;177:515–521. 12. Foo TK, Bernstein MA, Aisen AM. Improved ejection fraction and flow velocity estimates with use of view sharing and uniform repetition time excitation with fast cardiac techniques. Radiology. 1995;195: 471–478. 13. Kennedy K, Dick A, Drangova M, et al. Magnetic resonance measurements of coronary flow reserve in heart transplant recipients: an exploratory study of the relationship to coronary angiographic findings. J Cardiovasc Magn Reson. 2007;9:701–707. 14. Koskenvuo JW, Sakuma H, Niemi P, et al. Global myocardial blood flow and global flow reserve measurements by MRI and PET are comparable. J Magn Reson Imag. 2001;13:361–366. 15. Koskenvuo JW, Hartiala JJ, Knuuti J, et al. Assessing coronary sinus blood flow in patients with coronary artery disease: a comparison of phase-contrast MR imaging with positron emission tomography. Am J Roentgenol. 2001;177:1161–1166. 16. Ferrigno M, Hickey DD, Liner MH, Lundgren CEG. Cardiac performance in humans during breath-holding. J Appl Physiol. 1986;60:1871–1877. 17. Schwitter J, DeMarco T, Kneifel S, et al. Magnetic resonance-based assessment of global coronary flow and flow reserve and its relation to left ventricular functional parameters. Circulation. 2000;101:2696–2709. 18. Pelc NJ, Herfkens RJ, Shimakawa A, Enzmann DR. Phase contrast cine magnetic resonance imaging. Magn Reson Q. 1991;7:229–254.

19. Bailes DR, Gilderdale DJ, Bydder GM, Collins AG, Firmin DN. Respiratory ordered phase encoding (ROPE): a method for reducing respiratory motion artefacts in MR imaging. J Comput Assist Tomogr. 1985;9:835–838. 20. Waller BF, Schlant RC. Anatomy of the heart. In: Schlant RC, Alexander RW, eds. The Heart, Arteries and Veins. 8th ed. New York: McGraw-Hill; 1994:59–111. 21. Schwitter J, Kozerke S, Bremerich J, et al. Oral administration of 17bestradiol over 3 months without progestin co-administration does not improve coronary flow reserve in post-menopausal women: a randomised placebo-controlled cross-over CMR study. J Cardiovasc Magn Reson. 2007;9:665–672. 22. Lund GK, Wendland MF, Shimakawa A, et al. Coronary sinus flow measurement by means of velocity encoded cine MR imaging: validation by using flow probes in dogs. Radiology. 2000;217:487–493. 23. Lund GK, Watzinger N, Saeed M, et al. Chronic heart failure: global left ventricular perfusion and coronary flow reserve with velocityencoded cine MR imaging: initial results. Radiology. 2003;227: 209–215. 24. Aras A, Anik Y, Demirci A, et al. Magnetic resonance imaging measurement of left ventricular blood flow and coronary flow reserve in patients with chronic heart failure due to coronary artery disease. Acta Radiol. 2007;48:1092–1100. 25. Watzinger N, Lund GK, Saeed M, et al. Myocardial blood flow in patients with dilated cardiomyopathy: quantitative assessment with velocity-encoded cine magnetic resonance imaging of the coronary sinus. J Magn Reson Imaging. 2005;21:347–353. 26. Bogren HG, Buonocore MH. Measurement of coronary artery flow reserve by magnetic resonance velocity mapping in the aorta. Lancet. 1993;341:899–900. 27. Buonocore MH. Blood flow measurement using variable velocity encoding in the RR interval. Magn Reson Med. 1993;29:790–795. 28. Buonocore M. Estimation of total coronary artery flow using measurements of flow in the ascending aorta. Magn Reson Med. 1994;32: 602–611. 29. Sorenson MB, Fritz-Hansen T, Jensen HH, Pedersen AT, Ottesen B. Measurement of aortic blood flow by magnetic resonance below and above the origin of the coronary arteries in postmenopausal hormone replacement therapy. JCMR. 2004;6:637–644. 30. Burstein D. MR imaging of coronary artery flow in isolated and in vivo hearts. JMRI. 1991;1:337–346. 31. Edelman R, Mattle HP, Keefield J, Silver MS. Quantification of blood flow with dynamic MR imaging and presaturation bolus tracking. Radiology. 1989;171:551–556. 32. Chao H, Burstein D. Multibolus stimulated echo imaging of coronary artery flow. J Magn Reson Imaging. 1997;7:603–605. 33. Poncelet BP, Weisskoff RM, Weeden VJ, Brady TJ, Kantor H. Time of flight quantification of coronary flow with echo-planar MRI. Magn Reson Med. 1993;30:447–457. 34. Firmin DN, Poncelet BP. Echo-planar imaging of the heart. In: Schmitt F, Stehling MK, Turner R, eds. Echo-Planar Imaging: Theory, Technique and Application. Berlin: Springer-Verlag; 1998:389–418. 35. Edelman RR, Manning WJ, Gervino E, Li W. Flow velocity quantification in human coronary arteries with fast breath-hold MR angiography. J Magn Reson Imaging. 1993;3:699–703.

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throughout the cardiac cycle are unaffected because the through-plane movement of the vessel averages to zero. In addition, beat-to-beat variations in the position of the artery and of flow through the artery inevitably influence the results, and these variations may be exacerbated by breath holding, particularly for long periods. Therefore, a free breathing technique using navigator echoes may be the best approach and would also allow higher spatial and temporal resolution images, albeit with reduced scan efficiency and consequently longer scan duration. As discussed earlier, such a technique has been shown to significantly improve the accuracy of CMR velocity assessment

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36. Keegan J, Firmin D, Gatehouse PD, Longmore D. The application of breath hold phase velocity mapping techniques to the measurement of coronary artery blood flow velocity: phantom data and initial in vivo results. Magn Reson Med. 1994;31:526–536. 37. Hofman MBM, van Rossum AC, Sprenger M, Westerhof N. Assessment of flow in the right human coronary artery by magnetic resonance phase contrast velocity measurement: effects of cardiac and respiratory motion. Magn Reson Med. 1996;35:521–531. 38. Lenz GW, Haacke EM, White RD, et al. Retrospective respiratory gating: a review of technical aspects and future directions. Magn Reson Imaging. 1989;7:445–455. 39. Hofman MBM, Wickline SA, Lorenz CH. Quantification of in-plane motion of the coronary arteries during the cardiac cycle: implications for acquisition window duration for MR flow quantification. J Magn Reson Imaging. 1998;8:568–576. 40. Marcus JT, Smeenk HG, Kuijer JP, van der Geest RJ, Heethaar RM, van Rossum AC. Flow profiles in the left anterior descending and the right coronary artery assessed by MR velocity quantification: effects of through-plane and in-plane motion of the heart. J Comput Assist Tomogr. 1999;23:567–576. 41. Nagel E, Bornstedt A, Hug J, et al. Noninvasive determination of coronary blood flow velocity with magnetic resonance imaging: comparison of breath-hold and navigator techniques with intravascular ultrasound. Magn Reson Med. 1999;41:544–549. 42. Danias PG, McConnell MV, Khasigawala VC, et al. Prospective navigator correction of image position for coronary MR angiography. Radiology. 1997;203:733–736. 43. Meyer CH, Hu BS, Nishimura DG, Macovski A. Fast spiral coronary artery imaging. Magn Reson Med. 1992;28:202–213. 44. Keegan J, Gatehouse PD, Taylor AM, et al. Coronary artery imaging in a 0.5 Tesla scanner: implementation of real-time navigator echo controlled segmented k-space FLASH and interleaved spiral sequences. Magn Reson Med. 1999;41:392–399. 45. Keegan J, Gatehouse PD, Yang GZ, Firmin DN. Interleaved spiral cine coronary artery velocity mapping. Magn Reson Med. 2000; 43:787–792. 46. Keegan J, Gatehouse PD, Mohiaddin RH, Yang GZ, Firmin DN. Comparison of spiral and FLASH phase velocity mapping, with and without breath-holding, for the assessment of left and right coronary artery blood flow velocity. J Magn Reson Imaging. 2004;19:40–49. 47. Keegan J, Gatehouse PD, Yang GZ, Firmin DN. Spiral phase velocity mapping of left and right coronary artery blood flow: correction for through-plane motion using selective fat-only excitation. J Magn Reson Imaging. 2004;20:953–960. 48. Ofili EO, Labovitz AJ, Kern MJ. Coronary flow velocity dynamics in normal and diseased arteries. Am J Cardiol. 1993;71:3D–9D. 49. Hays AG, Kelle S, Hirsch GA, et al. Non-invasive measurement of coronary artery flow velocity at rest and during handgrip stress in healthy subjects using 3T MRI. In: Proceedings of the 11th Annual Scientific Sessions of the Society for Cardiovascular Magnetic Resonance, Los Angeles. 2008:103. 50. Yudilevich E, Stark H. Spiral sampling in magnetic resonance imaging: the effect of inhomogeneities. IEEE Trans Med Imaging. 1987;MI-6: 337–345. 51. Gatehouse PD, Keegan J, Firmin DN. Flow distortion and signal loss in spiral imaging. Magn Reson Med. 1999;41:1023–1031. 52. Clarke GD, Eckels R, Chaney C, et al. Measurement of absolute epicardial coronary artery blood flow and flow reserve with breathhold cine phase-contrast magnetic resonance imaging. Circulation. 1995;91:2627–2634. 53. Hundley GW, Lange RA, Clarke GD, et al. Assessment of coronary arterial flow and flow reserve in humans with magnetic resonance imaging. Circulation. 1996;93:1502–1508. 54. Sakuma H, Saeed M, Takeda K, et al. Quantification of coronary artery volume flow rate using fast velocity encoded cine MR imaging. Am J Roentgenol. 1997;168:1363–1367. 55. Sakuma H, Blake LM, Amidon TM, et al. Coronary flow reserve: noninvasive measurement in humans with breath-hold velocity encoded cine MR imaging. Radiology. 1996;198:745–750. 56. Karwatowski SP, Mohiaddin RH, Yang GZ, et al. Noninvasive assessment of regional left ventricular long axis motion using magnetic resonance velocity mapping in normal subjects. J Magn Reson Imaging. 1994;4:151–155. 57. Globits S, Sakuma H, Shimakawa A, et al. Measurement of coronary blood velocity during handgrip exercise using breath-hold velocity encoded cine magnetic resonance imaging. Am J Cardiol. 1997;79: 234–237.

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58. Shibata M, Sakuma H, Isaka N, et al. Assessment of coronary flow reserve with fast cine phase contrast magnetic resonance imaging: comparison with measurement by Doppler guidewire. J Magn Reson Imaging. 1999;10:563–568. 59. Bedaux WL, Hofman MB, Cock CC, Stoel MG, Visser CA, van Rossum AC. Magnetic resonance imaging versus Doppler guidewire in the assessment of coronary flow reserve in patients with coronary artery disease. Coron Artery Dis. 2002;13:365–372. 60. Sakuma H, Koskenvuo JW, Niemmi P, et al. Assessment of coronary flow reserve using fast velocity-encoded cine MR imaging. Validation study using positron emission tomography. Am J Roentgenol. 2000;175:1029–1038. 61. Davis CP, Liu P, Hauser M, et al. Coronary flow and coronary flow reserve measurements in humans with breath-hold magnetic resonance phase contrast velocity mapping. Magn Reson Med. 1997;37:537–544. 62. Bedaux WLF, Hofman MBM, Visser CA, van Rossum AC. Simultaneous non-invasive measurement of blood flow in the great cardiac vein and left anterior descending artery. J Cardiovasc Magn Reson. 2001;3:227–235. 63. Hofman MBM, Visser FC, van Rossum AC, Vink GQ, Sprenger M, Westerhof N. In vivo validation of magnetic resonance blood volume flow measurements with limited spatial resolution in small vessels. Magn Reson Med. 1995;33:778–784. 64. Polzin JA, Alley MT, Korosec FR, et al. A complex-difference phasecontrast technique for measurement of volume flow rates. J Magn Reson Imag. 1995;5:129–137. 65. Polzin JA, Korosec FR, Wedding KL, et al. Effects of through-plane myocardial motion on phase difference and complex difference measurements of absolute coronary artery flow. J Magn Reson Imaging. 1996;1:113–123. 66. Grist TM, Polzin JA, Bianco JA, Foo TKF, Bernstein MA, Mistretta CM. Measurement of coronary flow and flow reserve using magnetic resonance imaging. Cardiology. 1997;88:80–89. 67. Wedding KL, Grist TM, Folts JD, et al. Coronary flow and flow reserve in canines using MR phase difference and complex difference processing. Magn Reson Med. 1998;40:656–665. 68. Frayne R, Polzin JA, Mazaheri Y, Grist TM, Mistretta CA. Effect of and correction for in-plane myocardial motion on estimates of coronaryvolume flow rates. J Magn Reson Imaging. 1997;7:815–828. 69. Hundley GW, Hamilton CA, Clarke GD, Hillis DL, et al. Visualization and functional assessment of proximal and middle left anterior descending coronary stenoses in humans with magnetic resonance imaging. Circulation. 1999;99:3248–3254. 70. Hundley WG, Hillis LD, Hamilton CA, et al. Assessment of coronary arterial restenosis with phase-contrast magnetic resonance imaging measurements of coronary flow reserve. Circulation. 2000;101:2375–2385. 71. Saito Y, Sakuma H, Shibat M, et al. Assessment of coronary flow velocity reserve using fast velocity-encoded cine MRI for non-invasive detection of restenosis after coronary stent implantation. J Cardiovasc Magn Reson. 2001;3:209–214. 72. Nagel E, Thouet T, Klein C, et al. Noninvasive determination of coronary blood flow velocity with cardiovascular magnetic resonance in patients after stent deployment. Circulation. 2003;107:1738–1747. 73. Tang C, Blatter DD, Parker DL. Accuracy of phase contrast flow measurements in the presence of partial-volume effects. J Magn Reson Imaging. 1993;3:377–385. 74. Wolf RL, Ehman RL, Riederer SJ, Rossman PJ. Analysis of systematic and random error in MR volumetric flow measurements. Magn Reson Med. 1993;30:82–91. 75. Sondergaard L, Stahlberg F, Thomsen C, et al. Accuracy and precision of MR velocity mapping in measurement of stenotic cross-sectional area, flow rate and pressure gradient. J Magn Reson Imaging. 1993;3:433–437. 76. Arheden H, Saeed M, Tornqvist E, et al. Accuracy of segmented MR velocity mapping to measure small vessel pulsatile flow in a phantom simulating cardiac motion. J Magn Reson Imaging. 2001;13:722–728. 77. Clarke GD, Hundley WG, McColl RW, et al. Velocity-encoded, phasedifference cine MRI measurements of coronary artery flow: dependence of flow accuracy on the number of cine frames. J Magn Reson Imaging. 1996;6:733–742. 78. Jhooti P, Keegan J, Gatehouse PD, et al. 3D Coronary artery imaging with phase reordering for improved scan efficiency. Magn Reson Med. 1999;41:555–562. 79. Weiger M, Bornert P, Proksa R, Schaffter T, Haase A. A motionadapted gating based on k-space weighting for reduction of respiratory motion artefacts. Magn Reson Med. 1997;38:322–333.

Coronary Artery Bypass Graft Imaging and Assessment of Flow Constantin B. Marcu and Albert C. van Rossum

Since the introduction of coronary artery bypass grafting (CABG) by Favaloro in 1968, an increasing number of these surgical procedures have been performed worldwide.1 In the United States, an estimated 467,000 CABG procedures were performed on 268,000 patients in 2003.2 Generally, the saphenous vein is used for sequential grafting to distal branches of the right (RCA) and circumflex coronary artery (LCX) or to diagonal branches of the left anterior descending coronary artery (LAD). The left internal mammary artery (IMA) is frequently used as an arterial conduit to the LAD and its diagonal branches. Other arterial conduits include the right internal thoracic artery placed to the RCA or LAD, the right gastroepiploic artery, which may be placed to the RCA or the use of free radial artery grafts. The long-term results of aortocoronary bypass surgery depend largely on the maintenance of graft patency.3 About 25% of venous grafts occlude within 1 year of surgery, and half of these occur within 2 weeks of surgery. During the following 5 years, there is a 2% annual occlusion rate, which increases to 5% yearly thereafter. Thus, 50% to 60% of venous grafts are occluded after 10 years.4 The responsible mechanisms for venous graft occlusion are believed to be thrombosis in the early weeks after surgery, followed by intimal hyperplasia during the first year and accelerated atherosclerosis in the later stages. Atherosclerotic changes develop comparatively in a smaller percentage of patients with IMA grafts. As a result, in situ arterial grafts occlude less frequently, up to 5% in the first year and 20% to 30% after 10 years, leading to an improved long-term survival.5,6 Right IMA grafts appear to have a lower long-term patency rate compared to left IMA grafts.7 These figures imply a need for diagnostic modalities that can evaluate bypass grafts patency and function during postoperative follow-up. In many patients, these evaluations have to be made several times in a lifetime.

IMAGING MODALITIES CAPABLE OF EVALUATING GRAFTS Selective X-ray coronary angiography is the routine procedure and the gold standard for assessment of graft anatomy, but it is invasive and bears a limited risk associated with the use of ionizing radiation and iodine contrast material. An important advantage of selective angiography is the simultaneous assessment of the native coronary artery

system status. Also, using a Doppler-tipped guidewire, hemodynamic information about graft function can be obtained by measuring diastolic to systolic flow velocity ratios at rest, and flow velocity reserve after pharmacologically induced hyperemia.8 Noninvasive (cardiovascular magnetic resonance, or CMR) or semi-invasive (computed tomography, electron beam tomography) techniques have been developed for the direct evaluation of CABG patency.9–14 Two-dimensional (2D) Doppler echocardiography is restricted to evaluation of grafts placed on the LAD artery15–17 or through special transducers during the “open chest” intraoperative phase of CABG. A unique feature of CMR is that in addition to standard anatomic imaging, blood flow volume, and velocity can be quantified within the grafts. Thus, the true physiologic status of the functional unit represented by a graft and its recipient vessel can be determined noninvasively.18

CARDIOVASCULAR MAGNETIC RESONANCE OF BYPASS GRAFTS Over the years, several CMR techniques have been introduced to evaluate aortocoronary bypass grafts (Table 241).19 The assessment strategy may include either anatomic (angiographic) or hemodynamic (volume flow, velocities, flow reserve) evaluation or a combination of these modalities. In general, pulse sequences developed for imaging native coronary arteries can also be applied for imaging bypass grafts, but CABG imaging is associated with specific problems (difference in vessel anatomy and physiology of flow, presence of metallic vascular clips, postoperative mediastinal changes). The majority of clinical studies available report on imaging the proximal vein grafts. Proximal graft segments are less exposed to bulk cardiac motion than are distal segments or native coronary arteries, resulting in fewer motion artifacts, while the absence of direct contact with epicardial fat or myocardium results in higher contrast to surrounding tissues. However, graft stenosis often occurs at the site of anastomosis with the native vessel where CMR encounters artifacts and spatial resolution problems similar to those in imaging the native coronary arteries. Only few clinical CMR studies have addressed the imaging of arterial grafts, in which an additional problem in obtaining good image quality is the presence of artifacts associated with Cardiovascular Magnetic Resonance 329

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Table 24-1 Detection of Bypass Graft Patency According to Different CMR Techniques Reference

Technique 20

White et al. Rubinstein et al.21 Jenkins et al.22 Frija et al.23 White et al.24 Aurigemma et al.25 Galjee et al.26 Kessler et al.33 Vrachliotis et al.37 Wintersperger et al.36 Kalden et al.30 Bunce et al.39 Langerak et al.34

Number of Grafts

SE CMR SE CMR SE CMR SE CMR Cine CMR Cine CMR SE CMR Cine CMR Combined 3D navigator 3D CE MRA, ECG-triggered 3D CE MRA Non-ECG triggered HASTE 3D CE MRA, ECG-triggered SSFP 3D CE MRA ECG-triggered 3D navigator Two observers

65 44 60 52 28 45 98 19 44 76 59 79 56

Sensitivity (%) 91 92 90 98 93 88 98 98 98 87 93 95 95 93 84 85 65–83

Specificity (%) 72 85 90 78 86 100 85 88 76 100 97 81 93 93 45 73 80–100

Accuracy (%) 86 89 90 94 89 91 96 96 94 89 95 92 95 93 78 84 80

CE MRA, contrast-enhanced magnetic resonance angiography; ECG, electrocardiogram; HASTE, half-Fourier acquisition single-shot turbo spin echo sequence; SE CMR, spin echo cardiovascular magnetic resonance; SSFP, steady state free precession.

the use of metallic hemostatic clips for ITA grafts and the proximity of sternal wires to the graft course. Also, because arterial grafts are smaller and more tortuous vessels, they are harder to image than the much larger venous grafts.

BYPASS GRAFT ANATOMIC IMAGING TECHNIQUES Conventional Spin Echo and Gradient Echo Imaging The assessment of saphenous vein aortocoronary bypass graft patency has been a relatively early indication for CMR studies. Several groups have reported the feasibility of visualizing graft patency using conventional ECGtriggered multislice spin echo techniques.20–23 On spin echo images, patent grafts appear in consecutive imaging planes as conduits with a signal void, whereas stenotic grafts with slow flow or occluded grafts appear with intermediate signal intensity (Fig. 24-1). With X-ray angiography as the method of reference, the sensitivity of CMR in predicting graft patency ranged from 90% to 98% with a specificity of 72% to 90%. Using conventional gradient echo non-breath hold CMR, with a relatively long echo time (TE) and repetition time (TR), the sensitivity was in the same order of magnitude (88% to 98%), with a somewhat higher specificity (86% to 100%).24–26 On gradient echo images blood within patent grafts appears bright (Fig. 24-2). On spin echo images, signal voids from metal clips, stents, calcifications, and thickened pericardium can falsely mimic graft patency. These artifacts are larger and more easily recognized on gradient echo images than on spin echo images, thus decreasing the number of false positive patent grafts (increased specificity). 330 Cardiovascular Magnetic Resonance

On the other hand, one might expect the number of false positive occlusions to increase (decreased sensitivity).

Two-Dimensional Breath Hold Cardiovascular Magnetic Resonance Angiography This technique was first described for imaging of coronary arteries, but it can also be applied for visualization of bypass grafts.27 Within a breath hold of 16 to 20 heartbeats, a segmented gradient echo image is acquired with 4- or 5-mm slice thickness and an in-plane resolution of approximately 1.0 1.4 mm using a surface coil. To cover the three-dimensional (3D) course of a bypass graft, repetitive breath holding is necessary, which makes the technique highly patient-dependent. Especially when multiple grafts or sequential grafts are to be evaluated, the procedure is time-consuming. Hartnell and colleagues reported a high susceptibility to metallic materials used during CABG, which significantly affected the diagnostic accuracy.28 Post and colleagues used this technique to evaluate sequential grafts (Fig. 24-3).29 High sensitivity and specificity were found for predicting patency of proximal, mid, and distal graft segments to the branches of the RCA, but the accuracy decreased in distal segments to branches of the LAD and was poor in distal segments to the circumflex system. Another 2D breath hold approach uses a multislice halfFourier acquisition single-shot turbo spin echo sequence (HASTE). Seven T2-weighted images are generated within a breath hold of approximately 14 seconds, with a 1.3 1.4 mm in-plane resolution and 5-mm slice thickness. Using this technique, Kalden and colleagues reported a sensitivity and specificity of 95% and 93%, respectively, in predicting graft patency.30 Also a high percentage (83%) of distal graft anastomoses were revealed. Because the sequence is less

A

B

1 2

2

1

C

D

Figure 24-1 Series of four transverse spin echo images of a sequential vein graft to the obtuse marginal branch of the circumflex artery (1) and a second vein graft to the left anterior descending artery (2). Patent grafts show low signal intensity. A is the most superior image demonstrating the circumflex graft originating from the ascending aorta and overriding the main pulmonary artery. B shows the more descending course of the CX graft left of the pulmonary artery. C demonstrates the LAD graft originating from the more proximal part of the ascending aorta, whereas D indicates the course of both grafts at the level of the left atrium.

susceptible to artifacts induced by metallic implants than are gradient echo sequences, a remarkably high accuracy was found in detecting arterial graft patency, that is, a sensitivity of 90% and a specificity of 100%. However, these figures must be interpreted with care, since the pretest probability of occluded arterial grafts is generally low. In that study, it is unclear how many of the 14 reported ITA grafts were occluded. Personal experience obtained with the HASTE sequence indicates an excellent visualization of coronary arteries and bypass grafts but rather poor detection of disease (Fig. 24-4). Accordingly, Kalden and colleagues noted that only two out of eight hemodynamically significant graft stenoses were detected by using HASTE.

Three-Dimensional Respiratory Gated Cardiovascular Magnetic Resonance Angiography A 3D dataset of truly contiguous slices may be obtained with respiratory gated gradient echo techniques. Gating to the respiratory cycle is achieved by using navigators that monitor the diaphragm position.31 Magnetic resonance (MR) data acquired within a preset acceptance-window of

the respiration-induced diaphragm excursion are used for image reconstruction. The patient is allowed to breathe freely without the need for repetitive breath holding, at the expense of an increase in imaging time. Several refinements of this gating procedure have been developed for imaging of the coronary arteries.32 Kessler and colleagues used respiratory gated 3D MR angiography (MRA) with navigator guiding and retrospective MR data processing for imaging of bypass grafts.33 In a relatively small number of 19 grafts, 13 out of 15 patent grafts and 4 out of 4 occluded grafts were correctly classified. A study by Langerak and colleagues of 38 post-CABG patients (56 venous grafts) who underwent conventional angiography for recurrent angina, assessed the accuracy of contemporary high-resolution respiratory gated 3D MRA in detecting vein graft disease.34 The sequence used in the study included a T2-preparation prepulse (for muscle suppression), a fat suppression prepulse and had an acquisition window of 71 msec placed in mid-diastole and at end expiration. Fifty out of 56 grafts were adequately visualized with a spatial resolution of 0.7 1.0 1.5 mm. The study reported sensitivities of 73% and 83%, with specificities of 80% to 87% and 98% to 100% in detecting grafts with 70% diameter stenosis and complete occlusion, respectively. (The study excluded patients with graft stents.)

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RCA

A

P

LAD

CX

Figure 24-2 Transverse image obtained with gradient echo technique. Cross sections of patent grafts to the right coronary artery (RCA), left anterior descending coronary artery (LAD), and circumflex coronary artery (LCX) are indicated and demonstrate a high signal intensity. Source: Van Rossum AC, Galjee MA, Post JC, Visser CA. A practical approach to CMR of coronary artery bypass graft patency and flow. Int J Cardiac Imaging. 1997;13:199–204, with permission from Kluwer Academic Publishers.

Three-Dimensional ContrastEnhanced Breath Hold Cardiovascular Magnetic Resonance Angiography Breath hold contrast-enhanced MRA is a relatively new technique, first applied in imaging of the aorta.35 The T1-shortening effect of the contrast agent on the blood allows obtaining high vascular contrast, using short TR/ TE gradient echo sequences with high field-strength gradient systems. After intravenous injection of an interstitial contrast agent, preferably with the aid of an MR-compatible power injector, a 3D spoiled gradient echo sequence with short TR/TE (4.4/1.4 msec or even shorter) is applied. During a breath hold of approximately 30 seconds, a 3D volume slab of 24 to 32 contiguous slice partitions with a total thickness of 6 to 9 cm is imaged. Prior to acquiring the 3D volume slab, a single-slice 2D turbo gradient echo sequence is used to time the arrival in the aorta of a contrast agent test bolus. To maximize the contrastenhancing effect, acquisition of the central k-space lines of the 3D imaging data is set to coincide with peak contrast arrival. This is achieved by introducing a time delay between contrast injection and start of the imaging sequence (delay ¼ arrival time one-half or one-third of acquisition time). A spatial in-plane resolution of 1 1.5 mm with 2- to 3-mm section thickness is achieved depending on the field of view and matrix size, number of partitions, and slab thickness. Each partition of the 332 Cardiovascular Magnetic Resonance

3D acquisition yields a “source” image. Studies are evaluated by reviewing the source images and after postprocessing techniques such as maximum intensity projection and planar or curved reformatting (Figs. 24-5 and 24-6).19 Evaluation of aortocoronary bypass grafts using 3D contrast-enhanced MRA has been reported without and with ECG triggering.30,36,37 Sensitivity, specificity, and accuracy for predicting graft patency varied between 93% and 95%, 81% and 97%, and 92% and 95%, respectively (see Table 24-1). A lower specificity was found in the non-ECG-triggered study.36 Theoretically, one would expect ECG-triggered acquisitions to be superior for visualizing the sites of graft insertion on native coronary arteries. Although Kalden and colleagues specifically addressed this issue, they succeeded in pursuing distal anastomoses in only 64% of case.30 This might have been caused partly by the relatively long data collection window of 560 msec within each cardiac cycle, leading to residual blurring due to cardiac motion. The use of a shorter acquisition window of 120 msec, used for imaging coronary arteries of healthy volunteers,38 has not been implemented yet in imaging bypass grafts. Bunce and colleagues compared 3D contrast-enhanced MRA to the recently developed multislice balanced steady-state free precession (SSFP) sequence for the assessment of 56 venous and 23 arterial bypass grafts patency in 25 patients.39 SSFP had a temporal resolution of 336 msec with a spatial resolution of 2.7 1.4 4.0 mm, and 3D contrast-enhanced MRA had a temporal resolution of 440 msec with spatial resolution of 1.6 1.6 2.0 mm. The two methods had similar accuracy (78% versus 84%, p ¼ nonsignificant) for detection of coronary artery bypass graft patency, but there was a trend toward more false positive findings for occlusion and reduced visualization of arterial grafts with SSFP angiography.39

IMAGING STRATEGY The MR imaging strategy and image interpretation are facilitated when the surgical report is known prior to the MR procedure with respect to the number of grafts and insertion sites. Grafts descending to the perfusion area of the LCX, including anterolateral and obtuse marginal branches, generally originate most superiorly from the ascending aorta (Fig. 24-7).27 The graft to the perfusion area of the LAD, including diagonal branches, originates inferiorly compared with most LCX grafts. Both types of grafts cross the pulmonary artery trunk in a left lateral and inferior course. Then the LCX graft proceeds posteriorly and inferiorly, while the LAD graft goes left and anteriorly. Grafts to the posterior descending artery of the RCA generally have the lowest origin from the ascending aorta and run inferiorly on the lateral side of the right atrium. For assessment of patency and measurement of flow within a graft, CMR can best be performed at the proximal part of the graft. At this level, most grafts have a straight course, are least susceptible to motion, and are easily distinguished from native coronary vessels. The surface coil must be centered at the middle of the sternum, which is somewhat more superior than in a standard cardiac study.

24 CORONARY ARTERY BYPASS GRAFT IMAGING AND ASSESSMENT OF FLOW

A

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C Figure 24-3 Segmented gradient echo technique, 1 image per breath hold of 16 heartbeats. Patient with sequential graft from the aorta to the diagonal branch of the left anterior descending coronary artery (LAD) and obtuse-marginal branch of the left anterior descending coronary artery and left internal thoracic artery graft to the LAD. The tall arrow indicates bright segments of the sequential graft. The short arrow points at the dark signal loss of the ITA clip artifact. Source: Van Rossum AC, Bedaux WLF, Hofman MBM. Morphologic evaluation of coronary artery bypass conduits. J Magn Reson Imaging. 1999;10:734–740.

With the introduction of k-space segmented sequences and phased-array coils, images can now be obtained within a breath hold. Thus, respiratory motion is minimized and image quality is increased. Because the course of a graft is more or less predictable, the choice of imaging planes can be fairly well standardized. An effective approach is first to acquire a set of multislice transverse images covering the ascending aorta and superior part of the heart. In our experience, one or two breath hold multislice series using the HASTE sequence will be most informative (see Fig. 24-4). The images just superior of the pulmonary artery trunk will demonstrate in-plane views of proximal parts of LCX and LAD grafts, whereas the images at a lower level show cross sections of grafts to the left and right coronary arteries. Once the proximal course of the grafts has been localized, one may proceed with high-resolution single or multislice 2D imaging, in orientations following the more distal course of the grafts. The advantages of this approach are the short image reconstruction time and the immediate availability of the images. The disadvantage is that it requires patient cooperation and may be subject to misregistration due to inconsistent breath holding. Alternatively, 3D MRA techniques may be used to acquire larger imaging slabs covering the course of the grafts. The resolution is higher, and the acquisition

time is shorter. However, the reconstruction time is longer than in 2D techniques, and interpretation requires some form of postprocessing. Notwithstanding these limitations, the 3D approaches are likely to become the first choice with improving technology.

CMR QUANTIFICATION OF GRAFT FLOW AND FLOW RESERVE Flow velocity and volume flow in bypass grafts can be measured by applying velocity-encoded phase-contrast cine CMR sequences, thus allowing the assessment of graft function in addition to a morphologic evaluation (Fig. 24-8).40 Galjee and colleagues demonstrated that adequate velocity profiles throughout the cardiac cycle could be obtained in 85% of angiographically patent vein grafts, using nonbreath-hold MR velocity mapping.26 Graft flow velocity was characterized by a biphasic pattern, with one peak in systole and a second peak in diastole (Fig. 24-9).26 Similar findings have been obtained by using invasive Doppler guidewire approaches and transthoracic Doppler echocardiography.8,15 However, there are differences in resting Cardiovascular Magnetic Resonance 333

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Figure 24-4 Half-Fourier acquisition single-shot turbo spin echo (HASTE) sequence, four out of five images obtained in a single breath hold (A, B, C, D). Patient with single vein graft to the right coronary artery (1), sequential vein graft to the first diagonal branch of the left anterior descending coronary artery (LAD) and obtuse marginal branch of the circumflex coronary artery (2), and the left internal thoracic artery to the LAD (3). In panels A and B, the native left main coronary artery, the LAD, and the great cardiac vein are visualized. In panels C and D, the right coronary artery is seen inferior to the graft in the right atrioventricular groove.

phasic flow patterns between in situ IMA conduits and saphenous vein grafts that may have implications in the development of degenerative graft disease and long-term conduit patency.8 Thoracic artery grafts show a gradual longitudinal transition from predominantly systolic velocity proximally to predominantly diastolic velocity distally, while saphenous vein graft flow velocity patterns have a consistently diastolic predominance, both proximally and distally.8 CMR assessed volume flow of grafts with multiple sequential target vessel anastomoses significantly differed from single grafts.26 Sequential venous grafts have higher volume flow and velocities compared to single grafts, probably owing to a lower downstream resistance in sequential grafts. Volume flow and velocity can both be analyzed on CMR velocity maps. In the case of volume flow, the mean velocity is multiplied with the vessel area and all pixels over that area are included in the analysis. However, errors due to partial volume effects at the lumen edges may occur, and the analysis is more time consuming. Velocity analysis that measures velocity in the center of the lumen (over a 2 2 pixel area) is not affected by partial volume effects and 334 Cardiovascular Magnetic Resonance

is less time-consuming. Head-to-head comparisons have shown similar diagnostic accuracies in detection of significant (>70%) vein graft stenosis between volume flow and velocity analysis.41 Several studies, using breath hold segmented k-space sequences at baseline and during pharmacologic stress, have demonstrated the feasibility of measuring flow and flow velocity reserve in saphenous vein grafts and its potential to differentiate between patent and stenotic bypass grafts or runoff vessels.41–47 A report by Voigtla¨nder and colleagues demonstrated a flow velocity reserve of 2.6 1.5 versus 0.8 0.4 (p < 0.005) and a flow reserve of 2.9 1.9 versus 1.2 0.5 (p < 0.05) in 21 grafts without stenoses compared to 6 grafts with greater than 75% diameter stenoses.42 In a study by Bedaux and colleagues of 21 post-CABG patients (40 grafts: 18 single and 22 sequential) who underwent selective angiography due to recurrent angina pectoris, the combination of a basal, CMR-derived proximal segment graft flow volume of less than 20 mL/ min or a flow reserve of less than 2 had a sensitivity of 78% and a specificity of 80% in detecting graft stenosis or runoff vessel disease.43 Owing to the limited number

24 CORONARY ARTERY BYPASS GRAFT IMAGING AND ASSESSMENT OF FLOW

Figure 24-5 Gadolinium-enhanced three-dimensional breath hold magnetic resonance angiography, ECG-triggered. The six images on the left are several of the source images, demonstrating the proximal course of two sequential vein grafts to the left anterior descending coronary artery and circumflex coronary artery perfusion territory, respectively. The right image is obtained through postprocessing, using a maximum-intensity projection.

of patients, no differentiation in volume flow was possible between single and sequential grafts. A larger study by Langerak and colleagues of 69 post-CABG patients (166 grafts: 81 single vein, 44 sequential vein, and 41 arterial grafts) with recurrent angina who underwent CMR and selective angiography reported successful baseline and adenosine stress CMR flow scans in 80% of the grafts.44 The majority of unsuccessful flow scans (24 out of 26 grafts) were attributed to adenosine-related side effects. Peak velocities of CMR velocity maps were analyzed for single and sequential vein grafts separately, and a model using receiver operating characteristic (ROC) curve analysis and logistic regression to detect stenosis of 50% or more or 70% or more in the graft or runoff vessel was developed. Optimal cutoff points to differentiate single vein grafts with or without stenosis of 70% or more were 13.58 cm/sec for stress average peak velocity (sensitivity: 91%, specificity: 62%), 20.86 cm/ sec for stress diastolic peak velocity (sensitivity: 91%, specificity: 74%), and 1.43 for coronary velocity reserve (sensitivity: 91%, specificity: 78%). Higher cutoff points for stress average peak velocity (22.47 cm/sec; sensitivity 82%, specificity 62%) and diastolic peak velocity (38.67 cm/sec; sensitivity 86%, specificity 62%) were found to detect sequential vein grafts with 70% stenosis or more compared to

single vein grafts.44 A retrospective study by Salm and colleagues45 including 68 post-CABG patients (125 saphenous venous grafts) compared the method used by Bedaux and colleagues43 (CMR basal flow rate of <20 mL/min or a flow reserve <2) to the model described by Langerak and colleagues.44 The two methods demonstrated equivalent sensitivities (70% versus 74%), while the model using ROC curve analysis with logistic regression demonstrated a much higher specificity (68% versus 30%) for the detection of significant (50%) venous graft or target vessel stenosis.45 CMR has been compared with single-photon emission perfusion computed tomography (SPECT) for the evaluation of the hemodynamic significance of angiographic findings in bypass grafts. In a study including 25 post-bypass patients, flow velocity CMR was successful in 46 out of 57 grafts and demonstrated results similar to those of SPECT in 80% (k ¼ 0.61) of those grafts.46 The effects of percutaneous angioplasty on saphenous vein grafts function can be assessed by using flow CMR. In a small study of 15 venous grafts before and at least 6 weeks after percutaneous angioplasty, CMR has demonstrated a significant improvement in baseline average peak velocity (before: 9.2 6.6 cm/sec; after: 12.9 7.9 cm/sec; p ¼ 0.008) and stress average peak velocity (before: 12.9 6.3 cm/sec; after: 27.1 Cardiovascular Magnetic Resonance 335

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Figure 24-6 Patient with single vein graft inserting into the left anterior descending coronary artery (LAD) and sequential vein graft inserting on posterior descending and posterolateral branch of the right coronary artery. A, Curved planar reformat of gadoliniumenhanced three-dimensional magnetic resonance angiography of graft inserting on the LAD. The dashed line indicates the orientation of the plane perpendicular to the proximal graft segment, used for velocity-encoded cine cardiac magnetic resonance (CMR). B, X-ray angiography of distal graft segment demonstrating irregular aspect of luminal borders and tight insertion on LAD. C, Cross-sectional averaged velocities measured at multiple phases throughout the cardiac cycle using velocity-encoded cine CMR. Calculated volume-flow was 38 mL/min at rest and 68 mL/min during adenosine stress, yielding a flow reserve of 1.8. D, Curved planar reformat of proximal part of the sequential graft. The dashed line indicates the plane of CMR flow acquisition. E, X-ray angiography reveals a tight stenosis (arrow) in the graft segment between the two distal insertion sites. F, Volume flow calculated from MR flow velocity measurements was 50 mL/ min at rest and 51 mL/min during adenosine stress, yielding a flow reserve of 1.0. Source: Van Rossum AC, Bedaux WLF, Hofman MBM. Morphologic evaluation of coronary artery bypass conduits. J Magn Reson Imaging. 1999;10:734–740.

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13.9 cm/sec; p < 0.001).47 However, there was no improvement in velocity reserve, a result that is explained by persistent impairment of stress average peak velocity and not by an increased baseline peak velocity after intervention.47 By using non-breath-hold and breath hold techniques, functional results were also obtained in IMA grafts, despite problems with imaging artifacts due to metallic clips and sternal wires.48,49 The diastolic/systolic peak velocity ratio was found to be higher in IMA grafts than in native IMA. 336 Cardiovascular Magnetic Resonance

Findings in IMA grafts by Kawada and colleagues in 18 patients without and 5 patients with 75% or greater stenosis, indicate a higher sensitivity and specificity for detection of IMA graft stenosis using measurements of the mean flow rate and diastolic/systolic peak flow ratio at rest than using the flow reserve after administration of dipyridamole.50 A study by Ishida and colleagues including 26 patients with IMA grafts (with CMR flow performed successfully in the middle segment of 24 grafts) demonstrated similar sensitivities of 86% and specificities of 88% and 94% in detecting

CX graft RCA graft

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Sagittal plane Figure 24-7 Typical course of vein grafts from ascending aorta to coronary artery insertion sites, relative to sagittal and transverse imaging planes. CX, circumflex; LAD, left anterior descending coronary artery; RCA, right coronary artery. Source: Galjee MA, van Rossum AC, Doesburg T, van Eenige MJ, Visser CA. Value of magnetic resonance imaging in assessing patency and function of coronary artery bypass grafts: an angiographically controlled study. Circulation. 1996;93:660–666.

greater than 70% diameter stenosis for mean diastolic/systolic peak velocity ratio and mean baseline blood flow, respectively. No significant difference was observed in the flow reserve ratio between grafts with more than 70% stenosis and those without significant stenosis.51 CMR flow variables such as mean graft blood flow, diastolic/systolic flow profiles, and flow velocity reserve of bypass grafts are useful in the noninvasive assessment of graft and runoff vessel disease, and CMR could become an alternative to other methods for functional graft assessment such as SPECT or Doppler wire.

So far, CMR anatomic imaging of coronary artery bypass grafts has been limited, owing to the constraints of spatial resolution to demonstrating vessel patency versus total occlusion only. CMR measurements of blood flow at rest and under stress with calculation of the diastolic/systolic flow ratio and the coronary flow reserve have become helpful in addition to anatomic imaging to determine the functional status of a diseased bypass graft and its recipient coronary artery. Furthermore, most of the reported studies have been confined to imaging of proximal graft segments, and only few data are available with respect to assessing patency of segments beyond the first coronary anastomosis in sequential bypass grafts. Imaging of the distal graft segments requires a higher spatial resolution and signal-tonoise ratio than those obtained with currently available techniques. Another point of consideration is the problem associated with CMR of arterial grafts. IMA grafts are increasingly used because of the improved long-term survival but have been excluded from most CMR studies, owing to metallic clip and sternal wires artifacts. Flow reserve measurements proximal to the clip artifacts may be of diagnostic help or, alternatively, methods to obtain hemostasis without use of ferromagnetic clips might be useful. The anatomic imaging of bypass grafts after percutaneous angioplasty with stent deployment is limited by stent induced artifacts (most stents are made of surgical steel or tantalum with the latter demonstrating less signal void artifacts), although measurement of flow is possible in the nonstented regions.47,52 Integration of miniature radiofrequency coils into stents (transforming them into intravascular antennas coupled to a surface coil) allows the visualization of stent lumen.53 CMR-lucent stents are also under development.

CX LAD

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Figure 24-8 Oblique-sagittal image obtained with velocity-encoded phase-contrast technique to measure flow. A, Magnitude reconstruction of magnetic resonance (MR) signal depicting cross sections of two grafts anterior and superior of the main pulmonary artery. B, Corresponding reconstruction of the phase of the MR signal, where the brightness of each pixel is proportional to the flow-velocity and midgray equals zero flow velocity. Within the grafts, the bright signal (white) indicates high cross-sectional velocities. CX, circumflex; LA, left atrium; LAD, left anterior descending coronary artery; P, pulmonary artery. Source: Van Rossum AC, Galjee MA, Post JC, Visser CA. A practical approach to CMR of coronary artery bypass graft patency and flow. Int J Cardiac Imaging. 1997;13:199–204. Cardiovascular Magnetic Resonance 337

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LIMITATIONS

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Figure 24-9 Plots of volume-flow patterns in single and sequential saphenous vein grafts, normalized for heart rate. Typical patterns consist of a first flow peak in systole and a second peak in diastole. Source: Galjee MA, van Rossum AC, Doesburg T, van Eenige MJ, Visser CA. Value of magnetic resonance imaging in assessing patency and function of coronary artery bypass grafts: an angiographically controlled study. Circulation. 1996;93:660–666.

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However, even a clear demonstration of graft-segment patency or narrowing will often not suffice for clinical decision making. In most patients, there is a need to also know the status of the native coronary arteries. Narrowing of the recipient coronary artery may have developed beyond the anastomosis of a patent graft segment, and progression of disease in other native coronary arteries must be excluded. Thus, CMR does not eliminate the need for conventional X-ray coronary angiography when coronary reinterventions are under consideration.

INDICATIONS A clinical indication for CMR of bypass grafts may therefore exist only in patients in whom there is no immediate need of knowing the status of the native coronary arteries. Such a category consists, for example, of patients with chest pain shortly after CABG surgery. Also late after CABG surgery, information about graft patency and function might be helpful in deciding whether to postpone coronary angiography in patients with ambiguous anginal complaints. Noninvasive monitoring of flow parameters may then be useful to detect a gradual increase of graft stenosis and decide to proceed with conventional angiography and percutaneous intervention, before the onset of a total occlusion. Similarly, flow measurements could be used after bypass graft percutaneous angioplasty for the early detection of restenosis. Furthermore, CMR can be used as a screening procedure before angiography, providing a road map for the grafts to be visualized. This might considerably shorten the angiographic procedure with less exposure to ionizing radiation and lower total iodine contrast dose. A useful 338 Cardiovascular Magnetic Resonance

indication also appears to be the assessment of patency of grafts that are not visualized during conventional angiography. Although this often indicates a proximal occlusion of the graft, failure of the catheter to find an aortic graft anastomosis may result in a false diagnosis of graft occlusion. In cases of doubt, MRA will rapidly confirm or discard this diagnosis. Also, when angiography has demonstrated a graft stenosis, it can be helpful for further management to assess the flow reserve of a graft. There are also occasions when CMR is useful for the definition of the complications of vein grafting, such as aneurysm of the graft, and assisting in surgical management.54 Finally, vein graft imaging has a role in research of bypass techniques and has been used to demonstrate the efficacy of aprotinin55 during surgery and in comparison of techniques for vein graft stripping,56 by demonstrating patency rates without the need for invasive angiography, an important ethical and safety issue. Notwithstanding these indications, the majority of patients after CABG require evaluation with respect to percutaneous or surgical reintervention. Unless CMR will be able to provide more detailed information regarding the status of native coronary arteries, a wide application of the technique is unlikely to occur. Continuing improvement in CMR of coronary arteries and bypass grafts may be expected with new developments in scanner hardware and software and the use of contrast agents.

CONCLUSION There is clear evidence that conventional spin echo and gradient echo CMR is capable of assessing patency of coronary artery vein grafts. With more recently introduced

noninvasive monitoring of bypass graft narrowing progression before and even after percutaneous interventions.47 However, apart from a few exceptions, the majority of the patients undergo graft evaluation because they are considered for a reintervention by angioplasty or coronary artery bypass graft surgery. In these cases, information on the status of the native coronary arteries is required. A broader clinical use of CMR in the evaluation of patients with coronary artery bypass grafts may therefore be expected only with further improvement in CMR techniques for coronary angiography.

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to the left anterior descending coronary artery. Am J Cardiol. 1999;83:1157–1163. Galjee MA, van Rossum AC, Doesburg T, Hofman MBM, Falke THM, Visser CA. Quantification of coronary artery bypass graft flow by magnetic resonance phase velocity mapping. Magn Reson Imaging. 1996;14:485–493. Van Rossum AC, Bedaux WLF, Hofman MBM. Morphologic evaluation of coronary artery bypass conduits. J Magn Reson Imaging. 1999;10:734–740. White RD, Caputo GR, Mark AS, Modin GW, Higgins CB. Coronary artery bypass graft patency: noninvasive evaluation with MR imaging. Radiology. 1987;164:681–686. Rubinstein RI, Askenase AD, Thickman D, Feldman MS, Agarwal JB, Helfant RH. Magnetic resonance imaging to evaluate patency of aortocoronary bypass grafts. Circulation. 1987;76:786–791. Jenkins JPR, Love HG, Foster CJ, Isherwood I, Rowlands DJ. Detection of coronary artery bypass graft patency as assessed by magnetic resonance imaging. Br J Radiol. 1988;61:2–4. Frija G, Schouman-Claeys E, Lacombe P, Bismuth V, Ollivier JP. A study of coronary artery bypass graft patency using MR imaging. J Comput Assist Tomogr. 1989;13:226–232. White RD, Pflugfelder PW, Lipton MJ, Higgins CB. Coronary artery bypass grafts: evaluation of patency with cine MR imaging. AJR Am J Roentgenol. 1988;150:1271–1274. Aurigemma GP, Reichek N, Axel L, Schiebler M, Harris C, Kressel HY. Noninvasive determination of coronary artery bypass graft patency by cine magnetic resonance imaging. Circulation. 1989; 80:1595–1602. Galjee MA, van Rossum AC, Doesburg T, van Eenige MJ, Visser CA. Value of magnetic resonance imaging in assessing patency and function of coronary artery bypass grafts: an angiographically controlled study. Circulation. 1996;93:660–666. Manning WJ, Li W, Boyle NG, Edelman RR. Fat-suppressed breathhold magnetic resonance coronary angiography. Circulation. 1993;87:94–104. Hartnell GG, Cohen MC, Charlamb M, Saouaf R, Buff B, Finn JP. Segmented k-space magnetic resonance angiography for the detection of coronary artery bypass graft patency. Book of Abstracts of the 4th Meeting of the ISMRM, New York. 1996;1:178. Post JC, van Rossum AC, Bronzwaer JGF, de Cock CC, Visser CA. Magnetic resonance angiography of sequential aortocoronary bypass grafts. Circulation. 1997;96(Suppl):1–133. Kalden P, Kreitner KF, Wittlinger T, et al. Assessment of coronary artery bypass grafts: value of different breath-hold MR imaging techniques. Am J Roentgenol. 1999;172:1359–1364. Hofman MBM, Paschal CB, Li D, Haacke EM, van Rossum AC, Sprenger M. CMR of coronary arteries, 2D breath-hold versus 3D respiratory-gated acquisition. J Comput Assist Tomogr. 1995;19:56–62. Danias PG, McConnell MV, Khasgiwala VC, Chuang ML, Edelman RR, Manning WJ. Prospective navigator correction of image position for coronary MR angiography. Radiology. 1997;203:733–736. Kessler W, Achenbach S, Moshage W, et al. Usefulness of respiratory gated magnetic resonance coronary angiography in assessing narrowings > 50% in diameter in native coronary arteries and in aortocoronary bypass conduits. Am J Cardiol. 1997;80:989–993. Langerak SE, Vliegen HW, de Roos A, et al. Detection of vein graft disease using high-resolution magnetic resonance angiography. Circulation. 2002;105(3):328–333. Cardiovascular Magnetic Resonance 339

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breath hold 2D and contrast-enhanced 3D techniques, the predictive accuracy has further improved, with sensitivities and specificities in the 90% range. Limitations arise with regard to assessing obstructive disease and evaluating distal segments of sequential grafts, owing to insufficient spatial resolution, low signal-to-noise ratios, and cardiac motion. Imaging of arterial grafts is complicated by the sternal wire and metallic clip artifacts. Recently, significant progress has been made in the assessment of graft flow patterns and flow reserve using velocity-encoded cine CMR. Clinically, these functional measurements may become of use for the

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35. Prince MR, Narasimham DL, Stanley JC, et al. Breath-hold gadolinium-enhanced MR angiography of the abdominal aorta and its major branches. Radiology. 1995;197:785–792. 36. Wintersperger BJ, Engelmann MG, von Smekal A, et al. Patency of coronary artery bypass grafts: assessment with breath-hold contrastenhanced MR angiography – Value of a non-electrocardiographically triggered technique. Radiology. 1998;208:345–351. 37. Vrachliotis TG, Bis KG, Aliabadi D, Shetty AN, Safian R, Simonetti O. Contrast-enhanced breath-hold MR angiography for evaluating patency of coronary artery bypass grafts. Am J Roentgenol. 1997;168: 1073–1080. 38. Goldfarb JW, Edelman RR. Coronary arteries: breath-hold, gadolinium-enhanced, three-dimensional MR angiography. Radiology. 1998;206:830–834. 39. Bunce NH, Lorenz CH, John AS, Lesser JR, Mohiaddin RH, Pennell DJ. Coronary artery bypass graft patency: assessment with true fast imaging with steady-state precession versus gadoliniumenhanced MR angiography. Radiology. 2003;227(2):440–446. 40. Van Rossum AC, Galjee MA, Post JC, Visser CA. A practical approach to CMR of coronary artery bypass graft patency and flow. Int J Cardiac Imaging. 1997;13:199–204. 41. Salm LP, Langerak SE, Vliegen HW, et al. Blood flow in coronary artery bypass vein grafts: volume versus velocity at cardiovascular MR imaging. Radiology. 2004;232(3):915–920. 42. Voigtla¨nder T, Kreitner KF, Wittlinger T, Scharhag J, Abegunewardene N, Meyer J. MR measurement of flow reserve in coronary grafts [abstract]. J Cardiovasc Magn Reson. 1999;1:275. 43. Bedaux WLF, Hofman MBM, Vyt SLA, Bronzwaer JGF, Visser CA, Van Rossum AC. Assessment of coronary artery bypass graft disease using cardiovascular magnetic resonance determination of flow reserve. J Am Coll Cardiol. 2002;40(10):1848–1855. 44. Langerak SE, Vliegen HW, Jukema JW, et al. Value of magnetic resonance imaging for the noninvasive detection of stenosis in coronary artery bypass grafts and recipient coronary arteries. Circulation. 2003;107(11):1502–1508. 45. Salm LP, Vliegen HW, Langerak SE, et al. Evaluation of saphenous vein coronary artery bypass graft flow by cardiovascular magnetic resonance. J Cardiovasc Magn Reson. 2005;7(4):631–637.

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46. Salm LP, Bax JJ, Vliegen HW, et al. Functional significance of stenoses in coronary artery bypass grafts: evaluation by single-photon emission computed tomography perfusion imaging, cardiovascular magnetic resonance, and angiography. J Am Coll Cardiol. 2004;44 (9):1877–1882. 47. Langerak SE, Vliegen HW, Jukema JW, et al. Vein graft function improvement after percutaneous intervention: evaluation with MR flow mapping. Radiology. 2003;228(3):834–841. 48. Debatin JF, Strong JA, Sostman HD, et al. MR characterization of blood flow in native and grafted internal mammary arteries. J Magn Reson Imaging. 1993;3:443–450. 49. Sakuma H, Globits S, O’Sullivan M, et al. Breath-hold MR measurements of blood flow velocity in internal mammary arteries and coronary artery bypass grafts. J Magn Reson Imaging. 1996;6: 219–222. 50. Kawada N, Sakuma H, Cruz BC, et al. Noninvasive detection of significant stenosis in the coronary artery bypass grafts using fast velocity-encoded cine CMR [abstract]. J Cardiovasc Magn Reson. 1999;1:261. 51. Ishida N, Sakuma H, Cruz BP, et al. MR flow measurement in the internal mammary artery-to-coronary artery bypass graft: comparison with graft stenosis at radiographic angiography. Radiology. 2001;220 (2):441–447. 52. Nagel E, Thouet T, Klein C, et al. Noninvasive determination of coronary blood flow velocity with cardiovascular magnetic resonance in patients after stent deployment. Circulation. 2003;107(13): 1738–1743. 53. Quick HH, Ladd ME. Interventional MRA: concepts for active visualization of catheters and stents. Z Med Phys. 2003;13(3):188–192 (German). 54. Warner OJ, Ohri SK, Pennell DJ, Smith PLC. Magnetic resonance coronary artery imaging for redo cardiac surgery. Ann Thor Surg. 1996;62:1513–1516. 55. Bidstrup BP, Underwood SR, Sapsford RN. Effect of aprotinin (Trasylol) on aorto-coronary bypass graft patency. J Thorac Cardiovasc Surg. 1993;105:147–153. 56. O’Regan DJ, Borland JAA, Chester AH, Pennell DJ, Yacoub M, Pepper JR. Assessment of human long saphenous vein function with minimally invasive harvesting with the Mayo stripper. Eur J Cardiothor Surg. 1997;12:428–435.

Atherosclerotic Plaque Imaging: Aorta and Carotid James H. F. Rudd, Fabien Hyafil, Silvia Aguiar, and Zahi A. Fayad

Although death rates in the Western world have been consistently falling since the 1980s, atherosclerosis is now raging throughout the developing world.1 In 1996, more than 60% of the 15 million worldwide deaths due to circulatory diseases occurred in less developed nations. Atherosclerosis is an inflammatory disease that affects medium and large arteries beginning in the first decade of life. It has a predilection for certain arterial beds, the carotid artery and aorta being the most common sites of plaque formation.2 In the carotid circulation, the end result can be an ischemic cerebral insult causing either temporary (transient ischemic attack) or permanent (stroke) symptoms. If unchecked, aortic atherosclerosis can predispose to dissection, intramural hemorrhage, aneurysm formation, and downstream embolus.3

PATHOBIOLOGY OF ATHEROSCLEROSIS Atherosclerosis is characterized by the gradual accumulation of lipid, inflammatory cells, and connective tissue within the arterial wall. It is a chronic, progressive disease with a very long asymptomatic phase. The first abnormality is the fatty streak, caused by the collection of lipid and macrophages in the subendothelial space. Fatty streaks develop primarily in regions of endothelial dysfunction. Endothelial cells in these areas have a reduced production of nitric oxide as a result of being exposed to low wall shear stress.4 The major atherogenic risk factors, including cigarette smoking, hypertension, and diabetes mellitus all impair endothelial function.5 Both the barrier function and secretory capacity of the endothelium become affected. This manifests as an increase in permeability to blood-derived lipids and inflammatory cells. Once within the subendothelial space low-density lipoprotein (LDL) becomes oxidized and attracts monocytes by triggering the release of monocyte chemoattractant protein-1 from the overlying endothelial cells.6 Endothelial adhesion molecules, including vascular cell adhesion molecule-1 (VCAM), intercellular adhesion molecule-1, Eselectin, and P-selectin, facilitate the internalization of further monocytes. Once they have escaped the blood pool, monocytes transform into macrophages, and bind and internalize oxidative LDL via their scavenger receptors.7,8 Eventually, the subendothelial accumulation of modified LDL and macrophage-derived foam cells leads to the formation of the atheromatous lipid core.

Under the right conditions (presence of hypertension, diabetes, smoking, and other risk factors), the lipid core may become bounded on its luminal side by an endothelialized fibrous cap consisting of vascular smooth muscle cells (VSMCs) and connective tissue. VSMCs migrate from the arterial media and synthesize extracellular matrix components such as elastin and collagen, producing the fibrous cap. The cap also contains variable numbers of inflammatory cells, most importantly macrophages.9 As the plaque enlarges, the affected artery grows outward (by expansion of the external elastic lamina) so that lumen diameter and therefore blood flow are preserved (a process known as positive remodeling or the “Glagov effect.”10,11 As wall stress increases with outward remodeling, further expansion eventually becomes impossible, and the plaque starts to encroach into the lumen. This leads to symptoms by compromising blood flow. Mature plaques may also become calcified, a process that preferentially affects the intima of the artery. Very advanced plaques will also often be perforated by new blood vessels under the influence of angiogenic factors, a process called neovascularization that is similar to that which occurs within growing tumors. However, these small arteries are structurally fragile and have a tendency for spontaneous hemorrhage, which can destabilize the plaque, leading to clinical syndromes such as acute coronary syndrome.12,13 Mild and moderately obstructive atherosclerotic plaques may remain quiescent for decades.14 However, when they initiate clot formation in the vessel lumen, they can become life-threatening. This may occur as a result of fibrous cap rupture, with consequent exposure of the thrombogenic extracellular matrix of the cap and the tissue factor–rich lipid core to circulating blood. In addition, there can be erosion of the endothelial cell layer overlying the fibrous cap, again potentially leading to intravascular thrombosis. Endothelial erosion accounts for around 30% of plaque rupture events.15 Both forms of plaque disruption invariably lead to local platelet accumulation and activation with subsequent thrombus formation. The cellular and extracellular composition of the plaque is the primary determinant of plaque stability.16 Lesions with a large lipid core, a thin fibrous cap, a preponderance of inflammatory cells, and few VSMCs are at the highest risk of rupture. Inflammatory cells, particularly macrophages, produce metalloproteinase enzymes, which break down the matrix proteins in the fibrous cap. In addition, they secrete inflammatory cytokines, in particular interferon g, which inhibit VSMC proliferation and matrix Cardiovascular Magnetic Resonance 341

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CHAPTER 25

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production. Furthermore, VSMCs in the fibrous cap have a reduced proliferative capacity and a propensity for apoptosis.17 Consequently, inflammation within the plaques tends toward destruction of the fibrous cap and subsequent thrombosis, and there is a dynamic balance within the plaque between macrophages, which promote erosion and rupture of the fibrous cap, and VSMCs, which nourish and repair it. These processes are independent of plaque size. Consequently, small asymptomatic and angiographically minimal plaques can rupture, leading to major morbidity or a fatal clinical event. In contrast, some large, obstructive plaques will produce symptoms such as angina but are relatively stable with very slow progression. There is an urgent need to discriminate “stable” from potentially “unstable” lesions in clinical practice. This chapter will cover the role of cardiovascular magnetic resonance (CMR) imaging in this respect.

Table 25-1 Plaque Appearance on Multicontrast MRI CMR SEQUENCE Plaque Component

T1 Weighted

Proton Density Weighted

T2 Weighted

Recent thrombus Lipid Fibrous Calcium

þ to

to

to

þ

þ þ

to þ

Source: Adapted from Fuster V, Moreno PR, Fayad ZA, Corti R, and Badimon JJ. Atherothrombosis and high-risk plaque: part I: evolving concepts Journal of the American College of Cardiology 2005;46: 937–954.

IMAGING ATHEROSCLEROSIS The rationale behind the use of imaging in atherosclerosis is that the screening of high-risk patients might allow one to identify potentially dangerous (“vulnerable”) plaques before symptoms and complications develop.18 Early knowledge of disease might allow the implementation of pharmacologic therapy to halt or even reverse the process of plaque progression and possibly prevent rupture. The statin class of drugs have proved efficacious in this regard.19,20 Recently, rapid advances in CMR and the clarification of a number of potential cellular and molecular targets have pushed CMR to the top of the list of potential atherosclerosis imaging modalities.21

Multicontrast CMR Atherosclerotic plaques in the carotid artery are ideally suited for imaging with CMR. Their superficial location, without significant motion, presents less of a technical challenge than does imaging of the aorta or the coronary arteries. Multicontrast CMR of the plaque is based on successive high-resolution black-blood fast spin echo CMR sequences, obtained by nulling the signal of the flowing blood through preparatory pulses. Sequences with several different weightings are generally acquired (T1-weighted [T1W], T2-weighted [T2W], and proton-densityweighted [PDW]).22 Analysis of the varying signal intensities in each of these sequences allows reasonable differentiation of plaque components (lipid core, fibrous tissue, hemorrhage, fibrous cap, and degree of calcification) by their different relaxation properties on CMR (Table 251).23,24 Yuan and colleagues demonstrated that multicontrast CMR of human carotid arteries had sensitivity and specificity values of 85% and 92%, respectively, for the identification of a lipid core.25,26 Subsequent advances in CMR acquisition protocols have allowed acquisition times to be reduced manyfold.27 An axial section through the thorax reveals a complex aortic plaque in Figure 25-1. Figure 25-2 demonstrates a carotid plaque imaged at 3 Tesla, with corresponding T1W, T2W, and PDW images displayed. A final example of multicontrast 342 Cardiovascular Magnetic Resonance

Figure 25-1 An example of multicontrast plaque imaging with CMR. This transaxial section through the aorta demonstrates an eccentrically shaped plaque involving the lateral wall of the aorta.

CMR showing different plaque components and automatic segmentation of these plaques using a k-means cluster algorithm is shown in Figure 25-3.28 CMR has been validated against transesophageal echocardiography in the thoracic aorta, with a strong correlation shown between mean maximum plaque thickness measured by both modalities.29 Comparison with histology is also favorable in the carotid arteries.30 One potential application of CMR is the detection of subclinical atherosclerosis.31,32 It has proved useful in this respect in several large cohort studies. In asymptomatic subjects enrolled in the Framingham Heart Study (FHS), FHS coronary risk score was strongly associated with CMR evidence of aortic atherosclerosis.33,34 The prevalence and extent of aortic atherosclerosis significantly increased with age. In a substudy of the Multiethnic Study of Atherosclerosis (MESA), CMR measures of aortic wall thickness increased as a function of age, with males and African American participants having the greatest wall thickness.35 In a study of 102 patients undergoing coronary angiography, CMR aortic atherosclerotic plaques were

25 ATHEROSCLEROTIC PLAQUE IMAGING: AORTA AND CAROTID

PDW

T1W

T2W

Figure 25-2 An example of multicontrast plaque CMR at high (3-T) field strength. A nonstenotic plaque is well visualized in the left carotid artery. The fibrous cap and lipid core can be seen on the high-magnification images. PDW, proton density weighted; T1W, T1 weighted; T2W, T2 weighted.

detected with a higher frequency in active smokers and in those with high levels of LDL cholesterol, but the volume and area of aortic plaques correlated most strongly with age and the presence of hypertension. Atherosclerotic plaques located in the thoracic aorta were significantly associated with the presence of coronary artery disease.36 Taken together, these studies confirm the strong correlation between the presence of cardiovascular risk factors and the incidence of aortic atherosclerosis,37 with racial and population differences in the response to individual risk factors, which presumably have a genetic basis.35 In a cross-sectional study by Kim and colleagues among patients with type 1 diabetes, T2W fat-suppressed black-blood fast spin echo aortic CMR demonstrated increased thoracic and abdominal plaques among subjects with nephropathy as compared with nomoalbuminuria.38 A study by Underhill and colleagues performed carotid CMR in 191 participants.39 In male subjects, carotid wall thickness and lipid-rich necrotic core were increased among those with coronary artery disease. A benefit readily appreciated by pharmaceutical companies and others wishing to study the effects of drugs on plaque progression (and regression) is the low coefficient of variability of CMR plaque volume measurements.40–43

This allows for low patient numbers for studies investigating the effects of pharmacologic therapy on plaque volume. This was illustrated in two separate studies, in which highdose statin therapy was shown to be superior to low-dose therapy with regard to a reduction in atheroma burden, an effect that was demonstrated with only 20 patients in each group group.43–45 CMR has sufficient resolution to image the complications of atherosclerosis.46 A retrospective study imaged the carotid arteries of patients with a recent transient ischemic attack. In symptomatic patients, ruptured fibrous caps were detected with a much higher frequency (70%) in symptomatic patients as compared with only 9% of subjects with a fully intact thick fibrous cap, thus confirming the ability of CMR to illustrate the underlying pathology.47 Histopathologic studies suggest that intraplaque hemorrhage may also play a role in triggering of plaque rupture and growth.48,49 A first study proved that multicontrast CMR could accurately image intraplaque hemorrhage in carotid atherosclerotic plaques using T2*-weighted sequences.50 CMR can also distinguish between recent and remote hemorrhage on the basis of its methemoglobin content. In another prospective study, the finding of Cardiovascular Magnetic Resonance 343

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T1W

PDW

df

T2W

L - lumen df

nc - necrotic core L

lf

H - intraplaque hemorrhage

L lf

fc - fibrocellular tissue

H nc

fc

df - dense fibrous tissue

lf

H fc

nc

lf - loose fibrous tissue

RGB

Cluster

Figure 25-3 Multicontrast CMR of the carotid artery. Top row shows T1-weighted (T1W), proton-density-weighted (PDW) and T2-weighted (T2W) CMR of the same atherosclerotic plaque. The bottom left panel shows the RGB color composite image obtained by mapping the T1W image to the red, the PDW image to the green, and the T2W image to blue channel, respectively. The bottom right panel shows the plaque segmented automatically into its various components using an automated k-means cluster analysis algorithm. Intraplaque hemorrhage (H), necrotic core (nc), loose fibers (lf), dense fibers (df), and the fibrous cap (fc) can be clearly differentiated.

hemorrhage within carotid atherosclerotic plaques was associated with an accelerated increase in subsequent plaque volume over the next 18 months.51 This supports the hypothesis that intraplaque hemorrhage is a potent atherogenic stimulus. These studies show the value of plaque CMR for both generating and testing pathologic hypotheses about atherosclerosis.

Molecular CMR Imaging of Atherosclerosis Traditional plaque imaging modalities, such as computed tomography and multicontrast CMR, exploit anatomic variations between tissues to provide images. Molecular (or target-specific imaging) unites molecular and vascular biology with imaging modalities such as CMR, positron emission tomography (PET), and optical coherence tomography (OCT) to allow the study of biologic processes noninvasively and with high-spatial resolution.52–55 As we have seen, multicontrast CMR can characterize the various plaque components of intermediate to advanced atherosclerosis with a sensitivity in the range of 103 to 106 M.56 But the ability to deliver a larger payload of paramagnetic 344 Cardiovascular Magnetic Resonance

particles can improve the sensitivity of the technique, especially when trying to image elements of the plaque present at low concentrations (109 to 1013 M/g of tissue.57,58 Below, we will discuss a variety of molecular targets that may provide improved imaging of atherosclerotic plaque using CMR.

Endothelial Dysfunction Endothelial dysfunction is the earliest vascular abnormality seen in atherosclerosis. The ability to image the disease at this early stage might allow preemptive lifestyle or drug intervention to slow its progress. However, a great deal of the endothelium is dysfunctional even in patients with only risk factors for atherosclerosis, as well as those with the established disease. While plaques that are at risk of rupture may express adhesion molecules, those same molecules will be found throughout the vasculature and will therefore not be a specific marker for high-risk lesions. The ability to image vascular cell adhesion molecule has been elegantly demonstrated by Kelly and colleagues.59 They used a superparamagnetic fluorescent nanoparticle coupled to a payload peptide that was internalized by endothelial cells expressing VCAM. This was tested in

Angiogenesis Pathologic studies have revealed the central role of plaque neovascularization in contributing to atherosclerotic plaque vulnerability.63–65 The presence of neovessels is strongly associated with plaque inflammation, macrophage content, and likelihood of rupture,64 presumably by allowing an alternative route for entry of monocytes into the plaque. Gadolinium chelates represent the most commonly used CMR contrast agents for imaging plaque neovessels. These paramagnetic agents increase the luminal signal after intravenous injection and are therefore good candidates for measuring plaque neovasculature.66 Such techniques have been used in oncology imaging to quantify neovessels associated with tumors.67,68 While Kerwin and colleagues demonstrated a correlation between the increase in signal intensity in carotid atherosclerotic plaques with gadolinium-enhanced CMR and the extent of neovessels in plaques measured histologically, this imaging modality is not specific for neovascularization and is limited, because the vast majority of available gadolinium chelates rapidly distribute into the extracellular space.69 Further improvements in the quantification of plaque neovasculature with these agents will require the development of new compartmental kinetic models of the biodistribution of these contrast agents within atherosclerotic plaques.70 Alternatively, CMR contrast agents that remain within the bloodstream or diffuse more slowly into the extracellular space are currently being developed.71 For example, a novel agent targeting the integrin avb3 (specifically expressed on the endothelial surface of neovasculature) has been developed to identify regions in the vessel wall undergoing neovascularization. Winter and colleagues demonstrated in a rabbit model of atherosclerosis that regions of neovascularization in plaques had a 47% increase in CMR signal intensity after the injection of avb3-targeted nanoparticles.72 Another epitope that has been successfully exploited for imaging angiogenesis within plaque (at least in ApoE/ mice) is fibronectin. The binding of an antifibronectin antibody was confirmed both autoradiographically and by the use of a fluorescent probe and was specific for the vasovasorum of the atherosclerotic plaque.73

Thrombus Intraluminal thrombosis represents the final step in the evolution of vulnerable atherosclerotic plaque and is therefore a candidate target for novel specific CMR contrast agents.74 Histologic studies have demonstrated that superficial thrombus superimposed on a ruptured atherosclerotic plaque characterizes those plaques that are at high risk of ischemic events, presumably by implying previous asymptomatic rupture.75 Several approaches have been taken to image thrombus with CMR. Yu and colleagues used a gadolinium-loaded nanoparticle coupled to a fibrin antibody and tested this against in vitro thrombus.76 They were able to demonstrate significant signal enhancement and confirmed tight binding of the antibody with scanning electron microscopy. Winter and colleagues successfully employed a similar method but using a different nanoparticle construct.77 The histology of intracoronary thrombus aspirated from 211 patients with acute myocardial infarction has shown that over 50% of the culprit thrombi were at least days or weeks old.78 On the basis of this premise, a new fibrin-specific CMR contrast agent was designed and thrombus resulting from plaque rupture was identified in a rabbit model. In the 25 arterial thrombi that were induced by carotid crush injury, Botnar and colleagues demonstrated a sensitivity and specificity of 100% for in vivo thrombus detection.79 At our own institution, Sirol and colleagues used the same fibrin-specific CMR contrast agent (EP-1242, EPIX Pharmaceuticals, Cambridge, MA) in 12 guinea pigs to demonstrate that the signal intensity of the thrombus was increased by over fourfold compared to noncontrast images. The detection of thrombi improved from 42% pickup precontrast injection compared to 100% detection after injection.80 The ability to identify different thrombus components with molecular CMR may allow its age to be determined noninvasively. This has important clinical implications because histologic studies have shown that many microplaque rupture events with thrombus formation often predate the “catastrophic” rupture event that causes the clincal syndrome. The ability to detect these “early warnings” might allow intense therapy or device placement to avert symptoms. In this regard, our group has recently tested another experimental fibrin-targeted peptide (EP-2104R, EPIX Pharmaceuticals, Cambridge, MA) for thrombus detection, and compared it to CMR both without contrast and with gadolinium contrast CMR.81 Using this novel agent it was possible to discriminate between occlusive and non-occlusive thrombi, and also to track thrombus as it aged and became more organized by fibrous tissue infiltration. The new agent improved on the detection rate of both multicontrast and gadolinium-enhanced MR (Fig. 25-4).

Extracellular Matrix Other novel CMR contrast agents have been found to accumulate within atherosclerotic plaques. For example, gadofluorine M is a lipophilic, macrocyclic, water-soluble, gadolinium chelate complex with a perfluorinated side chain. Both Sirol and Barkhausen separately demonstrated that gadofluorine M significantly increased signal intensity Cardiovascular Magnetic Resonance 345

25 ATHEROSCLEROTIC PLAQUE IMAGING: AORTA AND CAROTID

ApoE/ mice that were fed a high-cholesterol diet and was highly specific for VCAM-expressing cells, with a gold standard of immunohistochemistry.58 In addition, an experimental study demonstrated that microparticles of iron oxide may be used as a functional CMR probe to reveal endothelial adhesion cells in mouse atherosclerosis.60 CMR imaging of the selectin family of molecules has also been attempted, at least in vitro. Kang and colleagues showed that a monoclonal antibody fragment tagged with iron oxide nanoparticles was specific for human endothelial cells expressing E-selectin in culture, with an increased binding of 200 times compared with controls.61 ICAM-1 receptors expressed on the cerebral arterial vascular endothelium have been imaged with CMR using antibodyconjugated paramagnetic liposomes.62 CMR provided sufficient signal enhancement to determine the areas of increased expression, and binding was verified by fluorescent histopathology.

B

C

T1W Non-CE CMR T2W Non-CE CMR EP-2104R

3 Signal intensity

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A

2.5 2 1.5 1 0.5 0 0

D

1

2

3

4

6

8

Weeks after thrombus induction

Figure 25-4 Transverse CMR images of a rabbit carotid artery 1 week after thrombus induction, imaged by using a double inversion recovery turbo spin echo sequence. T1W (A) and T2W (B) CMR were obtained without a contrast agent. The white arrow indicates the location of the thrombus. C, The T1W CMR obtained 30 minutes after EP-2104R (EPIX Pharmaceuticals, Cambridge, MA) injection. D, The relative signal intensity changes (mean standard deviation) over time for T1W (circles), T2W (squares), and after EP2104R injection (triangles). This gadolinium based, fibrin-targeted CMR contrast agent demonstrates significant enhancement of the thrombus compared to T1W CMR (p < .001). T1W, T1 weighted; T2W, T2 weighted. Source: Adapted from Sirol, Fuster, Badimon, Fallon, Moreno, Toussaint, & Fayad 2005b.

in rabbit aortic plaques compared to controls. A strong correlation was found between the intensity of CMR signal enhancement after the injection of gadofluorine and the presence of lipid-rich plaques on corresponding histologic sections.82,83 This suggests a high affinity of gadofluorine M for atherosclerotic plaque. There are emerging data that gadofluorine M is restricted to the extracellular space of plaques and may interact with resident proteins in the extracellular matrix milieu.

Plaque Inflammation The fibrous cap separates the artery lumen from the thrombogenic core of the plaque and as such is the final barrier to thrombus formation. Thus, identifying factors that lead to the disruption of the fibrous cap may aid in targeting therapy to patients who are at risk of plaque rupture. Matrix metalloproteinases (MMPs) are responsible for the degradation of proteins in the extracellular matrix, causing structural weakening of the cap. The role of MMPs in plaque instability and matrix remodeling in atherosclerotic plaque has been well described.18 Therefore, the ability to detect MMP activity in the fibrous cap with CMR may not only provide important information regarding risk of possible plaque rupture but also allow tracking of MMP inhibition with therapy.84,85 A gadolinium-based CMR contrast agent, P947, has been evaluated in vivo using ApoE/ mice and ex vivo in hyperlipidemic rabbits. The agent has affinity and specificity toward MMP-rich plaques with a preferential 346 Cardiovascular Magnetic Resonance

accumulation of contrast in atherosclerotic lesions compared with the nontargeted reference compound, Gd-DOTA.86 Resident plaque macrophages have been successfully imaged by using ultrasmall superparamagnetic particles of iron oxide (USPIO). These are are removed from the circulation by the reticuloendothelial system and accumulate in macrophages present in atherosclerotic plaques. Macrophages play a pivotal role in the destabilization of atherosclerotic plaques by secreting large quantities of fibrous cap–degrading MMPs, along with proinflammatory cytokines and tissue factor.87 Iron oxide contrast agents have superparamagnetic properties, that is, they decrease T2* relaxation time by generating heterogeneities in the local magnetic field and can be detected on CMR as signal voids on T2*-weighted sequences. In an early study, Weiss and coworkers administered gadolinium-DTPA (0.1 mmol/kg) and performed aortic and carotid CMR in 52 subjects 40 years of age and older, including 17 subjects with no risk factors for atherosclerosis.88 Twenty-two subjects had increased wall thickness, T2W signal intensity, and/or contrast enhancement. This group also had higher levels of interleukin-6, C-reactive protein, vascular cell adhesion molecule-1, and intercellular adhesion molecule. Kooi and colleagues studied 11 symptomatic patients scheduled for carotid endarterectomy with USPIO-enhanced CMR, found a 24% decrease in signal intensity on corresponding T2*-weighted sequences, and histologically verified uptake of USPIO in 75% of ruptured or ruptureprone lesions.89 Trivedi and colleagues expanded on this work and demonstrated that the optimum time for imaging symptomatic carotid plaque was between 24 and 36 hours after injection of USPIO.90,91 USPIO plaque imaging with CMR has been validated against histopathology at time points out to 8 weeks in an animal model. Iron staining closely matched that of macrophage distribution within the plaque, but interestingly, only a subset of smaller-sized macrophages actively accumulated USPIO.92,93 Image acquisition methods that render the superparamagnetic signal loss as positive enhancement promise to improve the detection of plaques in vivo.94 Finally, monocyte/macrophage recruitment into plaques after an inflammatory stimulus has been tracked by USPIO in a mouse model,95 which might be useful in assessing the antiatherogenic potential of new drugs. Macrophages within atherosclerotic plaques have also been targeted for imaging by the use of gadolinium-loaded immunomicelles. These agents, with diameters between 20 and 120 nm, are composed of phospholipids, a surfactant, and an aliphatic chain with Gd-DTPA attached at the polar head group. The polar head group of the aliphatic chain can be attached to antibodies directly or via a biotin-avidin bridge. Using this model, we have made micelles that have over 10,000 gadolinium ions on each micelle surface and the ability to specifically target the macrophage scavenger receptor. Promising work has demonstrated enhancement of murine atherosclerotic plaque on CMR using immunomicelles that target the MSR-A (Macrophage Scavenger Receptor-A).96 Our group has developed another type of imaging agent based on a recombinant high-density lipoprotein (rHDL) molecule that incorporates gadolinium-DTPA phospholipids.97 Natural HDL’s role in the body is that of

Phospholipid

Cholesterol ester

B

C

D

E

Triglyceride

Unesterified cholesterol Gd-DTPA-DMPE

A

Figure 25-5 A, This represents the reconstituted high-density lipoprotein (HDL)–like CMR contrast agent composed of an HDL-like particle and a phospholipid-based contrast agent (Gd-DTPA-DMPE). Transverse in vivo CMR images of the abdominal aorta in an 8-week-old mouse at 9.4 T before (B), 1 hour after (C ), 24 hours after (D), and 48 hours after (E) the injection of recombinant HDL-like nanoparticles are displayed. The insets denote the magnification of the aortic region. Source: Adapted from Frias, Williams, Fisher, & Fayad 2004.

removing lipid from atherosclerotic plaque and returning it to the liver (reverse cholesterol transport). Elevated levels of HDL are associated with a reduction in plaque rupture events, presumably because of this protective effect.98 The rHDL imaging agent has a small diameter (7 to 12 nm), allowing it to diffuse into atherosclerotic plaques, and by using endogenous transport molecules, it does not trigger any immune reaction. Atherosclerotic plaques had a 35% increase of CMR signal intensity 24 hours after particle injection in an ApoE knockout mouse model. Furthermore, fluorescent rHDL colocalized with macrophages present in atherosclerotic plaques with confocal microscopy. Figure 25-5 demonstrates the enhancement of atherosclerotic plaques after the injection of rHDL.

CONCLUSIONS AND FUTURE DIRECTIONS Thanks to the absence of ionizing radiation, CMR represents the imaging technology of choice for the noninvasive high-spatial resolution detection and serial monitoring of

atherosclerotic plaques in the carotid arteries and aorta. High image quality and sensitivity to small changes in plaque size mean that there is little variance between measurements, permitting small sample sizes to be used in comparative studies. Thus, multicontrast CMR is ideally suited for use in evaluation of novel antiatheroma drugs. The development of functional molecular imaging of atherosclerosis may also help to reveal the key pathologic steps that lead from a stable atherosclerotic plaque to an acute ischemic event. However, recent clinical studies have underscored the multiple locations of vulnerable and ruptured atherosclerotic plaques and the diffuse inflammation of the arterial tree in patients with acute ischemic events compared to stable patients.99 Therefore, the concept of detecting infrequent vulnerable atherosclerotic plaques with imaging and treating them individually has started to shift to a more global process of identifying vulnerable patients at high risk of acute clinical events, irrespective of the arterial location.100 In the future, CMR atherosclerosis imaging may help to focus individual evaluation of cardiovascular risk and to optimize antiatherosclerotic therapies.

References 1. British Heart Foundation Health Promotion Research Group. Coronary Heart Disease Statistics. London, UK: 2005. 2. Svindland A, Torvik A. Atherosclerotic carotid disease in asymptomatic individuals: an histological study of 53 cases. Acta Neurol Scand. 1988;78(6):506–517. 3. Sanz J, Fayad ZA. Imaging of atherosclerotic cardiovascular disease. Nature. 2008;451(7181):953–957.

4. Ku DN, Giddens DP, Zarins CK, Glagov S. Pulsatile flow and atherosclerosis in the human carotid bifurcation: positive correlation between plaque location and low oscillating shear stress. Arteriosclerosis. 1985;5(3):293–302. 5. Cunningham KS, Gotlieb AI. The role of shear stress in the pathogenesis of atherosclerosis. Lab Invest. 2005;85(1):9–23.

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ApoA-I

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6. Cushing SD, Berliner JA, Valente AJ, et al. Minimally modified low density lipoprotein induces monocyte chemotactic protein 1 in human endothelial cells and smooth muscle cells. Proc Natl Acad Sci U.S.A. 1990;87(13):5134–5138. 7. Hamilton JA, Myers D, Jessup W, et al. Oxidized LDL can induce macrophage survival, DNA synthesis, and enhanced proliferative response to CSF-1 and GM-CSF. Arterioscler Thromb Vasc Biol. 1999;19(1):98–105. 8. Ross R. Atherosclerosis: an inflammatory disease. N Engl J Med. 1999;340(2):115–126. 9. Zernecke A, Bernhagen J, Weber C. Macrophage migration inhibitory factor in cardiovascular disease. Circulation. 2008;117(12): 1594–1602. 10. Glagov S, Weisenberg E, Zarins CK, Stankunavicius R, Kolettis GJ. Compensatory enlargement of human atherosclerotic coronary arteries. N Engl J Med. 1987;316(22):1371–1375. 11. Keenan NG, Pennell DJ, Mohiaddin RH. Glagov remodeling in the atherosclerotic carotid artery by cardiovascular magnetic resonance. Heart. 2008;94(2):228. 12. Fuster V, Moreno PR, Fayad ZA, Corti R, Badimon JJ. Atherothrombosis and high-risk plaque: part I: evolving concepts. J Am Coll Cardiol. 2005;46(6):937–954. 13. Moreno PR, Purushothaman KR, Sirol M, Levy AP, Fuster V. Neovascularization in human atherosclerosis. Circulation. 2006;113 (18):2245–2252. 14. Juonala M, Viikari JS, Ronnema T, et al. Associations of dyslipidemias from childhood to adulthood with carotid intima-media thickness, elasticity, and brachial flow-mediated dilatation in adulthood. The Cardiovascular Risk in Young Finns Study. Atherosclerosis Thrombosis Vascular Biology. 2008;28 (Epub ahead of print) 15. Farb A, Burke AP, Tang AL, et al. Coronary plaque erosion without rupture into a lipid core: a frequent cause of coronary thrombosis in sudden coronary death. Circulation. 1996;93(7):1354–1363. 16. Davies MJ. Acute coronary thrombosis: the role of plaque disruption and its initiation and prevention. Eur Heart J. 1995;16(suppl L):3–7. 17. Bennett MR, Evan GI, Schwartz SM. Apoptosis of human vascular smooth muscle cells derived from normal vessels and coronary atherosclerotic plaques. J Clin Invest. 1995;95(5):2266–2274. 18. Rudd JH, Davies JR, Weissberg PL. Imaging of atherosclerosis: can we predict plaque rupture? Trends Cardiovasc Med. 2005;15(1): 17–24. 19. Cannon CP, Braunwald E, McCabe CH, et al. Intensive versus moderate lipid lowering with statins after acute coronary syndromes. N Engl J Med. 2004;350(15):1495–1504. 20. Schoenhagen P, Tuzcu EM, Apperson-Hansen C, et al. Determinants of arterial wall remodeling during lipid-lowering therapy: serial intravascular ultrasound observations from the Reversal of Atherosclerosis with Aggressive Lipid Lowering Therapy (REVERSAL) trial. Circulation. 2006;113(24):2826–2834. 21. Ibanez B, Vilahur G, Cimmino G, et al. Rapid change in plaque size, composition, and molecular footprint after recombinant apolipoprotein A-I Milano (ETC-216) administration: magnetic resonance imaging study in an experimental model of atherosclerosis. J Am Coll Cardiol. 2008;15(11):1104–1109. 22. Saam T, Hatsukami TS, Takaya N, et al. The vulnerable, or high-risk, atherosclerotic plaque: noninvasive MR imaging for characterization and assessment. Radiology. 2007;244(1):64–77. 23. Fayad ZA, Fuster V. Characterization of atherosclerotic plaques by magnetic resonance imaging. Ann N.Y Acad Sci. 2000;902:173–186. 24. Toussaint JF, LaMuraglia GM, Southern JF, Fuster V, Kantor HL. Magnetic resonance images lipid, fibrous, calcified, hemorrhagic, and thrombotic components of human atherosclerosis in vivo. Circulation. 1996;94(5):932–938. 25. Yuan C, Mitsumori LM, Beach KW, Maravilla KR. Carotid atherosclerotic plaque: noninvasive MR characterization and identification of vulnerable lesions. Radiology. 2001;221(2):285–299. 26. Yuan C, Mitsumori LM, Ferguson MS, et al. In vivo accuracy of multispectral magnetic resonance imaging for identifying lipid-rich necrotic cores and intraplaque hemorrhage in advanced human carotid plaques. Circulation. 2001;104(17):2051–2056. 27. Mani V, Itskovich VV, Szimtenings M, et al. Rapid extended coverage simultaneous multisection black-blood vessel wall MR imaging. Radiology. 2004;232(1):281–288. 28. Itskovich VV, Choudhury RP, Aguinaldo JG, et al. Characterization of aortic root atherosclerosis in ApoE knockout mice: high-resolution

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72. Winter PM, Morawski AM, Caruthers SD, et al. Molecular imaging of angiogenesis in early-stage atherosclerosis with alpha (v)beta3-integrin-targeted nanoparticles. Circulation. 2003;108 (18):2270–2274. 73. Matter CM, Schuler PK, Alessi P, et al. Molecular imaging of atherosclerotic plaques using a human antibody against the extra-domain B of fibronectin. Circ Res. 2004;95(12):1225–1233. 74. Spuentrup E, Botnar RM, Wiethoff AJ, et al. MR imaging of thrombi using EP-2104R, a fibrin-specific contrast agent: initial results in patients. Eur Radiol. 2008;18(9):1995–2005. 75. Virmani R, Kolodgie FD, Burke AP, Farb A, Schwartz SM. Lessons from sudden coronary death: a comprehensive morphological classification scheme for atherosclerotic lesions. Arterioscler Thromb Vasc Biol. 2000;20(5):1262–1275. 76. Yu X, Song SK, Chen J, et al. High-resolution MRI characterization of human thrombus using a novel fibrin-targeted paramagnetic nanoparticle contrast agent. Magn Reson Med. 2000;44(6):867–872. 77. Winter PM, Caruthers SD, Yu X, et al. Improved molecular imaging contrast agent for detection of human thrombus. Magn Reson Med. 2003;50(2):411–416. 78. Rittersma SZ, van der Wal AC, Koch KT, et al. Plaque instability frequently occurs days or weeks before occlusive coronary thrombosis: a pathological thrombectomy study in primary percutaneous coronary intervention. Circulation. 2005;111(9):1160–1165. 79. Botnar RM, Perez AS, Witte S, et al. In vivo molecular imaging of acute and subacute thrombosis using a fibrin-binding magnetic resonance imaging contrast agent. Circulation. 2004;109(16): 2023–2029. 80. Sirol M, Aguinaldo JG, Graham PB, et al. Fibrin-targeted contrast agent for improvement of in vivo acute thrombus detection with magnetic resonance imaging. Atherosclerosis. 2005;182(1):79–85. 81. Sirol M, Fuster V, Badimon JJ, et al. Chronic thrombus detection with in vivo magnetic resonance imaging and a fibrin-targeted contrast agent. Circulation. 2005;112(11):1594–1600. 82. Barkhausen J, Ebert W, Heyer C, Debatin JF, Weinmann HJ. Detection of atherosclerotic plaque with Gadofluorine-enhanced magnetic resonance imaging. Circulation. 2003;108(5):605–609. 83. Sirol M, Itskovich VV, Mani V, et al. Lipid-rich atherosclerotic plaques detected by gadofluorine-enhanced in vivo magnetic resonance imaging. Circulation. 2004;109(23):2890–2896. 84. Bremer C, Tung CH, Weissleder R. In vivo molecular target assessment of matrix metalloproteinase inhibition. Nat Med. 2001;7 (6):743–748. 85. Nighoghossian N, Derex L, Douek P. The vulnerable carotid artery plaque: current imaging methods and new perspectives. Stroke. 2005;36(12):2764–2772. 86. Lancelot E, Amirbekian V, Brigger I, et al. Evaluation of matrix metalloproteinases in atherosclerosis using a novel noninvasive imaging approach. Arterioscler Thromb Vasc Biol. 2008;28(3):425–432. 87. Libby P. Current concepts of the pathogenesis of the acute coronary syndromes. Circulation. 2001;104(3):365–372. 88. Weiss CR, Arai AE, Bui MN, et al. Arterial wall MRI characteristics are associated with elevated serum markers of inflammation in humans. J Magn Reson Imaging. 2001;14:698–704. 89. Kooi ME, Cappendijk VC, Cleutjens KB, et al. Accumulation of ultrasmall superparamagnetic particles of iron oxide in human atherosclerotic plaques can be detected by in vivo magnetic resonance imaging. Circulation. 2003;107(19):2453–2458. 90. Trivedi R, King-Im J, Gillard J. Accumulation of ultrasmall superparamagnetic particles of iron oxide in human atherosclerotic plaque. Circulation. 2003;108(19):e140. 91. Trivedi RA, King-Im JM, Graves MJ, et al. In vivo detection of macrophages in human carotid atheroma: temporal dependence of ultrasmall superparamagnetic particles of iron oxide-enhanced MRI. Stroke. 2004;35(7):1631–1635. 92. Hyafil F, Laissy JP, Mazighi M, et al. Ferumoxtran-10-enhanced MRI of the hypercholesterolemic rabbit aorta: relationship between signal loss and macrophage infiltration. Arterioscler Thromb Vasc Biol. 2006;26(1):176–181. 93. Yancy AD, Olzinski AR, Hu TC, et al. Differential uptake of ferumoxtran10 and ferumoxytol, ultrasmall superparamagnetic iron oxide contrast agents in rabbit: critical determinants of atherosclerotic plaque labeling. J Magn Reson Imaging. 2005;21(4):432–442. 94. Mani V, Briley-Saebo KC, Itskovich VV, Samber DD, Fayad ZA. Gradient echo acquisition for superparamagnetic particles with positive contrast (GRASP): sequence characterization in membrane and glass

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49. Kolodgie FD, Gold HK, Burke AP, et al. Intraplaque hemorrhage and progression of coronary atheroma. N Engl J Med. 2003;349 (24):2316–2325. 50. Chu B, Kampschulte A, Ferguson MS, et al. Hemorrhage in the atherosclerotic carotid plaque: a high-resolution MRI study. Stroke. 2004;35(5):1079–1084. 51. Takaya N, Yuan C, Chu B, et al. Presence of intraplaque hemorrhage stimulates progression of carotid atherosclerotic plaques: a highresolution magnetic resonance imaging study. Circulation. 2005;111 (21):2768–2775. 52. Wu JC, Bengel FM, Gambhir SS. Cardiovascular molecular imaging. Radiology. 2007;244(2):337–355. 53. Jaffer FA, Libby P, Weissleder R. Molecular imaging of cardiovascular disease. Circulation. 2007;116(9):1052–1061. 54. Wickline SA, Neubauer AM, Winter PM, Caruthers SD, Lanza GM. Molecular imaging and therapy of atherosclerosis with targeted nanoparticles. J Magn Reson Imaging. 2007;25(4):667–680. 55. Sosnovik DE, Nahrendorf M, Weissleder R. Molecular magnetic resonance imaging in cardiovascular medicine. Circulation. 2007; 115(15):2076–2086. 56. Aime S, Cabella C, Colombatto S, Geninatti CS, Gianolio E, Maggioni F. Insights into the use of paramagnetic Gd(III) complexes in MR-molecular imaging investigations. J Magn Reson Imaging. 2002;16(4):394–406. 57. Choudhury RP, Fuster V, Fayad ZA. Molecular, cellular and functional imaging of atherothrombosis. Nat Rev Drug Discov. 2004;3 (11):913–925. 58. Lipinski MJ, Fuster V, Fisher EA, Fayad ZA. Technology insight: targeting of biological molecules for evaluation of high-risk atherosclerotic plaques with magnetic resonance imaging. Nat Clin Pract Cardiovasc Med. 2004;1(1):48–55. 59. Kelly KA, Allport JR, Tsourkas A, Shinde-Patil VR, Josephson L, Weissleder R. Detection of vascular adhesion molecule-1 expression using a novel multimodal nanoparticle. Circ Res. 2005;96(3): 327–336. 60. McAteer MA, Schneider JE, Ali ZA, et al. Magnetic resonance imaging of endothelial adhesion molecules in mouse atherosclerosis using dual-targeted microparticles of iron oxide. Arterioscler Thromb Vasc Biol. 2008;28(1):77–83. 61. Kang HW, Josephson L, Petrovsky A, Weissleder R, Bogdanov Jr A. Magnetic resonance imaging of inducible E-selectin expression in human endothelial cell culture. Bioconjug Chem. 2002;13(1):122–127. 62. Sipkins DA, Gijbels K, Tropper FD, Bednarski M, Li KC, Steinman L. ICAM-1 expression in autoimmune encephalitis visualized using magnetic resonance imaging. J Neuroimmunol. 2000;104(1):1–9. 63. Kockx MM, Cromheeke KM, Knaapen MW, et al. Phagocytosis and macrophage activation associated with hemorrhagic microvessels in human atherosclerosis. Arterioscler Thromb Vasc Biol. 2003;23 (3):440–446. 64. Kumamoto M, Nakashima Y, Sueishi K. Intimal neovascularization in human coronary atherosclerosis: its origin and pathophysiological significance. Hum Pathol. 1995;26(4):450–456. 65. Moulton KS, Heller E, Konerding MA, Flynn E, Palinski W, Folkman J. Angiogenesis inhibitors endostatin or TNP-470 reduce intimal neovascularization and plaque growth in apolipoprotein E-deficient mice. Circulation. 1999;99(13):1726–1732. 66. Mulder WJ, Strijkers GJ, Vucic E, Cormode DP, Nicolay K, Fayad ZA. Magnetic resonance molecular imaging contrast agents and their application in atherosclerosis. Top Magn Reson Imaging. 2007;18 (5):409–417. 67. Jacobs MA, Barker PB, Argani P, Ouwerkerk R, Bhujwalla ZM, Bluemke DA. Combined dynamic contrast enhanced breast MR and proton spectroscopic imaging: a feasibility study. J Magn Reson Imaging. 2005;21(1):23–28. 68. Wang B, Gao ZQ, Yan X. Correlative study of angiogenesis and dynamic contrast-enhanced magnetic resonance imaging features of hepatocellular carcinoma. Acta Radiol. 2005;46(4):353–358. 69. Kerwin W, Hooker A, Spilker M, et al. Quantitative magnetic resonance imaging analysis of neovasculature volume in carotid atherosclerotic plaque. Circulation. 2003;107(6):851–856. 70. Lauffer RB, Parmelee DJ, Dunham SU, et al. MS-325: albumintargeted contrast agent for MR angiography. Radiology. 1998;207 (2):529–538. 71. Port M, Meyer D, Bonnemain B, et al. P760 and P775: MRI contrast agents characterized by new pharmacokinetic properties. MAGMA. 1999;8(3):172–176.

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superparamagnetic iron oxide phantoms at 1.5T and 3T. Magn Reson Med. 2006;55(1):126–135. 95. Litovsky S, Madjid M, Zarrabi A, Casscells SW, Willerson JT, Naghavi M. Superparamagnetic iron oxide-based method for quantifying recruitment of monocytes to mouse atherosclerotic lesions in vivo: enhancement by tissue necrosis factor-alpha, interleukin-1beta, and interferon-gamma. Circulation. 2003;107 (11):1545–1549. 96. Lipinski MJ, Fuster V, Fisher EA, Fayad ZA. Technology insight: targeting of biological molecules for evaluation of high-risk atherosclerotic plaques with magnetic resonance imaging. Nat Clin Pract Cardiovasc Med. 2004;1(1):48–55.

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97. Frias JC, Williams KJ, Fisher EA, Fayad ZA. Recombinant HDL-like nanoparticles: a specific contrast agent for MRI of atherosclerotic plaques. J Am Chem Soc. 2004;126(50):16316–16317. 98. Wilson PW, D’Agostino RB, Parise H, Sullivan L, Meigs JB. Metabolic syndrome as a precursor of cardiovascular disease and type 2 diabetes mellitus. Circulation. 2005;112(20):3066–3072. 99. Rioufol G, Finet G, Ginon I, et al. Multiple atherosclerotic plaque rupture in acute coronary syndrome: a three-vessel intravascular ultrasound study. Circulation. 2002;106(7):804–808. 100. Naghavi M, Libby P, Falk E, et al. From vulnerable plaque to vulnerable patient: a call for new definitions and risk assessment strategies: Part II. Circulation. 2003;108(15):1772–1778.

Atherosclerotic Plaque Imaging: Coronaries Won Yong Kim, David Maintz, Elmar Spuentrup, Susan B. Yeon, Warren J. Manning, and Rene´ M. Botnar

Atherothrombotic cardiovascular disease remains the leading cause of morbidity and mortality in the Western world, and it is rapidly becoming the number one killer in the developing countries.1 Atherosclerosis is a systemic and progressive disease involving the intima of large- and medium-sized arteries, including the aorta, carotid, coronary, and peripheral arteries. Atherothrombosis, defined as atherosclerotic plaque disruption with superimposed thrombosis, can progress to lifethreatening conditions, such as myocardial infarction or ischemic stroke. The concept of “vulnerable plaque” was introduced to distinguish thrombosis-prone plaques and plaques with a high probability of undergoing rapid progression. Studies have demonstrated widespread and diffuse coronary artery vessel wall inflammation in patients with unstable angina2 and acute myocardial infarction.3 Other studies4,5 have indicated that atherosclerotic plaque rupture occurs as a manifestation of a generalized and systemic vascular inflammatory process in combination with other culprit factors. Therefore, the term vulnerable patient has been suggested in recognition of contributing culprit factors including vulnerable blood (prone to thrombosis) and vulnerable myocardium (prone to fatal arrhythmia) that identifies high-risk patients with high likelihood of developing cardiovascular events in the near future.6 Integration of vascular biology with noninvasive imaging techniques in the intact human circulation should enhance our current understanding of the natural history, and of the pathophysiologic mechanisms of atherothrombosis. Noninvasive imaging techniques for atherothrombosis currently include cardiovascular magnetic resonance (CMR), multidetector computed tomography (MDCT), intravascular ultrasound, and ocular coherence tomography. Among these modalities, CMR has emerging as the most comprehensive noninvasive in vivo imaging modality for atherothrombosis,7–9a owing to its noninvasive character and lack of ionizing radiation exposure. With further development, CMR imaging of atherosclerosis and atherothrombosis in the wall of diseased vessels (carotid, thoracic and abdominal aorta, coronary vessels) may prove to be clinically beneficial in identifying vulnerable patients by guiding therapy to prevent progression to clinical disease.

NONINVASIVE DIAGNOSIS OF VULNERABLE PLAQUES On the basis of autopsy studies, a number of major and minor criteria for defining vulnerable plaques have been

proposed.6 These criteria of vulnerability include morphologic features (e.g., size of lipid core and thickness of fibrous cap) (Table 26-1) as well as markers of plaque activity (e.g., plaque inflammation, superficial platelet aggregation and fibrin deposition) (Table 26-2). Most of the proposed features of plaque vulnerability are based on cross-sectional and retrospective studies of culprit lesions. Noninvasive surrogate markers such as coronary vessel wall thickness and plaque burden together with markers of plaque activity could supplement or facilitate cardiovascular risk stratification especially in patients with intermediate cardiovascular risk. Noninvasive imaging techniques may be utilized for prospective outcome studies to determine whether identifying asymptomatic vulnerable patients may allow for primary interventions to reduce clinical cardiovascular events.

CHALLENGES IN CARDIOVASCULAR MAGNETIC RESONANCE CORONARY ARTERY PLAQUE IMAGING Coronary artery CMR10 and coronary artery plaque imaging share many technique obstacles. Both remain technically demanding, as respiratory motion and cardiac motion limit the useful data acquisition window. Coronary vessel wall imaging is even more demanding, owing to the thin vessel wall (coronary vessel wall thickness is approximately 0.5 to 2 mm), the tortuous three-dimensional (3D) course, and the close proximity to epicardial fat, coronary blood, and myocardium. Advanced motion compensation strategies to meet the challenges of coronary vessel wall imaging have been developed, allowing more consistent visualization of the proximal and middle portions of the native coronary vessel walls.

CARDIAC MOTION Owing to the rapid intrinsic cardiac motion during the cardiac cycle, synchronization of data acquisition to the R-wave of the electrocardiogram (ECG) is mandatory (See Chapter 21). Data acquisition is performed only Cardiovascular Magnetic Resonance 351

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CHAPTER 26

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Table 26-1 Criteria for Defining Vulnerable Plaque Major Criteria Active inflammation (monocyte/macrophage T-cell infiltration) Thin fibrous cap with large lipid core Endothelial denudation with superficial platelet aggregation Fissured plaque Stenosis diameter > 90% Minor Criteria Superficial calcified nodule Glistening yellow Intraplaque hemorrhage Endothelial dysfunction Outward (positive) remodeling (Glagov effect) Source: Modified from Naghavi M, Libby P, Falk E, Casscells SW, Litovsky S, Rumberger J, et al: From vulnerable plaque to vulnerable patient: a call for new definitions and risk assessment strategies: Part I. Circulation. 2003;108:1664–1672.

Table 26-2 Markers of Vulnerability at the Plaque Level Morphology/Structure Plaque cap thickness Plaque lipid core size Plaque stenosis (luminal narrowing) Remodeling (expansive versus constrictive remodeling) Color (yellow, glistening yellow, red, etc.) Collagen content versus lipid content, mechanical stability (stiffness and elasticity) Calcification burden and pattern (nodule versus scattered, superficial versus deep, etc.) Shear stress (flow pattern throughout the coronary artery) Activity/Function Plaque inflammation (macrophage density, rate of monocyte infiltration and density of activated T-cell) Endothelial denudation or dysfunction (local NO production, anti-/procoagulation properties of the endothelium) Plaque oxidative stress Superficial platelet aggregation and fibrin deposition (residual mural thrombus) Rate of apoptosis (apoptosis protein markers, coronary microsatellite, etc.) Angiogenesis, leaking vasa vasorum, and intraplaque hemorrhage Matrix-digesting enzyme activity in the cap (MMPs 2, 3, 9, etc.) Certain microbial antigens (e.g., heat shock protein 60, Chlamydia pneumoniae) Panarterial Transcoronary gradient of serum markers of vulnerability Total coronary calcium burden Total coronary vasoreactivity (endothelial function) Total arterial burden of plaque including peripheral (e.g., carotid IMT) IMT, intima medial thickness; MMP, matrix metalloproteinase; NO, nitric oxide. Source: Naghavi M, Libby P, Falk E, Casscells SW, Litovsky S, Rumberger J, et al: From vulnerable plaque to vulnerable patient: a call for new definitions and risk assessment strategies: Part I. Circulation. 2003;108:1664–1672.

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during a relatively short interval comprising less than 10% of the entire cardiac cycle in either end systole or middiastole (periods of relative myocardial diastasis).11,12 Subject-specific mid-diastolic acquisition usually provides the most optimal imaging in subjects with a heart rate less than 80/minute. For subjects with heart rates greater than 80/minute, minimal motion occurs usually during end systole. The timing of the subject-specific image acquisition can be assessed from a cine dataset in the four-chamber orthogonal to the mid-right coronary artery (RCA) view by visual inspection or automated analysis. These measures are essential to avoid cardiac motion–induced artifacts but markedly reduce scan time efficiency. To overcome this drawback, one could use a longer cardiac acquisition window with intra-R-R motion correction.13 Combining the motion-corrected segments produced a high-resolution image.13 The length of the mid-diastolic or end-systolic rest period (which typically varies from 50 to 350 msec) is inversely related to heart rate.11 Therefore, as with coronary MDCT methods, subjects with heart rates greater than 60/minute may benefit from beta-blockers to slow their heart rates and thereby increase the duration of their diastolic rest period.

RESPIRATORY MOTION During normal breathing, superior-inferior diaphragmatic excursion may exceed 20 mm, which translates to cardiac motion that greatly exceeds a multiple of the coronary vessel wall thickness.14 Therefore, either breath hold techniques or respiratory motion compensation during free breathing are required for coronary vessel wall imaging. Initial approaches for visualizing cross sections of the proximal coronary vessel wall used two-dimensional (2D) breath hold techniques with mid-diastolic image acquisition.15 This approach produces variable image quality, since some patients cannot tolerate breath holding for the required 15- to 20-second time period. Furthermore, the inadequate spatial coverage of 2D techniques limits the clinical potential, since evaluation of the entire proximal coronary vessel wall is likely to be important. To compensate for respiratory motion during a freebreathing examination, a 2D-selective pencil beam navigator is placed on the right hemidiaphragm. This information can be used to monitor respiration and to perform real-time gating and slice tracking.10 The use of navigator echoes that record and correct for diaphragmatic motion16 allows for free breathing and eliminates the time constraints of the breath hold approach thereby allowing for submillimeter spatial resolution.17,18 Similar to coronary artery CMR, coronary vessel wall imaging can be combined with the diaphragmatic navigator technique and be extended to 3D acquisitions providing higher signal-to-noise ratio (SNR) and improved vessel coverage (Fig. 26-1).19 The current implementation of respiratory navigators enables slice tracking only in the foot-head direction and accepts data only within a and fixed gating window, of typically 5 mm. Thus, free-breathing coronary artery CMR with respiratory navigator gating is hampered by prolonged scan time due to irregular breathing patterns, which result in low navigator efficiency. A more advanced respiratory motion compensation technique has

Slice 2

Slice 4

Slice 5

LAD

LV

Slice 3

Figure 26-1 Series of cross-sectional 3D proton density–weighted coronary vessel wall cardiovascular magnetic resonance (CMR) images (slice 1–5) in a healthy subject (60 years old, female) without a clinical history of coronary artery disease (CAD). Images were acquired perpendicular (gray box) to the course of the left anterior descending artery (LAD) using a spiral readout preceded by a respiratory navigator and a dual-inversion black-blood prepulse.19

been proposed, which enables the calibration of a 3D affine respiratory motion model to the individual motion pattern of the patient.20 Preliminary results indicate that this approach increases the scan efficiency without sacrificing image quality and therefore has the potential to be more robust for routine clinical usage.21

NONCONTRAST CARDIOVASCULAR MAGNETIC RESONANCE OF ATHEROSCLEROSIS Initial in vivo coronary vessel wall imaging studies used a 2D fat-suppressed fast spin echo (FSE) technique.15,22 Coronary blood signal was suppressed by using a double inversion prepulse23 leading to optimal contrast between lumen and vessel wall. This black-blood approach has been implemented both in a 2D breath hold mode15 and a free-breathing mode22 and allows for visualization of cross-sectional images of the left anterior descending artery (LAD) and RCA vessel wall in both healthy subjects and patients with coronary

artery disease (CAD). In-plane spatial resolution of these first implementations varied from 0.46 0.46 mm to 0.5 1 mm with a typical slice thickness of 3 to 5 mm. Coronary wall thickness was found to be higher in patients with CAD when compared to healthy subjects.15,22 Similarly, vessel wall thickening and signs of mural thrombosis were observed in patients with Kawasaki disease (Fig. 26-2).24 Because of the highly tortuous path of the coronary artery system, cross-sectional vessel wall imaging of the coronary arteries is time inefficient, making a vessel-targeted 3D approach more desirable. Such an approach was implemented by using a three-point plan scan method25 and combined with a modified black-blood prepulse (local inversion),19 which allows acquisition of 3D stacks along the major axis of the coronary artery system. This approach allowed for visualization of the proximal and middle portions of the RCA and LAD coronary artery wall with good contrast between coronary blood and the vessel wall.19 This 3D approach was implemented by using spiral19 and radial k-space sampling.26 While spiral imaging had the advantage of providing high SNR, radial imaging is less motion sensitive, thereby reducing residual motion artifacts (Fig. 26-3). In a clinical feasibility study, free breathing 3D black-blood coronary CMR identified an increased Cardiovascular Magnetic Resonance 353

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Slice 1

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Thrombus

RV

RV

RCA aneurysm

LV

LV

A

B

C

Figure 26-2 A 17-year-old subject with Kawasaki disease as shown on the X-ray angiogram (A). CMR vessel wall imaging (B, C) enabled visualization of the thickened aneurysmatic vessel wall and demonstrates an intraluminal mass (C) suggestive for an intra-arterial thrombus. LV, left ventricle; RCA, right coronary artery; RV, right ventricle.

RCA flow / R-R interval RCA wall

A

B

coronary vessel wall thickness with preservation of lumen size in patients with nonsignificant CAD, consistent with a “Glagov-type” outward arterial remodeling (Fig. 26-4).27 In a cross-sectional study of 136 asymptomatic type 1 diabetic patients with long-standing diabetes, CMR revealed greater coronary plaque burden and higher prevalence of coronary artery stenoses in diabetic patients with diabetic nephropathy than in patients with normoalbuminuria.28

CONTRAST-ENHANCED CARDIOVASCULAR MAGNETIC RESONANCE OF ATHEROSCLEROSIS Contrast-enhanced CMR using extracellular paramagnetic contrast agents have been studied for the characterization of carotid artery29–31 and aortic plaques32 in patients with cardiovascular disease. It was shown that extracellular gadolinium-based contrast agents accumulate both in the inflamed vessel wall31,33,34 and in fibrotic29–31 354 Cardiovascular Magnetic Resonance

Figure 26-3 A 68-year-old subject without a history of coronary artery disease. A, Inflow angiogram demonstrating right coronary artery (RCA) flow. RCA blood flow during one R-R interval was approximately 6 cm. B, Corresponding vessel wall image displaying the proximal and midRCA wall. The bright-appearing band results from the slice selective reinversion pulse. Source: Katoh M, Spuentrup E, Buecker A, et al: MR coronary vessel wall imaging: comparison between radial and spiral k-space sampling. J Magn Reson Imaging. 2006; 23(5):757–762.

atherosclerotic plaques. Work by Kerwin and colleagues35 revealed that kinetic modeling of dynamic contrastenhanced CMR provides an indication not only of the location but also of the amount of neovasculature in carotid artery plaques. Compared to previous double inversion recovery (IR) black-blood coronary vessel wall approaches, contrast-enhanced nonselective IR plaque imaging is expected to be less flow sensitive and relatively fast and may facilitate the assessment of the major coronary arteries within a scan time of 10 to 15 minutes. Furthermore, contrast-enhanced CMR vessel wall imaging has the potential to provide functional information (inflammation and neovascularization) in addition to vessel wall thickness as measured with dual IR black-blood techniques.35a In subjects with CAD, late gadolinium enhancement (LGE) CMR coronary plaque imaging enabled selective plaque visualization36,37 (Fig. 26-5). Gadolinium contrast uptake was more often observed in calcified plaques (53%) than in noncalcified plaques (27%) and segments without plaque (9%) when evaluated and compared to MDCT.37 The good correlation of coronary vessel wall enhancement with clinically evident CAD suggests that contrast uptake may be associated with increased vascular

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Figure 26-5 A 50-year-old subject without history of coronary artery disease. The coronary artery CMR (A, B) shows evidence of minor luminal narrowing in the left main coronary artery (dashed black (A) and white (B) arrows). C, The corresponding late gadolinium enhancement coronary vessel wall scan demonstrates contrast uptake with an eccentric appearing bright area suggestive for atherosclerotic plaque (white arrow). LAD, left anterior descending coronary artery; LCX, left circumflex coronary artery.

permeability (e.g., inflammation) and an increased distribution volume (e.g., fibrotic tissue) for gadolinium in the altered vessel wall. Since the LGE technique reduces the imaging task to assessment of the presence or absence of contrast uptake in the plaque, requirements on spatial

resolution are less stringent than for previously described approaches,15 which require very high spatial resolution to resolve plaque morphology.38 Furthermore, the “hot spot”–like approach simplifies image interpretation and may facilitate use in a clinical setting. Cardiovascular Magnetic Resonance 355

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Figure 26-4 X-ray angiography in two patients with (A) a focal 40% diameter stenosis (white arrow) and (C) minor (10% stenoses) luminal irregularities (white arrows) of the proximal right coronary artery (RCA). The corresponding black-blood 3D CMR vessel wall scans (B, D) demonstrate an irregularly thickened RCA wall (>2 mm) indicative of an increased atherosclerotic plaque burden. The inner and outer walls are indicated by the arrows. Source: Kim WY, Stuber M, Bornert P, Kissinger KV, Manning WJ, Botnar RM. Three-dimensional blackblood cardiac magnetic resonance coronary vessel wall imaging detects positive arterial remodeling in patients with nonsignificant coronary artery disease. Circulation. 2002;106: 296–299.

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Prospective studies to determine the predictive value of contrast-enhanced coronary plaque imaging are needed. Preliminary data in patients with acute myocardial infarction suggest that LGE CMR can monitor local inflammation.38a Angiographically normal segments showed no change in CNR within a week of infarction as compared with 3 months post-infarction. In comparison, there was a significant decrease in CNR in stenotic segments not revascularized. The decline paralleled the decline in C-reactive protein.

MOLECULAR CARDIOVASCULAR MAGNETIC RESONANCE OF ATHEROSCLEROSIS Different CMR probes have been developed to study various biologic processes (e.g., thrombosis, angiogenesis, inflammation, neoplasia) and diseases (e.g., cancer, cardiovascular disease, stroke, diabetes) by targeting a spectrum of molecular markers such as fibrin, selectins, and integrins. Molecular probes generally target either changes in receptor expression or alterations in metabolic processes.39 A molecular imaging agent is typically composed of a ligand such as an antibody, a short peptide or a sugar molecule, and a signal element (e.g., Gd3þ, iron oxides) that are attached to each other by a linker or spacer (Fig. 26-6). Larger molecular imaging probes often include a carrier or nanoparticle (e.g., liposomes, perfluorocarbon emulsions, cross-linked iron oxide) that can be loaded with several to a few thousand signal elements and ligands, thereby (1) Gd molecule attached to ligand (ligand = carrier)

increasing signal amplification and the affinity to the target of interest (see Fig. 26-6). The pharmacokinetics of those agents can be modulated by the size of the carrier, the number of ligands, or additional pharmacokinetic modulation units. Many such agents are in the preclinical stage. So far, only a few plaque-specific contrast agents (e.g., gadofluorin, iron oxides) have been investigated.

MOLECULAR CARDIOVASCULAR MAGNETIC RESONANCE OF THROMBOSIS Noninvasive visualization of evolving arterial thrombus may facilitate detection of unstable plaques and thrombosis burden in vulnerable patients. Recently, advances have been made with in vivo imaging of arterial thrombus by fibrin-binding molecular CMR contrast agents.40,41 Direct imaging of arterial thrombus by targeted or “molecular” contrast agents (which are engineered to bind to specific target molecules) is advantageous to classical noncontrast multispectral CMR, since the demand for high spatial resolution and motion compensation strategies is less stringent. Furthermore, even though several studies have shown high sensitivity of CMR to the detection of carotid and aortic thrombi,42–44 differentiation between complex atherosclerotic plaques and mural thrombosis remains difficult because of the complex composition (e.g., platelets, fibrin, and red blood cells) of thrombus and resultant complex CMR signal characteristics on T1-, T2-, and proton (2) Nanoparticle labeled with multiple Gd molecules and ligands (nanoparticle = carrier)

Signal element (e.g. gadolinium) Nanoparticle Linker Ligand Target Endothelium

Ligand: antibody, small peptides

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Carrier Liposomes, Perfluorocarbon emulsions Cross-linked iron oxide (CLIO)

Target: adhesion molecules, integrins, receptors, fibrin Figure 26-6 Schematic of molecular contrast agents targeted against endothelial activation. The basic components of a typical molecular contrast agent consist of a ligand that binds to a specific target and a signal element, which, in case of CMR, is made of a Gd3þ chelate or an iron oxide. These two basic components can be directly linked to each other, as in (1), or may be attached to or incorporated within a larger nanoparticle (carrier) as demonstrated in (2). 356 Cardiovascular Magnetic Resonance

Figure 26-7 A, Reformatted coronary artery CMR from a coronal 3D dataset shows subrenal aorta 20 hours after EP-1873 administration (EPIX Pharmaceuticals, Inc., Cambridge, MA) in a rabbit model of atherosclerosis and plaque rupture. Three well-delineated mural thrombi (arrows) can be observed, with good contrast between thrombus (numbered), arterial blood (dashed arrow), and vessel wall (solid arrow). The inplane view of the aorta allows simultaneous display of all thrombi, showing head, tail, length, and relative location. B–D, Corresponding cross-sectional views show good agreement with histopathology (E–G). Source: Botnar RM, Perez AS, Witte S, et al. In vivo molecular imaging of acute and subacute thrombosis using a fibrin-binding magnetic resonance imaging contrast agent. Circulation. 2004;109:2023–2029.

In another study by Sirol and colleagues,50,51 similar agents were used in an animal model of acute and chronic thrombosis. Signal intensity decreased with thrombus age, allowing differentiation between the acute and chronic stages. Combining the advent of fibrin-binding molecular CMR contrast agents and advances in coronary artery CMR techniques offers the potential for direct imaging of coronary thrombosis. The feasibility of this approach was demonstrated by using the gadolinium-based fibrin-binding contrast agent EP-2104R ((EPIX Pharmaceuticals, Inc., Cambridge, MA), in a swine model of native coronary thrombus (Fig. 26-8) and in-stent thrombosis using CMR-lucent stents.40 Potential applications for direct thrombus imaging include detection and evaluation of acute coronary syndromes and ischemic strokes. Platelets are also key to thrombus formation and play a role in the development of atherosclerosis. Though it has not yet been studied in the coronary arteries, von zur Muhlen and colleagues51a reported on a CMR contrast agent consisting of microparticles of iron oxide and single-chain antibody targeting ligand-induced binding sites on activated glycoprotein IIb/IIIa to image the carotid artery

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density–weighted images of arterial thrombi. The concept of target-specific imaging or molecular imaging was first introduced over a decade ago, and has since been further developed by Weissleder45,46 and others47 for MRI and for optical imaging in recent years. The advantage of CMR molecular plaque imaging is the potential for relatively high spatial resolution. The limitation is the inherently low sensitivity of CMR contrast enhancement technology, requiring a relatively high target molecule concentration (>50 100 mM Gd at target site; r1 ﬃ 21 mM1 s1 per Gd) for sufficient signal amplification. Initial attempts were made with targeting fibrin,47–49 which is abundant in arterial clots and therefore plays an important role in acute coronary syndromes and stroke. In vivo CMR of acute and subacute thrombosis following plaque rupture in an animal model of aortic atherosclerosis has been implemented by using a small-molecule fibrin-binding peptide derivative, EP-1873 (EPIX Pharmaceuticals, Inc., Cambridge, MA).41 This molecular agent allowed for imaging of large lumenencroaching thrombi as well as submillimeter mural thrombi with signal enhancement of the entire thrombus and excellent differentiation from the vessel wall (Fig. 26-7).

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Figure 26-8 In vivo CMR of in-stent thrombosis in a swine model of coronary thrombosis. A, D, Coronary artery CMR before (A) and after (D) injection of a fibrin-binding contrast agent, EP-2104R (EPIX Pharmaceuticals, Inc., Cambridge, MA). On both scans, no apparent thrombus is visible (arrows, circle). (B, E) Black-blood inversion recovery segmented k-space gradient echo CMR before (B) and after (E) contrast administration of EP-2104R (same view as A and D). After contrast injection (E), a bright area is readily visible (arrows, circle), consistent with the location of the in-stent thrombus. No apparent thrombus was visible on prethrombus (B) images (arrow, circle). (C, F) X-ray angiogram confirming CMR finding of in-stent thrombus in mid-LAD (circle). LAD, left anterior descending coronary artery; LCX, left circumflex coronary artery; LM, left main coronary artery. Source: Botnar RM, Buecker A, Wiethoff AJ, et al. In vivo molecular imaging of acute and subacute thrombosis using a fibrin-binding magnetic resonance imaging contrast agent. Circulation. 2004;110:1463–1466.

thrombi and atherosclerosis plaques in a mouse model of atherothrombosis studied at high field (9.4 T). In vivo visualization of platelet activation was demonstrated, along with signal loss after thrombolysis.

INFLAMMATION Vascular inflammation and associated endothelial activation are believed to play an integral role in the initiation and progression of atherosclerosis. Endothelial activation is characterized by the upregulation of leukocyte adhesion molecules (intercellular adhesion molecule-1 (ICAM-1) and vascular adhesion molecule-1 (VCAM-1), E-selectin, and P-selectin) on the endothelial cell surface.52 The adhesion molecules facilitate tethering, adhesion, and transendothelial migration of leukocytes, including monocytes. Differentiation of monocytes into macrophages and subsequent digestion of lipoproteins by macrophages occur in a later stage and eventually lead to the accumulation of lipid-filled macrophages, which are believed to be a precursor of rupture-prone vulnerable plaque. While the specific 358 Cardiovascular Magnetic Resonance

mechanisms by which the adhesion molecules contribute to this process have not yet been determined, studies of atherogenic mice with various adhesion molecule deficiencies have indicated some role for each of the inducible adhesion molecules, particularly VCAM-1 and P-selectin.52–54 Work by Sibson and colleagues55 and Barber and colleagues56 demonstrated that early endothelial activation occurs in focal ischemia in mice brains55 and in brain inflammation in rats (after IL-1b and TNF-a induced E- and P-selectin upregulation) using a novel MR contrast agent. This novel gadoliniumlabeled contrast agent, Gd-DTPA-B(sLex)A,57 consists of the Sialyl Lewisx (sLex) carbohydrate, which interacts with both E- and P-selectin. The relaxivity was measured as 3.5 mM1 sec1 at 1.5 T and thus is similar to gadolinium-DTPA.

ANGIOGENESIS Identification of angiogenesis may be useful in studying tumor growths and atherosclerotic plaque development. Integrins, such as aVb3, are overexpressed in activated neovascular endothelial cells, which are believed to play an

CLINICAL STUDIES Reproducibility studies have shown excellent results in measuring total plaque volumes of the thoracic and abdominal aorta as surrogate markers of atherosclerosis.61 Also, coronary plaque burden can be quantified with good reproducibility.62 Since CMR permits highly reproducible measures of aortic anatomy and atherosclerosis, serial studies to investigate the effect of, for example, statin therapy on regression of aortic plaque burden have been successfully performed.63

OUTLOOK With the broader availability of higher-field (>1.5 T) whole body CMR systems, in vivo coronary vessel wall imaging is likely to benefit from the expected near linear increase in SNR.64–66 Improved SNR may potentially improve spatial

resolution and/or reduce imaging time. In a preliminary study, we demonstrated the feasibility (see Fig. 26-5) of high-field in vivo coronary vessel wall imaging and found an SNR increase of approximately 50% compared to previous reports at 1.5 T.15,22 Other investigators67,68 have reported similar results. With continued advances in hardware technology and pulse sequence design, further advances in coronary plaque imaging should be expected in the years to come.

CONCLUSION Atherothrombosis, defined as atherosclerotic plaque disruption with superimposed thrombus formation, is the major cause of acute coronary syndromes and cardiovascular death. CMR is emerging as the most comprehensive noninvasive imaging technique for imaging of atherothrombosis in large- and medium-sized arteries, including the aorta and the carotid, coronary, and peripheral arteries. CMR imaging of coronary wall atherothrombosis is particularly challenging, owing to the small caliber of the vessels combined with respiratory and cardiac motion. Freebreathing 3D CMR coronary vessel wall imaging has enabled in vivo quantification of coronary plaque burden and remodeling as a marker of subclinical coronary artery disease. Molecular imaging utilizing target specific contrast agents such as fibrin-binding agents to detect arterial thrombus shows great promise as the new frontier in noninvasive imaging. Advances in molecular imaging and CMR techniques offer the potential for direct imaging of coronary thrombosis and in-stent thrombosis using fibrin-binding molecular CMR contrast agents. While the current role of noninvasive CMR imaging of atherothrombosis remains investigational, integration of vascular biology with CMR should enhance our understanding of the natural history of acute coronary syndromes and thereby facilitate strategies to prevent acute coronary syndromes and cardiovascular death in vulnerable patients.

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CHAPTER 27

Assessment of the Biophysical Mechanical Properties of the Arterial Wall Raad H. Mohiaddin

Arteries are elastic tubes whose diameter varies with the pulsating pressure. In addition, they propagate the pulse created by ejection of blood by the heart, at a velocity that is determined largely by the elastic properties of the arterial wall. The vascular wall can be deformed by pressure and shear stress forces exerted by the blood as well as the tethering imposed by the surrounding tissues. Biophysical mechanical properties of the arterial wall play an important role in the pathogenesis of cardiovascular diseases. Sclerosis (or stiffness), for example, is an important aspect of atherosclerotic vascular disease that can be demonstrated in experimental disease both in animals1 and in humans,2 and regression of the disease leads to reduced stiffness.3,4 Rupture of atherosclerotic plaque (a common mechanism of myocardial infarction) and aortic dissection can be viewed as mechanical failures in the diseased vessels. In addition, the effectiveness of interventional procedures such as angioplasty is often achieved by causing mechanical injury to the vessel wall, and the injury itself may lead to restenosis. A number of common conditions are associated with changes in arterial mechanical properties, although their importance is not always recognized. Systemic hypertension is almost always associated with altered mechanical properties of the peripheral vasculature. Though it is not clear which of the two is the inceptive event, one begets the other, fostering a vicious cycle. Compliance (the reciprocal of the resistance offered to deformation) of the proximal aorta is reduced as a result. This causes waves reflected from the periphery to return prematurely and coincide with the incident or forwardtraveling wave produced by the ventricle. This augments aortic systolic pressure, increases pulse pressure, and reduces diastolic pressure (as the reflected wave no longer contributes to diastolic pressure). Decreased subendocardial/subepicardial flow ratio has been demonstrated with reduced aortic compliance.5 Thus, decreased aortic distensibility may increase the risk of subendocardial ischemia in the presence of coronary artery stenosis, left ventricular hypertrophy, or both. Aortic compliance represents an important determinant of left ventricular performance. Measurement of ventricularvascular coupling takes into account the pulsatile load imposed on the left ventricle as well as the systemic vascular resistance.6 Metabolic disorders such as Ehlers-Danlos and Marfan syndromes,7 diabetes mellitus,8 familial 362 Cardiovascular Magnetic Resonance

hypercholesterolemia,9 and growth hormone deficiency10 are also known to alter arterial compliance. Similarly, the distensibility of the pulmonary arteries is reduced in pulmonary hypertension. Postmortem studies of externalized pulmonary arterial strips have shown that the extensibility of the pulmonary trunk is decreased in pulmonary arterial hypertension.11 The wall of an artery becomes less extensible, the more it is stretched from its natural length. An increased stretching of the circumference of the vessel will diminish the distensibility. When the pulmonary artery resistance increases in vivo, the vessels become more distended and less distensible. Indirect measurements of pulmonary artery compliance have suggested that pulmonary arterial distensibility decreases with rising pulmonary artery pressure.12,13 The ability of cardiovascular magnetic resonance (CMR) to image flow within any medium-sized vessel in any plane provides a unique opportunity to study the pulmonary arteries. Distensibility and flow have already been assessed in patients with pulmonary arterial hypertension.14,15 In this chapter, the clinical importance of arterial biophysical function and its assessment by CMR will be examined in detail as a complement to Chapter 26.

ARTERIAL STRUCTURE A normal artery consists of three morphologically distinct layers. The intima consists of a single continuous layer of endothelial cells bounded peripherally by a fenestrated sheet of elastic fibers. The media consists entirely of diagonally oriented smooth muscle cells, surrounded by variable amounts of collagen, elastin, and proteoglycans. The adventitia consists predominantly of fibroblasts intermixed with smooth muscle cells loosely arranged between bundles of collagen and proteoglycans. Each structural component has its own characteristic properties. Smooth muscle is the physiologically active element, and by contracting or developing force, it can alter the diameter of the vessel or the tension in the wall. The other components are essentially passive in their mechanical behavior. Elastin, which can be stretched to up to 300% of its length at rest without rupturing,16 behaves mechanically more closely to a linear elastic material such as rubber than other connective tissue components do. When elastin fibers are stretched and

DEFINITION OF VASCULAR WALL STIFFNESS Vascular mechanics have been described by using different elastic moduli and assumptions and for different purposes. Several approaches have been described that use clinically available methods for in vivo characterization of the stiffness of the vessel wall. The ability to measure vascular stiffness has been greatly improved by the recent advances in imaging, such as high-frequency ultrasound and CMR. The relationship between vascular wall deformation (strain) and the pressure exerted on the inner surface of the vascular wall (stress) is commonly used for the measurement of arterial wall biophysical properties (elastic modulus). A plethora of terminology for the description of different elastic moduli, which can be confusing, has been described.18 The pressure-strain elastic modulus of the arterial wall (Ep) described by Peterson and colleagues18a

is commonly used. This elastic modulus that applies to an open-ended vessel in the absence of reflection is defined as Ep ¼ 2DP/(DV/V). This is the inverse of the fractional distensibility DV/V of the arterial lumen per unit pulse pressure DP. Arterial compliance, C, which is defined as the change in volume DV per unit change in pressure DP, has been also used in the literature. It has been argued that this definition is appropriate to measurement of ventricular compliance and not to the compliance of an open-ended arterial segment. For the latter, the inverse of Peterson’s modulus was suggested 1/Ep. The average arterial compliance of a particular vessel pathway can also be determined by measuring the speed of propagation of the pulse in the vessel pathway. The velocity of such waves depends principally on the distensibility of the vessel wall. In real terms, this pulse is measurable by the disturbances in pressure, flow, or vessel diameter that it causes. The propagation of flow waves has not been studied as extensively as that of pressure waves, partly because, unlike flow, accurate methods of pulsatile pressure measurements have been available for a long time and partly because the distinction between flow wave velocity and blood velocity has not always been clearly recognized. Blood velocity means the speed of an average drop of blood, while flow wave velocity means the speed with which motion is transmitted. The wave velocity is usually much faster than that of the blood itself. While CMR is unable to assess pressure changes, alterations in the flow within (or diameter of) a vessel can be measured accurately.

MEASUREMENT OF ARTERIAL WALL STIFFNESS Arterial stiffness, which describes the resistance of arterial wall to deformation, is difficult to measure because of the complex mechanical behavior of arterial wall. As a result, a bewildering number of choices abound. Though smooth muscle tone is not considered, in vitro human arterial compliance has been measured from pressure-volume curves in postmortem arteries.19–22 In vivo estimation of arterial wall compliance is more difficult, however, and has been performed by using indirect and invasive techniques, including pulse wave velocity measurements in animals and in humans23,24 the pressure-radius relationship using the Peterson transformer coil in animals,25 X-ray contrast angiography in humans,26,27 and ultrasonography.28–30 The contributions of CMR to the assessment of arterial wall mechanics are discussed in the following paragraphs. Other noninvasive measures are forced to rely on the assessment of accessible and often superficial vessels. Under the assumption that central and peripheral arteries behave in a similar fashion, the properties of these arteries are then used as surrogates for those of central arteries. There is, however, considerable heterogeneity between peripheral and central sites. CMR circumvents this problem by allowing the study of central arteries. Its ability to identify anatomic landmarks suggests that reproducibility between studies should be improved, allowing more effective follow-up. Cardiovascular Magnetic Resonance 363

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released, they return promptly to their original state. Elastin fibers are important for maintaining normal pulsatile behavior, but they fracture at very low stresses and are probably much less important in determining the overall strength of the vessel wall. Collagen fibers, on the other hand, are much stiffer, but they are much stronger. The proportion of these components varies from artery to artery. In the thoracic aorta, the elastin forms 60% of total fibrous element, and collagen forms 40%. The collagen proportion increases with increasing distance from the heart, reaching 30% elastin and 70% collagen in the extrathoracic vessels.17 The collagen-to-elastin ratio increases with age, which is one reason why vascular stiffness increases with age. The human thoracic aorta is supplied by vasa vasorum and grows by increasing the number of lamellar units. The abdominal aorta, in comparison, is avascular, as it lacks vasa vasorum and grows by increasing the thickness of each lamellar unit. The avascular thickness and the elevated tension per lamellar unit of the abdominal aorta predispose it to atherosclerosis. The distensibility of a blood vessel depends on the proportions and interconnections of these materials and on the contractile state of the vascular smooth muscle. Elasticity is a material’s ability to return to its original shape and dimensions after deformation, the deformation being proportional to the force applied. This proportionality was first described by Hooke in 1676 and is known as Hooke’s law. The point at which Hooke’s law ceases to apply is known as the elastic limit, and when a solid has been deformed beyond this point, it cannot regain its original form and acquires a permanent distortion. With larger loads still, the yield point is reached when the deformation continues to increase without further load and usually rapidly leads to breakage. In purely elastic bodies, stress (the force per unit area that produces deformation) produces its characteristic strain (the deformation of a stressed object) instantaneously, and strain vanishes immediately on removal of the stress. Some materials, however, require a finite time to reach the state of deformation appropriate to the stress and a similar time to regain their unstressed shape. Blood vessels typically exhibit such behavior, which is called viscoelasticity.

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CARDIOVASCULAR MAGNETIC RESONANCE OF REGIONAL AORTIC COMPLIANCE CMR provides a direct, noninvasive way of studying local aortic compliance.31,32 High-resolution cine imaging or electrocardiogram-gated spin echo imaging in a plane perpendicular to the ascending and/or descending aorta allows measurement of aortic cross-sectional area during systole and diastole. Measurement of regional aortic compliance by CMR is calculated from the change in volume of an aortic segment and from aortic pulse pressure estimated by a sphygmomanometer at the level of the brachial artery. The lumen of the aorta is outlined manually on the computer screen to measure the change in aortic area (DA) between diastole and systole. Regional aortic compliance (C) (microl/mm Hg, m3/ Nm2) is calculated from the change in volume (DV ¼ DA slice thickness) of the aortic segment divided by the aortic pulse pressure (DP) measured by a sphygmomanometer (Fig. 27-1). Automatic measurement of aortic cross-sectional area is also possible.33 Other indices of aortic stiffness that can be derived from these measurements include distensibility, Peterson’s elastic modulus, and stiffness index b ([systolic blood pressure/diastolic blood pressure]/area strain).

The accuracy of the indirect measurement of the pressure change that is needed to compute compliance is limited, as it ignores the changes in the pressure wave as it propagates through the arterial tree (a process known as amplification). Further, it is important to obtain this pressure data on patients who are ideally lying in the cardiovascular magnetic resonance imaging (MRI) scanner using CMR compatible apparatus. Despite the limitations of the pressure measurement, there is a good correlation between measurement of local aortic compliance and measurement of global compliance from the speed of the propagation of the flow wave within the vessel.34

CARDIOVASCULAR MAGNETIC RESONANCE OF FLOW WAVE VELOCITY Flow wave velocity is defined as the speed with which a flow wave propagates along a vessel and is regarded as the purest measure of arterial stiffness. It is the quotient of distance traveled divided by the time taken for the flow wave to move between the two points and represents an average for that length of vessel (Figs. 27-2 and 27-3). The approach is dependent on assessment of path length

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Figure 27-1 A, Oblique sagittal image of the ascending aorta, arch and descending thoracic aorta showing the sites where flow wave velocity and regional compliance are measured. The oblique transverse plane shown in the top image is represented in diastole (B) and systole (C). This shows the change in area of a 22-year-old healthy subject. AA, ascending aorta; DA descending aorta. 364 Cardiovascular Magnetic Resonance

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traveled and accurate measurement of pulse arrival time. The latter requires recognition of equivalent features or points on leading edges of the proximal and distal flow waveforms (see Fig. 27-2), a process that is made complicated by alterations that occur in flow wave morphology and magnitude as it progresses down the vessel. Unlike noninvasive measurements that rely on linear, transcutaneous measurements, CMR makes no assumptions about the shape of the artery and can accurately measure the path length traveled. Mohiaddin and colleagues showed the feasibility of using CMR phase-shift velocity mapping to measure aortic flow wave velocity in humans.34 By taking advantage of the anatomy of the aorta, cine two-dimensional phase shift velocity maps were acquired with high temporal resolution in a single plane perpendicular to the ascending and descending aorta, and the time taken for the flow wave to travel between the two points was measured (Fig. 27-4). Instantaneous flow (in liters per second) in the midascending aorta and mid-descending aorta was calculated by multiplying the aortic cross-sectional area and the mean velocity within that area. Pulse wave velocity (PWV) was calculated in meters per second from the transit time (T) of the foot of the flow wave (see Fig. 27-4) and from the distance (D) between the two points obtained from an oblique sagittal image. The distance is determined manually on the computer screen by drawing a line in the center of the aorta joining the two points. In Figure 27-4, the foot of the flow wave was defined by extrapolation of the rapid upstroke of the flow wave to the baseline as opposed to the midpoint of the upslope method used in Figure 27-2. Others have used different MR flow imaging techniques to assess arterial compliance. Tarnawski and colleagues35

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Figure 27-3 Cine velocity mapping in a plane equivalent to that shown in Figure 27-1. The first frame was acquired 50 msec after the R-wave of the electrocardiogram and represents the onset of left ventricular systole. The velocity maps indicate zero velocity as medium gray, caudal velocities in the descending aorta as light gray to white, and cranial velocities as darker shades of gray to black, gray level intensity being proportional to velocity.

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Figure 27-4 The foot of the flow wave is defined by extrapolation of the rapid upstroke of the flow wave to the baseline, and this is performed for flow in both the ascending aorta and descending aorta. The transit time needed for the flow wave to propagate from a point in the midascending aorta to a point in the mid-descending aorta can then be measured. The distance around the arch from the plane in the ascending aorta to the plane in the descending aorta can be measured from the oblique sagittal image shown in Figure 27-1, and the flow wave velocity is calculated by division of this distance by the propagation time. A, Transit time from a normal subject with good compliance. B, Transit time from an elderly patient with poor compliance. Note that the transit time is significantly shorter in the patient with poor compliance. Source: Mohiaddin RH, Firmin DN, Longmore DB. Age-related changes of human aortic flow wave velocity measured non-invasively by magnetic resonance imaging. J Applied Physiol. 1993;74:492–497.

Figure 27-5 Fourier velocity measurements in a healthy subject. A, In this method, the magnetic resonance signal is obtained from a column aligned with the descending aorta. B, The position along the column is shown horizontally in each of the 12 cine frames. The velocity is shown by the vertical position of the signal. I, inferior; S, superior.

REFLECTED WAVES When the incident pulse wave from the ventricle reaches the periphery, it may be reflected and return toward the heart as a backward-running wave. Reflected waves express properties of the peripheral circulation; if the peripheral resistance is elevated, reflected waves will be greater in magnitude and will return sooner. Since its shape and magnitude will be defined by the complex interaction between forward incident wave and backward reflected waves, the measured flow wave will also be altered in pathologic conditions. As a result, the timing and location of measurements become important considerations in the CMR assessment of arterial properties. If work is concerned only with the structure of proximal arteries, measurement should be made as far from the periphery and as early in systole as possible to avoid the influence of reflections (see diagram). These requirements become more demanding when the available length of vessel is limited (e.g., in the pulmonary arteries) or baseline data particularly noisy. Conversely, study of the influence

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methods, the interleaved repeats needed to achieve high temporal resolution make them highly sensitive to physiologic variability. In vitro experiments showed CMR measurements of pulse wave velocity in a tube phantom to be very reproducible and in good agreement with pulse wave velocity measurements made with a pressure catheter.39 These methods have not been widely applied in clinical research, nor has their in vivo reproducibility been proven yet.

used the comb-excited Fourier velocity-encoded method, previously reported by Dumoulin and colleagues,36 to measure local arterial wave speed in the femoral artery in healthy men. In this method, simultaneous Fourier velocity-encoded data from multiple stations were acquired. The technique employs a comb excitation radiofrequency pulse that excites an arbitrary number of slices. This causes the signals from the spin in a particular slice to appear at a position in the phase-encoding direction, which is the sum of the spin velocity and an offset arising from the phase increment given to that excitation slice. Acquisition of spin velocity information occurs simultaneously for all slices, permitting the calculation of wave velocities arising from the pulsatile flow. Hardy and colleagues37 studied aortic flow wave velocity using a two-dimensional CMR selective excitation pulse to repeatedly excite a cylinder of magnetization in the aorta, with magnetization read out along the cylinder axis each time. A toggled bipolar flow-encoding pulse was applied prior to readout to produce a one-dimensional phase-contrast flow image. Cardiac gating and data interleaving were employed to improve the effective time resolution to 2 msec. Wave velocities were determined from the slope of the leading edge of flow measured on the resulting Mmode velocity image. Aortic pulse wave velocity was also measured by the same group using a combination of cylinder of magnetization with Fourier velocity encoding and readout gradients applied along the cylinder axis (aorta) (Fig. 27-5),38 with the advantage of eliminating partial volume effects that hindered their previous approach, but the Fourier method has the drawback that it is no longer in real time and errors occur, owing to accumulation of flow data over several (typically 16 to 32) cardiac cycles. For both

Mohiaddin and colleagues were the first to use CMR for measurement of aortic compliance.40 They demonstrated that aortic compliance in asymptomatic subjects falls with age and that patients with coronary artery disease have abnormally low compliance (Fig. 27-6). The results suggested a possible role for compliance in the assessment of cardiovascular fitness and the detection of coronary artery disease. Because there is overlap between normal compliance and compliance in patients with coronary artery disease above the age of 50 years, the test cannot have perfect sensitivity and specificity. Below the age of 50, however, there is much less overlap, and the test is more specific. Abnormally low aortic compliance has also been demonstrated in patients with aortic coarctation41 and in patients with Marfan syndrome.42 The same group also showed the feasibility of using CMR velocity mapping for measurement of aortic flow wave velocity. Aortic flow wave velocity increased linearly with age, and there was a significant difference between the youngest decade and the oldest

decade studied. Flow wave velocity was negatively correlated with regional ascending aortic compliance measured in the same subjects (Fig. 29-7). In a study employing a single-slice phase-contrast acquisition of the ascending, arch, proximal aorta, and distal thoracic aorta, Rogers and colleagues43 demonstrated an age-related increase in PWV among their cohort. In addition, among the older patients, stiffness increased, the more proximally the aorta was studied. The researchers ascribed these changes to the disproportionate effect of elastin degradation with age on the more proximal parts of the aorta, where the elastic to collagen ratio is at its greatest. Regression of atheroma with reduction of cholesterol levels is recognized to occur, but less is known about reversal of sclerosis. Noninvasive indices of sclerosis have largely

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Figure 27-6 Ascending aortic compliance displayed by using a logarithmic scale and plotted against age in three groups: normals, athletes, and patients with coronary artery disease (CAD). The 95% confidence limits are shown for the normals. The athletes’ compliance is abnormally high, and that in coronary disease patients is abnormally low. Source: Mohiaddin RH, Underwood SR, Bogren HG, Firmin DN, Klipstein RH, Rees RSO, Longmore DB. Regional aortic compliance studied by magnetic resonance imaging: the effects of age, training, and coronary artery disease. Br Heart J. 1989;62:90–96. 368 Cardiovascular Magnetic Resonance

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of reflected waves on measured flow waves might provide interesting insights into the nature of the more distal vessels.

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Figure 27-7 Flow wave velocity is directly proportional to age (A) and inversely proportional to regional aortic compliance (B). Source: Mohiaddin RH, Firmin DN, Longmore DB. Age-related changes of human aortic flow wave velocity measured non-invasively by magnetic resonance imaging. J Applied Physiol. 1993;74:492–497.

assessment of aortic area was made at the level of the pulmonary bifurcation by CMR and plotted against a surrogate for central aortic pressures (using a finger arterial blood pressure–derived brachial blood pressure). Using pressure-area lines, a subgroup with a transition point (a departure from a linear pressure-area relationship) was identified in 6 out of 32 patients. It was suggested that this point represented the recruitment of collagen to loadbearing elements. These patients demonstrated a trend toward reduced distensibility, though this relationship might have been stronger had the study group been larger. It was hypothesized that patients in whom the transition point is seen at higher blood pressures might experience greater improvements in distensibility with beta-blockade than those whose transition points occurred at lower blood pressures. Savolainen and colleagues51 used CMR and indirect brachial artery blood pressure measurements to assess aortic elastic modulus in patients with essential hypertension prior to therapy and after 3 weeks and 6 months of therapy with cilazapril (an angiotensin-converting enzyme inhibitor) or atenolol (a beta-1-adrenergic blocker). The authors concluded that 6 months of treatment with either cilazapril or atenolol reduces the stiffness of the ascending aorta in essential hypertension. No statistically significant differences between the effects of the two drugs were observed. Honda and associates used CMR to measure aortic distensibility in patients with systemic hypertension and demonstrated that the antihypertensive drugs nicardipine and alacepril have a beneficial effect on aortic distensibility.52 Resnick and colleagues53 assessed aortic distensibility, left ventricular mass index, abdominal fat (subcutaneous and visceral), and free magnesium levels in the brain and skeletal muscle by CMR. In patients with essential hypertension, the following were concluded: Systolic hypertension and increased left ventricular mass index may result from arterial stiffness; arterial stiffness may be one mechanism by which abdominal visceral fat contributes to cardiovascular risk; and decreased magnesium contributes to arterial stiffness in hypertension. Chelsky and coworkers54 used CMR to measure aortic compliance in nine premenopausal women before and after menotropin therapy. They demonstrated that a shortterm rise in estrogen induced by menotropin treatment was associated with an increase in aortic compliance. Aortic size was not significantly increased within this time frame. Bogren and colleagues55 used CMR to study pulmonary artery distensibility in healthy subjects (Fig. 29-8) and in patients with pulmonary arterial hypertension (Fig. 29-9). The distensibility was found to be significantly lower in pulmonary arterial hypertension than in normal subjects, but there was no age-related difference. The results also demonstrated a small retrograde flow (2%) in the pulmonary trunk of normal subjects close to the pulmonic valve. Antegrade plug flow occurred in most normal subjects but varied among individuals. There were also other variations in the flow pattern among normal individuals. All patients with pulmonary arterial hypertension had a markedly irregular antegrade and retrograde flow and a large retrograde flow (average 26%).

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been based on carotid ultrasound measurements. Forbat and colleagues44 measured aortic compliance, coronary calcification, and carotid intimal-medial thickness during reduction of cholesterol level in hypercholesterolemic patients with and without coronary artery disease. All received fluvastatin for 1 year. Aortic compliance was assessed by using CMR, and the coronary calcification score was determined by electron beam computed tomography. Carotid intimal-medial thickness was measured by carotid ultrasound. The authors showed an improvement in aortic compliance over 1 year, which indicates that the lipid changes induced by fluvastatin (an increase in highdensity lipoprotein level, decrease in very-low-density lipoprotein level, and improvement in low: high-density lipoprotein ratio) beneficially influenced vascular pathophysiology. In the patients who were studied with carotid ultrasound means, carotid intimal-medial thickness decreased from 1.09 to 0.87 mm (p ¼ 0.004), corroborating these results. Kupari and colleagues45 measured aortic elastic modulus by CMR in asymptomatic subjects and correlated these measurements with physical activity, ethanol consumption, systolic blood pressure, fasting blood lipids, and serum insulin. They showed that the average value of the ascending and descending aortic elastic modulus was associated positively and statistically significantly with blood pressure, physical inactivity, serum insulin, and high-density lipoprotein. The elastic modulus was associated negatively with the ratio of low-density lipoprotein cholesterol to highdensity lipoprotein cholesterol. No association between aortic elastic modulus and either smoking or ethanol consumption was demonstrated in this study. The same group demonstrated a higher aortic elastic modulus in patients with Marfan syndrome than in healthy subjects, indicating a relative decrease in the distensibility of the thoracic aorta.46 Kupari and colleagues also demonstrated that aortic flow wave velocity was more reproducible (interobserver and intraobserver) than measurement of the pulsatile aortic area change or the elastic modulus. However, interstudy reproducibility has not been tested.46 Adams and colleagues demonstrated abnormal aortic distensibility and stiffness index in patients with Marfan syndrome using CMR.47 These findings were confirmed by Groenink and associates using a CMR derived measure of distensibility and aortic PWV. Arrival time was determined by the point at which flow had reached half its maximum value.48 Beta-adrenergic blocking agents may reduce the rate of aortic root dilation and the development of aortic complications in patients with Marfan syndrome. This may be due to beta-blocker-induced changes in aortic stiffness. To investigate this, Groenink and colleagues used CMR to measure aortic distensibility and aortic pulse wave velocity to assess aortic stiffness in Marfan syndrome and healthy volunteers before and after beta-blocker therapy.49 They showed that in both groups, mean blood pressure decreased significantly but only the Marfan syndrome patients had a significant increase in aortic distensibility at multiple levels and a significant decrease in pulse wave velocity after beta-blocker therapy. The same group50 sought to explain why some of these patients responded to beta-blockers while others did not. Cross-sectional

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Figure 27-8 Right ventricular outflow tract image (A) and oblique sagittal and transverse image (B) of the main pulmonary artery, showing the sites where main pulmonary artery distensibility was measured. In the bottom row, the change in the main pulmonary artery crosssectional area between diastole (C) and systole (D) demonstrated a large change in a 25-year-old healthy subject.

Changes in the cross-sectional area of the pulmonary artery in a magnitude image from a phase velocity sequence have been measured to derive pulmonary artery PWV,56–58 Values for normal subjects and for patients with pulmonary hypertension have been derived.57 In the latter group, maximum and minimum values for mean pulmonary artery pressure (mPAP) were predicted that “framed” the actual mPAP reliability.58 Finally, Stefanides and colleagues have showed that aortic stiffness has prognostic power in determining the likelihood of future cardiac events (Fig. 29-10).59 This interesting study merits further examination, as it uses a simple marker of disseminated arterial disease that is simple to measure in large populations.

ASSESSMENT OF ENDOTHELIAL FUNCTION Brachial artery reactivity testing (BAR) has been proposed as a biomarker of endothelial function.60 In the normal

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subject, release of a previously occluded brachial blood pressure cuff results in postischemic hyperemia (Fig. 29-11) 61 and subsequent increased shear stress of the local arterial wall. This, in turn, causes the release of endothelium-derived nitric oxide, a potent vasodilator. The extent to which nitric oxide is released (itself dependent on the amount produced and stored locally) can be indirectly assessed by measuring the degree of vessel dilation that results from this provocation. Additionally, endotheliumindependent function can be assessed by using dilation mediated by glyceryl trinitrate (GTN) (normally administered sublingually). The noninvasive, uncomplicated and well-tolerated nature of this examination makes it an attractive endpoint in epidemiologic studies. This has been further strengthened as alterations in flow-mediated response (by ultrasound) have been shown to precede clinical manifestations of disease. These findings suggest a useful role for this biomarker as a screening tool in the future; allowing risk stratification and intervention at an earlier stage than might previously have been thought possible.

27 ASSESSMENT OF THE BIOPHYSICAL MECHANICAL PROPERTIES OF THE ARTERIAL WALL

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Figure 27-9 Right ventricular outflow tract image (A) and oblique sagittal and transverse image (B) of the main pulmonary artery, showing the sites where main pulmonary artery distensibility was measured. In the bottom row, the change in the main pulmonary artery crosssectional area between diastole (C) and systole (D) demonstrated little change in a patient with pulmonary hypertension.

While the peripheral location of the brachial artery allows ultrasound ready access to such a measure, CMR can also be employed in a similar fashion. In this way, BAR can be used as an adjunct, combining with more routine measures of ventricular and vascular function to provide a powerful means by which to evaluate patients in research (Fig. 29-12). CMR offers other advantages over ultrasonography for this measure as it is more reproducible and less operator dependent.62 This has logistical and economic consequences for researchers undertaking such work, as it allows sample size to be reduced without compromising the ability to identify statistically significant changes. Also CMR measures the true cross-sectional area, whereas ultrasound usually measures diameter only. In addition, assessment of reactive hyperemic response using real-time CMR flow imaging has been shown to be associated with cardiovascular increased risk.63 CMR has been used to demonstrate pertubations of endothelial function in a variety of different scenarios. Sorenson and coworkers64 were able to demonstrate impairment of flow-mediated dilation, whose degree was

inversely related to circulating levels of estradiol, in patients who were given depot medroxyprogesterone acetate when compared to controls. Wiesmann and colleagues64 showed that brachial artery flow–mediated dilation was significantly reduced in smokers in comparison with nonsmokers. As with Sorenson and colleagues’ work, impairment of dilation was endothelium-dependent only as degree of GTN-mediated dilation was similar in the two groups. Conversely, reduced endothelium-independent (but not endothelium-dependent) function has been demonstrated in young elite rowers.65

ARTERIAL WALL SHEAR STRESS The use of detailed anatomic model and boundary conditions, such as the inlet and outlet flow, captured in CMR are important for the generation of a patient-specific computational fluid dynamic (CFD) simulation. The CFD Cardiovascular Magnetic Resonance 371

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method allows the calculation of features and properties such as wall shear stress (WSS) and mass transfer rate, which are difficult to measure directly with imaging but are important to the understanding of basic hemodynamics. CFD is the technique of using numerical methods to solve equations that govern the fluid flow. The basic premise of CFD is to split the domain that is to be analyzed into small volumes or elements. For each of these, a set of partial differential equations is solved, which approximates a solution for the flow in order to achieve the basics of conservation of mass, momentum, and energy for each volume or element. The three basic equations are solved simultaneously, with any additional

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equations implemented in a particular model to obtain the flow velocities and pressure and therefore any derived quantities, such as shear stress. CFD has been extensively used to model physiologic flow and arterial WSS, particularly in the carotid bifurcation, in attempts to better understand the mechanism of atherosclerosis in this region66 and to give insight into the pathogenesis of arterial aneurysms.67 WSS is defined as the mechanical frictional force exerted on the vessel wall by flowing blood and is the product of the blood viscosity with the velocity gradient at the vessel wall. Low or oscillating arterial WSS is correlated with atherosclerosis.68 In the aorta, once an aneurysm is formed, fluctuation in blood flow within it can induce vibrations of the aneurysm wall that can contribute to progression and eventual rupture. Volume flow measured by CMR alone is not yet capable of sufficient spatial and temporal resolution for accurate measurement of WSS. Simulation with CFD provides more information than current diagnostic tools. In particular, it allows the determination of the shear stress level, which in turn helps in identifying regions that are susceptible to aneurysm formation, growth, and rupture. The only criterion that has been used so far for the selection of surgical patients has been the maximum diameter and the rate of change of aortic diameter.69 This is based on Laplace’s law for hollow circular pipes, which states that the maximum stress within the arterial wall is proportional to the radius. Works in literature suggest that the peak wall stress is a more reliable parameter for the assessment of the rupture risk of aortic aneurysms. Fillinger and colleagues reported that the peak wall stress in aneurysms has a higher sensitivity and specificity for predicting the rupture than maximum diameter.70 CMR and CFD are promising tools for stress and velocity patterns simulation using patient-specific flow condition. Figures 27-13 and 27-14 show how the stress pattern and values are influenced by the shape of the aneurysm and the presence of intraluminal thrombus.

Cuff release

Cuff inflated 1-minute gap

Baseline Cardiac cycles (R-R ≈ 900 msec)

Figure 27-11 Mean velocity in real time shown graphically for 72 cardiac cycles. The first 5 cardiac cycles are before cuff inflation. Imaging was suspended for the cuff inflation and a 5-minute delay. The next 10 cycles were acquired with the cuff inflated and show a shorter forward peak (i.e., reduced forward flow) in the waveform, although the peak velocity does not reduce. The occlusion cuff was placed more than 10 cm distal to the imaging plane. Note the increased reverse flow during the occlusion. After release of the cuff, the forward peak is longer, and the reverse flow changes to forward flow. Both effects recover to baseline after approximately 40 sec. Source: Mohiaddin RH, Gatehouse PD, Moon JCC, Youssuffidin M, Yang GZ, Firmin DN and Pennell DJ. Assessment of reactive hyperemia using real time zonal echo planar flow imaging. J Cardiovasc Magn Reson. 2002;4:283–287.

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B

C

D

27 ASSESSMENT OF THE BIOPHYSICAL MECHANICAL PROPERTIES OF THE ARTERIAL WALL

A

Figure 27-12 Assessment of brachial artery reactivity. The top panel demonstrates the arrangement of the surface coil and cuff on a subject’s right arm whose motion is restricted by sandbags. The transxial plane in which the brachial artery is at its most superficial (A) is used to prescribe a plane perpendicular to the brachial artery (B). Arterial reactivity is then assessed through peripheral reactive hyperemia (C) and sublingual glyceryl trinitrate (D). Source: Sorenson MB, Collins P, Ong PJL, et al. Long term use of contraceptive depot medroxyprogesterone acetate in young women impairs arterial endothelial function assessed by cardiovascular magnetic resonance. Circulation. 2002;106:1646–1651.

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A

B Smoothed Effective Stress RST CALC Time 1.000

64.50 57.00 49.50 42.00 34.50 27.00 19.50 12.00 4.50

C

Stress pattern at peak systole (units:kPa)

Figure 27-13 A, Surface-rendered contrast-enhanced magnetic resonance angiography in the left anterior oblique and right anterior oblique views in a patient with aortic arch aneurysm at the site of previous coarctation repair (left subclavian artery flap). B, The aortic arch morphology was segmented from the CMR images, and the inflow-outflow flow conditions were extracted from CMR flow mapping. C, Stress pattern at peak systole (units: kilopascals). Regions of high shear stress (arrows) are found in the aortic arch and at the entry of the aneurysmal bulge in the first model, with low values only in the distal region of the aneurysm.

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27 ASSESSMENT OF THE BIOPHYSICAL MECHANICAL PROPERTIES OF THE ARTERIAL WALL

A

B

Smoothed Effective Stress RST CALC Time 1.000

64.50 57.00 49.50 42.00 34.50 27.00 19.50 12.00 4.50

C Figure 27-14 A, Surface rendered contrast-enhanced magnetic resonance angiography in the coronal and sagittal views in a patient with atherosclerotic aneurysm in the distal descending thoracic aorta. B, The aneurysm contains a large crescent-shape mural thrombus. C, Stress pattern at peak systole (units: kilopascals). Low values of shear stress throughout the aneurysm are noted where the thrombus appears to absorb part of the pressure load resulting in low stress in the wall.

CONCLUSION Atherosclerosis consists of two components, the most often discussed being atherosis, relating to plaque genesis and composition. Sclerosis is often the forgotten partner in clinical practice, but measurement of vessel wall stiffening has now been shown to be useful both diagnostically and prognostically. Coupled with CMR’s powerful role in excluding secondary causes of hypertension and accurately assessing ventricular function, CMR is an ideal tool for its investigation, and further studies of the clinical role of parameters

of sclerosis can be expected. The capability of CMR to assess brachial artery reactivity and to measure WSS, in conjunction with CFD, will enhance further its role in the studies of the biophysical properties of the arterial wall.

ACKNOWLEDGMENTS The author acknowledges the helpful contributions of Peter Gatehouse, PhD, and William Bradlow, MD, in the preparation of this manuscript.

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References 1. Band W, Goedhard WJ, Knoop AA. Comparison of effects of high cholesterol intake on viscoelastic properties of the thoracic aorta in rats and rabbits. Atherosclerosis. 1973;18:163–172. 2. Banga I, Balo J. Elasticity of the vascular wall. 1 The elastic distensibility of the human carotid as a function of age and arteriosclerosis. Acta Physiol Acad Sci Hung. 1961;20–21:237–247. 3. Farrar DJ, Green HD, Wanger WD, Bond MG. Reduction in pulse wave velocity and improvement of aortic distensibility accompanying regression of atherosclerosis in Rhesus monkey. Circ Res. 1980;47:425–432. 4. Farrar DJ, Bond GM, Riley WA, Sawyer JK. Anatomic correlates of aortic pulse wave velocity and carotid artery elasticity during atherosclerosis progression and regression in monkeys. Circulation. 1991;83:1754–1763. 5. Ohtsuka S, Kakihana M, Watanabe H, et al. Chronically decreased aortic distensibility causes deterioration of coronary perfusion during increased left ventricular contraction. J Am Coll Cardiol. 1994;24:1406–1414. 6. Isnard RN, Pannier BM, Laurent S, et al. Pulsatile diameter and elastic modulus of the aortic arch in essential hypertension: a non-invasive study. J Am Coll Cardiol. 1989;13:399–405. 7. Handler CE, Child A, Light NM. Mitral valve prolapse, aortic compliance, and skin collagen in joint hypermotility syndrome. Br Heart J. 1985;54:501–508. 8. Lehman ED, Gosling RG, Sonksen PH. Arterial compliance in diabetes. Diabetes Care. 1986;9:27–31. 9. Lehman ED, Watts GF, Gosling RG. Aortic distensibility and hypercholesterolaemia. Lancet. 1992;340:1171–1172. 10. Lehman ED, Hopkins KD, Weissberger AJ, Gosling RG, Sonksen PH. Aortic distensibility in growth hormone deficient adults. Lancet. 1993;341:309. 11. Harris P, Heath D. The relation between structure and function in the blood vessels of the lung in pulmonary hypertension. In: Harris P, Heath D, eds. The Human Pulmonary Circulation. 3rd ed. London: Churchill Livingstone; 1986:284–297. 12. Reuben SR. Compliance of the human pulmonary arterial system in disease. Circ Res. 1971;29:40–50. 13. Harris P, Heath D, Apostolopoulos A. Extensibility of the pulmonary trunk in heart disease. Br Heart J. 1965;27:660–666. 14. Paz R, Mohiaddin RH, Longmore DB. Magnetic resonance assessment of pulmonary trunk: anatomy, flow, pulsatility and distensibility. Eur Heart J. 1993;14:1524–1530. 15. Mohiaddin RH, Paz R, Theodoropolus S, Firmin DN, Longmore DB, Yacoub MH. Magnetic resonance characterization of pulmonary arterial blood flow following single lung transplantation. J Thorac Cardiovasc Surg. 1991;101:1016–1023. 16. Mukherjee DP, Kagan HM, Jordan RE, Franzblau C. Effect of hydrophobic elastin ligands on the stress-strain properties of elastin fibers. Connect Tissue Res. 1976;4:177. 17. Harkness ML, Harkness RD, McDonald DA. The collagen and elastin content of the arterial wall in the dog. Proc Roy Soc [Biol]. 1957;146:541–551. 18. Lee RT, Kamm RD. Vascular mechanics for the cardiologist. J Am Coll Cardiol. 1994;23:1289–1295. 18a. Peterson LN, Jensen RE, Parnell R. Mechanical properties of arteries in vivo. Circ Res. 1968;8:622–639. 19. Bergel DH. The dynamic elastic properties of the arterial wall. J Physiol. 1961;156:458–469. 20. Hardung V. Vergleichende messungen der dynamischen elastizitat und viskositat von blutegfassen, kautschuk und synthetischen elastomeren. Helv Physiol Acta. 1953;11:194–211. 21. Learoyd BM, Taylor MG. Alterations with age in the visco-elastic properties of human arterial walls. Circ Res. 1966;18:278–292. 22. Remington JW. Hysteresis loop, behaviour of the aorta and other extensible tissues. Am. J Physiol. 1955;180:83–95. 23. Bramwell JC, Hill AV, McSwiney BA. The velocity of the pulse wave in man in relation to age as measured by hot-wire sphygmograph. Heart. 1923;10:233–255. 24. Hallock P. Arterial elasticity in man in relation to age as evaluated by the pulse wave velocity method. Arch Int Med. 1934;54:770–798. 25. Remington JW. Pressure-diameter relations of the in vivo aorta. Am J Physiol. 1962;203:440–448.

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26. Luchsinger PC, Sachs M, Patel D. Pressure-radius relationship in large blood vessels of man. Circ Res. 1962;11:885–887. 27. Stefanadis C, Stratos C, Boudoulas H, et al. Distensibility of the ascending aorta: comparison of invasive and non-invasive techniques in healthy men and in men with coronary artery disease. Eur Heart J. 1990;11:990–996. 28. Gosling RG, King DH. Arterial assessment by Doppler-shift ultrasound. Proc Roy Soc Med. 1974;67:447–449. 29. Kok WEM, Peters RJG, Prins MH, et al. Contribution of age and intimal lesion morphology to coronary artery wall mechanics in coronary artery disease. Clin Sci. 1995;89:239–246. 30. Hanrath P, Heinzt B, Dahl V, et al. Evaluation of segmental elastic properties of the aorta in normotensive and medically treated patients by intravascular ultrasound. In: Boudoulas P, Toutouzas PK, Wooley C, eds. Functional Abnormality of the Aorta. New York: Futura Publishing Company; 1996:221–231. 31. Mohiaddin RH, Underwood SR, Bogren HG, et al. Regional aortic compliance studied by magnetic resonance imaging: the effects of age, training, and coronary artery disease. Br Heart J. 1989;62:90–96. 32. Chien D, Saloner D, Laub G, Anderson CM. High resolution cine MRI of vessel distension. J Comput Assist Tomogr. 1994;18:576–580. 33. Rueckert D, Burger P, Yang GZ, Forbat SM, Mohiaddin RH. Automatic tracking of the aorta in cardiovascular MR images using deformable models. IEEE Trans Med Imaging. 1997;16:581–590. 34. Mohiaddin RH, Firmin DN, Longmore DB. Age-related changes of human aortic flow wave velocity measured non-invasively by magnetic resonance imaging. J Appl Physiol. 1993;74:492–497. 35. Tarnawski M, Cybulski G, Doorly D, Dumoulin C, Darrow R, Caro C. Noninvasive determination of local wavespeed and distensibility of the femoral artery by comb-excited Fourier velocity-encoded magnetic resonance imaging: measurements on athletic and nonathletic human subjects. Heart Vessels. 1994;9:194–201. 36. Dumoulin CL, Doorly DJ, Caro CG. Quantitative measurement of velocity at multiple positions using comb excitation and Fourier velocity encoding. Magn Reson Med. 1993;29:44–52. 37. Hardy CJ, Bolster BD, McVeigh ER, Adams WJ, Zerhouni EA. A onedimensional velocity technique for NMR measurement of aortic distensibility. Magn Reson Med. 1994;31:513–520. 38. Hardy CJ, Bolster Jr BD, McVeigh ER, Iben IE, Zerhouni EA. Pencil excitation with interleaved Fourier velocity encoding: NMR measurement of aortic distensibility. Magn Reson Med. 1996;35:814–819. 39. Bolster Jr BD, Atalar E, Hardy CJ, McVeigh ER. Accuracy of arterial pulse-wave velocity measurement using MR. J Magn Reson Imaging. 1998;8:878–888. 40. Mohiaddin RH, Underwood SR, Bogren HG, et al. Regional aortic compliance studied by magnetic resonance imaging: the effects of age, training, and coronary artery disease. Br Heart J. 1989;62:90–96. 41. Rees RSO, Somerville J, Ward C, et al. Magnetic resonance imaging in late post-operative assessment of coarctation of the aorta. Radiology. 1989;173:499–502. 42. Manzara CC, Mohiaddin RH, Pennell DJ, et al. Magnetic resonance assessment of thoracic aorta in Marfan’s syndrome. American Heart Association, Dallas. Circulation. 1990;82(suppl III):497(Abstract). 43. Rogers WJ, Hu YL, Coast D, et al. Age-associated changes in regional aortic pulse wave velocity. JACC. 2001;38(4):1123–1129. 44. Forbat SM, Naoumova RP, Sidhu PS, et al. The effect of cholesterol reduction with fluvastatin on aortic compliance, coronary calcification and carotid intimal-medial thickness: a pilot study. J Cardiovasc Risk. 1998;5:1–10. 45. Kupari K, Hekali P, Keto P, et al. Relation of aortic stiffness to factors modifying the risk of atherosclerosis in healthy persons. Arterioscler Thromb. 1994;14:386–394. 46. Kupari K, Keto P, Hekali P, Poutanen V, Savolainen A, StandertskjoldNordenstam CG. Cine magnetic resonance imaging in the assessment of aortic distensibility. In: Boudoulas P, Toutouzas PK, Wooley C, eds. Functional Abnormality of the Aorta. New York: Futura Publishing Company; 1996:247–268. 47. Adams JN, Brooks M, Redpath TW, et al. Aortic distensibility and stiffness index measured by magnetic resonance imaging in patients with Marfan’s syndrome. Br Heart J. 1995;73:265–269.

59. Stefanides C, Dernellis J, Tsiamis E, et al. Aortic stiffness as a risk factor for recurrent acute coronary events in patients with ischaemic heart disease. Eur Heart J. 2000;21:390–396. 60. Deanfield J, Donald A, Ferri C, et al. Working Group on Endothelin and Endothelial Factors of the European Society of Hypertension. Endothelial function and dysfunction. Part II: Association with cardiovascular risk factors and diseases. A statement by the Working Group on Endothelins and Endothelial Factors of the European Society of Hypertension. J Hypertens. 2005;23(2):233–246. 61. Mohiaddin RH, Gatehouse PD, Moon JCC, et al. Assessment of reactive hyperemia using real time zonal echo planar flow imaging. J Cardiovasc Magn Reson. 2002;4:283–287. 62. Sorenson MB, Collins P, Ong PJL, et al. Long term use of contraceptive depot medroxyprogesterone acetate in young women impairs arterial endothelial function assessed by cardiovascular magnetic resonance. Circulation. 2002;106:1646–1651. 63. Schwitter J, Oelhafen M, Wyss BM, et al. 2D spatially selective real time magnetic resonance imaging for the assessment of microvasculature function and its relation to the cardiovascular risk profile. J Cardiovasc Magn Reson. 2006;8:759–769. 64. Wiesmann F, Petersen SE, Leeson PM, et al. Global impairment of brachial, carotid, and aortic vascular function in young smokers: direct quantification by high-resolution magnetic resonance imaging. J Am Coll Cardiol. 2004;16;44(10):2056–2064. 65. Petersen SE, Wiesmann F, Hudsmith LE, et al. Functional and structural vascular remodeling in elite rowers assessed by cardiovascular magnetic resonance. J Am Coll Cardiol. 2006;48(4):790–797. 66. Milner JS, Moore JA, Rutt BK, Steinman DA. Hemodynamics of human carotid artery bifurcations: computational studies with models reconstructed from magnetic resonance imaging of normal subjects. J Vasc Surg. 1998;28(1):143–156. 67. Borghi A, Wood NB, Mohiaddin RH, Xu XY. 3D geometric reconstruction of thoracic aortic aneurysms. Biomed Eng Online. 2006;5:59. 68. Malek AM, Alper SL, Izumo S. Hemodynamic shear stress and its role in atherosclerosis. JAMA. 1999;282:2035–2042. 69. Elefteriades JA. Natural history of thoracic aortic aneurysms: indications for surgery and surgical versus nonsurgical risks. Ann Thorac Surg. 2002;74:S1877–S1880. 70. Fillinger MF, Marra PS, Raghavan ML, Kennedy EF. Prediction of rupture risk in abdominal aortic aneurysm during observation: wall stress versus diameter. J Vasc Surg. 2003;37(4):724–732.

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48. Groenink M, de Roos A, Mulder BJ, et al. Biophysical properties of the normal-sized aorta in patients with Marfan syndrome: evaluation with MR Flow mapping. Radiology. 2001;219:535–540. 49. Groenink M, de Roos A, Mulder BJ, Spaan JA, van der Wall EE. Changes in aortic distensibility and pulse wave velocity assessed with magnetic resonance imaging following beta-blocker therapy in the Marfan syndrome. Am J Cardiol. 1998;82:203–208. 50. Nollen GJ, Westerhof BE, Groenink M, Osnabrugge A, Van der Wall EE, Mulder BJM. Aortic pressure-area relation in Marfan patients with and without B blocking agents: a new invasive approach. Heart. 2004;90:314–318. 51. Savolainen A, Keto P, Poutanen VP, et al. Effects of angiotensin-converting enzyme inhibition versus beta-adrenergic blockade on aortic stiffness in essential hypertension. J Cardiovasc Pharmacol. 1996;27:99–104. 52. Honda T, Hamada M, Shigematsu Y, Matsumoto Y, Matsuoka H, Hiwada K. Effect of antihypertensive therapy on aortic distensibility in patients with essential hypertension: comparison with trichlormethiazide, nicardipine and alacepril. Cardiovasc Drugs Ther. 1999;13:339–346. 53. Resnick LM, Militianu D, Cunnings AJ, Pipe JG, Evelhoch JL, Soulen RL. Direct magnetic resonance determination of aortic distensibility in essential hypertension: relation to age, abdominal visceral fat, and in situ intracellular free magnesium. Hypertension. 1997;30 (3 Pt 2):654–659. 54. Chelsky R, Wilson RA, Morton MJ, et al. Alteration of ascending thoracic aorta compliance after treatment with menotropin. Am J Obstet Gynecol. 1997;176:1255–1259. 55. Bogren HG, Klipstein RH, Mohiaddin RH, et al. Pulmonary artery distensibility and blood flow patterns: a magnetic resonance study of normal subjects and of patients with pulmonary arterial hypertension. Am Heart J. 1989;118:990–999. 56. Laffon E, Bernard V, Montaudon M, Marthan R, Barat JL, Laurent F. Tuning of pulmonary arterial circulation evidenced by MR phase mapping in healthy volunteers. J Appl Physiol. 2001;90(2):469–474. 57. Laffon E, Laurent F, Bernard V, De Boucaud L, Ducassou D, Marthan R. Noninvasive assessment of pulmonary arterial hypertension by MR phase-mapping method. J Appl Physiol. 2001;90 (6):2197–2202. 58. Laffon E, Vallet C, Bernard V, et al. A computed method for noninvasive MRI assessment of pulmonary arterial hypertension. J Appl Physiol. 2004;96(2):463–468.

Cardiovascular Magnetic Resonance Assessment of Right Ventricular Anatomy and Function Alicia M. Maceira and Dudley J. Pennell

Accurate noninvasive assessment of right ventricular (RV) mass and systolic function is important in several pathologies, such as grown-up congenital heart disease, pulmonary hypertension, and arrhythmogenic RV cardiomyopathy. This chapter aims to summarize the features of the normal right ventricle, briefly describe cardiovascular magnetic resonance (CMR) techniques for assessing RV dimensions and function, and give reference values for the assessment of the right ventricle.

NORMAL RIGHT VENTRICULAR ANATOMY The right ventricle is a thin, highly trabeculated structure that is triangular in form and, on gross inspection, appears to be wrapped around the left ventricle. The anterosuperior wall of the right ventricle is rounded and convex, its inferior surface is flattened and forms a small part of the diaphragmatic surface of the heart, and its posterior wall is formed by the ventricular septum, which bulges into the right ventricle, owing to the much greater left ventricular (LV) systolic pressure,1 so a transverse section of the cavity presents a semilunar outline. The right ventricle has a continuum of muscle bands that rotate by approximately 160 from the epicardium to the endocardium.2 The principal axis of these fibers is oblique to the long axis of the right ventricle. In the normal adult, the total RV free wall mass is 26 5 g/m2. The right ventricle has several distinctive features. In its upper left portion, there is a conical pouch called the conus arteriosus or infundibulum, from which the pulmonary artery arises. A tendinous band connects the posterior surface of the conus arteriosus to the aorta. Also, the RV wall is thinner than the LV wall, the proportion between them being as 1 to 3;3 it is thickest at the base and gradually becomes thinner toward the apex. The whole inner surface except the conus arteriosus is covered by more or less prominent muscular columns called trabeculae carneae and from some of them (papillary muscles), the chordae tendinae connect the myocardium to the tricuspid valve, which is more apically placed than the septal leaflet of

the mitral valve. Finally, a muscular band frequently extends from the base of the anterior papillary muscle to the ventricular septum. This band is considered to prevent overdistension of the ventricle and is called the moderator band.4 The depictions of the moderator band, the infundibulum, and the different levels of insertion of the tricuspid and mitral septal leaflets are important diagnostic features for identification of the right ventricle, which can be difficult in some congenital cardiomyopathies.

Importance of Assessing Right Ventricular Dimensions and Function The measurement of RV dimensions, morphology, and function is important in several situations, such as congenital heart disease, LV heart failure, pulmonary hypertension, pulmonary embolism, valvular heart disease, lung disease, and arrhythmogenic RV cardiomyopathy. RV failure may result from conditions that lead to impaired RV contractility, such as RV infarction, right-sided cardiomyopathies, or severe sepsis; RV pressure overload, including pulmonic stenosis, pulmonary primary hypertension, and pulmonary hypertension with left heart disease, lung disease, or thromboembolic disease; and RV volume overload, for instance, tricuspid regurgitation. Many disorders, such as corrected and uncorrected adult congenital heart disease and intracardiac shunts, may result in right ventricle failure through a combination of pathophysiologic mechanisms. Also, decompensated right ventricle (both acute and acute-on-chronic) is increasing as the prevalence of predisposing conditions grows.5 The prognostic value of RV function has been shown in several conditions such as left heart failure due to coronary or valvular heart disease,6,7 pulmonary hypertension,8 or myocardial infarction.9 Thus, the early detection of RV dysfunction can have an impact on therapeutic decision making and on prognosis. Finally, improved understanding of the RV response to pressure and volume overload might lead to more optimal surgical and medical treatments.

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CHAPTER 28

RIGHT VENTRICULAR AND CONGENITAL HEART DISEASE

Techniques for Assessing Right Ventricular Dimensions and Function Angiography used to be the gold standard for assessment of RV volumes and regional and global function. But this technique is invasive, involves ionizing radiation and the use of potentially nephrotoxic contrast, and is not as accurate as CMR.10 Echocardiography and radionuclide ventriculography have been used for the assessment of RV dimensions and function. More recently, “nongeometric” techniques such as three-dimensional (3D) echocardiography, CMR, and multidetector-row computed tomography (CT) permit accurate assessment of RV volumes, function, and mass.

Echocardiography Echocardiography is the most frequently used technique for assessing the right ventricle. and it can be used bedside in very ill patients. It provides information about RV morphology, dimensions, septum convexity, function, tricuspid regurgitation, and estimates of pulmonary arterial pressure and RV pressure.11 But the assessment of the right ventricle with echocardiography has several limitations. First, the location of the right ventricle behind the sternum restricts the window that can be accessed by the ultrasound beam. Second, the complex shape and thin walls of the right ventricle make it necessary to image the right ventricle from several projections. Third, the thick trabeculations in the chamber may be confused with a thrombus, a tumor, or hypertrophic cardiomyopathy. Finally, there is a lack of accurate mathematical models to quantify RV mass and volumes with M-mode or twodimensional (2D) echocardiography, as quantitative measurements are based on geometric assumptions that do not apply to the right ventricle. Other indicators of RV function are Doppler-derived indices such as the myocardial performance index,12 tissue Doppler measurements of myocardial velocities and time intervals, and strain and strain rate measurements of contractility.13 Transesophageal echocardiography is another echocardiographic method of RV assessment, but it is semi-invasive, is not well suited for evaluation of anteriorly positioned RVs, and requires special skills. Three-dimensional echocardiography has emerged as a more accurate and reproducible approach to ventricular quantitation, mainly by avoiding the use of geometric assumptions of the ventricular shape. Three methods have been proposed for the acquisition of temporal and positional data: real-time volumetric scanning, the use of positional locators or free-hand scanning, and rotational systems.14 Real-time 3D echocardiography is an on-line acquisition of a 3D dataset of the heart without the need for electrocardiographic and respiratory gating, which has a great potential for immediate assessment of ventricular function. However, these methods need a stable cardiac rhythm and constant cardiac function during image acquisition, and there are other practical problems, such as full cardiac visualization, good-quality endocardial border recognition for manual endocardial 382 Cardiovascular Magnetic Resonance

tracing, and time consumption. Faster data acquisition, acceleration of data processing and reconstruction, use of automatic border detection algorithms, improvement of spatial resolution, and development of stable intravenous contrast agents that enhance the endocardial delineation are being developed that should provide automatic, even on-line, volume measurement. Threedimensional echocardiography has been compared with CMR for the evaluation of RV function, and improved results in comparison with 2D echocardiography have been obtained.15,16 Nonetheless, 3D echocardiography has been used mainly for the left ventricle, and little has been reported on 3D echocardiography of the RV.

Radionuclide Angiography This technique provides a reliable quantitative measurement of ventricular function not based on geometric assumptions with good agreement with CMR.17 This technique works well for the left ventricle but not so well for the right ventricle, owing to problems such as the limited count numbers in this chamber and the overlap of other cardiac chambers.18–20 It also has disadvantages, such as poor resolution compared to other imaging modalities, the use of ionizing radiation, and the need for an adequate bolus injection for first pass studies and a regular rhythm. Therefore, it has been of limited use for the study of the right ventricle so far.

Multislice Computed Tomography Multislice CT (MSCT) is emerging as an alternative technique, especially for patients with implantable devices (a contraindication for CMR). However, MSCT uses ionizing radiation and potentially nephrotoxic contrast and requires a low and stable heart rate for image acquisition.21

CMR CMR has some important advantages over other imaging techniques, which have led to the growing enthusiasm for its use. CMR offers accurate and reproducible tomographic, static, or cine images of high spatial and temporal resolution in any desired plane without exposure to contrast agents or ionizing radiation. It allows the acquisition of true RV short axis images encompassing the entire RV with high spatial and temporal resolution, thereby providing highly accurate and reproducible quantitative RV mass and functional data regardless of its position in the thorax.22–25 Nowadays, this technique is considered the gold standard for quantitative assessment of RV volume, mass, and function.

Imaging Strategies for Cardiovascular Magnetic Resonance of the Right Ventricle Before the study begins, it is essential to obtain an accurate electrocardiographic gating with minimal ectopy and to instruct the patient in breath holding. Sometimes oxygen may be applied to improve breath hold length. Ventricular

ectopy can be a problem, mainly in patients with congenital heart disease or with suspicion of arrhythmogenic RV cardiomyopathy (ARVC). If this condition is present, pretreatment with an antiarrhythmic agent should be considered. Spin echo (black-blood) sequences (including turbo spin echo, half-Fourier acquisition single-shot turbo spin echo, or spin echo-echo planar imaging) are used for anatomic assessment (Fig. 28-1) as well as to rule out possible fatty replacement/infiltration of the RV free wall as can be seen in RV dysplasia.26 Functional evaluation of the RV is performed by using gradient echo (white-blood) cine imaging. In the past, to achieve full 3D coverage of the ventricle with conventional nonsegmented free breathing gradient echo cine sequences, a total scanning time of 30 minutes or more was required. On modern scanners with segmented fast imaging, a single cine can be acquired in just one breath hold of about 8 to 10 sec, allowing the whole stack of images to be acquired in 5 to 10 minutes,27 thereby reducing breathing and movement artifact (Fig. 28-2). Moreover, real-time steady state free precession (SSFP) imaging can acquire the whole stack in just one breath hold with acceptable accuracy and image quality.28 In patients who are unable to hold their breath consistently, solutions using the same sequence with more signal averages or combined with navigator echo are successful with free breathing,29 with a slight increase in the scanning time. The RV is well shown in the transaxial plane from the tricuspid valve to the apex. This plane has been used for analysis of RV function and has been shown to yield good

agreement with pulmonary flow and LV stroke volume30,31 as well as better interobserver and intraobserver reproducibility than the short axis plane32 (Fig. 28-3). However, it can be problematic because it is subject to partial volume effects in the anterior and inferior walls. The question has been raised as to whether the RV volumes should always be measured in the axial orientation. Yet the interstudy reproducibility of RV measurements in the short axis orientation is good,33 and in practice, this orientation allows both the LV and RV dimensions to be measured simultaneously. Simpson’s rule is used for measuring volumes and function. A standardized method of combined ventricular functional analysis is described in Chapters 11 and 14. Briefly, a stack of contiguous tomographic slices is acquired that encompass the entire ventricles. It is critical to ensure that the most basal part of the free wall of the right ventricle in diastole is included in the most basal cine slice, as it can be easily truncated. Manual or semiautomatic planimetry of the endocardial borders of each ventricle at both end-diastole and end-systole and of the epicardial borders at enddiastole is done. Care must be taken to exclude the right and left atria as they come into the basal imaging planes during systole. The ventricular volume is equal to the sum of the endocardial areas multiplied by the distance between the centers of each slice (Fig. 28-4). The stroke volume is equal to the difference between the end-diastolic volume (EDV) and the end-systolic volume (ESV). The ejection fraction (EF) is calculated as the stroke volume divided by the EDV. The mass is calculated as the volume Cardiovascular Magnetic Resonance 383

28 CARDIOVASCULAR MAGNETIC RESONANCE ASSESSMENT OF RIGHT VENTRICULAR ANATOMY AND FUNCTION

Figure 28-1 Short axis and transaxial spin echo images of the heart in a healthy subject. The right ventricular free wall is perfectly depicted between the black-blood pool and the black pericardium. Note that there is epicardial fat deposition around the right coronary artery (anterior atrioventricular groove), left anterior descending artery (anterior interventricular groove), and left circumflex (posterior atrioventricular groove). It is important to know the normal patterns of fat distribution to prevent false positive reading of scans in assessing patients for possible fat infiltration in arrhythmogenic right ventricular cardiomyopathy.

RIGHT VENTRICULAR AND CONGENITAL HEART DISEASE

RV

LV

Figure 28-2 SSFP cine images in eight short axis planes from cines acquired from base to apex. The series of short axis slices begins in the atrioventricular groove in diastole and goes through the ventricles down to the apex, thus covering the whole left ventricle and right ventricle. In the most basal slice (upper left), the pulmonary artery will appear in systole because of atrioventricular ring descent. The high contrast between blood and myocardium allows the drawing of the endocardial and epicardial borders for volume and mass determinations. For the images to be accurate and reproducible, it is important to use always the same slice thickness and the same gap between slices. Also, retrospective gating should be used in order to acquire the whole cardiac cycle. LV, left ventricle; RV, right ventricle.

PA PV

RV

A Figure 28-3 A, To plan the transversal slices, a cine sequence through the right ventricular (RV) outflow tract aligned in an oblique sagittal plane should be acquired. This view is useful for examining the function of the right ventricular free wall from the apex to the pulmonary valve (PV) and visualizing pulmonary regurgitation. (Continued) 384 Cardiovascular Magnetic Resonance

PA

RV LV

RA LA

B Figure 28-3—Cont’d B, Multiple transaxial steady-state free precession cine sequences in contiguous planes should be acquired from the pulmonary valve level down to the inferior right ventricular wall, using the right ventricular outflow tract view to pilot them. These planes are invaluable for examining regional right ventricular wall motion. Again, the slice thickness and gap between slices should be consistent among studies. Two features of the normal right ventricle are seen in this transversal stack: the muscular outflow tract and the more apical insertion of the tricuspid valve in the septum compared with the septal leaflet of the mitral valve. Ao, aorta; LA, left atrium; LV, left ventricle; PA, pulmonary artery; RA, right atrium.

of tissue occupied by the free wall multiplied by an assumed density of 1.05 g/cc. The volumes obtained by this method are independent of geometric assumptions and dimensionally accurate.34 Mass measurements agree well with autopsy studies.35,36 RV parameters obtained with this technique are reproducible, with an interobserver variability of 6.3% for EDV, 8.6% for ESV, 7% for stroke volume, 4.4% for EF, and 7.8% for RV mass and an intraobserver variability of 3.6% for EDV, 6.5% for ESV, 5.9% for stroke volume, 4% for EF, and 5.7% for RV mass.37 With semiautomated analysis, all images throughout the cardiac cycle can be planimetered, and diastolic function parameters can also be calculated. If necessary, other CMR techniques should be used in the assessment of the right ventricle. Flow velocity maps allow for accurate assessment of pulmonary valve regurgitation (regurgitant fraction) and systemic-to-pulmonary flow ratio. Magnetic resonance angiography must be done in case assessment of great vessel anatomy is advisable, such as in a number of congenital cardiac conditions.38 CMR with late gadolinium enhancement can detect myocardial fibrosis in both ischemic and nonischemic cardiomyopathies.39,40

NORMAL RIGHT VENTRICULAR VOLUMES AND SYSTOLIC FUNCTION There have been a number of human series describing the normal characteristics of RV size and function using

autopsy,41,42 echocardiography,43–45 X-ray angiography,19 radionuclide angiography,46 CMR,38,47–51 and ultrafast CT.52,53 We have reported on RV reference parameters for mass, volumes, and systolic and diastolic function using cine SSFP CMR techniques and analysis from 120 healthy adult subjects.37 These gender-specific data are summarized in Tables 28-1 to 28-5. We observed that many clinical parameters of RV volumes and systolic and diastolic function are significantly dependent on gender, age, and body surface area (BSA). On multivariable analysis, BSA was found to significantly influence RV mass, EDV, ESV, stroke volume, and early tricuspid peak filling rate (PFRE). Gender had a significant independent influence on absolute and normalized RV mass, EDV, and stroke volume (Figs. 28-5 and 28-6). It was also an independent predictor of absolute and normalized active tricuspid peak filling rate (PFRA, PFRA/BSA). There was a significant decrease with age of normalized RV mass and of absolute and normalized EDV and ESV in both males and females. There was a significant increase with age in absolute right ventricular ejection fraction (RVEF) in both males and females and a significant increase in normalized EF in males. For diastolic function, absolute and normalized PFRE decreased significantly with age in males and females, while absolute and normalized PFRA increased in males. Accordingly, PFRE/ PFRA decreased significantly. In a multivariable analysis, age was an independent predictor of absolute and normalized ventricular mass and volumes (EDV, ESV, stroke volume, EDV/BSA, ESV/BSA, stroke volume/BSA) and of EF. It was also an independent predictor of diastolic variables (PFRE, PFRA, PFRE/ PFRA, PFRE/EDV, PFRA/EDV, PFRE/BSA, PFRA/BSA). Thus, the interpretation Cardiovascular Magnetic Resonance 385

28 CARDIOVASCULAR MAGNETIC RESONANCE ASSESSMENT OF RIGHT VENTRICULAR ANATOMY AND FUNCTION

Ao

RIGHT VENTRICULAR AND CONGENITAL HEART DISEASE

RV

LV

A

B Figure 28-4 Summation of discs is used to quantify ventricular volumes and mass. The ventricular volume is equal to the sum of the endocardial areas multiplied by the distance between the centers of each slice, both for the end-diastolic volume (A) and the end-systolic volume (B). The ventricular mass is calculated as the sum of the epicardial minus the endocardial areas multiplied by the distance between the centers of each slice, as shown in A. LV, left ventricle; RV, right ventricle.

Table 28-1 SSFP Right Ventricular Volumes, Systolic Function and Mass (Absolute and Normalized to Body Surface Area) by Age Decile (Mean, 95% Confidence Interval): Males 20–29 Years

30–39 Years

40–49 Years

50–59 Years

60–69 years

70–79 Years

160 (111–210) 55 (25–85) 106 (72–140) 66 (53–79) 65 (37–94)

155 (105–205) 50 (20–80) 105 (71–139) 68 (55–81) 63 (35–92)

150 (100–200) 46 (16–76) 104 (70–138) 70 (57–83) 62 (33–90)

EDV (mL) SD 25.4 ESV (mL) SD 15.2 SV (mL) SD 17.4 EF (%) SD 6.5 Mass (g) SD 14.4

Absolute Values 177 (127–227) 171 (121–221) 68 (38–98) 64 (34–94) 108 (74–143) 108 (74–142) 61 (48–74) 63 (50–76) 70 (42–99) 69 (40–97)

166 (116–216) 59 (29–89) 107 (73–141) 65 (52–77) 67 (39–95)

EDV/BSA (mL/m2) SD 11.7 ESV/BSA (mL/m2) SD 7.4 SV/BSA (mL/m2) SD 8.2 EF/BSA (%/m2) SD 4 Mass/BSA (g/m2) SD 6.8

Normalized to Body Surface Area 91 (68–114) 88 (65–111) 35 (21–50) 33 (18–47) 56 (40–72) 55 (39–71) 32 (24–40) 32 (25–40) 36 (23–50) 35 (22–49)

(BSA) 85 (62–108) 30 (16–45) 55 (39–71) 33 (25–41) 34 (21–48)

82 28 54 34 33

(59–105) (13–42) (38–70) (26–42) (20–46)

79 25 53 35 32

(56–101) (11–40) (37–69) (27–42) (19–45)

75 23 52 35 31

(52–98) (8–37) (36–69) (27–43) (18–44)

BSA, body surface area; EDV, end-diastolic volume; EF, ejection fraction; ESV, end-systolic volume; SD, standard deviation; SV, stroke volume.

386 Cardiovascular Magnetic Resonance

20–29 Years

30–39 Years

40–49 Years

50–59 Years

60–69 years

70–79 Years

EDV (mL) SD 21.6 ESV (mL) SD 13.3 SV (mL) SD 13.1 EF (%) SD 6 Mass (g) SD 10.6

Absolute Values 142 (100–184) 136 (94–178) 55 (29–82) 51 (25–77) 87 (61–112) 85 (59–111) 61 (49–73) 63 (51–75) 54 (33–74) 51 (31–72)

130 (87–172) 46 (20–72) 84 (58–109) 65 (53–77) 49 (28–70)

124 (81–166) 42 (15–68) 82 (56–108) 67 (55–79) 47 (26–68)

117 (75–160) 37 (11–63) 80 (55–106) 69 (57–81) 45 (24–66)

111 (69–153) 32 (6–58) 79 (53–105) 71 (59–83) 43 (22–63)

EDV/BSA (mL/m2) SD 9.4 ESV/BSA (mL/m2) SD 6.6 SV/BSA (mL/m2) SD 6.1 EF/BSA (%/m2) SD 5.2 Mass/BSA (g/m2) SD 5.2

Normalized to Body Surface Area 84 (65–102) 80 (61–98) 32 (20–45) 30 (17–43) 51 (39–63) 50 (38–62) 37 (27–47) 38 (27–48) 32 (22–42) 30 (20–40)

76 27 49 38 29

(57–94) (14–40) (37–61) (28–49) (19–39)

72 24 48 39 27

(53–90) (11–37) (36–60) (29–49) (17–37)

68 21 46 40 26

(49–86) (8–34) (34–58) (30–50) (16–36)

64 19 45 41 24

(45–82) (6–32) (33–57) (31–51) (14–35)

BSA, body surface area; EDV, end-diastolic volume; EF, ejection fraction; ESV, end-systolic volume; SD, standard deviation; SV, stroke volume.

Table 28-3 SSFP Right Ventricular Diastolic Function and Atrioventricular Plane Descent (Absolute and Normalized Values) by Age Decile (Mean, 95% Confidence Interval): Males 20–29 Years

30–39 Years

40–49 Years

50–59 Years

60–69 years

70–79 Years

PFRE (mL/s) SD137 PFRA (mL/s) SD 175 PFRE/PFRA SD* 0.49 Septal AVPD (mm) SD 4.1 Lateral AVPD (mm) SD 4.4

Absolute Values 545 (277–814) 491 (223–760) 366 (23–709) 413 (70–756) 1.6 (0.6–2.5) 1.2 (0.3–2.2) 16 (8–24) 15 (7–24) 23 (14–32) 23 (14–31)

438 (169–706) 461 (118–804) 1.0 (0.0–1.9) 15 (7–23) 22 (14–31)

384 (116–652) 508 (165–852) 0.7 ( 0.2–1.7) 14 (6–22) 22 (13–30)

330 (62–599) 556 (213–899) 0.6 ( 0.4–1.5) 14 (6–22) 21 (13–30)

276 (8–545) 604 (260–947) 0.5 ( 0.5–1.4) 13 (5–21) 21 (12–29)

PFRE/BSA (mL/s/m2) SD 71 PFRE/EDV/s SD 0.75 PFRA/BSA (mL/s/m2) SD 94 PFRA/EDV/s SD 1.07 Septal AVPD/long length (%) SD 4.5 Lateral AVPD/long length (%) SD 4.1

Normalized Values 280 (142–419) 252 (114–390) 3.1 (1.6–4.6) 2.8 (1.4–4.3) 190 (6–374) 213 (29–397) 2.1 (0.0–4.2) 2.5 (0.4–4.6) 18 (9–27) 18 (9–27) 23 (15–31) 23 (15–31)

224 (85–362) 2.6 (1.1–4.1) 236 (52–420) 2.9 (0.8–4.9) 17 (9–26) 23 (15–31)

195 (57–334) 2.3 (0.9–3.8) 259 (75–443) 3.2 (1.1–5.3) 17 (8–26) 23 (15–31)

167 (29–306) 2.1 (0.6–3.6) 283 (98–467) 3.6 (1.5–5.7) 17 (8–26) 23 (15–31)

139 (1–277) 1.9 (0.4–3.3) 306 (122–490) 4.0 (1.9–6.1) 16 (8–25) 23 (15–31)

A, active; AVPD, atrioventricular plane descent; BSA, body surface area; E, early; PFR, peak filling rate; SD, standard deviation; SD*, standard deviation of log transformed data.

Table 28-4 SSFP Right Ventricular Diastolic Function and Atrioventricular Plane Descent (Absolute and Normalized Values) by Age Decile (Mean, 95% Confidence Interval): Females 20–29 Years PFRE (mL/s) SD 117 PFRA (mL/s) SD 153 PFRE/PFRA SD* 0.46 Septal AVPD (mm) SD 3.0 Lateral AVPD (mm) SD 3.5

30–39 Years

40–49 Years

50–59 Years

60–69 Years

70–79 Years

Absolute Values 471 (241–701) 419 (189–649) 368 (137–598) 316 (86–546) 264 (34–494) 213 ( 17–443) 355 (54–656) 360 (59–660) 365 (64–665) 370 (69–670) 374 (74–675) 379 (79–680) 1.6 (0.7–2.5) 1.3 (0.4–2.2) 1.0 (0.1–1.9) 0.8 ( 0.1–1.7) 0.7 ( 0.2–1.6) 0.5 ( 0.4–1.4) 16 (10–22) 15 (9–20) 13 (7–19) 12 (6–18) 11 (5–17) 10 (4–16) 22 (15–29) 21 (14–28) 21 (14–28) 20 (13–27) 20 (13–27) 19 (12–26)

Normalized Values 278 (145–411) 247 (114–380) 216 (83–349) PFRE/BSA (mL/s/m2) SD 68 3.4 (1.8–5.1) 3.1 (1.5–4.8) 2.8 (1.2–4.5) PFRE/EDV/s SD 0.85 211 (36–386) 212 (37–388) 214 (39–389) PFRA/BSA (mL/s/m2) SD 89 2.4 (0.4–4.4) 2.6 (0.6–4.6) 2.8 (0.8–4.8) PFRA/EDV/s SD 1.03 Septal AVPD/long length (%) SD 3.9 19 (11–27) 18 (11–26) 17 (10–25) Lateral AVPD/long length (%) SD 4.0 24 (16–32) 24 (16–32) 24 (16–32)

185 (52–318) 2.5 (0.9–4.2) 215 (40–390) 3.0 (1.0–5.0) 17 (9–24) 24 (16–32)

153 (20–286) 2.2 (0.6–3.9) 217 (42–392) 3.2 (1.2–5.2) 16 (8–23) 24 (16–32)

122 ( 11–255) 1.9 (0.3–3.6) 218 (43–393) 3.4 (1.4–5.4) 15 (7–22) 24 (16–31)

A, active; AVPD, atrioventricular plane descent; BSA, body surface area; E, early; PFR, peak filling rate; SD, standard deviation; SD*, standard deviation of log transformed data.

Cardiovascular Magnetic Resonance 387

28 CARDIOVASCULAR MAGNETIC RESONANCE ASSESSMENT OF RIGHT VENTRICULAR ANATOMY AND FUNCTION

Table 28-2 SSFP Right Ventricular Volumes, Systolic Function and Mass (Absolute and Normalized to Body Surface Area) by Age Decile (Mean, 95% Confidence Interval): Females

RIGHT VENTRICULAR AND CONGENITAL HEART DISEASE

Table 28-5 SSFP Right Ventricular Summary Data for All Ages (Mean Standard Deviation, 95% Confidence Interval) EDV (mL) EDV/BSA (mL/m2) ESV (mL) ESV/BSA (mL/m2) SV (mL) SV/BSA (mL/m2) EF (%) EF/BSA (%/m2) Mass (g) Mass/BSA (g/m2) PFRE (mL/s) PFRE/BSA (mL/m2) PFRE/EDV/s PFRA (mL/s) PFRA/BSA (mL/m2) PFRA/EDV/s PFRE/PFRA Septal AVPD (mm) Septal AVPD/long length (%) Lateral AVPD (mm) Lateral AVPD/long length (%)

All

Males

Females

144 23 (98–190) 78 11 (57–99) 50 14 (22–78) 27 7 (13–41) 94 15 (64–124) 51 7 (37–65) 66 6 (54–78) 36 5 (27–45) 48 13 (23–73) 31 6 (19–43) 371 125 (126–615) 202 69 (67–337) 2.6 0.8 (1.0–4.1) 429 168 (99–759) 233 93 (50–415) 3.0 1.0 (1.0–5.1) 0.9 0.47 ( 0.1–1.8) 14 3.6 (6–21) 17 4.2 (9–25) 21 3.9 (13–29) 23 4.0 (15–31)

163 25 (113–213) 83 12 (60–106) 57 15 (27–86) 29 7 (14–43) 106 17 (72–140) 54 8 (38–70) 66 6 (53–78) 34 4 (26–41) 66 14 (38–94) 34 7 (20–47) 405 137 (137–674) 207 70 (68–345) 2.4 0.75 (1.0–3.9) 489 175 (146–833) 250 94 (66–434) 3.1 1.0 (1.0–5.2) 0.8 0.49 ( 0.1–1.8) 15 4.1 (6–23) 17 4.5 (8–26) 22 4.4 (13–30) 23 4.1 (15–31)

126 21 (84–168) 73 9 (55–92) 43 13 (17–69) 25 7 (12–38) 83 13 (57–108) 48 6 (36–60) 66 6 (54–78) 39 5 (29–49) 48 11 (27–69) 28 5 (18–38) 337 117 (107–567) 197 68 (64–330) 2.7 0.85 (1.0–4.3) 368 153 (67–668) 215 89 (40–390) 2.9 1.0 (0.9–5.0) 0.9 0.46 (0.0–1.8) 13 3.0 (7–19) 17 3.9 (9–25) 21 3.5 (14–27) 24 4.0 (16–32)

A, active; AVPD, atrioventricular plane descent; BSA, body surface area; E, early; EDV, end-diastolic volume; EF, ejection fraction; ESV, end-systolic volume; PFR, peak filling rate; SV, stroke volume.

of RV parameters in borderline clinical cases, especially in arrhythmogenic RV cardiomyopathy, cardiovascular shunting, and adult congenital heart disease, should be referred to age-, gender-, and BSA-normalized values to determine normality or severity of abnormality. Another CMR approach that may be particularly valuable for quantifying regional RV free wall systolic function is myocardial tagging,54–59 a technique that enables the clinician to assess the complex mechanism of myocardial contraction and to quantify myocardial strain. Klein and colleagues analyzed the RV free wall motion and contraction in humans with CMR tagging.54 In this study percent segmental shortening (PSS) was obtained to measure the amount of contraction and a vector analysis was used to show the trajectory of the RV free wall in systole. PSS increased through time to an average of 12% across all segments (inferior, mid, and superior wall) at the base, 14% at the mid-ventricle, and 16% at the apex, with a wave of motion toward the septum and outflow tract. Naito and colleagues determined PSS only at the midventricle55 and found a PSS of 6.7% in the superior wall segment and 20% for the midwall segment. Fayad and colleagues reported similar PSS values: 24.7% in the midwall segment of the midventricular slice and 28.7% in the midwall segment of the apical slice.56 A 3D reconstruction of RV contraction with CMR tagging58 shows a primary contraction of the RV tangential to its own surface plane and a circumferential contraction as the RV moves apically, with a twisting motion similar to that described by Klein and colleagues.54 Still, while CMR tagging seems a promising method for the assessment of regional function, especially with stress techniques in ischemic heart disease, it remains to be seen whether it provides clinically relevant information beyond that provided by standard cine CMR. Further studies are needed to better define the clinical role of CMR tagging. 388 Cardiovascular Magnetic Resonance

CARDIOVASCULAR MAGNETIC RESONANCE ASSESSMENT OF RIGHT VENTRICULAR ANATOMY AND FUNCTION IN DISEASE Right Ventricular Assessment in Heart Failure In left heart failure, atrial pressure rises, forcing open a number of pulmonary capillaries. When all reserve capillaries are open, the increase in pulmonary pressure leads to an increased load on the RV. Therefore, the function of the RV during exercise in heart failure is important.60 The prognostic value of RV function in advanced heart failure of various causes has been reported;61,62 therefore, the estimation of RV function is now warranted in the standard evaluation of patients with heart failure, since it is helpful in the clinical assessment and prognostic stratification of such patients. Di Salvo and colleagues studied 67 patients with heart failure who had been referred for cardiac transplantation with ischemic (46%) or dilated (54%) cardiomyopathy.63 An RVEF of 35% or more at rest and with exercise predicted overall survival. Maximal oxygen consumption was also predictive of survival, with a modest correlation between RVEF and maximal oxygen consumption. Also, in patients with moderate heart failure, de Groote and colleagues62 showed that three variables, NYHA classification, percent of maximal predicted VO2, and RVEF, were independent predictors of both survival and event-free cardiac survival. Left ventricular EF and peak VO2 normalized to

RV end-systolic volume/BSA−−Females (mL/m2)

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Age (years) Figure 28-5 Normal values for right ventricular (RV) end-diastolic volume, end-systolic volume, mass, and parameters of diastolic function for females normalized to body surface area (BSA). Cardiovascular Magnetic Resonance 389

28 CARDIOVASCULAR MAGNETIC RESONANCE ASSESSMENT OF RIGHT VENTRICULAR ANATOMY AND FUNCTION

FEMALES 120

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RIGHT VENTRICULAR AND CONGENITAL HEART DISEASE

MALES

+95% CI 2 Mean

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Age (years) Figure 28-6 Normal values for right ventricular (RV) end-diastolic volume, end-systolic volume, mass, and parameters of diastolic function for males normalized to body surface area (BSA). 390 Cardiovascular Magnetic Resonance

Right Ventricular Assessment in Ischemic Heart Disease Isolated RV infarction is relatively rare, but concurrent RV infarction in the setting of an inferior infarction due to proximal right coronary occlusion64,65 occurs in up to half of LV infarctions.66 RV infarction can be detected and evaluated in extent by using late enhancement gadolinium CMR. RV necrosis causes a loss of contractile mass, and if the inferior interventricular septum is involved, there is also a loss of septal augmentation of RV function. The existence of RV dysfunction in patients with inferior myocardial infarction is associated with high rates of morbidity and mortality.65–67 Cardiogenic shock is also more frequent if the RV is involved in inferior infarctions.66 Reperfusion of acute RV infarcts by primary angioplasty has been shown to greatly improve RV function.68 These studies also highlight the importance of CMR in characterizing the RV in the setting of acute infarction, not only for RV mass and function quantification but also for detection and quantification of necrosis with late gadolinium techniques.

Right Ventricular Assessment in Arrhythmogenic Right Ventricular Cardiomyopathy In ARVC, normal myocardium is replaced by fibrofatty tissue, and electrical instability develops. This disorder usually involves the right ventricle, but the left ventricle and septum may also be affected. After hypertrophic heart disease, ARVC is the number one cause of sudden cardiac death in young people, especially athletes. Evident forms of the disease are straightforward to diagnose on the basis of a series of diagnostic criteria proposed by the International Task Force for Cardiomyopathy69; however, the diagnosis of early and mild forms of the disease often is difficult. CMR is regarded as the best imaging technique for detecting RV structural and functional abnormalities.70 The main advantage of CMR is the possibility of planning any desired view so that regions such as the outflow tract, which are hard to visualize with other techniques, can be assessed with great precision. The most common findings in ARVC are RV wall motion abnormalities and dilation.71 CMR can also detect fatty infiltration; however, this alone does not allow a definitive diagnosis of ARVC, as fatty infiltration occurs in a high proportion of healthy people, particularly elderly subjects.

Several authors have reported variable diagnostic sensitivity of RV fatty infiltration in patients with ARVC. The presence of fatty infiltration is often associated with RV structural abnormalities or RV wall motion abnormalities. It has been suggested that the findings of studies of fatty infiltration vary greatly from observer to observer72; therefore, diagnosis should be done carefully by experts to avoid erroneous diagnosis of ARVD, particularly as fatty infiltration is a major criterion. Fat-suppressed sequences may be useful for confirming the diagnosis in doubtful cases. Diagnosis of ARVC should always be made according to the criteria proposed by the International Task Force and never according to findings from a single test such as CMR.70

Congenital Heart Disease A comprehensive summary of the role of CMR in congenital heart disease is detailed in Chapters 29 and 30. Assessment of RV size, location, and connections as well as of function and pulmonary flow are important.73 In congenital heart disease, the right ventricle may support the pulmonary (subpulmonary right ventricle) or the systemic circulation (systemic right ventricle). In many of these patients, RV dysfunction develops and leads to considerable morbidity and mortality. Therefore, RV function in certain conditions needs close surveillance and timely and appropriate intervention to optimize outcomes. Many of these patients have come into the adult age, and this has created a patient population in which the right ventricle is often the center of attention. Despite major progress being made, assessing the RV in either the subpulmonary or the systemic circulation remains challenging, often requiring a multi-imaging approach. CMR is of use not only in the assessment of the anatomy and physiology of CHD but also, in some cases, in the risk stratification.39,74

Pulmonary Hypertension and Lung Transplantation The fact that CMR is a radiation-free, highly accurate and reproducible technique for quantitative assessment of RV mass and volume makes it the most appropriate imaging modality for serial studies on the same patient. Cine gradient echo CMR has been used to study pulmonary hypertension,75,76 its response to therapy77 and the progression of RV failure in this condition.75 CMR has been used to confirm the diagnosis of cor pulmonale through increased RV mass measurements above 60 g.78 Pulmonary flow patterns are known to be abnormal in pulmonary hypertension,79 and this may affect RV afterload, while diastolic function has also been found to be abnormal in pulmonary fibrosis by using tricuspid flow patterns.80 CMR has been used to determine the time course of changes in ventricular mass and function after lung transplantation.81–83 Frist and colleagues observed that RVEF normalized in the early postlung transplantation period.81 Other early changes included a decrease in RV end-diastolic volume to below normal levels, with persistence at this level even in the late

Cardiovascular Magnetic Resonance 391

28 CARDIOVASCULAR MAGNETIC RESONANCE ASSESSMENT OF RIGHT VENTRICULAR ANATOMY AND FUNCTION

body weight had no predictive value. The event-free survival rates from cardiovascular mortality and urgent transplantation at 1 year were 80%, 90%, and 95% in patients with an RVEF less than 25%, with an RVEF of 25% or more and less than 35% and with an RVEF of 35% or more, respectively. At 2 years, survival rates were 59%, 77%, and 93%, respectively, in the same subgroups. To date, we are unaware of any published data on the prognostic value of CMR-derived indices of RV function, but these data suggest that CMR quantitative RV volumetric assessment should be measured on studies that are performed on patients with heart failure.

RIGHT VENTRICULAR AND CONGENITAL HEART DISEASE

studies. RV mass also regressed early but remained increased in comparison to healthy control subjects. The authors concluded that RV anatomic normalization occurred later than functional normalization, and RV mass remained increased. Fayad and colleagues also studied RV tagging in patients with chronic pulmonary hypertension.84 Regional short axis shortening was reduced in patients in comparison to healthy controls, and the greatest reductions in shortening were found in the outflow tract and basal septal region.

CONCLUSION The role of the RV in acquired and congenital heart disease is being increasingly recognized. CMR is a highly accurate, reproducible, and versatile technique that is considered the ideal imaging modality for the comprehensive evaluation of RV dimensions and global and regional function. However, further clinical studies are needed to establish standards for the best use of CMR for predicting patient outcome and the role of serial evaluation in the case of known RV dysfunction.

References 1. Guyton AC. The pulmonary circulation. In: Textbook of Medical Physiology. 7th ed. Philadelphia: WB Saunders; 1986:287–294. 2. Armour JA, Randall WC. Structural basis for cardiac function. Am J Physiol. 1970;218:1517–1523. 3. Lorenz CH, Walker ES, Morgan VL, Klein SS, Graham TP. Normal human right and left ventricular mass, systolic function and gender differences by cine magnetic resonance imaging. J Cardiovasc Magn Reson. 1999;1:7–22. 4. Lewis WH. Gray’s anatomy of the human body. 20th ed. Philadelphia. 2000. 5. Piazza G, Goldhaber SZ. The acutely decompensated right ventricle. Chest. 2005;128:1836–1852. 6. Ghio S, Gavazzi A, Campana C, et al. Independent and additive prognostic value of right ventricular systolic function and pulmonary artery pressure in patients with chronic heart failure. J Am Coll Cardiol. 2001;37:183–188. 7. Meluzin J, Spinarova L, Hude P, et al. Prognostic importance of various echocardiographic right ventricular functional parameters in patients with symptomatic heart failure. J Am Soc Echocardiogr. 2005;18:435–444. 8. Gavazzi A, Ghio S, Scelsi L, et al. Response of the right ventricle to acute pulmonary vasodilation predicts the outcome in patients with advanced heart failure and pulmonary hypertension. Am Heart J. 2003;145:310–316. 9. Zornoff LA, Skali H, Pfeffer MA, et al. SAVE Investigators. Right ventricular dysfunction and risk of heart failure and mortality after myocardial infarction. J Am Coll Cardiol. 2002;39:1450–1455. 10. Rumberger JA, Behrenbeck T, Bell MR, et al. Determination of ventricular ejection fraction: a comparison of available imaging methods: the cardiovascular imaging working group. Mayo Clin Proc. 1997;72:860–870. 11. Baker BJ, Scovil JA, Kane JJ, Murphy ML. Echocardiographic detection of right ventricular hypertrophy. Am Heart J. 1983;505:611–614. 12. Kaul S, Tei C, Hopkins JM, Shah PM. Assessment of right ventricular function using two-dimensional echocardiography. Am Heart J. 1984;107:526–531. 13. Borges AC, Knebel F, Eddicks S, et al. Right ventricular function assessed by two-dimensional strain and tissue Doppler echocardiography in patients with pulmonary arterial hypertension and effect of vasodilator therapy. Am J Cardiol. 2006;98:530–534. 14. Gopal AS, Keller AM, Shen Z, et al. Three-dimensional echocardiography: in vitro and in vivo validation of left ventricular mass and comparison with conventional echocardiographic methods. J Am Coll Cardiol. 1994;24:504–513. 15. Fujimoto S, Mizuno R, Nagakawa Y, Dohi K, Nakano H. Estimation of the right ventricular volume and ejection fraction by transthoracic three-dimensional echocardiography: a validation study using magnetic resonance imaging. Int J Cardiac Imaging. 1998;14:385–390. 16. Vogel M, Gutberlet M, Dittrich S, Hosten N, Lange PE. Comparison of transthoracic three dimensional echocardiography with magnetic resonance imaging in the assessment of right ventricular volume and mass. Heart. 1997;78:127–130. 17. Hornung TS, Anagnostopoulos C, Bhardwaj P, et al. Comparison of equilibrium radionuclide ventriculography with cardiovascular magnetic resonance for assessing the systemic right ventricle after Mustard or Senning procedures for complete transposition of the great arteries. Am J Cardiol. 2003;92:640–643.

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18. Ohsuzu F, Handa S, Kondo M, et al. Thallium-201 myocardial imaging to evaluate right ventricular overloading. Circulation. 1980;61: 620–625. 19. Pietras RJ, Kondos GT, Kaplan D, Lam W. Comparative angiographic right and left ventricular volumes. Am Heart J. 1985;109:321–326. 20. Dehmer GJ, Firth DG, Hillis LD, Nicod P, Willerson JT, Lewis SE. Non geometric determination of right ventricular volumes from equilibrium blood pool scans. Am J Cardiol. 1982;49:79–84. 21. Schmermund A, Rensing BJ, Sheedy PF, Rumberger JA. Reproducibility of right and left ventricular volume measurements by electronbeam CT in patients with congestive heart failure. Int J Card Imaging. 1998;14:201–209. 22. Pignatelli RH, McMahon CJ, Chung T, Vick 3rd GW. Role of echocardiography versus MRI for the diagnosis of congenital heart disease. Curr Opin Cardiol. 2003;18:357–365. 23. Grothues F, Moon JC, Bellenger NG, Smith GS, Klein HU, Pennell DJ. Interstudy reproducibility of right ventricular volumes, function, and mass with cardiovascular magnetic resonance. Am Heart J. 2004;147:218–223. 24. Katz J, Whang J, Boxt LM, Barst RJ. Estimation of right ventricular mass in normal subjects and in patients with pulmonary hypertension by nuclear magnetic resonance imaging. J Am Coll Cardiol. 1993;21:1475–1478. 25. Lorenz CH, Walker ES, Graham TP, Powers TA. Right ventricular performance and mass in adults late after atrial repair of transposition of the great arteries using cine magnetic resonance imaging. Circulation. 1995;92(suppl II):233–239. 26. Sen-Chowdhry S, Lowe MD, Sporton SC, McKenna WJ. Arrhythmogenic right ventricular cardiomyopathy: clinical presentation, diagnosis, and management. Am J Med. 2004;117:685–695. 27. Maceira AM, Prasad SK, Khan M, Pennell DJ. Reference right ventricular systolic and diastolic function normalized to age, gender, and body surface area from steady-state free precession cardiovascular magnetic resonance. Eur Heart J. 2006;27:2879–2888. 28. Lee VS, Resnick D, Bundy JM, Simonetti OP, Lee P, Weinreb JC. Cardiac function: MR evaluation in one breath hold with real-time true fast imaging with steady-state precession. Radiology. 2002;222: 835–842. 29. Bellenger NG, Gatehouse PD, Rajappan K, Keegan J, Firmin DN, Pennell DJ. Left ventricular quantification in heart failure by CMR using prospective respiratory navigator gating: comparison with breath-hold acquisition. J Magn Reson Imaging. 2000;11:411–417. 30. Helbing WA, Rebergen SA, Maliepaard C, et al. Quantification of right ventricular function with magnetic resonance imaging in children with normal hearts and with congenital heart disease. Am Heart J. 1995;130:828–837. 31. Jauhiainen T, Jarvinen VM, Hekali PE, Poutanen VP, Penttila A, Kupari M. MR Gradient echo volumetric analysis of human cardiac casts: focus on the right ventricle. J Comput Assist Tomogr. 1998;22:899–903. 32. Alfakih K, Plein S, Bloomer T, Jones T, Ridgway J, Sivananthan M. Comparison of right ventricular volume measurements between axial and short axis orientation using steady-state free precession magnetic resonance imaging. J Magn Reson Imaging. 2003;18:25–32. 33. Grothues F, Moon JC, Bellenger NG, Smith GS, Klein HU, Pennell DJ. Interstudy reproducibility of right ventricular volumes, function and mass with cardiovascular magnetic resonance. Am Heart J. 2004;147:218–223.

56. Fayad ZA, Kraitchman DL, Ferrari VA, Axel L. Right ventricular regional function in normal subjects using magnetic resonance tissue tagging. In: Book of Abstracts: Society of Magnetic Resonance. Berkeley, CA:1994:1504. 57. Young AA, Fayad ZA, Axel L. Right ventricular midwall surface motion and deformation using magnetic resonance tagging. Am J Physiol. 1996;271:H2677–H2688. 58. Young AA, Cowan BR, Occleshaw CJ, Oxenham HC, Gentles TL. Temporal evolution of left ventricular strain late after repair of coarctation of the aorta using 3D MR tissue tagging. J Cardiovasc Magn Reson. 2002;4:233–243. 59. Menteer J, Weinberg PM, Fogel MA. Quantifying regional right ventricular function in tetralogy of Fallot. J Cardiovasc Magn Reson. 2005;7:753–761. 60. Brieke A, DeNofrio D. Right ventricular dysfunction in chronic dilated cardiomyopathy and heart failure. Coronary Artery Disease. 2005;16: 5–11. 61. Meluzin J, Spinarova L, Hude P, et al. Prognostic importance of various echocardiographic right ventricular functional parameters in patients with symptomatic heart failure. J Am Soc Echocardiogr. 2005;18:435–444. 62. de Groote P, Millaire A, Foucher-Hossein C, et al. Right ventricular ejection fraction is an independent predictor of survival in patients with moderate heart failure. J Am Coll Cardiol. 1998;32:948–954. 63. Di Salvo TG, Mathier M, Semigran MJ, Dec GW. Preserved right ventricular ejection fraction predicts exercise capacity and survival in advanced heart failure. J Am Coll Cardiol. 1995;25:1143–1153. 64. Zehender M, Kasper W, Kauder E, et al. Right ventricular infarction as an independent predictor of prognosis after acute inferior myocardial infarction. N Engl J Med. 1993;328:981–988. 65. Bueno H, Lopez-Palop R, Bermejo J, Lopez-Sendon JL, Delcan JL. Inhospital outcome of elderly patients with acute inferior myocardial infarction and right ventricular involvement. Circulation. 1997;96:436–441. 66. Kinch JW, Ryan TJ. Right ventricular infarction. N Engl J Med. 1994;330:1211–1217. 67. Moazami N, Hill L. Right ventricular dysfunction in patients with acute inferior MI: role of RV mechanical support. Thorac Cardiovasc Surg. 2003;51:290–292. 68. Bowers TR, O’Neill WW, Grines C, Pica MC, Safian RD, Goldstein JA. Effect of reperfusion on biventricular function and survival after right ventricular infarction. N Engl J Med. 1998;338:933–940. 69. Diagnosis of arrhythmogenic right ventricular dysplasia/cardiomyopathy. Task Force of the Working Group Myocardial and Pericardial Disease of the European Society of Cardiology and of the Scientific Council on Cardiomyopathies of the International Society and Federation of Cardiology. Br Heart J. 1994;71(3):215–218. 70. Pennell DJ, Sechtem UP, Higgins CB, et al. European Society of Cardiology; Society for Cardiovascular Magnetic Resonance. Clinical indications for cardiovascular magnetic resonance (CMR): Consensus Panel report. J Cardiovasc Magn Reson. 2004;6:727–765. 71. Tandri H, Calkins H, Nasir K, et al. Magnetic resonance imaging findings in patients meeting task force criteria for arrhythmogenic right ventricular dysplasia. J Cardiovasc Electrophysiol. 2003;14:476–482. 72. Bluemke DA, Krupinski EA, Ovitt T, et al. MR imaging of arrhythmogenic right ventricular cardiomyopathy: morphologic findings and interobserver reliability. Cardiology. 2003;99:153–162. 73. Rebergen SA, Ottenkamp J, Doornbos J, van der Wall EE, Chin JG, de Roos A. Postoperative pulmonary flow dynamics after Fontan surgery: assessment with nuclear magnetic resonance velocity mapping. J Am Coll Cardiol. 1993;21:123–131. 74. Babu-Narayan SV, Kilner PJ, Li W, et al. Ventricular fibrosis suggested by cardiovascular magnetic resonance in adults with repaired tetralogy of fallot and its relationship to adverse markers of clinical outcome. Circulation. 2006;113:405–413. 75. Boxt LM, Katz J, Kolb T, Czegledy FP, Barst RJ. Direct quantitation of right and left ventricular volumes with nuclear magnetic resonance imaging in patients with primary pulmonary hypertension. J Am Coll Cardiol. 1992;19:1508–1515. 76. Saito H, Dambara T, Aiba M, Suzuki T, Kira S. Evaluation of cor pulmonale on a modified short-axis section of the heart by magnetic resonance imaging. Am Rev Respir Dis. 1992;146:1576–1581. 77. Wilkins MR, Paul GA, Strange JW, et al. Sildenafil versus Endothelin Receptor Antagonist for Pulmonary Hypertension (SERAPH) study. Am J Respir Crit Care Med. 2005;171:1292–1297. 78. Pattynama PMT, Willems LNA, Smit AH, van der Wall EE, de Roos A. Early diagnosis of cor pulmonale with MR imaging of the right ventricle. Radiology. 1992;182:375–379.

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34. Longmore DB, Klipstein RH, Underwood SR, et al. Dimensional accuracy of magnetic resonance in studies of the heart. Lancet. 1985;1:1360–1362. 35. Cutrone JA, Georgiou D, Khan S, et al. Comparison of electron beam computed tomography scanning and magnetic resonance quantification of right ventricular mass: validation with autopsy weights. Acad Radiol. 1996;3:395–400. 36. Sinitsyn VE, Mareeva GI, Galakhov IE, Veselova SP. The determination of ventricular myocardial mass by magnetic resonance tomography. Med Radiol. 1990;35:9–11. 37. Maceira AM, Prasad SK, Khan M, Pennell DJ. Reference right ventricular systolic and diastolic function normalized to age, gender and body surface area from steady state free precession cardiovascular magnetic resonance. Eur Heart J. (Accepted). 38. Prasad SK, Soukias N, Hornung T, et al. Role of magnetic resonance angiography in the diagnosis of major aortopulmonary collateral arteries and partial anomalous pulmonary venous drainage. Circulation. 2004;109:207–214. 39. Babu-Narayan SV, Goktekin O, Moon JC, et al. Late gadolinium enhancement cardiovascular magnetic resonance of the systemic right ventricle in adults with previous atrial redirection surgery for transposition of the great arteries. Circulation. 2005;111:2091–2098. 40. Sen-Chowdhry S, Prasad SK, McKenna WJ. Arrhythmogenic right ventricular cardiomyopathy with fibrofatty atrophy, myocardial oedema, and aneurysmal dilation. Heart. 2005;91:784. 41. Fulton RM, Hutchinson EC, Morgan-Jones A. Ventricular weight in cardiac hypertrophy. Br Heart J. 1952;4:413–420. 42. Hangartner JRW, Marley NJ, Whitehead A, Thomas AC, Davies MJ. The assessment of cardiac hypertrophy at autopsy. Histopathology. 1985;9:1295–1306. 43. Daniels SR, Meyer RA, Liang Y, Bove KE. Echocardiographically determined left ventricular mass index in normal children, adolescents and young adults. J Am Coll Cardiol. 1988;12:703–708. 44. Am K, Gopal AS, King DL. Left and right atrial volume by freehand three-dimensional echocardiography: in vivo validation using magnetic resonance imaging. Eur J Echocardiogr. 2000;1:55–65. 45. Schvartzman PR, Fuchs FD, Mello AG, Coli M, Schvartzman M, Moreira LB. Normal values of echocardiographic measurements: a population-based study. Arq Bras Cardiol. 2000;75:107–114. 46. Kjaer A, Lebech AM, Hesse B, Petersen CL. Right-sided cardiac function in healthy volunteers measured by first-pass radionuclide ventriculography and gated blood-pool SPECT: comparison with cine MRI. Clin Physiol Funct Imaging. 2005;25:344–349. 47. Lorenz CH, Walker ES, Morgan VL, Klein SS, Graham Jr TP. Normal human right and left ventricular mass, systolic function, and gender differences by cine magnetic resonance imaging. J Cardiovasc Magn Reson. 1999;1:7–21. 48. Rominger MB, Bachmann GF, Pabst W, Rau WS. Right ventricular volumes and ejection fraction with fast cine MR imaging in breath-hold technique: applicability, normal values from 52 volunteers and evaluation of 352 adult cardiac patients. J Magn Reson Imaging. 1999;10:908–918. 49. Alfakih K, Plein S, Thiele H, Jones T, Ridgway JP, Sivananthan MU. Normal human left and right ventricular dimensions for MRI as assessed by turbo gradient echo and steady-state free precession imaging sequences. J Magn Reson Imaging. 2003;17:323–329. 50. Beygui F, Furber A, Delepine S, et al. Routine breath-hold gradient echo MRI-derived right ventricular mass, volumes and function: accuracy, reproducibility and coherence study. Int J Cardiovasc Imaging. 2004;20:509–516. 51. Hudsmith LE, Petersen SE, Francis JM, Robson MD, Neubauer S. Normal human left and right ventricular and left atrial dimensions using steady state free precession magnetic resonance imaging. J Cardiovasc Magn Reson. 2005;7:775–782. 52. Hajduczok ZD, Weiss RM, Stanford W, Marcus ML. Determination of right ventricular mass in humans and dogs with ultrafast cardiac computed tomography. Circulation. 1990;82:202–212. 53. Wachspress JD, Clark NR, Untereker WJ, Kraushaar BT, Kurnik PB. Systolic and diastolic performance in normal human subjects as measured by ultrafast computed tomography. Cathet Cardiovasc Diagn. 1988;15:277–283. 54. Klein SS, Graham TP, Lorenz CH. Noninvasive delineation of normal right ventricular contractile motion with MRI myocardial tagging. Ann Biomed Eng. 1998;26:756–763. 55. Naito H, Arisawa J, Yamagami H, Kozuka T, Tamura S. Assessment of right ventricular regional contraction and comparison with left ventricle in normal humans: a cine magnetic resonance study with presaturation myocardial tagging. Br Heart J. 1995;74:186–191.

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79. Bogren HG, Klipstein RH, Mohiaddin RH, et al. Pulmonary artery distensibility and blood flow patterns: a magnetic resonance study of normal subjects and of patients with pulmonary arterial hypertension. Am Heart J. 1989;118:990–999. 80. Kroft LJ, Simons P, van Laar JM, de Roos A. Patients with pulmonary fibrosis: cardiac function assessed with MR imaging. Radiology. 2000;216:464–471. 81. Frist WH, Lorenz CH, Walker ES, et al. MRI complements standard assessment of right ventricular function after lung transplantation. Ann Thorac Surg. 1995;60:268–271.

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82. Moulton JM, Creswell LL, Ungacta FF, Downing SW, Szabo BA, Pasque MK. Magnetic resonance imaging provides evidence for remodeling of the right ventricle after single-lung transplantation for pulmonary hypertension. Circulation. 1996;94(suppl II):312–319. 83. Lorenz CH, Loyd JE, Klein SS, et al. Characterization of different time courses of left and right ventricular recovery after lung transplantation. J Am Coll Cardiol. 1997;29(Suppl A):23A. 84. Fayad ZA, Ferrari VA, Kraitchman DL, et al. Right ventricular regional function using MR tagging: normals versus chronic pulmonary hypertension. Magn Reson Med. 1998;39:116–123.

Cardiovascular Magnetic Resonance of Simple Congenital Cardiovascular Defects Arno A. W. Roest, Lucia J. M. Kroft, and Albert de Roos

The evaluation of congenital heart and large vessel disease is one of the well-established clinical applications of cardiovascular magnetic resonance (CMR). Accurate determination of cardiac anatomy and function is crucial for patient management at initial diagnosis, during intervention, and at follow-up after repair of cardiovascular malformations. The incidence of moderate to severe forms of congenital heart disease (CHD) is estimated to be 5 to 12 per 1000 live births; worldwide, 1.5 million children are born each year with a congenital cardiac malformation.1 The prevalence increases to 19 per 1000 live births if bicuspid aortic valves are included and to 75 per 1000 live births if tiny ventricular septal defects (VSDs) and other forms of trivial lesions are taken into account.2 The estimated distribution of various types of congenital cardiac malformations is listed in Table 29-1. The initial diagnosis of CHD is most often made with transthoracic echocardiography. Before correction, CMR is especially useful to evaluate complex vascular malformations, assess possible airway compression, provide tissue characterization of cardiac tumors,3 and assess the spatial relationships in complex CHD.4 Furthermore, left-to-right shunting of blood frequently occurs in congenital heart defects, such as VSD, patent ductus arteriosus (PDA), atrial septal defect (ASD), aortopulmonary window, and partial or total anomalous pulmonary venous return. Frequently used methods to evaluate shunt volume, such as invasive oximetry, first-pass radionuclide angiography, and echocardiography, have limitations,5 whereas CMR is safe, does not expose the patient to radiation or iodinated contrast, and is accurate and reproducible for quantifying left-to-right shunting.6 Most types of CHD require surgical or catheter-based intervention.7 Since the introduction of various types of interventions for CHD, the survival rate of these patients has increased dramatically. In developed countries, more than 85% of infants with CHD now reach adulthood.8 In the planning of interventional procedures, such as closure of an ASD, CMR is a noninvasive and accurate method for anatomic delineation9,10 and can be used in the stratification of patients for either interventional or surgical correction.9–12 The long-term outcome of patients with corrected or palliated CHD is determined by residua (preoperative abnormality intentionally unaffected by intervention), sequelae (unintended but foreseen result of intervention), and complications after surgical intervention.7 Timely detection of

these morphologic and functional abnormalities requires accurate and preferably noninvasive imaging methods. CMR is ideally suited to assess morphologic and functional abnormalities after correction or palliation of CHD because this technique is not hampered by anatomic limitations and does not use ionizing radiation. CMR can provide detailed anatomic information and quantitative data on vascular stenosis and valvular function and can accurately assess the dimensions and function of the left ventricle (LV) and right ventricle (RV). Especially in corrected CHD, echocardiography may be hampered by the presence of scar tissue, rib and chest deformations, and interposed lung tissue. Cardiac catheterization using X-ray is not suited for routine follow-up after correction because of its invasiveness and limitations or repeated radiation exposure. Especially in patients with CHD, multiple cardiac catheterizations may lead to high radiation exposure, a risk factor for the development of cancer.13 The further development and application of interventional CMR-guided cardiac catheterization will further decrease the use of radiation in patients with a congenital cardiac defect.14,15

CARDIOVASCULAR MAGNETIC RESONANCE IN PEDIATRIC PATIENTS The duration of a “typical” CMR study is 45 to 60 minutes during which the patient is located within the bore of the scanner. Claustrophobia is less common in children than in adults,16 and most children older than 10 years can hold still for 60 minutes and can cooperate with breath hold maneuvers. For children 5 to 10 years of age, lying still is accomplished by proper instructions and by allowing a parent to remain with them in the scanner room. Breath holding, however, is difficult at this age, and free breathing techniques, in combination with navigator gated respiratory techniques, can be used. In patients younger than 5 years, some form of anesthesia is needed to perform a CMR study. The application of sedation has been proposed as a safe form of “conscious” anesthesia,17–19 although some advocate general anesthesia to protect the airway and control respiration.20 Specialized personnel who are familiar with providing anesthesia to Cardiovascular Magnetic Resonance 395

29 CARDIOVASCULAR MAGNETIC RESONANCE OF SIMPLE CONGENITAL CARDIOVASCULAR DEFECTS

CHAPTER 29

Ventricular septal defect* Secundum atrial septal defect* Patent ductus arteriosus* Pulmonary valve stenosis* Coarctation of the aorta* Tetralogy of Fallot Aortic valve stenosis* Transposition of the great arteries Atrioventricular septal defect Hypoplastic left heart Hypoplastic right heart

31% 7.5% 7.1% 7.0% 5.6% 5.5% 4.9% 4.5% 4.4% 3.1% 2.4%

(CE-MRA) is particularly of use in the evaluation of the thoracic vasculature at initial diagnosis as well as after correction of a congenital vascular malformation, such as coarctation of the aorta.25 CMR-guided cardiac catheterization with intervention is a new, exciting area and is especially applicable in the catheter-based closure of ASDs,26– 28 in the nonsurgical replacement of semilunar valves,29 and in guiding intervention for native or recurrent coarctation of the aorta.30–32 This chapter reviews the value of CMR for the evaluation of the most common simple congenital cardiovascular defects.

*Cardiovascular magnetic resonance evaluation discussed in this chapter.

Ventricular Septal Defect patients with CHD and CMR-compatible respiratory equipment and monitoring systems are needed to perform CMR studies with general anesthesia. An advantage of using general anesthesia is that the CMR examination is not limited to free breathing techniques because ventilation is controlled by the anesthesiologist and suspended respiration studies can be performed.3 Few reports exist on normal values for cardiac dimensions and function obtained with CMR in healthy children. Because the evaluation of CHD and large vessel disease is one of the well-established clinical applications of CMR, more studies must be performed to establish reliable information on normal cardiac dimensions and function in children.21

CARDIOVASCULAR MAGNETIC RESONANCE TECHNIQUES IN CONGENITAL HEART DEFECTS Spin echo CMR is a black-blood technique and is used to assess the cardiac and vascular anatomy under investigation, whereas gradient recalled echo (GRE) and steady-state free precession (SSFP) CMR are bright-blood techniques often used for assessment of LV and RV function or to study flow phenomena across stenoses and to depict vascular disease. Flow mapping is useful to quantify flow in large vessels as well as across valves and is useful to quantify valvular regurgitation. It can also be used to quantify flow velocity for calculation of pressure gradients in case of a stenosed vascular segment or valve. Flow measurements can be readily performed in vascular areas that may not routinely be accessible by Doppler echocardiography. CMR can directly measure flow in the aorta and pulmonary circulation, thereby allowing quantification of shunt lesions that manifest themselves by a discrepancy between aortic and pulmonary flow. For example, atrial left-to-right shunts can be assessed by quantifying the stroke volume in the aorta and pulmonary artery, thereby allowing direct shunt quantification with high precision and accuracy.22–24 The use of contrast-enhanced magnetic resonance angiography 396 Cardiovascular Magnetic Resonance

The most common CHD malformation and the major cause of left-to-right shunts, a VSD can occur as an isolated anomaly or in combination with other cardiac malformations, such as coarctation of the aorta, tetralogy of Fallot, double-outlet RV, or truncus arteriosus. VSDs are classified according to the part of the RV surface of the interventricular septum in which they are located.33 The four parts of the ventricular septum are the inlet septum; the muscular, or trabecular, septum; the outlet septum; and the membranous septum (Fig. 29-1). A defect can occur in each part of the ventricular septum. Most VSDs are small and result in a small left-to-right shunt without any signs and symptoms other than a systolic murmur. VSDs often become smaller or close spontaneously. At least 70% of defects that are present at birth will spontaneously close, usually within the first years of life.34 With a larger VSD, high pulmonary vascular resistance in the first weeks of life will prevent significant left-to-right shunting of blood. As pulmonary vascular resistance declines, the left-to-right shunt will increase and tachypnea, dyspnea, and feeding difficulties will develop. Such infants often present 2 to 6 weeks after birth. Indications for interventional or surgical closure and the age at which

PA 4 Outlet

*3 Trabecular TV

RIGHT VENTRICULAR AND CONGENITAL HEART DISEASE

Table 29-1 Distribution of Types of Congenital Heart Disease in Liveborn Children

1 2 Inlet

Figure 29-1 Most common locations of ventricular septal defect shown over the four parts of the interventricular septum, viewed from the right ventricular septal surface.* Membranous part of the septum. 1, inlet defect; 2, trabecular or muscular defect; 3, perimembranous defect; 4, outlet defect, also called supracristal or subaortic defect; PA, pulmonary artery; TV, tricuspid valve.

Figure 29-2 Transverse spin echo cardiovascular magnetic resonance image at the level of the aortic root in a patient with corrected tetralogy of Fallot. Note the position of the aortic orifice, overriding the ventricular septum and the pericardial patch (arrow), closing the outlet ventricular septal defect. (From Roest AA, Helbing WA, van der Wall EE, de Roos A. Postoperative evaluation of congenital heart disease by magnetic resonance imaging. J Magn Reson Imaging. 1999;10:656–666, with permission.)

Figure 29-3 Sagittal gradient recalled echo cardiovascular magnetic resonance image obtained during systole in a patient after correction of coarctation of the aorta. The area of signal loss (arrowhead) in the right ventricle from the interventricular septum is caused by turbulent flow across a perimembranous ventricular septal defect. (From Roest AA, Helbing WA, van der Wall EE, de Roos A. Postoperative evaluation of congenital heart disease by magnetic resonance imaging. J Magn Reson Imaging. 1999; 10:656–666, with permission.)

Estimation of the Qp:Qs ratio with echocardiography is not reliable because it is highly operator dependent.36 In addition, it is difficult to assess the cross-sectional area of the vessels throughout the cardiac cycle, the limited sample volume does not necessarily represent mean velocity across the vessels, and the main pulmonary arteries are sometimes difficult to assess.22 Catheterization in combination with oximetry can be used; however, this remains an invasive procedure that exposes the infant or young child to potentially harmful ionizing radiation. Radionuclide angiocardiography is restricted to simple shunt lesions with normal ventricular function37 and also exposes the young child to radiation. Several reports have shown that CMR is ideal for the measurement of flow volume through the aorta and pulmonary trunk using velocity-encoded CMR from which the shunt ratio can be extracted (Fig. 29-4).22–24 Earlier reports assessed shunt volume in adult patients. In pediatric patients, higher peak velocities and blood pulsation rates may influence velocity and flow quantification.38 Data show that CMR accurately provides quantitative shunt volume data in the pediatric population.37,39 Advances in interventional cardiology have allowed for transcatheter closure of VSDs.40,41 CMR-guided catheterization15 and intervention are now feasible,26–28,32 and the combination of X-ray fused with CMR (XMR) has proven beneficial in the closure of perimembranous VSD.42 In swine models, the use of XMR-guided antegrade catheter crossing and closure of the VSD is both easier and faster, and is associated with reduced radiation compared with conventional techniques.42 Cardiovascular Magnetic Resonance 397

29 CARDIOVASCULAR MAGNETIC RESONANCE OF SIMPLE CONGENITAL CARDIOVASCULAR DEFECTS

the correction is performed are based on the clinical status of the patient. Although the magnitude of the left-to-right shunt is important, there is controversy about the threshold Qp:Qs ratio (i.e., flow volume through the pulmonary circulation divided by flow volume through the systemic circulation) at which correction is indicated. Ratios greater than 2:1 are generally accepted as an indication for intervention. However, a chronic moderate left-to-right shunt of greater than 1.5:1 to 1.8:1 also justifies closure because the LV is subjected to volume overload for a longer period and may remain dysfunctional after the defect is closed.34 Therefore, precise estimation of shunt flow is essential in the management of patients with a VSD. In addition, for surgical management of complex cardiac malformations, it is of utmost importance to know the spatial relationship between the VSD and the orifices of the great arteries. Depending on the location of the VSD, surgical management can vary from a simple patch, inserted from the right atrium (Fig. 29-2), to a complicated biventricular correction.4,35 For anatomic delineation of VSDs, particularly those associated with complex malformations, CMR is superior to two-dimensional echocardiography, especially in the visualization of the spatial relationship with surrounding structures, the atrioventricular valves and great arteries. To determine the location of the VSD, a black-blood fast spin echo CMR imaging sequence is used and images are made in the three orthogonal planes. The transverse plane is the most useful for the identification of all types of VSDs.33 When precise information is needed about the shape and dimensions of the VSD and the spatial relationship between the VSD and great arteries, these images should be completed with en face images of the VSD, made by planning a series of slices parallel to the VSD.35 A bright-blood GRE CMR imaging sequence is used to visualize a jet caused by turbulent blood flow through the VSD (Fig. 29-3). As mentioned previously, precise shunt quantification is essential in the management of patients with a VSD.

250

Flow volume (mL/sec)

RIGHT VENTRICULAR AND CONGENITAL HEART DISEASE

Figure 29-4 By subtracting the aortic flow curve from the pulmonary flow curve, one can quantify the amount of shunting in each heartbeat. The Qp:Qs ratio in this case was 1.7. (From Roest AA, Helbing WA, van der Wall EE, de Roos A. Postoperative evaluation of congenital heart disease by magnetic resonance imaging. J Magn Reson Imaging. 1999; 10:656–666, with permission.)

300

Pulmonary trunk 105 mL

200

Ascending aorta 63 mL

150 100 50 0

−100

0 −50

100

200

300

400

500

600

700

Time after R wave (msec)

−50

Atrial Septal Defect An ASD is a common cause of left-to-right intracardiac shunting. Patients with an ASD are generally asymptomatic through infancy and childhood. When symptoms occur in infancy, patients present with frequent respiratory infections, symptoms of volume overload of the pulmonary circulation, and sometimes congestive heart failure. Later in childhood, a murmur is often heard because of high pulmonary artery flow, and patients are referred for cardiac evaluation. After the age of 20 years, symptoms are more common and include dyspnea on exertion, fatigue, palpitations, sustained atrial arrhythmias,43 and cryptogenic stroke.44 Several types of ASDs are recognized (Fig. 29-5). The most common type is the ostium secundum or fossa ovalis defect, which is located in the central part of the atrial septum (Fig. 29-6C). The sinus venosus defect is situated high or low on the septum near the entrance of, respectively, the superior vena cava or inferior vena cava into the right atrium (RA). This type may be associated with anomalous drainage of the right upper lobe pulmonary vein into the RA (see Fig. 29-6A and B).45 The ostium primum defect is situated low in the atrial septum, close to the atrioventricular valves. This defect belongs to the spectrum of atrioventricular septal defects referred to as endocardial cushion defects and is accompanied by abnormal position and structure of the atrioventricular valves. Transthoracic echocardiography is the first-line clinical tool for detecting an ASD and is likely more sensitive for detecting a very small ASD. Transesophageal echocardiography is a moderately invasive but even more accurate diagnostic tool. CMR can be useful, especially in older patients, who may have suboptimal echocardiographic windows.43 The presence of an ASD can be established definitively by using cine GRE CMR to identify the ASD jet across the septum, with visualization of a low signal jet on the right side of the interatrial septum.46 The definition of such flow-related signal void can be enhanced by a spatial saturation slab at the inflow region of the ASD.47 The diameter of the defect 398 Cardiovascular Magnetic Resonance

SVC

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Figure 29-5 Most common locations of atrial septal defect, viewed from the right atrium. 1, ostium primum defect (also known as atrioventricular septal defect); 2, ostium secundum defect, or fossa ovalis defect; 3, sinus venosus defect; 4, coronary sinus defect; IVC, inferior vena cava; SVC, superior vena cava; TV¼tricuspid valve.

can be derived from the maximum width of the transseptal flow acquired with multiple parallel and intersecting CMR acquisitions.48 Especially phase contrast CMR is used to determine the size of the ASD, whereas spin echo CMR techniques tend to overestimate the size of the defect.9 The size and morphologic features of the ASD and its anatomic

RV

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Figure 29-6 Axial fast spin echo images in a 26-year-old woman with partial abnormal pulmonary venous return showing the right upper lobe pulmonary vein (RULPV) draining into the superior vena cava (SVC; A). A large sinus venosus type atrial septal defect (ASD) is present high in the atrial septum (*; B). At the lower level, a fossa ovale, or ostium secundum, ASD is shown (arrow; C). A thin membrane may be present at this fossa ovale level despite visual interruption of the atrium septum, as shown in these images. Sizing of the ASD is best performed with velocity encoded or cine gradient recalled echo cardiovascular magnetic resonance. AA, ascending aorta; L, left; LV, left ventricle; P, posterior; RA, right atrium; RPA, right pulmonary artery; RV, right ventricle; RVOT, right ventricular outflow tract.

relationships are more important because percutaneous closure of ostium secundum ASDs is now routinely performed. The location, size of the defect, and age of the patient are the major determining factors for the choice between sternotomy/surgical correction and percutaneous interventional closure.49 Transthoracic echocardiography is not always accurate in the precise evaluation of the ASD.10 Transesophageal echocardiography is more accurate, but is moderately invasive.50 CMR is an accurate and noninvasive method for evaluation of the size and morphologic features of ASDs. Furthermore, CMR provides precise information about the surrounding structures, such as the atrioventricular valves and the entrance of the systemic and pulmonary veins.9–12 In children in whom transesophageal echocardiography was inconclusive for deciding on percutaneous or surgical closure of the ASD, CMR provided additional information, thereby assisting in the planning of interventional ASD closure.10 In adults with a patent foramen ovale and cryptogenic ischemic events, CMR currently appears to be inferior to transesophageal echocardiography in the detection of agitated-saline-demonstrated right-to-left shunt shunting and identification of an atrial septal aneurysm.50 The advantage of CMR in evaluating patients with leftto-right shunt at the atrial level is the ability to assess the functional significance by determining flow through the shunt. In ASDs, the Qp:Qs ratio and shunt flow can be extracted from the discrepancy between RV and LV stroke volumes and by measuring aortic and pulmonary flow with flow mapping (Fig. 29-7).22,24,37,51,52 This provides for independent confirmation of the size of the intracardiac shunt (in the absence of significant mitral regurgitation and tricuspid regurgitation). Direct measurement of flow across the ASD is also possible.37,48,53 Furthermore, the effects of shunting through an ASD on RV size can be readily quantified by CMR. The use of CMR-guided transcatheter closure of ASDs has been reported in animal models using real-time CMR.26–28 This is a promising application of CMR to further decrease the use of ionizing radiation in patients with CHD. The acute effect of transcatheter closure of an ASD

on atrial and ventricular parameters can be assessed by performing CMR shortly before and after the intervention.54 Within 24 hours after closure of an ASD, the mean size of the RA and RV decreases, whereas the LA and LV volume remains unchanged.54

[mL/sec] 0

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Figure 29-7 Same patient as shown in Figure 29-6, with partial abnormal pulmonary venous return and atrial septal defect, both contributing to a left-to-right shunt. Graphic reconstruction representing flow over one heart phase through the aortic valve (A) and through the pulmonary valve (B). The shunt size was calculated by dividing the flow through the pulmonary valve, representing the pulmonary circulation Qp, by flow over the aortic valve, representing the systemic flow Qs. Qp was 13.0 L/min and Qs was 4.6 L/min. The shunt size was therefore 2.8 (13.0/4.6). Cardiovascular Magnetic Resonance 399

29 CARDIOVASCULAR MAGNETIC RESONANCE OF SIMPLE CONGENITAL CARDIOVASCULAR DEFECTS

SVC RULPV

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After closure of the ASD, CMR can be used for the evaluation of cardiac function. If a septal occluder has been used for ASD closure, CMR can clearly show the location of the device, its effect on systemic and pulmonary venous return, and the presence of valve insufficiency.55 Furthermore, CMR can be used for follow-up of ventricular volumes and function after surgical or interventional ASD closure.56–58,58a

Patent Ductus Arteriosus The ductus arteriosus is the connecting vessel between the pulmonary trunk and the descending aorta. In utero, most of the RV stroke volume bypasses the still unexpanded lungs and enters the systemic circulation through the ductus arteriosus. In the vast majority of infants, the ductus arteriosus closes within the first week of life. There is delayed closure of the ductus arteriosus in preterm infants. Spontaneous closure occurs in most preterm infants in the first month of life. In term infants, spontaneous closure of a PDA is rare because of abnormalities in the structure of the ductus.59 Interventional closure of a PDA is needed when the shunt flow is hemodynamically significant and is also recommended in the case of a small ductus without hemodynamic significance because of the increased risk of infective endocarditis.59 A PDA results in left-to-right shunt and therefore volume overload of the pulmonary circulation and LV, which may lead to pulmonary hypertension. The amount of shunting depends on the ductus size and the difference between pulmonary vascular resistance and systemic vascular resistance.59 Although a PDA is a frequently encountered CHD malformation, there are few reports on evaluation with CMR.60–64 Transthoracic echocardiography is the first-line clinical tool for detecting a PDA, especially in infants and young children. In adult patients, transthoracic echocardiography may be less sensitive than CMR.60 Using CEMRA, a PDA is readily seen.61,62 Measurement of shunt volume can be performed by assessing LV and RV stroke volumes with cine GRE or SSFP CMR or by evaluation of flow volume over the pulmonary and aortic valves with flow mapping.61 The Qp:Qs ratio is less than 1.0 in patients with a PDA. Furthermore, CMR can be used to assess the complications of chronic left-to-right shunting caused by a PDA, such as dilation of the pulmonary arteries, LV dysfunction as a result of chronic volume overload, and dilation and hypertrophy of the RV secondary to pulmonary artery hypertension.60,63 Similar results appear to be available at 3 Tesla.64a

poststenotic dilation is frequently observed. A coarctation can occur in isolation, but can be associated with VSDs, aortic valve stenosis, mitral valve anomalies, double-outlet RV, transposition of the great arteries, hypoplastic left heart syndrome, and tricuspid atresia.65 In severe coarctation, signs of cardiac failure and decreased or absent femoral pulsations develop within the first weeks of life as the ductus arteriosus closes. In milder coarctation, after closure of the ductus arteriosus, cardiovascular adaptation will occur, including LV hypertrophy and cavity dilation, increased sympathetic activation causing hypertension, and development of a collateral circulation to bypass the coarctation.65 Currently, CMR is considered the standard noninvasive technique for the evaluation of native and repaired aortic coarctation.66,67 Associated abnormalities, such as arch hypoplasia, bicuspid aortic valve, and VSD, can also be assessed with CMR.68 CMR readily identifies the site and extent of coarctation, involvement of arch vessels, poststenotic dilation, and dilated collateral vessels (Figs. 29-8 to 29-10).69,70 CE-MRA has emerged as a valuable technique to assess the thoracic aorta and is now routinely performed in older children and adults, resulting in high-quality aortograms with the aid of an intravenous infusion of gadolinium to shorten the arterial blood T1 relaxation time (see Figs. 29-9 and 29-10).25 Velocity encoded CMR measurement of peak jet velocity across the coarctation provides comparable estimates of the pressure gradient to those obtained from continuous wave Doppler echocardiography (by application of the modified Bernoulli formula).71,72 Significant narrowing will impair blood flow into the descending aorta; therefore, collateral vessels are required

Coarctation of the Aorta Coarctation of the aorta most commonly occurs as a discrete stenosis of the proximal descending aorta, just opposite the (former) insertion of the ductus arteriosus (juxtaductal location). Most commonly, coarctation of the aorta is congenital, although coarctation of the aorta can occur after Takayasu arteritis. The gross morphologic features of the coarctation may vary from a discrete narrowing to long segment stenosis. Distal to the coarctation, 400 Cardiovascular Magnetic Resonance

Figure 29-8 Oblique sagittal fast spin echo image of a 10-yearold boy showing severe aortic coarctation (arrow) at the classic position in the aorta distal from the origin of the left subclavian artery.

B

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Figure 29-9 Same patient as shown in Figure 29-8. Contrast-enhanced magnetic resonance angiography after injection of gadolinium chelate. There are several postprocessing options for displaying the site of coarctation. A, Maximum intensity projection. B and C, Shaded surface display. Note the large collateral arteries entering the descending aorta distal from the coarctation (arrows).

A

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Figure 29-10 Same patient as shown in Figure 29-8. A, Volume rendering image. B and C, Shaded surface display. Changing the window width and window level settings allows more collateral arteries to be observed. Large collateral arteries include the internal mammarian arteries to the abdomen (double arrows in A and C) and large thoracic wall collateral artery (single arrow in A, B, and C) entering an intercostal artery (B). Note the large intercostal arteries (B). Also note other collateral flow pathways to the lower body, such as internal mammarian artery flow via the epigastric arteries (arrowheads in C).

to reestablish aortic flow distal to the coarctation. The intercostal, lateral thoracic, internal mammary, anterior spinal, and epigastric arteries can all serve as collateral pathways. Flow mapping has proven to be a valuable adjunct to assess the severity of coarctation by measuring flow in the proximal and distal descending thoracic aorta (Figs. 29-11 and

29-12).73,74 This method for assessing the severity of coarctation is based on the increase of flow in the distal aorta above the diaphragm with regard to flow near the coarctation site. Measurement within the coarctation tends to overestimate flow, whereas measurement of flow immediately above and below the coarctation yields similar findings Cardiovascular Magnetic Resonance 401

29 CARDIOVASCULAR MAGNETIC RESONANCE OF SIMPLE CONGENITAL CARDIOVASCULAR DEFECTS

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Figure 29-11 Same patient as shown in Figure 29-8. A, Maximum intensity projection image. B, Double oblique transverse gradient recalled echo cardiovascular magnetic resonance at the aortic level indicated with phase and modulus images just proximal to the coarctation (upper panel), just distal to the coarctation (middle panel), and in the abdominal aorta (lower panel). C, Graphic reconstructions representing flow during the cardiac cycle at the corresponding levels, where flow volumes are calculated per minute. The flow volume at the level immediately before (and after) the coarctation was 0.6 L/min. Flow volume at the level of the abdominal aorta was 2.6 L/min. This indicates that almost 80% of the blood flow through the descending aorta was derived from collateral artery flow (2.6 0.6)/2.6. The flow volume measured just above the aortic valve was 5.0 L/min.

(see Fig. 29-11).74 Retrograde flow in collateral channels will increase distal aortic flow, depending on the severity of coarctation, thereby providing a direct estimate of the hemodynamic severity of the coarctation. Combining anatomic and flow techniques, CMR provides a sensitive and specific test for predicting a catheterization gradient of greater than 20 mm Hg, which is considered an important criterion for a significant coarctation.75 Repair of native coarctation of the aorta can be performed either surgically or with catheter-based balloon angioplasty and stenting. Comparison of the long-term outcome of native coarctation in children has shown that balloon angioplasty is associated with a higher incidence of aneurysm formation and iliofemoral artery injury compared with surgery.76 After initial coarctation repair, restenosis and aneurysm formation frequently occur. CMR guidance can be used to monitor for these complications and to guide subsequent repair.30,32,77 Because CMR is superior to Doppler echocardiography in evaluating patients with a corrected coarctation,78 it is ideally suited for long-term follow-up. CMR is now considered the standard method for the evaluation of children and adults with corrected coarctation of the aorta because there is no exposure to ionizing radiation.79–81 Spin echo CMR can be used to assess the anatomy of the aorta, but the addition of GRE and SSFP cine CMR as well as CEMRA provides more detailed information than spin echo CMR.78,80,82 Furthermore, velocity encoded CMR, preferably in-plane as well as through-plane, provides an 402 Cardiovascular Magnetic Resonance

accurate estimate of the gradient over a restenosis.78,83 After repair, information on collateral flow, as assessed with CMR, proved more accurate than arm/leg blood pressure gradient in the assessment of the hemodynamic severity of restenosis and may be helpful for planning treatment options and monitoring patient outcome.84

SE/M SL7

Figure 29-12 Transverse spin echo cardiovascular magnetic resonance image at the level of the atrioventricular valves in a patient with tricuspid valve atresia. The single-headed arrow indicates the right coronary artery surrounded by fat, located at the site of the atretic tricuspid valve.

Valvular Heart Disease Valvular abnormalities are frequently seen in patients with CHD, either as a congenital malformation or as a result of treatment. CMR can be used for evaluation of valvular anatomy and function. Valve abnormalities frequently seen are atresia of one of the valves (Fig. 29-12), Ebstein anomaly, bicuspid semilunar valve (Figs. 29-13 and 29-14), single atrioventricular valve, and congenital mitral valve malformations. Functional valve abnormalities occur either congenitally or because of intervention. Both valvular stenosis and regurgitation can be quantified by velocity encoded CMR. Estimation of the severity of stenosis is possible using velocity encoded CMR.86 As discussed previously, measurement of peak velocity, preferably in the orthogonal plane,

A

allows estimation of the pressure gradient over the valve by applying the modified Bernoulli equation.66,87 In the case of insufficiency of a valve, the volume of forward and backward flow can be measured, allowing quantification of regurgitation volume (Fig. 29-15).88 A major difficulty associated with using CMR in the evaluation of valvular function is movement of the valve during the cardiac cycle, influencing flow and velocity measurements.89,90 This problem can be overcome by applying dedicated scanning protocols that allow more accurate measurements.91–93 The most frequently encountered valvular abnormality is bicuspid aortic valve, occurring in 1% of the adult population.94 Complications of bicuspid aortic valve are aortic valve stenosis and regurgitation, progressive aortic dilation, aortic aneurysm formation, and aortic dissection. Dilation of the aortic root is associated with intrinsic aortic wall pathology in patients with bicuspid aortic valve.94 CMR can be used to evaluate valvular competency and the effect of the bicuspid valve on LV volume and systolic function.94 Furthermore, aortic distensibility and pulse wave velocity can be

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C Figure 29-13 Double oblique sagittal gradient recalled echo cardiovascular magnetic resonance in a 23-year-old woman at the aortic valve level showing the slit-like opening of the bicuspid aortic valve. A, Modulus image. B, Corresponding phase image. C, Coronal fast spin echo image showing poststenotic dilation of the ascending aorta (AA). Also, the left ventricle (LV) was dilated with a calculated end-diastolic volume of 285 mL. F, feet; L, left; PT, pulmonary trunk; RA, right atrium; RV, right ventricle. Cardiovascular Magnetic Resonance 403

29 CARDIOVASCULAR MAGNETIC RESONANCE OF SIMPLE CONGENITAL CARDIOVASCULAR DEFECTS

In conclusion, the combination of clinical assessment and CMR in every patient after surgical or interventional repair of coarctation of the aorta is the most cost-effective way to diagnose complications.85

RIGHT VENTRICULAR AND CONGENITAL HEART DISEASE

assessed by CMR. Reduced aortic elasticity and aortic dilation have been shown to correlate with aortic valve regurgitation and LV hypertrophy in patients with nonstenotic bicuspid aortic valves.94 In patients with aortic regurgitation, quantification of regurgitant volume is superior to indices of LV volume and systolic function for identification of patients requiring surgery.95 The pulmonary valve is frequently affected in CHD, and the application of dedicated CMR protocols with long and short axis views allows detailed anatomic and functional evaluation of this valve and the influence of dysfunction on the RV (Fig. 29-14).96 Additional information on all sources of blood supply to the lungs is essential in the evaluation of pulmonary stenosis and pulmonary atresia. Multiplanar GRE and SSFP cine CMR and CE-MRA are fast and accurate techniques for this purpose (see Fig. 29-14C).97 As a result of stenosis or regurgitation, the cardiac chamber situated upstream of the abnormal valve will react to the increased workload. Pressure overload

A

resulting from stenosis causes myocardial hypertrophy and eventually ventricular dilation (see Fig. 29-14D), whereas chronic regurgitation causes chamber enlargement because of volume overload. With GRE and SSFP cine CMR, the severity and progression of these secondary abnormalities can be analyzed by measuring LV and RV volumes, wall thickness, and mass.88 After valve replacement, CMR can be used to evaluate the amount of residual regurgitation or stenosis and the effect of replacement on cardiac function.98 Nonsurgical valve replacement has been introduced for correction of stenosed or insufficient pulmonary valves.99,100 CMR proved to be essential in the selection of patients for percutaneous pulmonary valve replacement.101 In a swine model, it was shown that CMR can be used to guide stenting of the pulmonary and aortic valve and the pulmonary arteries, with immediate postinterventional evaluation of flow within the stent.29,102,103 In the future, CMR may be used for routine guidance and follow-up of these interventional procedures.

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Figure 29-14 Gradient recalled echo cardiovascular magnetic resonance images in a 22-year-old man with an isolated pulmonary valve stenosis caused by a bicuspid pulmonary valve (arrow, axial image A). A turbulent flow jet caused by this valve is visible in the main pulmonary artery (arrow, axial image B). Note the asymmetrically developed branch pulmonary arteries: the large pulmonary artery on the left and the hypoplastic pulmonary artery on the right. Contrast-enhanced magnetic resonance angiography shows asymmetrical pulmonary perfusion with preferential flow to the left lung and an apical perfusion defect in the right lung (coronal image C). D, Gradient recalled echo imaging, axial orientation. End-diastolic (a) and end-systolic (b) images at the midventricular level. Note the severely hypertrophied right ventricle (RV) with the interventricular septum bulging toward the left ventricle (LV). RA, right atrium.

404 Cardiovascular Magnetic Resonance

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Figure 29-15 Follow-up cardiovascular magnetic resonance study in a 19-year-old woman with corrected tetralogy of Fallot and pulmonary regurgitation. A, Graphic reconstruction representing flow over one heart phase through the pulmonary valve. Forward flow during systole was 100 mL, and regurgitant flow during diastole was 53 mL. End-diastolic forward flow was 9 mL because of a restrictive right ventricle. Total regurgitant fraction was 53/109 ¼ 49%. B and C, Double oblique transverse gradient recalled echo magnetic resonance phase and corresponding modulus images on pulmonary valve level showing the forward flow as a bright area and the regurgitant flow as a dark area measured at the level just above the pulmonary valve in the main pulmonary artery (PA). AA, ascending aorta.

CONCLUSION In conclusion, CMR has emerged as an indispensable tool in the management of patients with CHD. CMR in pediatric patients requires dedicated personnel and equipment, but when available, CMR is a valuable tool in the diagnosis and follow-up of patients with CHD. Several CMR techniques are available for the accurate delineation of anatomy of the heart and the great vessels, allowing determination of the location of intracardiac shunts and vascular abnormalities and their relationship to surrounding

structures. Additional to providing morphologic information, CMR has the unique capability of quantification of ventricular function and blood flow velocity and volume, allowing quantification of intra- and extracardiac shunting, valvular function, and ventricular performance. Because of its non-invasive, non-ionizing nature, CMR is ideally suited for monitoring of cardiovascular function during follow-up or to evaluate the effect of intervention in patients with CHD. Advances in interventional cardiology and XMR will further expand the application of CMR in patients with CHD.

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[mL/sec]

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40.

41. 42.

43. 44. 45. 46.

47. 48. 49. 50.

51. 52.

53.

54.

55. 56.

57.

58.

58a. 59. 60.

resonance velocity mapping: a validation study. Circulation. 2004;109(16):1987–1993. Holzer R, Balzer D, Cao QL, Lock K, Hijazi ZM. Device closure of muscular ventricular septal defects using the Amplatzer muscular ventricular septal defect occluder: immediate and mid-term results of a U.S. registry. J Am Coll Cardiol. 2004;43(7):1257–1263. Knauth AL, Lock JE, Perry SB, et al. Transcatheter device closure of congenital and postoperative residual ventricular septal defects. Circulation. 2004;110(5):501–507. Ratnayaka K, Raman VK, Faranesch AR, et al. Antegrade percutaneous closure of membraneous ventricular septal defect using X-ray fused with magnetic resonance imaging. JACC Cardiovasc Imaging 2008;2:224–230. Latson LA. Atrial septal defect. In: Moller JH, Hofman JIE, eds. Pediatric Cardiovascular Medicine. 1st ed. 2000:311–322. Wechsler LR. PFO and stroke: what are the data? Cardiol Rev. 2008;16 (1):53–57. Valente AM, Sena L, Powell AJ, Del Nido PJ, Geva T. Cardiac magnetic resonance imaging evaluation of sinus venosus defects: comparison to surgical findings. Pediatr Cardiol. 2007;28(1):51–56. Theissen P, Sechtem U, Mennicken U, Hilger HH, Schicha H. [Noninvasive diagnosis of atrial septal defects and anomalous pulmonary venous return using magnetic resonance tomography]. Nuklearmedizin. 1989;28(5):172–180. Hartnell GG, Sassower M, Finn JP. Selective presaturation magnetic resonance angiography: new method for detecting intracardiac shunts. Am Heart J. 1993;126(4):1032–1034. Holmvang G. A magnetic resonance imaging method for evaluating atrial septal defects. J Cardiovasc Magn Reson. 1999;1(1):59–64. Ferreira SM, Ho SY, Anderson RH. Morphological study of defects of the atrial septum within the oval fossa: implications for transcatheter closure of left-to-right shunt. Br Heart J. 1992;67(4):316–320. Nusser T, Hoher M, Merkle N, et al. Cardiac magnetic resonance imaging and transesophageal echocardiography in patients with transcatheter closure of patent foramen ovale. J Am Coll Cardiol. 2006;48:322–329. Brenner LD, Caputo GR, Mostbeck G, et al. Quantification of left to right atrial shunts with velocity-encoded cine nuclear magnetic resonance imaging. J Am Coll Cardiol. 1992;20(5):1246–1250. Sieverding L, Jung WI, Klose U, Apitz J. Noninvasive blood flow measurement and quantification of shunt volume by cine magnetic resonance in congenital heart disease: preliminary results. Pediatr Radiol. 1992;22(1):48–54. Thomson LEJ, Crowley AL, Heitner JF, et al. Direct en face imaging of secundum atrial septal defects by velocity-encoded cardiovascular magnetic resonance in patients evaluated for possible transcatheter closure. Circ Cardiovasc Imaging. 2008;1:31–40. Burgstahler C, Wohrle J, Kochs M, et al. Magnetic resonance imaging to assess acute changes in atrial and ventricular parameters after transcatheter closure of atrial septal defects. J Magn Reson Imaging. 2007;25(6):1136–1140. Lapierre C, Raboisson MJ, Miro J, Dahdah N, Guerin R. Evaluation of a large atrial septal occluder with cardiac MR imaging. Radiographics. 2003;23:S51–S58. Bolz D, Lacina T, Buser P, Buser M, Guenthard J. Long-term outcome after surgical closure of atrial septal defect in childhood with extensive assessment including MRI measurement of the ventricles. Pediatr Cardiol. 2005;26(5):614–621. Schoen SP, Kittner T, Bohl S, et al. Transcatheter closure of atrial septal defects improves right ventricular volume, mass, function, pulmonary pressure, and functional class: a magnetic resonance imaging study. Heart. 2006;92(6):821–826. Weber M, Dill T, Deetjen A, et al. Left ventricular adaptation after atrial septal defect closure assessed by increased concentrations of N-terminal pro-brain natriuretic peptide and cardiac magnetic resonance imaging in adult patients. Heart. 2006;92 (5):671–675. Teo KSL, Dundon BK, Molaee P, et al. Percutaneous closure of atrial septal defects leads to normalization of atrial and ventricular volumes. J Cardiovasc Magn Resonan. 2008;10:55. Gersony WM, Apfel HD. Patent ductus arteriosus and other aortopulmonary anomalies. In: Moller JH, Hofman JIE, eds. Pediatric Cardiovascular Medicine. 1st ed. 2000:323–334. Schmidt M, Theissen P, Deutsch HJ, Erdmann E, Schicha H. Magnetic resonance imaging of ductus arteriosus Botalli apertus in adulthood. Int J Cardiol. 1999;68(2):225–229.

82. Riquelme C, Laissy JP, Menegazzo D, et al. MR imaging of coarctation of the aorta and its postoperative complications in adults: assessment with spin-echo and cine-MR imaging. Magn Reson Imaging. 1999;17 (1):37–46. 83. Henk CB, Grampp S, Koller J, et al. Elimination of errors caused by first-order aliasing in velocity encoded cine-MR measurements of postoperative jets after aortic coarctation: in vitro and in vivo validation. Eur Radiol. 2002;12(6):1523–1531. 84. Araoz PA, Reddy GP, Tarnoff H, Roge CL, Higgins CB. MR findings of collateral circulation are more accurate measures of hemodynamic significance than arm-leg blood pressure gradient after repair of coarctation of the aorta. J Magn Reson Imaging. 2003;17(2):177–183. 85. Therrien J, Thorne SA, Wright A, Kilner PJ, Somerville J. Repaired coarctation: a “cost-effective” approach to identify complications in adults. J Am Coll Cardiol. 2000;35(4):997–1002. 86. Kilner PJ, Manzara CC, Mohiaddin RH, et al. Magnetic resonance jet velocity mapping in mitral and aortic valve stenosis. Circulation. 1993;87(4):1239–1248. 87. Caruthers SD, Lin SJ, Brown P, et al. Practical value of cardiac magnetic resonance imaging for clinical quantification of aortic valve stenosis: comparison with echocardiography. Circulation. 2003;108(18):2236–2243. 88. Niezen RA, Helbing WA, van der Wall EE, van der Geest RJ, Rebergen SA, de Roos A. Biventricular systolic function and mass studied with MR imaging in children with pulmonary regurgitation after repair for tetralogy of Fallot. Radiology. 1996;201(1):135–140. 89. Reid SA, Walker PG, Fisher J, et al. The quantification of pulmonary valve haemodynamics using MRI. Int J Cardiovasc Imaging. 2002;18 (3):217–225. 90. Chatzimavroudis GP, Walker PG, Oshinski JN, Franch RH, Pettigrew RI, Yoganathan RI. Slice location dependence of aortic regurgitation measurements with MR phase velocity mapping. Magn Reson Med. 1997;37(4):545–551. 91. Kozerke S, Schwitter J, Pedersen EM, Boesiger P. Aortic and mitral regurgitation: quantification using moving slice velocity mapping. J Magn Reson Imaging. 2001;14(2):106–112. 92. Friedrich MG, Schulz-Menger J, Poetsch T, Pilz B, Uhlich F, Dietz R. Quantification of valvular aortic stenosis by magnetic resonance imaging. Am Heart J. 2002;144(2):329–334. 93. Westenberg JJ, Doornbos J, Versteegh MI, et al. Accurate quantitation of regurgitant volume with MRI in patients selected for mitral valve repair. Eur J Cardiothorac Surg. 2005;27(3):462–466. 94. Grotenhuis HB, Ottenkamp J, Westenberg JJ, Bax JJ, Kroft LJ, de Roos A. Reduced aortic elasticity and dilatation are associated with aortic regurgitation and left ventricular hypertrophy in nonstenotic bicuspid aortic valve patients. J Am Coll Cardiol. 2007;49(15):1660–1665. 95. Myerson SG, Karamitsos TD, Francis JM, Banning AP, Neubauer S. Quantifying aortic regurgitation with CMR can predict patients requiring aortic valve surgery. J Cardiovasc Magn Reson. 2008. 10(suppl). 96. Kivelitz DE, Dohmen PM, Lembcke A, et al. Visualization of the pulmonary valve using cine MR imaging. Acta Radiol. 2003;44(2):172–176. 97. Greil GF, Powell AJ, Gildein HP, Geva T. Gadolinium-enhanced threedimensional magnetic resonance angiography of pulmonary and systemic venous anomalies. J Am Coll Cardiol. 2002;39(2):335–341. 98. Vliegen HW, van SA, de Roos A, et al. Magnetic resonance imaging to assess the hemodynamic effects of pulmonary valve replacement in adults late after repair of tetralogy of Fallot. Circulation. 2002;106 (13):1703–1707. 99. Khambadkone S, Bonhoeffer P. Nonsurgical pulmonary valve replacement: why, when, and how? Catheter Cardiovasc Interv. 2004;62(3):401–408. 100. Khambadkone S, Coats L, Taylor A, et al. Percutaneous pulmonary valve implantation in humans: results in 59 consecutive patients. Circulation. 2005;112(8):1189–1197. 101. Schievano S, Coats L, Migliavacca F, et al. Variations in right ventricular outflow tract morphology following repair of congenital heart disease: implications for percutaneous pulmonary valve implantation. J Cardiovasc Magn Reson. 2007;9(4):687–695. 102. Kuehne T, Yilmaz S, Meinus C, et al. Magnetic resonance imaging-guided transcatheter implantation of a prosthetic valve in aortic valve position: feasibility study in swine. J Am Coll Cardiol. 2004;44(11):2247–2249. 103. Kuehne T, Saeed M, Reddy G, et al. Sequential magnetic resonance monitoring of pulmonary flow with endovascular stents placed across the pulmonary valve in growing swine. Circulation. 2001;104(19): 2363–2368.

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61. Wang ZJ, Reddy GP, Gotway MB, Yeh BM, Higgins CB. Cardiovascular shunts: MR imaging evaluation. Radiographics. 2003;23:S181–S194. 62. Goitein O, Fuhrman CR, Lacomis JM. Incidental finding on MDCT of patent ductus arteriosus: use of CT and MRI to assess clinical importance. AJR Am J Roentgenol. 2005;184(6):1924–1931. 63. Frydrychowicz A, Bley TA, Dittrich S, Hennig J, Langer M, Markl M. Visualization of vascular hemodynamics in a case of a large patent ductus arteriosus using flow sensitive 3D CMR at 3T. J Cardiovasc Magn Reson. 2007;9(3):585–587. 64. Rigatelli G, Zamboni A, Cardaioli P. Three-dimensional rotational digital angiography in a complicated case of patent ductus arteriosus transcatheter closure. Catheter Cardiovasc Interv. 2007;70(6):900–903. 64a. Frydrychowicz A, Bley TA, Dittrich S, et al. Visualization of vascular hemodynamics in a case of a large patent ductus arteriosus using flow sensitive 3D CMR at 3T. J Cardiovasc Magn Reson. 2007;9:585–587. 65. Rocchini AP. Coarctation of the aorta and interrupted aortica arch. In: Moller JH, Hofman JIE, eds. Pediatric Cardiovascular Medicine. 1st ed. 2000:567–593. 66. Varaprasathan GA, Araoz PA, Higgins CB, Reddy GP. Quantification of flow dynamics in congenital heart disease: applications of velocityencoded cine MR imaging. Radiographics. 2002;22(4):895–905. 67. Konen E, Merchant N, Provost Y, McLaughlin PR, Crossin J, Paul NS. Coarctation of the aorta before and after correction: the role of cardiovascular MRI. AJR Am J Roentgenol. 2004;182(5):1333–1339. 68. Haramati LB, Glickstein JS, Issenberg HJ, Haramati N, Crooke GA. MR imaging and CT of vascular anomalies and connections in patients with congenital heart disease: significance in surgical planning. Radiographics. 2002;22(2):337–347. 69. von Schulthess GK, Higashino SM, Higgins SS, Didier D, Fisher MR, Higgins CB. Coarctation of the aorta: MR imaging. Radiology. 1986;158(2):469–474. 70. Simpson IA, Chung KJ, Glass RF, Sahn DJ, Sherman FS, Hesselink J. Cine magnetic resonance imaging for evaluation of anatomy and flow relations in infants and children with coarctation of the aorta. Circulation. 1988;78(1):142–148. 71. Mohiaddin RH, Kilner PJ, Rees S, Longmore DB. Magnetic resonance volume flow and jet velocity mapping in aortic coarctation. J Am Coll Cardiol. 1993;22(5):1515–1521. 72. Oshinski JN, Parks WJ, Markou CP, et al. Improved measurement of pressure gradients in aortic coarctation by magnetic resonance imaging. J Am Coll Cardiol. 1996;28(7):1818–1826. 73. Steffens JC, Bourne MW, Sakuma H, O’Sullivan M, Higgins CB. Quantification of collateral blood flow in coarctation of the aorta by velocity encoded cine magnetic resonance imaging. Circulation. 1994;90(2):937–943. 74. Holmqvist C, Stahlberg F, Hanseus K, et al. Collateral flow in coarctation of the aorta with magnetic resonance velocity mapping: correlation to morphological imaging of collateral vessels. J Magn Reson Imaging. 2002;15(1):39–46. 75. Nielsen JC, Powell AJ, Gauvreau K, Marcus EN, Prakash A, Geva T. Magnetic resonance imaging predictors of coarctation severity. Circulation. 2005;111(5):622–628. 76. Cowley CG, Orsmond GS, Feola P, McQuillan L, Shaddy RE. Longterm, randomized comparison of balloon angioplasty and surgery for native coarctation of the aorta in childhood. Circulation. 2005;111(25):3453–3456. 77. Saeed M, Henk CB, Weber O, et al. Delivery and assessment of endovascular stents to repair aortic coarctation using MR and X-ray imaging. J Magn Reson Imaging. 2006;24(2):371–378. 78. Didier D, Saint-Martin C, Lapierre C, et al. Coarctation of the aorta: pre and postoperative evaluation with MRI and MR angiography; correlation with echocardiography and surgery. Int J Cardiovasc Imaging. 2006;22(3–4):457–475. 79. Schmidt M, Theissen P, Klempt G, et al. Long-term follow-up of 82 patients with chronic disease of the thoracic aorta using spin-echo and cine gradient magnetic resonance imaging. Magn Reson Imaging. 2000;18(7):795–806. 80. Bogaert J, Kuzo R, Dymarkowski S, et al. Follow-up of patients with previous treatment for coarctation of the thoracic aorta: comparison between contrast-enhanced MR angiography and fast spin-echo MR imaging. Eur Radiol. 2000;10(12):1847–1854. 81. Hager A, Kaemmerer H, Leppert A, et al. Follow-up of adults with coarctation of the aorta: comparison of helical CT and MRI, and impact on assessing diameter changes. Chest. 2004;126(4): 1169–1176.

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CHAPTER 30

Cardiovascular Magnetic Resonance in Complex Congenital Heart Disease Jens Bremerich, Rolf Wyttenbach, Peter T. Buser, and Charles B. Higgins

Cardiovascular magnetic resonance (CMR) is an attractive tool for noninvasive imaging of congenital heart disease (CHD) because it provides comprehensive cardiac morphologic and functional data without ionizing radiation, a particularly valuable attribute when considering imaging in children and young adults.1 The number of patients with CHD is continuously increasing as a result of extended survival produced by palliative and corrective surgical procedures. The role of CMR in the assessment of CHD is primarily defined by the use and limitations of two-dimensional transthoracic echocardiography (TTE), which has become the accepted standard for evaluation of most forms of CHD. Advantages of CMR compared with TTE include the following: (1) The entire thorax can be imaged with sequential high-resolution tomograms; (2) the information represents a continuous three-dimensional (3D) dataset that can be obtained in any plane; (3) the technique provides superior depiction of the central pulmonary arteries,2 including the systemic and pulmonary veins3 as well as the aorta.4 The major requirement for evaluation of CHD is the precise depiction of cardiovascular anatomy. Electrocardiographically (ECG) gated spin echo CMR can be used with high diagnostic accuracy for assessment of the morphologic features in simple and complex CHD.5 Cine gradient echo CMR permits accurate measurement of functional parameters such as left ventricular (LV) and right ventricular (RV) stroke volume, ejection fraction, regional wall motion, and wall thickening. Real-time imaging methods are becoming increasingly popular.6 Moreover, cine CMR can detect the jet flow associated with valvular lesions and intracardiac shunts. Operatorindependent isotropic 3D cine sequences are now available for assessment of the morphologic features and function in any plane reconstructed after the scan.7 With velocity encoded CMR, measurement of flow velocity and flow volume is possible, allowing quantification of pulmonary artery flow, shunt lesions, valvular regurgitation, and ventricular filling. These CMR techniques are ideally suited for sequential follow-up studies of surgically treated patients.8 Contrast-enhanced 3D magnetic resonance angiography (MRA) provides precise depiction of vessels such as pulmonary artery stenoses in tetralogy of Fallot9 and extracardiac thoracic vessels.10 Reconstruction of targeted maximum 408 Cardiovascular Magnetic Resonance

intensity projections (MIP) and multiplanar images (MPR) in any plane allows for precise assessment of the morphologic features of stenosis and measurement of the cross-sectional area. In aortic coarctation, the extent of collateral circulation through the intercostal and mammarian arteries provides an estimate of the hemodynamic relevance of the stenosis. Vascular function and abnormal connections are evaluated with contrast-enhanced dynamic timeresolved MRA.11 Tissue characterization is obtained through different sequence weighting and preparation pulses. Late gadolinium enhancement is visible in scar tissue or inflammation 20 minutes after injection of contrast agent. Inversion recovery sequences are applied to minimize signal from normal myocardium and to optimize contrast. After surgical repair of tetralogy of Fallot, dilation and scar formation in the RV outflow tract adversely affect RV hemodynamics.12,13 This chapter discusses the CMR appearance of cardiovascular morphology with the use of a segmental approach that represents the most rational way to evaluate complex CHD. The reader is also referred to consensus documents regarding the use of CMR in CHD.14 The segmental approach is based on morphologic identification of the great arteries, atria, and ventricles, and the visceroatrial relationship as well as the type of connection among these structures. There is also a brief review of the morphologic and functional evaluation of complex CHD with CMR pre- and postoperatively.

ATRIAL MORPHOLOGY AND DETERMINATION OF SITUS Axial spin echo CMR images extending from the base of the heart to the dome of the liver show the segmental cardiovascular anatomy. A segmental approach is based on localization of the three cardiac segments (atria, ventricles, and great arteries), the type of atrioventricular (AV) and ventriculoarterial (VA) connections, and the detection of associated anomalies (e.g., shunts, valve atresia).15,16 This segmental approach to CHD provides a precise description of the cardiac morphologic features, allowing accurate diagnosis of CHD with CMR.17 Atrial situs solitus is the normal situation in which the morphologic right (systemic venous) atrium is positioned

VENTRICULAR MORPHOLOGY AND ISOMERISM Transverse images at the midventricular level permit definition of the ventricular loop. The morphologic left ventricle (LV) and morphologic right ventricle (RV) have different anatomic characteristics that allow their differentiation. The morphologic RV can be recognized by its anterior and right-sided location (D-ventricular loop). The RV has prominent septomarginal trabecula passing from the apical portion of the ventricular septum to the anterior wall of the morphologic RV. The septal leaflet of the tricuspid valve is attached more anteriorly toward the cardiac apex than is the septal leaflet of the mitral valve. The depiction of different levels of insertion of the septal leaflets is an important diagnostic feature for distinction of the ventricles.19 Probably the most reliable sign for defining the RV is the presence of an infundibulum, or conus, that separates the tricuspid and pulmonary valves. This characteristic feature of the RV is best appreciated on axial images. The morphologic LV is located posteriorly and to the left. The septal leaflet of the mitral valve is located more distant from the cardiac apex compared with the tricuspid valve. The anatomic LV has a smoothly contoured apical portion of the ventricular septum and is characterized by a lack of complete muscular infundibulum. Therefore, the morphologic LV is characterized by a direct (fibrous) continuity between the mitral and aortic valves. The relationship of the great vessels can be determined by axial sections through the base of the heart. The ascending aorta can be identified by its continuity with the aortic arch and the brachiocephalic arteries. The main pulmonary artery is characterized by its bifurcation into the pulmonary arteries. Normally, the aorta lies posterior and to the right of the pulmonary trunk, and both vessels are similar in size.

The term isomerism means that both atria have features of the RA or the LA. In general, both atria develop with the same side as the thoracic and abdominal viscera (visceral-atrial rule). Therefore, bilateral left pulmonary artery anatomy with the artery passing over the left bronchus indicates left-sided isomerism (bilateral left-sidedness). These findings are best shown on coronal sections. Conversely, bilateral right pulmonary anatomy indicates rightsided isomerism (bilateral right-sidedness). The former is usually associated with polysplenia and the latter with asplenia. A previous report showed that gated spin echo CMR is highly accurate for determination of relationships among the great arteries, visceroatrial situs, and type of ventricular loop in patients with CHD.5 In another study, surgical planning was altered in some patients with heterotaxy syndrome, based on additional data supplied by CMR and not available by TTE or invasive cardiac catheterization.20

ABNORMALITIES OF THE ATRIOVENTRICULAR CONNECTION Once the atrial and ventricular morphology is determined, the next step is to determine whether the AV connection is concordant or discordant. Concordance is defined by a connection between the morphologic RA and the morphologic RV or between the morphologic LA and the morphologic LV, regardless of the positions of the atria and ventricles. Thus, discordance means that the morphologic RA drains to the morphologic LV and the morphologic LA drains to the morphologic RV, again regardless of the chamber position within the chest. An example of AV (and VA) discordance is corrected transposition (L-transposition), which is described later. AV discordance in conjunction with VA concordance is a rare malformation called isolated ventricular inversion. Other abnormalities of AV connections in which the terms concordant and discordant are not appropriate include the following: (1) double-inlet AV connection in which both atria are connected to a single ventricular chamber; (2) straddling AV valves in which one of the valves overlies the septum and drains blood into both ventricles; and (3) AV valve atresia in which one of the valves does not form and the AV ring may be replaced by fat (Fig. 30-1).21

ABNORMALITIES OF VENTRICULOARTERIAL CONNECTIONS Transposition Transposition of the great arteries (TGA) is one of the most common cyanotic forms of CHD. CMR has been shown to define the pathoanatomy of transposition and other abnormalities of VA connections accurately in several studies.5,17,19,22 In complete transposition, or D-transposition Cardiovascular Magnetic Resonance 409

30 CARDIOVASCULAR MAGNETIC RESONANCE IN COMPLEX CONGENITAL HEART DISEASE

on the right of the spine and the morphologic left (pulmonary venous) atrium is positioned to the left of the spine. In atrial situs inversus, the mirror image of the normal situation occurs, with the anatomic right atrium (RA) on the left side and the anatomic left atrium (LA) on the right side. The morphologic RA has distinctive anatomic characteristics, such as the broad-based, triangular appendage, whereas the morphologic LA contains an appendage with a narrow ostium and a more tubular configuration.18 The RA can also be identified by its connection to the inferior vena cava. In virtually all individuals, the side of the inferior vena cava defines the side of the RA. Furthermore, the situs of the atria can be defined by the visceral situs because they are nearly always concordant.16 Thus, the morphologic RA is determined by the side of the short main bronchus and the liver, whereas the long main bronchus, spleen, stomach, and aorta define the morphologic LA. In situs solitus, the left pulmonary artery courses cranially over the left main bronchus and the right pulmonary artery runs anterior and slightly inferior to the right bronchus.

RIGHT VENTRICULAR AND CONGENITAL HEART DISEASE

A

Figure 30-1 Electrocardiogram (ECG)-gated spin-echo transverse CMR at the level of the cardiac base (A) demonstrates side-by-side relationship of the aorta (A) and the main pulmonary artery (P) in a patient with DORV. Note the complete muscular ring surrounding both great arteries. Image B shows a stenosis of the left pulmonary artery (straight arrow) and an occluded right pulmonary artery (curved arrow).

P

A

B

show the anterior aorta arising from the morphologic RV and the posterior pulmonary artery arising from the morphologic LV.17 In patients with levo-TGA, also known as congenitally corrected TGA, the developing embryonic cardiac tube bends initially toward the left, rather than to the right, called an L-loop. In L-transposition, both AV and VA connections are discordant. This results in an aorta that is positioned anterior and leftward toward the pulmonary artery (Fig. 30-3). In addition, the aorta arises from the left-sided morphologic right ventricle and the pulmonary artery from the right-sided morphologic LV because the position of the ventricles is inverted (L-loop) in L-transposition. Therefore,

of the great arteries, VA discordance exists in the presence of AV concordance. The ventricles receive blood from the correct atrium, although the pulmonary artery is connected to the LV and the aorta is connected to the RV. CMR of the normal anatomy of the great artery at the base of the heart shows the aorta posterior and to the right of the pulmonary artery. Contrary to the normal anatomy, the transaxial images at the base of the heart in TGA show the anterior position of the aorta relative to the main pulmonary artery (Fig. 30-2). The aorta is located to the right of the pulmonary artery in D-transposition and to the left of the pulmonary artery in L-transposition. In patients with transposition of the great vessels, sagittal images can clearly

Figure 30-2 Complex congenital heart disease shown on T1-weighted fast spin echo (A) and gradient recalled echo (B to D) images. DTransposition with aorta (Ao) arising from the right ventricle (RV), which is connected to the left ventricle (LV) through a large ventricle septum defect (arrow in B). There is atresia of the tricuspid valve (D). Blood from the right atrium (RA) is directed through a large atrial septal defect to the left atrium (LA). Moreover, this patient has a left superior vena cava (asterisk in C and D) draining into the coronary sinus. Ao, aorta.

Ao

RV LV

A

B

Ao

RV RA

*

C

410 Cardiovascular Magnetic Resonance

LV LA

*

D

Ao Ao LA RV RV

B

A

RVOT

Ao RA

LA

C

D

the severity of valvular and subvalvular stenosis as well as pulmonary artery stenosis or atresia can be determined with CMR.23 Reversal in muscle thickness and shape in the RV compared with the LV, which is characteristic of transposition, can also be shown by CMR. In this respect, CMR is now regarded as the most accurate method to quantify ventricular mass. Consequently, cine CMR can be used to determine LV mass in children or adults in whom an arterial switch procedure is under consideration (Table 30-1).

systemic venous blood is pumped to the lungs by the LV through the pulmonary arteries, and oxygenated blood is pumped to the systemic circulation by the RV through the aorta. These anatomic features can be assessed readily by transverse spin echo CMR. In the coronal plane, the left-sided ascending aorta typically forms the upper heart border with L-transposition. Furthermore, CMR has the capability to show anomalies associated with TGA. Axial cine CMR images can assess the presence and severity of atrial and ventricular septal defects (VSDs). Additionally,

Table 30-1 Surgical Procedures Used to Treat Congenital Heart Disease Name

Description

Indication

Jatene

Arterial switch (Ao , PA) Reanastomose coronaries Atrial switch Conduit RV ) PA Anastomosis/conduit RA ) PA Anastomosis ascending Ao ) right PA Anastomosis descending Ao ) left PA Shunt subclavian artery ) PA Anastomosis SVC ) PA

TGA

Mustard, Senning Rastelli Fontan Waterston-Cooley Potts Blalock-Taussig Glenn

TGA Pulmonary atresia, TGA with PA stenosis, infundibular stenosis RVOT Tricuspid stenosis/atresia, hypoplastic RV, solitary ventricle TOF TOF TOF Tricuspid atresia, solitary ventricle, hypoplastic RV, septum intact

Ao, aorta; PA, pulmonary artery; RA, right atrium; RV, right ventricle; RVOT, right ventricular outflow tract; SVC, superior vena cava; TGA, transposition of the great arteries; TOF, tetralogy of Fallot.

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30 CARDIOVASCULAR MAGNETIC RESONANCE IN COMPLEX CONGENITAL HEART DISEASE

Figure 30-3 Congenitally corrected transposition (L-transposition) on steady-state free precession (SSFP) images. Blood from the left atrium (LA) is directed through the right ventricle (RV) to the aorta (Ao) as shown in A and B. Blood from the right atrium (RA) is directed through the left ventricle (arrow in D) to the pulmonary trunc. Note the aorta (C) arises from the right ventricular outflow tract (RVOT), which is characterized by a complete muscular ring.

RIGHT VENTRICULAR AND CONGENITAL HEART DISEASE

semilunar valves (uncommitted VSD). Double-outlet RV may be associated with valvular or subvalvular pulmonic stenosis. Especially if this arises in combination with a subaortic VSD, differentiation from tetralogy of Fallot may be difficult clinically and angiographically. Axial CMR images provide direct visualization of this type of double-outlet right ventricle by showing a complete circle of muscle separating the aortic from the mitral valve rather than indirectly gathering its presence from the distance between the aortic and mitral valve, as shown on left ventriculography.30 CMR can be used to determine the size and location of the VSD relative to the great arteries and to define subpulmonary or subaortic stenosis, the spatial relationships of the great vessels, and the status of the pulmonary arteries and the aortic arch (Fig. 30-4). These features are important for clinical and surgical management. Additionally, double-outlet right ventricle may be associated with atresia of the right AV valve. Because this condition will affect surgical repair, it is crucial to identify both AV valves in the axial plane.31

Double-Outlet Right Ventricle Double-outlet right ventricle is defined as an abnormal VA connection in which more than half of both the aorta and the pulmonary artery arise from the morphologic RV.24,25 On axial CMR images, double-outlet RV is typically characterized by side-by-side positioning of the great arteries at the semilunar valve level, with the aorta to the right of the pulmonary artery, although this relationship may be variable.26,27 An additional important feature of this condition is that neither semilunar valve is in direct fibrous continuity with the mitral valve. There is a complete rim of muscle separating both semilunar valves from the anterior mitral valve leaflet. On transverse CMR images, this side-by-side positioning of two muscular circles in the outflow region of the RV is diagnostic of double-outlet RV (see Fig 30-1). Coronal images define the side-by-side relationship of the aorta and the pulmonary artery at the level of the semilunar valves and their origin from the RV. The LV is shown to be separated from the semilunar valves. The only outlet for the LV becomes the requisite VSD. To plan a surgical repair adequately, it is important to determine the relative relationship of the VSD to the great vessels.28,29 This can usually be assessed on axial images. The VSD may be localized near the aortic valve (subaortic VSD), the pulmonic valve (subpulmonic VSD; Bing-Taussig anomaly), or both semilunar valves (doubly committed VSD). It may also be found remote from both

Truncus Arteriosus Persistent truncus arteriosus results from failure of division of the embryonic truncus into a separate aorta and pulmonary artery. This abnormality is infrequent, representing 0.4% of all cases of CHD.32 On axial, sagittal, and coronal

TA

RV LV

*

A

B

Ao W

Ao

C

412 Cardiovascular Magnetic Resonance

D

Figure 30-4 Pulmonary atresia with truncus arteriosus. Axial T1- weighted fast spin echo CMR (A) shows large ventricular septal defect between right (RV) and left (LV) ventricles. Axial gradient recalled echo image cranial to the ventricles (B) shows one large truncus arteriosus (TA) arising from both ventricles. No central pulmonary arteries can be identified. Palliation was done in childhood with a Waterston anastomosis (W in C) connecting the right pulmonary arteries to the ascending aorta (Ao). The Potts anastomosis connects the left pulmonary arteries to the descending aorta (arrow), but shows a stenosis. Contrast enhanced time resolved dynamic 2D MR angiography (D) shows perfusion of the right lung (arrowheads) but hypoperfusion of the left lung, confirming the hemodynamic relevance of the stenosis.

TETRALOGY OF FALLOT Consecutive transverse tomograms (spin echo; 5- to 7-mm slice thickness) through the entire heart and pulmonary hili show the following: (1) RV hypertrophy; (2) unequal division of the outflow tracts, with an enlarged and anteriorly displaced aorta; (3) membranous VSD; and (4) multilevel narrowing of the infundibulum, pulmonary annulus, main, and central pulmonary arteries. The infundibulum and pulmonary annulus are best shown on sagittal tomograms. The degree of pulmonary stenosis varies, and in extreme cases, the pulmonary trunk may not be identifiable. Severe pulmonary stenosis and atresia are usually associated with numerous large collateral channels arising from the aorta, principally, the descending aorta, and proceeding to the pulmonary hili.35 These vessels may be seen on flowsensitive transverse gradient echo CMR at the level of the carina or on coronal images. GRE techniques provide a bright signal from flowing blood. Stenoses of the central pulmonary arteries are frequent in tetralogy of Fallot and not uncommonly remain after initial surgical correction. These stenoses are best shown on a set of very thin tomograms with 3-mm slice thickness acquired in a plane parallel to the long axis of the right and left pulmonary arteries. The image plane should be parallel to the long axis of the right or left pulmonary artery. Most adult patients with tetralogy of Fallot have already undergone one or more corrective surgical procedures. CMR is an ideal technique for monitoring these patients after surgery. Cine CMR is used to monitor RV volume, mass, and ejection fraction.36 Velocity encoded cine CMR has been used to monitor pulmonary regurgitation, which occurs in most patients after total correction of the anomaly. One study has shown that the severity of pulmonary regurgitation, as quantified by velocity

encoded CMR, correlates with RV volume, mass, and ejection fraction.37 Failure of the RV can occur in tetralogy of Fallot. RV mass or functional parameters, such as ejection fraction and stroke volume, can be readily calculated from a stack of consecutive cine gradient echo CMR images acquired in the short axis of the heart and covering the entire RV. RV mass calculated from such CMR images has been shown to correlate with the width of the QRS complex. A wide QRS complex is a harbinger of RV arrhythmias.

EBSTEIN ANOMALY OF THE TRICUSPID VALVE Ebstein anomaly is an uncommon congenital developmental abnormality of the tricuspid valve that has a wide spectrum of pathologic anatomy. Although the diagnosis of Ebstein anomaly is usually straightforward and is based on transthoracis echocardiography findings, CMR may be helpful in defining the pathologic anatomy and tailoring the surgical method for each patient (Figs. 30-5 and 30-6). Axial CMR images are most informative. The tricuspid valve is almost always dysplastic in addition to being abnormally inserted. In fact, the dominant feature is dysplasia of the valve rather than displacement.38 The anterior leaflet usually attaches normally to the AV junction and is enlarged,39 and these features can be shown on CMR images in most cases.40 Distally, the anterior leaflet may be attached to an abnormal anterolateral papillary muscle. Therefore, it may be mobile, may have a continuous muscular connection with restricted mobility, or may be broadly plastered to the anterior wall of the RV and therefore would not be distinguishable on CMR.41,42 Septal and posterior leaflets are displaced downward in Ebstein anomaly,42 and they are best assessed on axial and coronal images, respectively.43 Both leaflets, however, may be deficient or absent in Ebstein anomaly and therefore would not be detectable on CMR.42,44 Planning of surgical repair must include assessment of the morphologic and functional features of the atrialized and hemodynamically effective portion of the RV. This assessment is best achieved with coronal T1-weighted spin echo and cine CMR. Reconstruction of the tricuspid valve with a prosthetic ring and vertical plication of the RA and the AV annulus may be complemented by an additional plication of the atrialized ventricle, depending on the size and function of the atrialized ventricle.44 Moreover, cine CMR enables assessment of tricuspid regurgitation, tricuspid stenosis, and shunts through the atrial septal defect (see Fig. 30-5).

COMPLEX VENTRICULAR ABNORMALITIES (SINGLE VENTRICLE) Compared with invasive angiography, CMR has many advantages for the definition of segmental anatomy and other defects in patients with a suspected diagnosis of single ventricle.45 The specific goals of CMR imaging in complex ventricular abnormalities are: (1) determination of Cardiovascular Magnetic Resonance 413

30 CARDIOVASCULAR MAGNETIC RESONANCE IN COMPLEX CONGENITAL HEART DISEASE

images, a single large vessel can be seen arising from the expected position of the semilunar valves just above the VSD. The truncus arteriosus gives rise to the aorta and the pulmonary and coronary arteries. Based on the configuration of the pulmonary arteries, three different types of truncus arteriosus are recognized. The most common, type I, is characterized by a short main pulmonary artery arising from the truncus. Bidirectional shunting occurs at the truncus, resulting in early cyanosis and eventual congestive heart failure. Axial T1-weighted CMR images can define the anatomy of the truncus and the pulmonary arteries.4 Sagittal and coronal planes are useful for demonstrating the origin of the pulmonary arteries from the truncus. For precise assessment of pulmonary artery caliber, a reduction of slice thickness to 3 mm is needed. Furthermore, CMR can show ventricular size and wall thickness as well as associated abnormalities, such as VSD, right aortic arch, or interrupted aortic arch. Axial and sagittal CMR tomograms are most effective for postoperative evaluation of the caliber of the conduit between the right ventricle and the pulmonary artery (Rastelli procedure). Narrowing of the conduit, stenosis at the origin of the right or left pulmonary artery, and pseudoaneurysm are complications that can be detected readily by CMR.33,34 Truncal insufficiency, which is common, may be detected by the use of cine CMR.

RIGHT VENTRICULAR AND CONGENITAL HEART DISEASE

*

A

B

RV RA LV RV

C

D

Ao

PT

E

F

Figure 30-5 Large atrial septal defect. Conventional X-rays show enlarged pulmonary (arrowhead) in PA (A) and enlarged right ventricle and right pulmonary artery in lateral view (B). Ebstein anomaly shown on axial T1-weighted fast spin echo (A, B) and gradient echo (C, D) images. Axial gradient recalled echo CMR (C) shows large atrial septal defect (arrow in C) and enlarged right atrium (RA) and ventricle (RV), also shown on short axis view (D). Velocity encoded magnitude (E) and phase (F) image show enlarged pulmonary trunc (PT). Through-plane flow measurement in the ascending aorta (circle) and pulmonary trunc yield a left to right shunt of 1:5.

visceral situs; (2) assessment of the type of ventricular loop, the morphologic features of the predominant ventricle (right, left, or primitive), and the position of the rudimentary ventricle; (3) definition of the AV and VA connections; (4) determination of the size of interventricular communication; and (5) definition of the connections of the systemic and pulmonary veins and arteries. Transverse gated spin echo tomograms with 7-mm slice thickness are, in general, the most useful for evaluation of complex cardiac anomalies.46,47 AV connections can be identified as double-inlet, absent left AV, or absent right AV connections. Stenoses or regurgitation can be detected on 414 Cardiovascular Magnetic Resonance

cine bright blood gradient echo CMR as flow void caused by spin dephasing in turbulent blood flow. After identification of the AV connection, the ventricular morphology must be determined. Distinction between dominant LV or dominant RV is usually possible by transverse and coronal CMR. When there is no detectable muscle separating either AV valve from the adjacent semilunar valve, the chamber is considered an LV. The position of a rudimentary RV is usually anterior and superior to the dominant ventricle, whereas the rudimentary LV is usually posterior and inferior to the dominant ventricle. A dominant LV is most common in adult patients. A primitive type of single ventricle has morphologic

* ** RA

A

B

*

C

features characteristic of neither the RV nor the LV. The communication between the dominant and the rudimentary ventricle can be assessed on both T1-weighted spin echo and cine gradient echo CMR.

POSTOPERATIVE EVALUATION Many patients with complex CHD have survived to adulthood after various palliative and corrective procedures (see Table 30-1). These patients require frequent monitoring at regular intervals with imaging studies. Transthoracic echocardiography is usually the initial technique used for this purpose, but it is not as ideal in adults as in children and is less quantitative. Moreover, many surgical procedures involve supracardiac as well as cardiac structures. Because the supracardiac structures are sometimes not well depicted by echocardiography, CMR is increasingly recognized as more effective for postoperative monitoring of older children and adults after complex surgical procedures.48 Several reports confirm the effectiveness of CMR for postoperative evaluation of complex CHD, suggesting that this noninvasive technique may obviate the serial use of invasive studies for postoperative follow-up in many cases.34,49,50 CMR not only is capable of visualizing cardiac

D

and extracardiac morphology, but also can quantify blood flow in the pulmonary arteries and conduits. Compared with echocardiography, CMR has the advantage of superior demonstration of conduits and anastomoses at the level of the great arteries. Furthermore, it is unaffected by postsurgical changes or graft material that can make echocardiography difficult.51 In addition, CMR has been found to be effective for monitoring pulmonary arterial status postoperatively and to be superior to echocardiography for evaluation of the pulmonary arteries.52,53 Many adult patients with TGA were treated with Mustard and Senning procedures (see Fig. 30-2). This atrial switch procedure is accomplished with a complex baffle constructed in the atria to direct blood flow from the pulmonary veins across the tricuspid valve into the RV and then to the aorta. Blood from the superior vena cava and the inferior vena cava flows through the baffle and then across the mitral valve into the pulmonary artery. On postoperative follow-up after a Mustard procedure, CMR in the coronal plane is particularly useful for evaluation of the superior systemic venous channel, which is the most common site of systemic venous obstruction. The sagittal and transaxial planes can show obstruction of pulmonary venous return or narrowing of the connection between the dorsal and ventral parts of the pulmonary venous atrium. In both procedures, the RV continues to work against systemic load. This may eventually result in RV systolic Cardiovascular Magnetic Resonance 415

30 CARDIOVASCULAR MAGNETIC RESONANCE IN COMPLEX CONGENITAL HEART DISEASE

Figure 30-6 Ebstein anomaly shown on axial T1-weighted fast spin echo (A and B) and gradient echo (C and D) images. Anterior displacement of the septal leaflet (arrow in A) of the tricuspid valve apically of the normal tricuspid plane (dashed line) results in the following changes: The functional right atrium comprises the morphologic right atrium (RA) plus the atrialised right ventricle (asterisks). The functional right ventricle is small (asterisk), which explains why such patients frequently suffer from right heart failure. Apical displacement of the septal leaflet of the tricuspid valve also leads to tricuspid insufficiency shown by a flow void on gradient echo images (arrow in D).

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dysfunction and tricuspid insufficiency because the RV and the tricuspid valve are not structured for systemic pressure load.54 Consequently, cine CMR can be used to monitor RV volume, mass, and ejection fraction.55 This method is also used to detect and estimate the severity of tricuspid regurgitation. Function of the systemic ventricle may also be monitored by through-plane velocity encoded cine measurements perpendicular to the ascending aorta.56 Currently, the favored procedure for TGA is the Jatene, or arterial switch, procedure, in which the aorta and the pulmonary artery are transected above the sinus portion and switched to redirect blood flow. The coronary arteries are then transplanted onto the neoaorta. This is usually done in the neonatal period. However, many older children or adults previously treated with the Senning or Mustard procedure are now candidates for the arterial switch procedure to avert or relieve RV pressure overload or failure. CMR has a major role in the long-term follow-up of patients after surgical repair of TGA.50,57 CMR can assess various complications after an atrial switch procedure for d-TGA, including baffle leaks, systemic or pulmonary venous obstructions, LV outflow tract stenosis, tricuspid regurgitation, and coronary anatomy.58 Transverse and sagittal spin echo CMR is used for assessment of the great vessel anatomy after a Jatene procedure. Because this procedure is frequently performed in the neonatal period, thin sections of 3-mm thickness are preferable. Postoperatively, the position of the aorta, posterior to the main pulmonary artery, occasionally results in proximal stenosis of the right or left pulmonary artery. This stenosis can be clearly seen on axial CMR images. In addition, CMR can visualize other complications after a Jatene procedure, such as narrowing of the RV outflow region, dilation of the aortic root, and supravalvular aortic stenosis.34,57 Compared with transthoracic echocardiography, CMR has been shown to be virtually equal in depicting stenoses of the RV outflow region. However, compared with echocardiography, CMR was superior in the detection of proximal pulmonary artery stenosis in patients who underwent the Jatene procedure (41% vs. 94%).59 Furthermore, CMR is unaffected by postsurgical changes, graft material, or an inadequate acoustic window, which may render echocardiography difficult. The Fontan procedure consists of an anastomosis between the RA or right atrial appendage and the main pulmonary artery.60 Numerous variations of the Fontan procedure have been devised. Many patients initially have a bidirectional Glenn shunt (SVC-to-right pulmonary artery anastomosis) followed months or years later by placement of a conduit from the inferior vena cava to the right or left pulmonary artery. Therefore, systemic venous blood is forwarded directly to the pulmonary circulation, bypassing the functional single ventricle, which is functioning as a systemic pumping chamber. The concomitant ASD is also closed. The major indications for Fontan procedure are tricuspid atresia or severe stenosis, single ventricle, and hypoplastic left heart. CMR can be used to show the size of the atriopulmonary connection, assess the distribution of pulmonary perfusion, and recognize the presence of obstruction.61,62 Axial and coronal images are usually effective for this purpose. Besides obstruction of the conduit, complications of the Fontan procedure include residual ASD, systemic venous hypertension, and thrombosis.63 The former can be diagnosed with cine CMR, and the latter may 416 Cardiovascular Magnetic Resonance

result in right atrial enlargement, venous stasis, pleural and peritoneal effusion, and edema. Severe right atrial enlargement may even compress the right pulmonary veins at the entrance to the LA.34 Determination of the size of the pulmonary arteries is important in patients undergoing Fontan reconstruction because artery size is considered a major indicator of prognosis. CMR has been shown to be useful in determining pulmonary artery size in these patients and has been found to be superior to echocardiograpy.54 Surgical repair of tetralogy of Fallot is achieved by closing the VSD and enlarging the pulmonary outflow tract with patches. Axial CMR scans can show abnormalities in the RV outflow tract, including residual stenosis or aneurysmal patch dilation. In the case of concomitant pulmonary artery atresia, surgical repair is more complex, necessitating systemic-to-pulmonary shunts to allow blood flow to the lungs and to promote the growth of pulmonary vessels. The subclavian-to-pulmonary artery shunt (Blalock-Taussig), which was used in the past, has been largely replaced by the modified Blalock shunt, representing a graft connecting the aorta or brachiocephalic artery with the pulmonary artery. An earlier report showed the usefulness of electrocardiographically gated CMR to assess the size, course, and patency of Blalock-Taussig, Glenn, and aortopulmonary shunts.50 Coronal and axial CMR imaging planes are particularly useful for demonstrating systemicto-pulmonary shunts as well as potential complications, including stenosis or occlusion of the shunt. Although long-term results after surgical repair of tetralogy of Fallot are good, most patients have some degree of pulmonary regurgitation. Long-standing pulmonary regurgitation may result in severe arrhythmia and sudden death. Moreover, exercise capacity is often diminished in patients with severe dilation of the RV. Recent studies showed a 30% decrease in RV dilation and improved systolic function after valve replacement in adults who had undergone repair of tetralogy of Fallot.64 These authors suggest the use of axial cine CMR to evaluate RV function and double oblique velocity encoded CMR perpendicular to the vessel to assess flow measurement. Through-plane flow is encoded up to 200 cm/sec. Recurrence of pulmonary regurgitation after surgery seems to predict reduced recovery of RV systolic function.65 Repair of pulmonary atresia may also be accomplished by placing a valve conduit between the RV and the main pulmonary artery (Rastelli conduit). Sagittal CMR images are the most effective way to visualize the proximal anastomosis to the RV and the distal anastomosis to the pulmonary artery. Possible complications include false aneurisms at the anastomosis and stenosis or kinking of the conduit. Whereas spin echo CMR is used for morphologic evaluation of postoperative CHD, CMR techniques, such as cine steady state free precession CMR and velocity encoded cine CMR, allow functional assessment of surgical baffles, conduits, and valvular function. In patients who had undergone a Mustard or Senning procedure for TGA, cine CMR was able to evaluate pulmonary and systemic venous connections as well as RV systolic function and tricuspid and mitral regurgitation.66–69 Flow quantification with velocity encoded cine CMR improved the evaluation of venoatrial connections after a Mustard or Senning procedure.70,71 In addition, CMR velocity mapping has been used successfully to assess tricuspid volume flow in patients after Mustard or

EVALUATION OF FUNCTION IN CONGENITAL HEART DISEASE Cine gradient echo CMR and velocity encoded cine CMR are attractive methods for quantifying ventricular function and volumetric flow, respectively, in patients with untreated or repaired CHD. Standard or breath hold (segmented kspace) cine CMR provides sequential images through the cardiac cycle, and the images can be viewed as a cine loop. Typically, cine CMR uses short repetition times (20 to 35 msec), short echo times (4 to 20 msec), and low flip angles (35 to 60 ) to acquire 16 phases spaced evenly throughout the cardiac cycle.75 Cine CMR images show high signal intensity in areas of normal blood flow. However, turbulent flow, which may occur in stenosis, regurgitation, or shunt lesions, causes a signal loss within the blood pool, rendering these lesions readily visible on cine CMR. In addition, LV and RV mass and systolic function can be measured precisely with 3D cine CMR images (see Fig. 30-4). Unlike transthoracic echocardiography or invasive left ventriculography, CMR does not rely on geometric assumptions or calculations based on partial sampling of the cardiac volume. This 3D dataset from end-diastolic and end-systolic tomograms encompassing both ventricles allows measurement of LV and RV volume, mass, stroke volume, and ejection fraction. In contrast to typical adult cardiologic studies, the RV is often of particular interest in CHD. Double-oblique short axis tomograms are used to quantify biventricular volume and global function, with inclusion of trabeculations and papillary muscles in the RV cavity suggested for enhanced reproducibility.69 End-diastolic and end-systolic measurements acquired through the entire stack of images provide end-diastolic (EDV) and end-systolic volume (ESV), stroke volume (SV ¼ EDV ESV), and ejection fraction (EF ¼ SV/EDV) for both the LV and the RV. In normal individuals, LV stroke volume and RV stroke volume are the same.76 Therefore, differences in ventricular stroke volume can be used to quantify valvular regurgitation and shunt lesions.77 For example, in pulmonary or tricuspid regurgitation, the difference between RV stroke volume and LV stroke volume corresponds to regurgitant volume. LV stroke volume is greater than RV stroke volume in patients with patent ductus arteriosus, aortic regurgitation, and mitral regurgitation. Measurements of volume and function of the ventricles have been shown to be highly reproducible on sequential studies in patients with morphologically normal and abnormal LV.78 Thus, cine CMR is a highly attractive method for detecting changes in ventricular volume and function over time.

Velocity encoded cine CMR provides direct measurement of aortic and pulmonary artery flow and therefore measures the effective stroke volume of both ventricles. In the absence of valvular regurgitation, this method can be used to determine intracardiac shunt volume. For example, in ASD, partial anomalous pulmonary venous connection, or VSD, the difference between RV stroke volume and LV stroke volume is the left-to-right shunt volume. Likewise, velocity encoded CMR can be used to quantify shunt volume in patent ductus arteriosus by calculating the difference between net forward flow in the ascending aorta and that in the main pulmonary artery. Velocity encoded cine CMR has been used successfully in patients with various congenital shunt lesions.79 The accuracy and reproducibility of this method for measuring Qp:Qs in left-to-right shunts have been assessed previously.80,81 Velocity encoded CMR can measure blood flow in the main pulmonary artery as well as separately in the left and right pulmonary arteries. This is one of the few techniques with the capability to quantify left and right pulmonary flow separately.81,82 This method can be used to determine the percentage of blood flow to each lung and assess the hemodynamic significance of stenoses of the left or right pulmonary artery. Tricuspid regurgitation and shunts may be detected and quantified during the same examination.83 In patients who have undergone a Mustard or Senning procedure, abnormal tricuspid flow patterns have been found to precede RV systolic dysfunction.84 Velocity encoded cine CMR can also be used to determine the severity of coarctation by quantifying collateral flow in the descending aorta. Measurements of volume flow are performed in the proximal descending aorta just below the site of the coarctation and in the distal descending aorta near the diaphragm. In healthy subjects, flow volume decreases steadily from the proximal to the more distal parts of the descending aorta because of antegrade flow into the intercostal arteries. In patients with hemodynamically significant coarctation of the aorta, the normal flow pattern in branches of the descending aorta is reversed, resulting in an increase in flow from the proximal to the distal descending aorta because of retrograde collateral flow mainly through intercostal arteries.85 Velocity encoded CMR can also be used for identification of patients with a mismatch between the severity of anatomic obstruction and collateral flow, which may be of importance for planning surgery.86,87 Velocity encoded cine CMR has been used successfully to demonstrate abnormalities in volume flow in the descending aorta in adult patients after repair of coarctation. These abnormalities are probably related to resistance to flow imposed by the coarctation segment and may represent an additional index for monitoring the hemodynamic significance of coarctation of recoarctation in patients before and after intervention.87,88 Thus, cine CMR allows for a comprehensive evaluation of aortic coarctation by determining the location and severity of stenosis and pressure gradients across the coarctation segment.85–88

CONCLUSION Because of the progress in surgery in recent decades, an increasing number of patients with complex CHD survive to adulthood after having undergone previous palliative Cardiovascular Magnetic Resonance 417

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Senning procedure, and it often showed abnormal tricuspid flow patterns.72 Velocity encoded cine CMR also provides accurate information about pulmonary flow volume and velocity after Fontan procedure. Velocity encoded cine CMR can be used to assess the volume of retrograde flow in patients after a Fontan procedure.73 Velocity encoded CMR has also been used to quantify the volume of pulmonary regurgitation after patch repair of tetralogy of Fallot74 and to estimate pressure gradients across ventriculopulmonary (Rastelli) conduits.

RIGHT VENTRICULAR AND CONGENITAL HEART DISEASE

or corrective surgery. Other patients with milder form of CHD may become symptomatic only as adults. Patients with CHD often require numerous follow-up studies for evaluation of morphology and function. CMR imaging is a noninvasive method that provides excellent anatomic detail of cardiac structures and the great vessels, even when echocardiography access to the area of interest is limited. With cine CMR, it is possible to quantify valvular and shunt lesions and to measure ventricular volume without making any geometric assumptions. Velocity encoded CMR can be

used to measure flow velocity and volume in the heart and the great arteries. CMR is particularly useful for postoperative follow-up in patients with repaired CHD, especially after placement of intra-atrial baffles or extracardiac conduits. Furthermore, CMR can assess RV function, which is of particular interest in many patients with CHD and which may be difficult for other imaging modalities. Thus, CMR is emerging as the noninvasive imaging technique of choice that allows comprehensive assessment of cardiovascular anatomy and function in patients with complex CHD.

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19. Guit GL, Bluemm R, Rohmer J, et al. Levotransposition of the aorta: identification of segmental cardiac anatomy using MR imaging. Radiology. 1986;161:376. 20. Geva T, Wesley Vick III G, Wendt RE, Rockey R. Role of spin echo and cine magnetic resonance imaging in presurgical planning of heterotaxy syndrome. Circulation. 1994;90:348. 21. Fletcher BD, Jacobstein MD, Abramowsky CR, et al. Right atrioventricular valve atresia: anatomic evaluation with MR imaging. Am J Roentgenol. 1987;148:671. 22. Higgins CB, Byrd III BF, Farmer DW, et al. Magnetic resonance imaging in patients with congenital heart disease. Circulation. 1984;70:851. 23. Kersting-Sommerhoff BA, Sechtem UP, Higgins CB. Evaluation of pulmonary artery supply by nuclear magnetic resonance imaging in patients with pulmonary atresia. J Am Coll Cardiol. 1989;11:166. 24. Lev M, Bharati S, Meng CCL, et al. A concept of double-outlet right ventricle. J Thorac Cardiovasc Surg. 1972;64:271. 25. Wilcox BR, Ho SY, Macartney FJ, et al. Surgical anatomy of doubleoutlet right ventricle with situs solitus and atrioventricular concordance. J Thorac Cardiovasc Surg. 1981;82:405. 26. Niezen RA, Beekman RP, Helbing WA, van der Wall EE, de Roos A. Double outlet right ventricle assessed with magnetic resonance imaging. Int J Card Imaging. 1999;15:323–329. 27. Mayo JR, Roberson D, Sommerhoff B, Higgins CB. MR imaging of double outlet right ventricle. J Comput Assist Tomogr. 1990;14:336. 28. Patrick DL, McGoon DC. An operation for double outlet right ventricle with transposition of the great arteries. J Cardiovasc Surg. 1968;9:537. 29. Sridaromont S, Feldt RH, Ritter DG, et al. Double outlet right ventricle: hemodynamic and anatomic correlations. Am J Cardiol. 1976;38:85. 30. Higgins CB. Congenital heart disease. In: Higgins CB, Hricak H, Helms CA, eds. Magnetic Resonance Imaging of the Body. 3rd ed. Philadelphia: Lippincott-Raven; 1996:461. 31. Smith JRWL, Stanford W, Skorton DJ, Wolf GL. Assessment of congenital heart disease by nuclear magnetic resonance imaging. In: Skorton DJ, ed. Marcus Cardiac Imaging: A Companion to Braunwald’s Heart Disease. 2nd ed., vol. 2. Philadelphia: W.B. Saunders Company; 1996:886. 32. Donnelly LF, Higgins CB. MR imaging of conotruncal abnormalities. AJR Am J Roentgenol. 1996;166:925. 33. Murashita T, Hatta E, Imamura M, Yasuda K. Giant pseudoaneurysm of the right ventricular outflow tract after repair of truncus arteriosus: evaluation by MR imaging and surgical approach. Eur J Cardiothorac Surg. 2002;22:849–851. 34. Kersting-Sommerhoff BA, Seelos KC, Hardy C, et al. Evaluation of surgical procedures for cyanotic congenital heart disease by using MR imaging. AJR Am J Roentgenol. 1990;155:259. 35. Berry BE, McGoon DC, Ritter DG, Davis GD. Absence of anatomic origin from heart of pulmonary arterial supply: clinical application of classification. J Thorac Cardiovasc Surg. 1974;68:119. 36. van Straten A, Vliegen HW, Hazekamp MG, de Roos A. Right ventricular function late after total repair of tetralogy of Fallot. Eur Radiol. 2005;15:702–707. 37. Niezen RA, Helbing WA, van der Wall EE, et al. Biventricular systolic function and mass studied with MR imaging in children with pulmonary regurgitation after repair for tetralogy of Fallot. Radiology. 1996;201:135. 38. Becker AE, Becker MJ, Edwards JE. Pathologic spectrum of dysplasia of the tricuspid valve: features in common with Ebstein’s malformation. Arch Pathol. 1971;91:167. 39. Anderson KR, Lie JT. Pathologic anatomy of Ebstein’s anomaly of the heart revisited. Am J Cardiol. 1978;41:739.

66. Fontan F, Deville C, Quaegebeur J, et al. Repair of tricuspid atresia in 100 patients. J Thoracic Cardiovasc Surg. 1983;85:647. 67. Rees S, Sommerville J, Warnes C, et al. Comparison of magnetic resonance imaging with echocardiography and radionuclide angiography in assessing cardiac function and anatomy following Mustard’s operation for transposition of the great arteries. Am J Cardiol. 1988;61: 1316. 68. Chung KJ, Simpson IA, Glass RF, et al. Cine magnetic resonance imaging after surgical repair in patients with transposition of the great arteries. Circulation. 1988;77:104. 69. Winter MM, Bernink JP, Groenink M, et al. Evaluating the systemic right ventricle by CMR: the importance of consistent and reproducible delineation of the cavity. J Cardiovasc Magn Reson. 2008;10:40. 70. Varaprasathan GA, Araoz PA, Higgins CB, Reddy GP. Quantification of flow dynamics in congenital heart disease: applications of velocityencoded cine MR imaging. Radiographics. 2002;22:895–905. 71. Sampson C, Kilner PJ, Hirsch R, et al. Venoatrial pathways after the Mustard operation for transposition of the great arteries: anatomic and functional MR imaging. Radiology. 1994;193:211. 72. Rebergen SA, Helbing WA, van der Wall EE, et al. MR velocity mapping of tricuspid flow in healthy children and in patients who have undergone Mustard or Senning repair. Radiology. 1995;194:505. 73. Rebergen SA, Ottenkamp J, Doornbos J, et al. Postoperative pulmonary flow dynamics after Fontan surgery assessment with nuclear magnetic resonance velocity mapping. J Am Coll Cardiol. 1993;21:123. 74. Rebergen SA, Chin JGJ, Ottenkamp J, et al. Pulmonary regurgitation in the late postoperative follow-up of tetralogy of Fallot: volumetric quantitation by nuclear magnetic resonance velocity mapping. Circulation. 1993;88:2257. 75. Sechtem U, Pflugfelder PW, White RD, et al. Cine MR imaging: potential for the evaluation of cardiovascular function. AJR Am J Roentgenol. 1987;148:239–246. 76. Wang ZJ, Reddy GP, Gotway MB, Yeh BM, Higgins CB. Cardiovascular shunts: MR imaging evaluation. Radiographics. 2003;23:S181–S194. 77. Sechtem U, Pflugfelder PW, Gould RG, et al. Measurement of right and left ventricular volumes in healthy individuals with cine MR imaging. Radiology. 1987;163:697. 78. Semelka RC, Tomei E, Wagner S, et al. Interstudy reproducibility of dimensional and functional measurements between cine magnetic resonance studies in the morphologically abnormal left ventricle. Am Heart J. 1990;119:1376. 79. Rees S, Firmin D, Mohiaddin R, et al. Application of flow measurements by magnetic resonance velocity mapping to congenital heart disease. Am J Cardiol. 1989;64:953. 80. Brenner LD, Caputo GR, Mostbeck GH, et al. Quantification of left-toright atrial shunts with velocity-encoded cine nuclear magnetic resonance imaging. J Am Coll Cardiol. 1992;20:1246. 81. Roman KS, Kellenberger CJ, Farooq S, MacGowan CK, Gilday DL, Yoo SJ. Comparative imaging of differential pulmonary blood flow in patients with congenital heart disease: magnetic resonance imaging versus lung perfusion scintigraphy. Pediatr Radiol. 2005;35:295–301. 82. Sieverding L, Jung WI, Klose U, Apirz J. Noninvasive blood flow measurement and quantification of shunt volume by cine magnetic resonance in congenital heart disease. Pediatr Radiol. 1992;22:48. 83. Theissen P, Kaemmerer H, Sechtem U, et al. Magnetic resonance imaging of cardiac function and morphology in patients with transposition of the great arteries following Mustard procedure. Thorac Cardiovasc Surg Suppl. 1991;39:221. 84. Araoz PA, Reddy GP, Tarnoff H, Roge CL, Higgins CB. MR findings of collateral circulation are more accurate measures of hemodynamic significance than arm-leg blood pressure gradient after repair of coarctation of the aorta. J Magn Reson Imaging. 2003;17:177–183. 85. Nishimura RA, Housmans PR, Hatle LK, Tajik AJ. Assessment of diastolic function of the heart: background and current applications of Doppler echocardiography. I. Physiologic and pathophysiologic features. Mayo Clin Proc. 1989;64:71. 86. Steffens JC, Bourne MW, Sakuma H. Quantification of collateral blood flow in coarctation of the aorta by velocity encoded cine magnetic resonance imaging. Circulation. 1994;90:937. 87. Mohiaddin RH, Kilner PJ, Rees S, Longmore DB. Magnetic resonance volume flow and jet velocity mapping in aortic coarctation. J Am Coll Cardiol. 1993;22:1515. 88. Oshinski JN, Parks WJ, Markou CP, et al. Improved measurement of pressure gradients in aortic coarctation by magnetic resonance imaging. J Am Coll Cardiol. 1996;28:1818.

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40. Link KM, Herrera MA, D’Souza VJ, Formanek AG. MR imaging of Ebstein anomaly: results in four cases. AJR Am J Roentgenol. 1988;150:363. 41. Markiewicz W, Sechtem U, Higgins CB. Evaluation of the right ventricle by magnetic resonance imaging. Am Heart J. 1987;113:8. 42. Choi YH, Park JH, Choe YH, Yoo SJ. MR imaging of Ebstein’s anomaly of the tricuspid valve. AJR Am J Roentgenol. 1994;163:539. 43. Lev M, Liberthson RR, Joseph RH, et al. The pathologic anatomy of Ebstein’s disease. Arch Pathol. 1970;90:334. 44. Carpentier A, Chauvaud S, Mace L, et al. A new reconstructive operation for Ebstein’s anomaly of the tricuspid valve. J Thorac Cardiovasc Surg. 1988;96:92. 45. Fogel MA. Is routine cardiac catheterization necessary in the management of patients with single ventricles across staged Fontan resonstruction? No! Pediatr Cardiol. 2005;26:154–158. 46. Kersting-Sommerhoff BA, Diethelm L, Stanger P, et al. Evaluation of complex ventricular anomalies with magnetic resonance imaging. Am Heart J. 1990;120:133. 47. Higgins CB, Byrd BF, Farmer DW, et al. Magnetic resonance imaging in patients with congenital heart disease. Circulation. 1984;70:851. 48. Fogel MA, Hubbard A, Weinberg PM. A simplified approach for assessment of intracardiac baffles and extracardiac conduits in congenital heart surgery with two- and three-dimensional magnetic resonance imaging. Am Heart J. 2001;142:1028–1036. 49. Higgins CB, Byrd BF, McNamara MT, et al. Magnetic resonance imaging of the heart: a review of the experience in 172 subjects. Radiology. 1985;155:671. 50. Jacobstein MD, Fletcher BD, Nelson AD, et al. Magnetic resonance imaging: evaluation of palliative systemic-pulmonary artery shunts. Circulation. 1984;70:650. 51. Soulen RL, Donner RM, Capitanio M. Postoperative evaluation of complex congenital heart disease by magnetic resonance imaging. Radiographics. 1987;7:975. 52. Sampson C, Martinez J, Rees S, et al. Evaluation of Fontan’s operation by magnetic resonance imaging. Am J Cardiol. 1990;65:819. 53. Fogel MA, Donofrio MT, Ramaciotti C, et al. Magnetic resonance and echocardiographic imaging of pulmonary artery size throughout stages of Fontan reconstruction. Circulation. 1994;90:2927. 54. Duerinckx AJ, Wexler L, Banerjee A, et al. Postoperative evaluation of pulmonary arteries in congenital heart surgery by magnetic resonance imaging: comparison with echocardiography. Am Heart J. 1994;128:1139. 55. Hornung TS, Kilner PJ, Davlouros PA, Grothues F, Li W, Gatzoulis MA. Excessive right ventricular hypertrophic response in adults with the Mustard procedure for transposition of the great arteries. Am J Cardiol. 2002;90:800–803. 56. Laffon E, Jimenez M, Latrabe V, et al. Quantitative MRI comparison of systemic hemodynamics in Mustard/Senning repaired patients and healthy volunteers at rest. Eur Radiol. 2004;14:875–880. 57. Mee RB. Severe right ventricular failure after Mustard or Senning operation. Two stage repair: pulmonary artery banding and switch. J Thorac Cardiovasc Surg. 1986;92:385. 58. Taylor AM, Dymarkowski S, Hamaekers P, et al. MR coronary angiography and late-enhancement myocardial MR in children who underwent arterial switch surgery for transposition of great arteries. Radiology. 2005;234:542–547. 59. Blankenberg F, Rhee J, Hardy C, et al. MRI vs echocardiography in the evaluation of the Jatene procedure. J Comput Assist Tomogr. 1994;18:749. 60. Fontan F, Baudet E. Surgical repair of tricuspid atresia. Thorax. 1971;26:240. 61. Fratz S, Hess J, Schwaiger M, Marinoff S, Stern HC. More accurate quantification of pulmonary blood flow by magnetic resonance imaging than by lung perfusion scintigraphy in patients with fontan circulation. Circulation. 2002;106:1510–1513. 62. Weir RA, Steedman T, Hillis WS, Swan L. Relief of Fontan obstruction demonstrated non-invasively by cardiac magnetic resonance imaging. Int J Cardiol. 2008;127:e167–e169. 63. Casolo G, Rega L, Gensini GF. Detection of right atrial and pulmonary artery thrombosis after the Fontan procedure by magnetic resonance imaging. Heart. 2004;90:825. 64. Vliegen HW, Van Straten A, de Roos A, et al. Magnetic resonance imaging to assess the hemodynamic effects of pulmonary valve replacement in adults late after repair of tetralogy of Fallot. Circulation. 2002;106:1703–1707. 65. Van Straten A, Vliegen HW, Hazekamp MG, et al. Right ventricular function after pulmonary valve replacement in patients with tetralogy of Fallot. Radiology. 2004;233:824–829.

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CHAPTER 31

Complex Congenital Heart Disease: Infant and Pediatric Patients Tal Geva, Adam L. Dorfman, and Andrew J. Powell

The role of cardiovascular magnetic resonance (CMR) in the evaluation of infants and children with complex congenital heart disease (CHD) is increasing as a result of continued technologic advances. Improvements in hardware and the development of new, highly efficient imaging techniques have allowed for sufficient spatial and temporal resolution to evaluate cardiac anatomy and function in pediatric patients comprehensively, despite these patients’ small body size and rapid heart rate. Although transthoracic echocardiography provides the necessary diagnostic information in most patients in this age group, CMR offers an important alternative in certain circumstances: (1) when transthoracic echocardiography is incapable of providing the required diagnostic information; (2) when clinical assessment and other diagnostic tests are inconsistent; (3) as an alternative to diagnostic cardiac catheterization, with its associated risks and higher cost; and (4) to obtain diagnostic information for which CMR offers unique advantages. It is helpful to note that complex CHD has no precise accepted definition. For the purposes of this chapter, the definition of complex CHD includes conotruncal anomalies and single-ventricle heart disease. Conotruncal anomalies refer to a group of congenital defects involving the outflow tracts of the heart and the great vessels. Single-ventricle heart disease refers to a heterogeneous group of anomalies in which one of the two ventricular sinuses is absent (anatomic single ventricle) or to hearts with complex anatomy in which biventricular physiology cannot be attained (functional single ventricle). The special considerations that pertain to patient preparation, sedation, and monitoring in pediatric patients undergoing CMR are discussed in Chapter 9.

TETRALOGY OF FALLOT Tetralogy of Fallot (TOF) is the most common type of cyanotic CHD, with an incidence of 356 per million live births.1 Although TOF involves several anatomic components, the anomaly is believed to result from a single developmental anomaly, underdevelopment of the subpulmonary infundibulum (conus).2,3 The anatomy is characterized by infundibular and valvar pulmonary stenosis associated with anterior, superior, and leftward deviation of the infundibular (conal) septum; hypoplasia of the pulmonary valve annulus and 420 Cardiovascular Magnetic Resonance

thickened leaflets; ventricular septal defect (VSD); and overriding of the aortic valve above the ventricular septum. The degree of right ventricular outflow tract (RVOT) obstruction varies from mild to complete (i.e., TOF with pulmonary atresia). The size of the mediastinal pulmonary arteries (PAs) varies considerably. Although in some patients they can be dilated (e.g., TOF with absent pulmonary valve syndrome), more commonly, their diameter ranges from normal to hypoplastic. In some patients, the mediastinal PAs are discontinuous or absent. In patients with pulmonary atresia or diminutive or absent branch PAs, pulmonary blood flow may come from a patent ductus arteriosus, from collateral vessels arising from the aorta or its branches, or from both sources. The VSD in TOF is usually located between the malaligned conal septum superiorly and the muscular septum inferiorly (termed conoventricular septal defect4). The aortic valve is rotated clockwise (as viewed from the apex) and is positioned above the ventricular septal crest, committing to both the left ventricle (LV) and the right ventricle (RV). In 5% to 6% of patients with TOF, a major coronary artery crosses the RVOT.5 Most commonly, the left anterior descending (LAD) coronary artery originates from the right coronary artery (RCA) and traverses the infundibular free wall before reaching the anterior interventricular groove. Preoperative identification of a major coronary artery crossing the RVOT is important to avoid inadvertent damage to the coronary artery during surgical relief of RVOT obstruction. Additional cardiovascular and noncardiac anomalies can be associated with TOF.6 Although the clinical presentation and course of patients with TOF can vary, most patients have cyanosis during the first year of life. Some patients with mild or no RVOT obstruction are not cyanotic at birth (“pink TOF”) and may exhibit signs and symptoms of pulmonary overcirculation similar to patients with a large VSD. As these patients grow, the subpulmonary infundibulum becomes progressively obstructive and cyanosis ensues.7 Surgical repair of TOF is usually performed during the first year of life, often during the first 6 months.8 A typical repair includes patch closure of the VSD and relief of the RVOT obstruction using a combination of resection of obstructive muscle bundles and an overlay patch. When the pulmonary valve annulus is moderately or severely hypoplastic, the RVOT patch is extended across the pulmonary valve annulus into the main pulmonary artery (MPA),

Cardiovascular Magnetic Resonance Evaluations Tetralogy of Fallot is the most frequent diagnosis among patients referred for CMR evaluation at Children’s Hospital Boston. Unlike infants in whom transthoracic echocardiography generally provides all of the necessary diagnostic information for surgical repair,5,16 CMR assumes an increasing role in adolescents and adults with TOF, in whom the acoustic windows are frequently limited. CMR is useful in both pre- and postoperative assessment of TOF, but the focus of the examination is different.

Preoperative Cardiovascular Magnetic Resonance Assessment In most patients with unrepaired TOF, the central question for the CMR examination is to delineate all sources of pulmonary blood flow, including the PAs, aortopulmonary

collaterals, and ductus arteriosus. Several studies have shown that black-blood spin echo and two-dimensional (2D) gradient recalled echo (GRE) cine CMR techniques provide excellent imaging of the central PAs and major aortopulmonary collaterals.17–20 However, these CMR techniques require relatively long scan times for complete anatomic coverage, and small vessels (<2 mm) may not be detected. Furthermore, these 2D techniques are not optimal for imaging long and tortuous blood vessels, some of which arise from the brachiocephalic arteries or the abdominal aorta. Three-dimensional (3D) gadolinium contrast-enhanced magnetic resonance angiography (CE-MRA) is ideally suited to image these vessels (Fig. 31-1). Compared with conventional X-ray angiography, CE-MRA is highly accurate in depicting all sources of pulmonary blood supply in patients with complex pulmonary stenosis or atresia, including infants with multiple small aortopulmonary collaterals.21,22 Electrocardiographically (ECG) triggered cine steadystate free precession (SSFP) is used to assess ventricular dimensions and function, the RVOT, and valve function. When the origins and proximal course of the left and right coronary arteries are not known from other imaging studies, they should be evaluated by a sequence designed for coronary imaging. Particular attention is paid to the exclusion of a major coronary artery crossing the RVOT.

Postoperative Cardiovascular Magnetic Resonance Assessment Cardiovascular magnetic resonance has been used extensively for assessment of postoperative TOF.23,24 Quantitative assessment of LV and RV volume and systolic function is a key element of CMR evaluation in patients with repaired TOF. The degree of RV dysfunction is an important determinant of clinical status late after TOF

DAo

A

B

Figure 31-1 Multiple aortopulmonary collateral vessels in a newborn with tetralogy of Fallot and pulmonary atresia evaluated by gadoliniumenhanced three-dimensional magnetic resonance angiography. A, Subvolume maximum intensity projection image in the coronal plane showing multiple aortopulmonary collateral vessels arising from the descending aorta. B, Subvolume maximum intensity projection image in the transverse plane showing aortopulmonary collateral vessels arising from the descending aorta (DAo) and splitting into two large branches, one to the left lung and one to the right lung (arrows). Cardiovascular Magnetic Resonance 421

31 COMPLEX CONGENITAL HEART DISEASE: INFANT AND PEDIATRIC PATIENTS

sacrificing much of the valve mechanism and leading to pulmonary regurgitation. In patients with TOF and pulmonary atresia, or when a major coronary artery crosses the RVOT, a conduit—either a homograft or a prosthetic tube—is placed between the RVOT and the PAs. Most patients with repaired TOF have residual hemodynamic abnormalities, primarily because of right ventricular dilation from chronic pulmonary regurgitation. Other sequelae include RV pressure overload from RVOT or pulmonary arterial obstruction, RV dysfunction, tricuspid regurgitation, LV dilation from a residual shunt or a VSD, and aortic dilation. Conduction and rhythm abnormalities are another major source of late morbidity and mortality in this growing patient population.9–15

RIGHT VENTRICULAR AND CONGENITAL HEART DISEASE

MPA

RPA

* Ao RA

LV LPA

A

B

Figure 31-2 Three-dimensional contrast-enhanced magnetic resonance angiography in an infant with a large pseudoaneurysm of the main pulmonary artery (MPA) after tetralogy of Fallot repair. A, Coronal plane maximum intensity projection image showing the pseudoaneurysm (arrow) originating from the MPA through a narrow neck (*). B, Oblique transverse maximum intensity projection image showing the aneurysm and branch pulmonary arteries. Ao, ascending aorta; LPA, left pulmonary artery; LV, left ventricle; RA, right atrium; RPA, right pulmonary artery.

and is also closely associated with LV dysfunction.25,26 Many studies have shown that the degree of pulmonary regurgitation measured by velocity encoded cine (VEC) CMR is closely associated with the degree of right ventricular dilation.27–30 Another factor that affects RV function is the presence and extent of an aneurysm in the RVOT (Fig. 31-2).31 Taken together with clinical assessment and electrophysiologic data, information derived from CMR on pulmonary regurgitation fraction, LV and RV size and function, the presence and extent of an RVOT aneurysm, and the presence of branch PA stenosis is used to direct clinical care in patients with repaired TOF. Another technique that is increasingly being used in patients with repaired CHD is late gadolinium enhancement (LGE) for assessment of myocardial fibrosis. LGE is common late after TOF repair, and a greater extent of RV involvement is associated with older age, worse symptoms, exercise intolerance, right ventricular dysfunction, and clinical arrhythmia.32 The goals of CMR examination, therefore, include: (1) quantitative assessment of LV and RV volume, mass, stroke volume, and ejection fraction; (2) imaging the anatomy of the RVOT, PAs, aorta, and aortopulmonary collaterals; and (3) quantification of pulmonary regurgitation, tricuspid regurgitation, cardiac output, and pulmonary-to-systemic flow ratio. These objectives can be achieved by the following protocol: Axial, sagittal, and coronal scout images ECG gated cine SSFP sequences in the two-chamber and four-chamber planes ECG gated cine SSFP sequence in the short axis plane across the ventricles from base to apex for quantitative assessment of ventricular dimensions and function ECG gated cine SSFP sequence parallel to the RVOT and PAs 3D CE-MRA ECG gated VEC CMR perpendicular to the main PA ( branch PAs), ascending aorta, and mitral and tricuspid valves LGE to evaluate for the presence of scar tissue 422 Cardiovascular Magnetic Resonance

Whenever possible, SSFP cine sequences are acquired with breath holding. Fast (turbo) spin echo (FSE) with double inversion recovery sequence may be used to minimize artifacts from metallic implants, when present.

TRANSPOSITION OF THE GREAT ARTERIES Transposition of the great arteries (TGA) is defined as discordant connections between the ventricles and the great arteries; the aorta arises from the RV and the MPA arises from the LV. There are several anatomic types of TGA, depending on the visceroatrial situs (solitus or inversus) and the type of ventricular loop (D or L).33 The most common type of TGA is visceroatrial situs solitus (S), ventricular D-loop (D), and dextro-malposition of the aortic valve relative to the pulmonary valve (D). This anatomic arrangement can be summarized as {S,D,D} TGA. Note that the term D-TGA is nonspecific because the D might relate to the ventricular loop or to the spatial position of the aortic and pulmonary valves. This ambiguity can be avoided by using the term D-loop TGA. The incidence of D-loop TGA is estimated at 303 per million live births.1 The principal physiologic abnormality in D-loop TGA is that systemic venous blood returns to the aorta and oxygenated pulmonary venous blood returns to the lungs, resulting in profound hypoxemia. Consequently, survival is dependent on communications that allow mixing of blood between the systemic and pulmonary circulations. The most common sites of shunting are through a ductus arteriosus, atrial septal defect, or VSD. Associated anomalies include VSD in approximately 45% of patients, coarctation or interrupted aortic arch in approximately 12%, pulmonary stenosis in approximately 5%, RV hypoplasia in approximately 4%, and juxtaposition of the atrial appendages in approximately 2%.34 Surgical management of D-loop TGA in the 1960s and 1970s consisted mostly of an atrial switch procedure—the Senning or Mustard procedure. In both procedures, systemic and pulmonary venous blood is redirected within

of the left ventricular outflow tract (LVOT) and RVOT for obstruction; and (5) detection of aortopulmonary collateral vessels and other associated anomalies. Data suggest that including RV trabeculations and papillary muscles as ventricular volume (rather than wall/mass) results in shorter analysis time and better interobserver reproducibility for ejection fraction.59a In patients with RV dysfunction, LGE can be used to detect myocardial fibrosis, which has been shown to be associated with right ventricular dysfunction.60 The response of the systemic RV to pharmacologic stress (dobutamine) or to exercise can be tested by CMR, but the clinical utility of this information awaits further study.57,59 The major objectives of CMR evaluation of postoperative atrial switch can be achieved with the following protocol: Axial, sagittal, and coronal scout images ECG gated cine SSFP sequence in the axial or fourchamber plane, with multiple contiguous slices from the level of the diaphragm to the level of the transverse arch (provides dynamic imaging of the venous pathways, qualitative assessment of ventricular function, AV valve regurgitation, and imaging of the great arteries) Based on the previous sequence, ECG gated cine SSFP sequence in multiple oblique coronal planes parallel to the superior vena cava and inferior vena cava pathways to obtain long axis images ECG gated cine SSFP sequence in the plane across the ventricles from base to apex for quantitative assessment of LV and RV volume and systolic function (Fig. 31-3) Free breathing (navigator-gated), ECG triggered, isotropic, 3D SSFP sequence (Fig. 31-4)61 3D CE-MRA ECG gated VEC CMR perpendicular to the mitral/tricuspid valves, MPA, and ascending aorta; additional VEC

Cardiovascular Magnetic Resonance Evaluations Cardiovascular magnetic resonance is seldom requested for preoperative assessment of infants with D-loop TGA because transthoracic echocardiography usually provides all of the necessary diagnostic information.34 In postoperative TGA, CMR assumes an increasing role because of its ability to evaluate most clinically relevant issues noninvasively.51–59

RV

LV

POSTOPERATIVE ATRIAL SWITCH The goals of CMR evaluation of postoperative atrial switch include: (1) quantitative evaluation of the size and function of the systemic RV52; (2) imaging of the systemic and pulmonary venous pathways for obstruction or baffle leak; (3) assessment of tricuspid valve regurgitation; (4) evaluation

Figure 31-3 Electrocardiographically gated steady-state free precession cinecardiovascular magnetic resonance in the short axis plane for assessment of biventricular volume, mass, and function in a patient with Senning procedure. Note the dilated and hypertrophied systemic right ventricle (RV) and the thin-walled, lowpressure subpulmonary left ventricle (LV). Cardiovascular Magnetic Resonance 423

31 COMPLEX CONGENITAL HEART DISEASE: INFANT AND PEDIATRIC PATIENTS

the atria so that the pulmonary venous blood reaches the tricuspid valve, RV, and aorta, whereas the systemic venous blood reaches the mitral valve, LV, and PAs. The main technical difference between these two procedures is that in the Mustard procedure, the surgeon uses pericardium to redirect blood flow and in the Senning procedure the surgeon uses native atrial tissue.35 The main drawbacks of the atrial switch procedures include RV (systemic ventricle) dysfunction, sinus node dysfunction, atrial arrhythmias, obstruction of the systemic or pulmonary venous pathways, and baffle leaks.36–38 Beginning in the late 1970s and rapidly gaining popularity in the 1980s, the arterial switch procedure largely replaced the atrial switch procedures.39,40 The advantages of the arterial switch procedure over the atrial switch procedures include establishment of the LV as the systemic ventricle and avoidance of extensive suture lines in the atria. Recent data on late outcome of the arterial switch procedure continue to show excellent overall survival with low morbidity.41–46 The Rastelli procedure is another surgical option for patients with associated subvalvar and valvar pulmonary stenosis and a VSD. It consists of patch closure of the VSD so that the LV outflow is directed to the aortic valve and placement of a conduit between the RV and the PAs. The second most common type of TGA is visceroatrial situs solitus (S), L-ventricular loop (L), and levomalposition of the aortic valve relative to the pulmonary valve (L). This anatomic arrangement can be summarized as {S,L,L} TGA, or L-loop TGA.47–50 It is also known as physiologically corrected TGA because the systemic venous return reaches the pulmonary circulation through the right-sided LV and the pulmonary venous return reaches the aorta through the left-sided RV. Associated anomalies include tricuspid valve abnormalities (e.g., Ebsteins anomaly), RV hypoplasia, VSD, subvalvar and valvar pulmonary stenosis, dextrocardia, and conduction abnormalities, including complete heart block. Outcome is determined primarily by the associated lesions and RV (systemic ventricle) function.47–50

RIGHT VENTRICULAR AND CONGENITAL HEART DISEASE

Figure 31-4 Free breathing, navigator-gated, electrocardiographically-triggered, isotropic, three-dimensional steady-state free precession in a patient with Mustard palliation of transposition of the great arteries. The three-dimensional data volume is reformatted off-line in multiple user-defined planes. A, Transverse view of the pulmonary venous pathway. *Systemic venous atrium; PVA, pulmonary venous atrium. B, Oblique view showing the superior vena cava (SVC). C, Oblique sagittal view of the left (LV) and right ventricular (RV) outflow tracts. *Main pulmonary artery; Ao, aorta.

RV PVA LV

*

A

SVC

PVA LV

B

Ao

*

RV

LV

C CMR sequences may be obtained to evaluate specific areas suspected of having an obstruction62 LGE to evaluate the presence of scar tissue Whenever possible, SSFP cine sequences are acquired with breath holding. FSE with double inversion recovery sequence may be used to minimize artifacts from metallic implants. 424 Cardiovascular Magnetic Resonance

The long-term concerns in patients after the arterial switch procedure relate primarily to the technical challenges of the procedure, including transfer of the coronary arteries from the native aortic root to the neoaortic root (native pulmonary root) and the transfer of the PAs anterior to the neo-ascending aorta (Fig. 31-5). Some of the postoperative complications include stenosis of the neo-main PA,

LPA

LCA MPA

DOUBLE-OUTLET RIGHT VENTRICLE RV

LV

Figure 31-5 Three-dimensional contrast-enhanced magnetic resonance angiography in a postoperative arterial switch procedure for D-loop transposition of the great arteries. The main pulmonary artery (MPA) is anterior to the ascending aorta. Note the diffuse hypoplasia of the left pulmonary artery (LPA). The proximal left coronary artery (LCA) originates from the neoaortic root. LV, left ventricle; RV, right ventricle.

branch PA stenosis, dilation of the neoaortic root, semilunar valve regurgitation, and coronary artery stenosis.34,40,41,45,63 Consequently, the goals of CMR evaluation of postoperative arterial switch include: (1) evaluation of global and regional LV and RV size and function; (2) evaluation of the LVOT and RVOT for obstruction; (3) qualitative estimation of RV systolic pressure based on the configuration of the interventricular septum; (4) imaging of the great vessels with emphasis on evaluation of the PAs for stenosis and the aortic root for dilation; (5) definition of the coronary arteries; and (6) detection of aortopulmonary collaterals and other associated anomalies. The role of myocardial perfusion and viability imaging in this population needs further study.64 These objectives can be achieved by the following protocol: Axial, sagittal, and coronal scout images ECG gated cine SSFP sequence in the axial plane with multiple contiguous slices from the midventricular level to the level of the transverse arch (provides axial dynamic imaging of the outflow tracts and the great arteries and qualitative assessment of ventricular function and mitral/tricuspid valve regurgitation) ECG gated cine SSFP sequence in the coronal or oblique sagittal planes parallel to the LVOT and RVOT ECG gated cine SSFP sequences in the two- and fourchamber planes followed by a stack across the ventricles from base to apex for quantitative assessment of LV and RV volume and systolic function Free-breathing (navigator-gated), ECG triggered, 3D coronary MRI 3D CE-MRA

Double-outlet right ventricle (DORV) is defined as a specific type of ventriculoarterial alignment in which both great vessels arise from the RV or from the infundibulum. The incidence of DORV is estimated at 127 per million live births.1 It is important to recognize the wide spectrum of anatomic and physiologic variations that share this type of ventriculoarterial alignment. The clinical course and management of patients with DORV are dictated in large part by the size and location of the VSD in relation to the semilunar valves, the anatomy of the infundibulum and the semilunar valves, the position of the infundibular septum, the size of the left and right ventricular sinuses, and the anatomy of the mitral/tricuspid valves. The LV can be of normal size, hypoplastic, or absent. The RV is usually of normal size, but in rare circumstances, it can be hypoplastic or even absent (double-outlet infundibulum). Both semilunar valves can be patent, but stenosis or atresia is relatively common. The mitral or tricuspid valve may have attachments in both ventricles (straddling), and this abnormality is particularly important with regard to surgical planning. The ultimate goal of surgical management of DORV is to align the LV with systemic outflow and the RV with pulmonary outflow. In DORV with a subaortic VSD, the LV can be aligned with the aorta by placing a patch on the right ventricular aspect of the defect, leaving the aortic valve on the left ventricular side. Resection of RVOT obstruction, with or without an outflow patch, may be necessary in patients with subvalvar or valvar pulmonary stenosis, analogous to TOF repair. In DORV with a subpulmonary VSD (Taussig-Bing variety), the VSD is closed with a patch that directs the blood from the LV to the pulmonary valve, accompanied by an arterial switch procedure. Concomitant repair of an aortic arch anomaly (e.g., hypoplasia, interruption, or coarctation) is often required as well. More complex forms of DORV, with heterotaxy syndrome, severe hypoplasia or absence of one of the ventricular sinuses, major straddling of an atrioventricular valve, or mitral atresia, are palliated as a functional single ventricle with an eventual Fontan procedure. Late complications in patients with repaired DORV are relatively common and vary with the underlying anatomy, physiology, and surgical repair. Subaortic stenosis can develop after the LV is baffled to the aorta.65 Complications after repair of DORV with pulmonary outflow tract obstruction are similar to those seen after TOF repair, including Cardiovascular Magnetic Resonance 425

31 COMPLEX CONGENITAL HEART DISEASE: INFANT AND PEDIATRIC PATIENTS

ECG gated VEC CMR perpendicular to the main and branch PAs and the ascending aorta; additional VEC CMR sequences may be obtained to evaluate specific areas suspected of having an obstruction62 LGE to evaluate the presence of scar tissue Whenever possible, SSFP cine sequences are acquired with breath holding. FSE with double inversion recovery sequence may be used to minimize artifacts from metallic implants. Pharmacologic stress testing, either adenosine or dobutamine, may be performed for evaluation of myocardial ischemia.

RIGHT VENTRICULAR AND CONGENITAL HEART DISEASE

chronic pulmonary regurgitation, dilation and dysfunction of the RV, and arrhythmia. Aortic arch obstruction can be found in patients after aortic coarctation or interrupted aortic arch repair. Those who undergo an arterial switch procedure may have the same problems described for D-loop TGA.

Preoperative Cardiovascular Magnetic Resonance Assessment Because echocardiography is usually sufficient for diagnosis and surgical planning in most newborns or infants with DORV, CMR is seldom requested for preoperative evaluation in this age group. Exceptions include patients with complex anomalies of the aortic arch, PAs, or aortopulmonary collaterals, and systemic or pulmonary venous anomalies that are not completely delineated by echocardiography. Several investigators demonstrated the use of CMR for assessment of the relationship between the great vessels and the VSD as well as the position of the great vessels in relation to the conal septum.66–70 The imaging strategy is tailored to address the specific clinical question. In general, 3D CE-MRA is particularly helpful for evaluation of great vessel anatomy. Intracardiac anatomy is assessed by SSFP cine CMR and by free breathing, navigator-gated, ECG triggered, isotropic, 3D SSFP (Fig. 31-6).61 In patients with metallic implants, FSE with double inversion recovery is helpful for assessment of both intra- and extracardiac anatomy.

Ao

MPA

RV

LA

Figure 31-6 Free breathing, navigator-gated, electrocardiographically triggered, isotropic, three-dimensional steadystate free precession imaging in a patient with a double-outlet right ventricle (RV). The ascending aorta (Ao) and main pulmonary artery (MPA) originate from the RV. LA, left atrium. 426 Cardiovascular Magnetic Resonance

Postoperative Cardiovascular Magnetic Resonance Assessment The role of CMR after DORV repair increases as patients grow and their acoustic windows become progressively more limited. The examination strategy is tailored based on the underlying anatomy, the procedure performed, and the specific clinical and other diagnostic findings. Although no single generic imaging protocol covers all possible scenarios after DORV repair, certain patterns are recognized. Patients with a “TOF-like” DORV repair have similar longterm sequelae as those after TOF repair, and the CMR examination protocol is comparable (discussed earlier). Similarly, in patients with Taussig-Bing-type DORV who have undergone an arterial switch procedure, the postoperative issues are similar to those encountered after the arterial switch procedure for TGA (discussed earlier). The importance of modifying the examination protocol to address anatomic and functional abnormalities specific to the individual patient cannot be overstated. This requires ongoing evaluation of the imaging data as the examination proceeds because unsuspected abnormalities may only be detected during the scan.

TRUNCUS ARTERIOSUS Truncus arteriosus is an uncommon conotruncal anomaly with a reported incidence of only 94 per million live births.1 It is defined by the presence of a single artery arising from the heart, with a single semilunar valve giving rise to the coronary arteries, aorta, and at least one branch PA. Van Praagh and Van Praagh71 modified the original classification of Collett and Edwards, as follows (Fig. 31-7)72: Type I: The branch PAs arise from a short main PA. Type II: The branch PAs arise directly from the arterial trunk through separate orifices. Type III: Only one branch PA arises from the ascending segment of the trunk. Collateral vessels usually supply the contralateral lung (Fig. 31-8). Type IV: Truncus arteriosus with aortic arch hypoplasia, coarctation, or interruption (usually type B, distal to the left common carotid artery). In this anatomic variation, there is usually a well-formed main PA and a small ascending aorta (Fig. 31-9). In most cases, there is a subtruncal VSD over which the truncal valve sits, similar to TOF. Rarely, the ventricular septum is intact. The conal septum is usually absent and the truncal valve is in direct fibrous continuity with the mitral valve. In rare circumstances, the truncal valve may be supported by a complete infundibulum and relate exclusively to the RV. The truncal valve is most commonly tricommissural, with the bicommissural type the next most common, followed by the quadricommissural type. The valve leaflets can be thickened and redundant with stenosis, regurgitation, or both. Associated cardiovascular and noncardiac anomalies are frequent. Examples of associated cardiovascular anomalies include multiple VSDs, partial and complete atrioventricular canal defects, mitral atresia, mitral stenosis, hypoplastic LV, double-inlet LV, tricuspid atresia, straddling tricuspid

A1

A2

A3

A4

valve, Ebsteins malformation, heterotaxy syndrome, aberrant origin of the right or left subclavian artery, coarctation of the aorta, secundum atrial septal defect, partially and totally anomalous pulmonary venous connections, left superior vena cava to the coronary sinus, retroaortic innominate vein, and left PA sling.73 Various noncardiac anomalies have been described in patients with truncus arteriosus. DiGeorge syndrome, velocardiofacial syndrome, and chromosome 22q11 deletion are frequently associated. A large series found 22q11 deletion in 34.5% of patients with truncus arteriosus.74 Most patients with truncus arteriosus are diagnosed early in life by transthoracic echocardiography, which is sufficient for surgical planning in almost all cases. Typically, this lesion is repaired shortly after diagnosis by closing the VSD with a patch so that the truncal valve is aligned with the LV (becoming the neoaortic valve), and the PAs are detached from the arterial trunk and connected to the RV with a valved homograft. Surgical repair of the truncal valve for stenosis or regurgitation is uncommon during initial repair. The surgical mortality rate is low and has improved with the overall advances in the surgical management of infants. The use of a nongrowing homograft in infancy makes additional procedures inevitable as patients grow. Important residual lesions after truncus arteriosus repair include progressive stenosis and regurgitation of the RV-to-PA homograft, branch PA stenosis, and regurgitation or stenosis of the neoaortic (truncal) valve. Aortic arch obstruction can complicate the course of patients after coarctation or interrupted aortic arch repair.

Preoperative Cardiovascular Magnetic Resonance Assessment Cardiovascular magnetic resonance is rarely requested for preoperative evaluation in an infant with truncus arteriosus because echocardiography is almost always adequate.75 Exceptions include complex aortic arch or pulmonary venous anomalies that require further delineation and the occasional older patient with an unrepaired truncus arteriosus (see Figs. 31-8 and 31-9).

Postoperative Cardiovascular Magnetic Resonance Assessment The role of CMR in patients with repaired truncus arteriosus increases with patient age. The anatomic and functional issues in these patients are similar to those encountered in patients with repaired TOF, especially those with TOF and pulmonary atresia. Neoaortic valve dysfunction and aortic arch obstruction are additional issues that may require investigation. Therefore, the goals of CMR examination after truncus repair include: (1) quantitative assessment of LV and RV volume, function, and mass; (2) measurement of pulmonary and neoaortic valve regurgitation; (3) imaging of the RVOT, the homograft, and the branch PAs; (4) assessment of residual shunts; and (5) imaging of the aortic Cardiovascular Magnetic Resonance 427

31 COMPLEX CONGENITAL HEART DISEASE: INFANT AND PEDIATRIC PATIENTS

Figure 31-7 Classification of truncus arteriosus. A1, The branch pulmonary arteries arise from a short main pulmonary artery; A2, The branch pulmonary arteries arise directly from the arterial trunk through separate orifices; A3, Only one branch pulmonary artery arises from the ascending segment of the trunk. Collateral vessels usually supply the contralateral lung; A4, Truncus arteriosus with aortic arch hypoplasia, coarctation, or interruption (usually type B, distal to the left common carotid artery).

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INTERRUPTED AORTIC ARCH

Ao

LPA

Tr

Figure 31-8 Cardiovascular magnetic resonance evaluation of truncus arteriosus with an absent right pulmonary artery (truncus arteriosus type A3 of Van Praagh). Three-dimensional reconstruction of contrast-enhanced magnetic resonance angiography showing the origin of the left pulmonary artery (LPA) from the truncal root (Tr). Ao, aortic arch.

arch and isthmus. These objectives can be achieved with modifications of the protocol described earlier for TOF and individualized for the patient’s anatomic and hemodynamic issues.

Interruption of the aortic arch is an uncommon CHD malformation characterized by anatomic discontinuity between segments of the aortic arch. The prevalence of interruption of the aortic arch is only 19 per million live births, or 1.3% of infants with CHD in the New England Regional Infant Cardiac Program.76 This condition should be distinguished from aortic arch atresia, in which there is anatomic continuity between the arch segments through a fibrous strand but the aortic lumen is completely obstructed. Because of their identical hemodynamic consequences, both conditions will be discussed together. The classification proposed by Celoria and Patton in 1959 is widely used (Fig. 31-10).77 Type A denotes interruption distal to the left subclavian artery, type B interruption between the left common carotid and the left subclavian arteries, and type C interruption between the common carotid arteries. Type B is the most common anatomic variation, accounting for approximately 62% of cases of interruption of the aortic arch, with type A accounting for 37%, and type C accounting for 1%. Other morphologic variations associated with interruption of the aortic arch include aberrant origin of the right subclavian artery from the proximal descending aorta (found in 50% of patients with type B interruption but only in a minority of those with type A interruption). Other rare variations include interruption of the right aortic arch78 and interruption of the cervical arch.79 Survival of patients with interruption of the aortic arch depends on a patent ductus arteriosus (PDA). Intravenous administration of prostaglandin E begins immediately once the diagnosis is suspected and is followed by surgical repair. In most institutions, the preferred surgical approach is direct anastomosis of the interrupted (or atretic) aortic segments. When the distance between the

LCCA LSCA Ao Tr

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428 Cardiovascular Magnetic Resonance

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Figure 31-9 Three-dimensional reconstruction contrast-enhanced magnetic resonance angiography in a newborn with truncus arteriosus and type B interrupted aortic arch (truncus type A4 of Van Praagh). A, Anterior view showing the ascending aorta (Ao), which gives rise to the right innominate artery and the left common carotid artery. The main pulmonary artery (MPA) continues as a large patent ductus arteriosus (PDA), which supplies the descending aorta. B, Posterior view showing the interruption between the left common carotid artery (LCCA) and the left subclavian artery (LSCA), which arises from the proximal descending aorta. LPA, left pulmonary artery; RPA, right pulmonary artery; Tr, truncal root.

RSCA

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interrupted aortic arch segments is large, homograft augmentation may be added to the arch reconstruction. The use of a tubular conduit to bridge the arch segments is usually reserved for unusually long segment interruptions or for reoperations. In patients with an associated VSD, the defect is closed at the time of arch repair. In patients with type B interruption with posterior malalignment of the conal septum and markedly hypoplastic LVOT, the VSD is baffled to the pulmonary valve, the main PA is transected and anastomosed to the ascending aorta, and a conduit (usually a valved homograft) is placed between the RV and the PAs.

Cardiovascular Magnetic Resonance Evaluation

CE-MRA (Fig. 31-11).82 The use of gradient recalled echo cine and black-blood FSE imaging is tailored to the clinical and imaging issues of individual patients. Evaluation of the intracardiac anatomy is usually not necessary because the information should be available from echocardiography.

Isthmus PDA

DAo

Transthoracic echocardiography is usually adequate for the preoperative diagnosis of interrupted aortic arch and associated anomalies.80 CMR is used in select patients in whom the anatomy is not clearly defined with echocardiography.81,82 CMR assumes a larger role in patients with repaired interruption of the aortic arch as they grow and their acoustic windows become restricted.

Preoperative Cardiovascular Magnetic Resonance Assessment The goal of CMR examination is to delineate the anatomy of the aortic arch and the branching pattern of the brachiocephalic arteries. It is important to evaluate the vascular anatomy fully to exclude associated anomalies (e.g., systemic and pulmonary venous anomalies). The most robust and time-efficient technique to achieve these goals is 3D

Figure 31-11 Subvolume maximum intensity projection threedimensional contrast-enhanced magnetic resonance angiography showing type A interrupted aortic arch distal to the left subclavian artery. Note the short segment of luminal discontinuity (arrow) between the aortic isthmus and the descending aorta (DAo). PDA, patent ductus arteriosus. Cardiovascular Magnetic Resonance 429

31 COMPLEX CONGENITAL HEART DISEASE: INFANT AND PEDIATRIC PATIENTS

Figure 31-10 Classification of interrupted aortic arch. Ao, aorta; LCCA, left common carotid artery; LSCA, left subclavian artery; MPA, main pulmonary artery; PDA, patent ductus arteriosus; RCCA, right common carotid artery; RSCA, right subclavian artery.

RIGHT VENTRICULAR AND CONGENITAL HEART DISEASE

Postoperative Cardiovascular Magnetic Resonance Assessment The goal of CMR examination after interruption surgery is to evaluate residual or recurrent anatomic and hemodynamic problems. Often the focus is on imaging of the aortic arch and the repair site for evaluation of obstruction or aneurysm formation (Fig. 31-12). However, other abnormalities, such as LVOT obstruction, aortic valve stenosis or regurgitation, and residual VSD, as well as LV volume and function should be examined as well. The hemodynamic severity of residual or recurrent aortic arch obstruction can be assessed based on body surface area-adjusted smallest cross-sectional area of the aortic arch or isthmus (from CE-MRA) and the heart-rate-adjusted mean deceleration rate in the descending aorta (from VEC CMR), as described by Nielsen and colleagues83 These objectives can be achieved with the following protocol: Axial, sagittal, and coronal scout images ECG gated cine SSFP sequences in the two- and fourchamber planes ECG gated cine SSFP sequence in the short axis plane across the ventricles from base to apex for quantitative assessment of ventricular volume, function, and mass ECG gated cine SSFP sequence parallel to the LVOT ECG gated cine SSFP sequence in the long axis of the aortic arch ECG gated FSE in the long axis of the aortic arch (optional) ECG gated VEC CMR perpendicular to the ascending and descending aorta; additional flow measurements are obtained based on clinical relevance (e.g., assessment of aortic regurgitation) 3D CE-MRA

* DAo Ao MPA

A

B

Figure 31-12 Three-dimensional reconstruction contrastenhanced magnetic resonance angiography in postoperative type B interrupted aortic arch repaired with a conduit (*) between the distal ascending aorta (Ao) and the proximal descending aorta (DAo). A, Anterior view. B, Left posterior view. MPA, main pulmonary artery. 430 Cardiovascular Magnetic Resonance

Measurement of blood pressure in the upper and lower extremities at the time of CMR provides additional useful information provided the upper extremity artery is proximal to the site of obstruction.

SINGLE VENTRICLE The normal human heart is composed of three chambers at the ventricular level (between the atrioventricular [AV] valves and the semilunar valves): LV sinus, RV sinus, and infundibulum. From an anatomic standpoint, single ventricle is defined as a circumstance in which one of the two ventricular sinuses is absent. The infundibulum is always present and has been described using various terms, such as infundibular outlet chamber, rudimentary or hypoplastic RV, and rudimentary chamber. As defined earlier, single ventricle accounts for approximately 1% of cases of CHD, with a median incidence of 85 per million live births.1 There are other congenital cardiac anomalies in which the anatomy precludes establishment of biventricular physiology. These conditions, which are also often treated with one of the modifications of the Fontan procedure, are often grouped under the term functional single ventricle or functional univentricular heart. Examples include tricuspid atresia, mitral atresia, unbalanced common AV canal defect, pulmonary atresia with intact ventricular septum and diminutive RV, and tricuspid valve. From an anatomic perspective, there are two types of single ventricle: 1. Single LV (Fig. 31-13): Several anatomic types of single LV are recognized. Common to all is absence of the RV sinus and the presence of an LV and an infundibulum. The different anatomic types of single LV vary according to the type of ventricular loop present and the types of AV and ventriculoarterial alignment. The communication between the LV and the infundibulum is termed the bulboventricular foramen. 2. Single RV (Fig. 31-14): In hearts with a single RV, the ventricular mass consists of the RV sinus and the infundibulum, together forming a common chamber. The septal band is present, indicating the location of a ventricular septum, but there is no grossly recognizable left ventricular sinus on the other side of the septum. Several anatomic types of single RV are recognized. (1) In double-inlet RV, both AV valves open into the RV, the tricuspid valve attaches in the inlet (sinus) portion of the RV, and the mitral valve attaches in the outflow or infundibulum. Both valves exhibit attachments to the septal band. (2) In common-inlet single RV, a single AV valve connects both atria with the RV. The morphology is often that of a tricuspid valve, but the presence of an ostium primum defect and the alignment of both atria with the single RV indicate that this is a common AV valve mimicking a tricuspid valve. This type of single RV usually occurs in association with visceral heterotaxy and asplenia. Single-ventricle physiology is characterized by complete mixing of systemic and pulmonary venous return flow. The proportion of the ventricular output distributed to the pulmonary or systemic vascular bed is determined by the relative resistance to flow in the two circuits.

Inf RA

TV

*

MV TV

MV

LA

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B

Figure 31-14 Electrocardiographically gated steady-state free precession in a patient with a double-inlet single right ventricle. Note the absence of a ventricular septum.

Cardiovascular Magnetic Resonance 431

31 COMPLEX CONGENITAL HEART DISEASE: INFANT AND PEDIATRIC PATIENTS

Figure 31-13 Electrocardiographically gated steady-state free precession cine cardiovascular magnetic resonance in a patient with a double-inlet single left ventricle. A, Transverse plane. B, Plane. Note that both the mitral (MV) and tricuspid valves (TV) enter the left ventricle, which is characterized by a smooth septal surface (arrow). The bulboventricular foramen (*) provides an exit from the left ventricle to the infundibulum (Inf). LA, left atrium; RA, right atrium.

RIGHT VENTRICULAR AND CONGENITAL HEART DISEASE

The goals of current surgical therapy in patients with single ventricle are to separate the systemic and pulmonary circulations and eliminate volume overload on the ventricle. These goals are often achieved with staged palliative procedures, leading to one of the modifications of the Fontan procedure (Fig. 31-15). The principal aim of the procedure is to divert the systemic venous return from the inferior vena cava and superior vena cava to the PAs, separating the poorly oxygenated systemic venous return from the oxygenated pulmonary venous return. Since it was first described in 1971,84 the Fontan procedure has undergone multiple modifications, including direct anastomosis of the right atrial appendage to the main PA (called atriopulmonary anastomosis),85 RA-to-RV conduit, lateral tunnel between the inferior vena cava and the undersurface of the ipsilateral branch of the PA,86 fenestration of the lateral tunnel baffle,87 and an extracardiac conduit between the inferior vena cava and the ipsilateral branch of the PA.88 With refinement of the criteria for patient selection, management strategies, and surgical techniques, the shortand medium-term results of the Fontan procedure have improved gradually.89 Nevertheless, this growing patient

population continues to be at risk for complications, such as systemic ventricular dysfunction, thromboembolism, dilation of the systemic venous atrium, obstruction of the Fontan pathways, PA stenosis, compression of the pulmonary veins, mitral/tricuspid valve regurgitation, protein-losing enteropathy, and arrhythmias.89–95 Therefore, prompt detection of these complications is an important element of managing these patients.

Cardiovascular Magnetic Resonance Evaluations Cardiovascular magnetic resonance is particularly well suited for the detailed investigation of the divergent and often complex anatomy of single-ventricle heart disease (see Figs. 31-13 and 31-14). Although echocardiography is typically adequate for initial diagnosis in the neonate with a single ventricle, CMR is assuming an increasingly important role in the evaluation of patients before and after later stages of palliation.

Atriopulmonary Fontan

Classic Fontan

Fenestrated Fontan

Extracardiac Fontan

Figure 31-15 Four variations of the Fontan procedure. 432 Cardiovascular Magnetic Resonance

Staged palliation of single-ventricle disease often includes superior cavopulmonary anastomosis (bidirectional Glenn procedure), typically performed at 4 to 6 months of age. Traditionally, this procedure is preceded by diagnostic cardiac catheterization for anatomic and physiologic information. CMR offers an alternative for cardiac catheterization in selected cases (Fig. 31-16).96,96a Brown and associates, in a prospective randomized clinical trial, showed that CMR is a safe, effective, and less costly alternative to routine catheterization in the evaluation of selected patients before bidirectional Glenn procedure.97 The goals of CMR before bidirectional Glenn procedure include: (1) performing quantitative assessment of LV and RV mass, volume, and function; (2) imaging the anatomy of the pulmonary veins, PAs, and aortic arch; (3) imaging

Figure 31-16 Three-dimensional contrast-enhanced magnetic resonance angiography before bidirectional Glenn shunt in infants with palliated hypoplastic left heart syndrome. A, Subvolume maximum intensity projection in an oblique transverse plane showing a conduit (*) from the right ventricle (RV) to the pulmonary arteries. B, Three-dimensional volume reconstruction in the same patient showing the conduit (arrow) as well as the anastomosis between the main pulmonary artery (MPA) and the reconstructed neoaorta. C, Oblique coronal subvolume maximum intensity projection image in a patient with a modified right Blalock-Taussig shunt. The shunt (arrow) extends from the right innominate artery (RIA) to the right pulmonary artery. Note the hypoplasia of the right pulmonary artery. LIV, left innominate vein. D, Oblique sagittal subvolume maximum intensity projection image of a reconstructed neoaorta (Neo-Ao). Az, azygos vein; DAo, descending aorta; RPA, right pulmonary artery; SVC, superior vena cava.

RV Neo-Ao

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Cardiovascular Magnetic Resonance 433

31 COMPLEX CONGENITAL HEART DISEASE: INFANT AND PEDIATRIC PATIENTS

the anatomy of any previous surgical intervention (e.g., aortic arch reconstruction, atrial septectomy); (4) evaluating for the presence of aortopulmonary or venovenous collateral vessels; and (5) quantifying mitral/tricuspid or semilunar valve regurgitation. These objectives can be achieved by the following protocol: Sagittal and coronal scout images ECG gated cine SSFP sequence in the axial plane for evaluation of the anatomy, with particular attention to the atrial septum, pulmonary veins, and PAs ECG gated cine SSFP sequence of the systemic ventricle in its long axis, followed by cine imaging from the level of the AV valve to the apex ECG gated cine SSFP sequence in an oblique sagittal plane to image the aortic arch ECG gated FSE with double inversion recovery to image the branch PAs and aortic arch (optional) ECG gated velocity encoded cine CMR sequences perpendicular to the mitral/tricuspid valves, ascending aorta, descending aorta, and (optional) branch PAs 3D CE-MRA LGE for evaluation of endocardial fibroelastosis or myocardial fibrosis (Fig. 31-17)98

Cardiovascular Magnetic Resonance Evaluation During Staged Palliation of Single Ventricle

RIGHT VENTRICULAR AND CONGENITAL HEART DISEASE

LA LV

Figure 31-17 Myocardial late gadolinium enhancement in an infant with severe aortic valve stenosis who underwent prenatal balloon aortic valvuloplasty. The circumferential subendocardial enhancement in the left ventricle (LV) is consistent with endocardial fibroelastosis. LA, left atrium.

Post-Fontan Cardiovascular Magnetic Resonance Several reports have used CMR as an investigational tool to study blood flow dynamics within the Fontan pathways and delineate the distribution of inferior and superior caval flow to each lung.99–102 Myocardial tagging has proved an important investigational tool in the evaluation of myocardial mechanics in patients with functional single-ventricle and Fontan circulation, demonstrating

asynchrony and impaired regional wall motion.103 The clinical utility of CMR in patients with the Fontan circulation increases as these patients grow and their acoustic windows become more restricted (Fig. 31-18). The goals of CMR examination in patients with the Fontan circulation include: (1) assessment of the pathways from the systemic veins to the PAs for obstruction and thrombus; (2) detection of Fontan baffle fenestration or leaks; (3) evaluation of the pulmonary veins for compression; (4) systemic ventricular volume, mass, and function; (5) imaging of the systemic ventricular outflow tract for obstruction; (6) quantitative assessment of the atrioventricular and semilunar valves for regurgitation; (7) imaging of the aorta for obstruction or an aneurysm; and (8) detection of aortopulmonary, systemic venous, or systemic-to-pulmonary venous collateral vessels. A representative imaging protocol begins with localizing images and continues with the following sequences: ECG gated cine SSFP sequence in the axial plane from the level of the confluence of the inferior vena cava and hepatic veins to the level of the cranial end of the superior vena cava Additional cine SSFP sequences may be obtained in the coronal or oblique planes to image areas of the Fontan pathways or other areas that are suspected to be abnormal based on the previous sequence Free breathing (navigator-gated), ECG triggered, isotropic, 3D SSFP sequence (optional)61 ECG gated cine SSFP sequence of the systemic ventricle in its long axis, followed by cine imaging from the level of the mitral/tricuspid valves to the apex VEC CMR sequences perpendicular to the mitral/ tricuspid valves and systemic arterial root for flow quantification. Based on clinical relevance, additional flow measurements are obtained in selected areas of the Fontan pathways and systemic veins to evaluate the pulmonary-to-systemic flow ratio and differential pulmonary blood flow

Figure 31-18 Evaluation of Fontan pathways with threedimensional contrast-enhanced magnetic resonance angiography. A, Subvolume maximum intensity projection image in the coronal plane showing a lateral tunnel (LT) extending from the inferior vena cava (IVC) to the pulmonary arteries. B, Subvolume maximum intensity projection image in the transverse plane showing an atriopulmonary anastomosis. AAo, ascending aorta; Ao, aorta; DAo, descending aorta; LPA, left pulmonary artery; RA, right atrium; RPA, right pulmonary artery; SVC, superior vena cava.

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434 Cardiovascular Magnetic Resonance

B

ACKNOWLEDGMENT The authors thank the physicians, technologists, nurses, and support personnel in the Cardiovascular MRI Program at Children’s Hospital Boston for their help. Parts of the text and figures of this chapter have been included in other chapters recently written by the authors on MRI evaluation of congenital heart disease.

References 1. Hoffman JI, Kaplan S. The incidence of congenital heart disease. J Am Coll Cardiol. 2002;39:1890–1900. 2. Van Praagh R, Van Praagh S, Nebesar RA, Muster AJ, Sinha SN, Paul MH. Tetralogy of Fallot: underdevelopment of the pulmonary infundibulum and its sequelae. Am J Cardiol. 1970;26:25–33. 3. Van Praagh R. Etienne-Louis Arthur Fallot and his tetralogy: a new translation of Fallot’s summary and a modern reassessment of this anomaly. Eur J Cardiothorac Surg. 1989;3:381–386. 4. Van Praagh R, Geva T, Kreutzer J. Ventricular septal defects: how shall we describe, name and classify them? J Am Coll Cardiol. 1989;14:1298–1299. 5. Need LR, Powell AJ, del Nido P, Geva T. Coronary echocardiography in tetralogy of fallot: diagnostic accuracy, resource utilization and surgical implications over 13 years. J Am Coll Cardiol. 2000;36: 1371–1377. 6. Marino B, Digilio MC, Grazioli S, et al. Associated cardiac anomalies in isolated and syndromic patients with tetralogy of Fallot. Am J Cardiol. 1996;77:505–508. 7. Geva T, Ayres NA, Pac FA, Pignatelli R. Quantitative morphometric analysis of progressive infundibular obstruction in tetralogy of Fallot: a prospective longitudinal echocardiographic study. Circulation. 1995;92:886–892. 8. Kaulitz R, Jux C, Bertram H, Paul T, Ziemer G, Hausdorf G. Primary repair of tetralogy of fallot in infancy: the effect on growth of the pulmonary arteries and the risk for late reinterventions. Cardiol Young. 2001;11:391–398. 9. Saul JP, Alexander ME. Preventing sudden death after repair of tetralogy of Fallot: complex therapy for complex patients. J Cardiovasc Electrophysiol. 1999;10:1271–1287. 10. Kugler JD. Predicting sudden death in patients who have undergone tetralogy of fallot repair: is it really as simple as measuring ECG intervals? J Cardiovasc Electrophysiol. 1998;9:103–106. 11. Bricker JT. Sudden death and tetralogy of Fallot: risks, markers, and causes. Circulation. 1995;92:158–159. 12. Gatzoulis MA, Till JA, Somerville J, Redington AN. Mechanoelectrical interaction in tetralogy of Fallot: QRS prolongation relates to right ventricular size and predicts malignant ventricular arrhythmias and sudden death. Circulation. 1995;92:231–237. 13. Berul CI, Hill SL, Geggel RL, et al. Electrocardiographic markers of late sudden death risk in postoperative tetralogy of Fallot children. J Cardiovasc Electrophysiol. 1997;8:1349–1356. 14. Hokanson JS, Moller JH. Significance of early transient complete heart block as a predictor of sudden death late after operative correction of tetralogy of Fallot. Am J Cardiol. 2001;87:1271–1277. 15. Hamada H, Terai M, Jibiki T, Nakamura T, Gatzoulis MA, Niwa K. Influence of early repair of tetralogy of fallot without an outflow patch on late arrhythmias and sudden death: a 27-year follow-up study following a uniform surgical approach. Cardiol Young. 2002;12:345–351. 16. Mackie AS, Gauvreau K, Perry SB, del Nido PJ, Geva T. Echocardiographic predictors of aortopulmonary collaterals in infants with tetralogy of fallot and pulmonary atresia. J Am Coll Cardiol. 2003; 41:852–857. 17. Vick 3rd GW, Wendt 3rd RE, Rokey R. Comparison of gradient echo with spin echo magnetic resonance imaging and echocardiography in the evaluation of major aortopulmonary collateral arteries. Am Heart J. 1994;127:1341–1347. 18. Powell AJ, Chung T, Landzberg MJ, Geva T. Accuracy of MRI evaluation of pulmonary blood supply in patients with complex pulmonary stenosis or atresia. Int J Card Imaging. 2000;16:169–174.

19. Holmqvist C, Hochbergs P, Bjorkhem G, Brockstedt S, Laurin S. Preoperative evaluation with MR in tetralogy of fallot and pulmonary atresia with ventricular septal defect. Acta Radiol. 2001;42:63–69. 20. Beekman RP, Beek FJ, Meijboom EJ. Usefulness of MRI for the preoperative evaluation of the pulmonary arteries in tetralogy of Fallot. Magn Reson Imaging. 1997;15:1005–1015. 21. Geva T, Greil GF, Marshall AC, Landzberg M, Powell AJ. Gadoliniumenhanced 3-dimensional magnetic resonance angiography of pulmonary blood supply in patients with complex pulmonary stenosis or atresia: comparison with x-ray angiography. Circulation. 2002;106:473–478. 22. Prasad SK, Soukias N, Hornung T, et al. Role of magnetic resonance angiography in the diagnosis of major aortopulmonary collateral arteries and partial anomalous pulmonary venous drainage. Circulation. 2004;109:207–214. 23. Geva T, Sahn DJ, Powell AJ. Magnetic resonance imaging of congenital heart disease in adults. Progress in Pediatric Cardiology. 2003;17: 21–39. 24. Helbing WA, de Roos A. Clinical applications of cardiac magnetic resonance imaging after repair of tetralogy of Fallot. Pediatr Cardiol. 2000;21:70–79. 25. Geva T, Sandweiss BM, Gauvreau K, Lock JE, Powell AJ. Factors associated with impaired clinical status in long-term survivors of tetralogy of Fallot repair evaluated by magnetic resonance imaging. J Am Coll Cardiol. 2004;43:1068–1074. 26. Knauth AL, Gauvreau K, Powell AJ, et al. Ventricular size and function assessed by cardiac MRI predict major adverse clinical outcomes late after tetralogy of Fallot repair. Heart. 2008;94:211–216. 27. Roest AA, Helbing WA, Kunz P, et al. Exercise MR imaging in the assessment of pulmonary regurgitation and biventricular function in patients after tetralogy of fallot repair. Radiology. 2002;223: 204–211. 28. Rebergen SA, Chin JG, Ottenkamp J, van der Wall EE, de Roos A. Pulmonary regurgitation in the late postoperative follow-up of tetralogy of Fallot: volumetric quantitation by nuclear magnetic resonance velocity mapping. Circulation. 1993;88:2257–2266. 29. Niezen RA, Helbing WA, van Der Wall EE, van Der Geest RJ, Vliegen HW, de Roos A. Left ventricular function in adults with mild pulmonary insufficiency late after Fallot repair. Heart. 1999;82:697–703. 30. Niezen RA, Helbing WA, van der Wall EE, van der Geest RJ, Rebergen SA, de Roos A. Biventricular systolic function and mass studied with MR imaging in children with pulmonary regurgitation after repair for tetralogy of Fallot. Radiology. 1996;201:135–140. 31. Davlouros PA, Kilner PJ, Hornung TS, et al. Right ventricular function in adults with repaired tetralogy of Fallot assessed with cardiovascular magnetic resonance imaging: detrimental role of right ventricular outflow aneurysms or akinesia and adverse right-to-left ventricular interaction. J Am Coll Cardiol. 2002;40:2044–2052. 32. Babu-Narayan SV, Kilner PJ, Li W, et al. Ventricular fibrosis suggested by cardiovascular magnetic resonance in adults with repaired tetralogy of Fallot and its relationship to adverse markers of clinical outcome. Circulation. 2006;113:405–413. 33. Van Praagh R. The importance of segmental situs in the diagnosis of congenital heart disease. Semin Roentgenol. 1985;20:254–271. 34. Blume ED, Altmann K, Mayer JE, Colan SD, Gauvreau K, Geva T. Evolution of risk factors influencing early mortality of the arterial switch operation. J Am Coll Cardiol. 1999;33:1702–1709. 35. Levinsky L, Srinivasan V, Alvarez-Diaz F, Subramanian S. Reconstruction of the new atrial septum in the Senning operation: new technique. J Thorac Cardiovasc Surg. 1981;81:131–134.

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3D CE-MRA of the Fontan pathways, systemic and pulmonary veins, and aorta Whenever possible, SSFP cine sequences are acquired with breath holding. FSE with double inversion recovery sequence may be used to minimize artifacts from metallic implants, which are common in this patient population.104 Additional sequences, such as myocardial perfusion or LGE viability, and myocardial tagging can be performed when the information provided by these sequences is needed.

RIGHT VENTRICULAR AND CONGENITAL HEART DISEASE

36. Myridakis DJ, Ehlers KH, Engle MA. Late follow-up after venous switch operation (Mustard procedure) for simple and complex transposition of the great arteries. Am J Cardiol. 1994;74:1030–1036. 37. Redington AN, Rigby ML, Oldershaw P, Gibson DG, Shinebourne EA. Right ventricular function 10 years after the Mustard operation for transposition of the great arteries: analysis of size, shape, and wall motion. Br Heart J. 1989;62:455–461. 38. Deanfield J, Camm J, Macartney F, et al. Arrhythmia and late mortality after Mustard and Senning operation for transposition of the great arteries: an eight-year prospective study. J Thorac Cardiovasc Surg. 1988;96:569–576. 39. Van Praagh R, Jung WK. The arterial switch operation in transposition of the great arteries: anatomic indications and contraindications. Thorac Cardiovasc Surg. 1991;39(suppl 2):138–150. 40. Wernovsky G, Jonas RA, Colan SD, et al. Results of the arterial switch operation in patients with transposition of the great arteries and abnormalities of the mitral valve or left ventricular outflow tract. J Am Coll Cardiol. 1990;16:1446–1454. 41. Rehnstrom P, Gilljam T, Sudow G, Berggren H. Excellent survival and low complication rate in medium-term follow-up after arterial switch operation for complete transposition. Scand Cardiovasc J. 2003; 37:104–106. 42. Kramer HH, Scheewe J, Fischer G, et al. Long term follow-up of left ventricular performance and size of the great arteries before and after one- and two-stage arterial switch operation of simple transposition. Eur J Cardiothorac Surg. 2003;24:898–905. 43. Prifti E, Crucean A, Bonacchi M, et al. Early and long term outcome of the arterial switch operation for transposition of the great arteries: predictors and functional evaluation. Eur J Cardiothorac Surg. 2002;22:864–873. 44. Dunbar-Masterson C, Wypij D, Bellinger DC, et al. General health status of children with D-transposition of the great arteries after the arterial switch operation. Circulation. 2001;104:I138–I142. 45. Losay J, Touchot A, Serraf A, et al. Late outcome after arterial switch operation for transposition of the great arteries. Circulation. 2001;104:I121–I126. 46. Mahle WT, McBride MG, Paridon SM. Exercise performance after the arterial switch operation for D-transposition of the great arteries. Am J Cardiol. 2001;87:753–758. 47. Van Praagh R, Papagiannis J, Grunenfelder J, Bartram U, Martanovic P. Pathologic anatomy of corrected transposition of the great arteries: medical and surgical implications. Am Heart J. 1998;135:772–785. 48. Beauchesne LM, Warnes CA, Connolly HM, Ammash NM, Tajik AJ, Danielson GK. Outcome of the unoperated adult who presents with congenitally corrected transposition of the great arteries. J Am Coll Cardiol. 2002;40:285–290. 49. Colli AM, de Leval M, Somerville J. Anatomically corrected malposition of the great arteries: diagnostic difficulties and surgical repair of associated lesions. Am J Cardiol. 1985;55:1367–1372. 50. Van Praagh R. What is congenitally corrected transposition? N Engl J Med. 1970;282:1097–1098. 51. Chung KJ, Simpson IA, Glass RF, Sahn DJ, Hesselink JR. Cine magnetic resonance imaging after surgical repair in patients with transposition of the great arteries. Circulation. 1988;77:104–109. 52. Lorenz CH, Walker ES, Graham Jr TP, Powers TA. Right ventricular performance and mass by use of cine MRI late after atrial repair of transposition of the great arteries. Circulation. 1995;92:II233–II239. 53. Hardy CE, Helton GJ, Kondo C, Higgins SS, Young NJ, Higgins CB. Usefulness of magnetic resonance imaging for evaluating great-vessel anatomy after arterial switch operation for D-transposition of the great arteries. Am Heart J. 1994;128:326–332. 54. Beek FJ, Beekman RP, Dillon EH, et al. MRI of the pulmonary artery after arterial switch operation for transposition of the great arteries. Pediatr Radiol. 1993;23:335–340. 55. Theissen P, Kaemmerer H, Sechtem U, et al. Magnetic resonance imaging of cardiac function and morphology in patients with transposition of the great arteries following Mustard procedure. Thorac Cardiovasc Surg. 1991;39(suppl 3):221–224. 56. Rees S, Somerville J, Warnes C, et al. Comparison of magnetic resonance imaging with echocardiography and radionuclide angiography in assessing cardiac function and anatomy following Mustard’s operation for transposition of the great arteries. Am J Cardiol. 1988; 61:1316–1322. 57. Tulevski II, Lee PL, Groenink M, et al. Dobutamine-induced increase of right ventricular contractility without increased stroke volume in

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adolescent patients with transposition of the great arteries: evaluation with magnetic resonance imaging. Int J Card Imaging. 2000;16: 471–478. Tulevski II, van der Wall EE, Groenink M, et al. Usefulness of magnetic resonance imaging dobutamine stress in asymptomatic and minimally symptomatic patients with decreased cardiac reserve from congenital heart disease (complete and corrected transposition of the great arteries and subpulmonic obstruction). Am J Cardiol. 2002;89: 1077–1081. Roest AA, Lamb HJ, van der Wall EE, et al. Cardiovascular response to physical exercise in adult patients after atrial correction for transposition of the great arteries assessed with magnetic resonance imaging. Heart. 2004;90:678–684. Winter MM, Bernink FJP, Groenink M, et al. Evaluating the systemic right ventricle by CMR: the importance of consistent and reproducible delineation of the cavity. J Cardiovasc Magn Reson. 2008;10:40. Babu-Narayan SV, Goktekin O, Moon JC, et al. Late gadolinium enhancement cardiovascular magnetic resonance of the systemic right ventricle in adults with previous atrial redirection surgery for transposition of the great arteries. Circulation. 2005;111:2091–2098. Sorensen TS, Korperich H, Greil GF, et al. Operator-independent isotropic three-dimensional magnetic resonance imaging for morphology in congenital heart disease: a validation study. Circulation. 2004;110:163–169. Videlefsky N, Parks WJ, Oshinski J, et al. Magnetic resonance phaseshift velocity mapping in pediatric patients with pulmonary venous obstruction. J Am Coll Cardiol. 2001;38:262–267. Pasquali SK, Hasselblad V, Li JS, Kong DF, Sanders SP. Coronary artery pattern and outcome of arterial switch operation for transposition of the great arteries: a meta-analysis. Circulation. 2002;106: 2575–2580. Taylor AM, Dymarkowski S, Hamaekers P, et al. MR coronary angiography and late-enhancement myocardial MR in children who underwent arterial switch surgery for transposition of great arteries. Radiology. 2005;234:542–547. Belli E, Serraf A, Lacour-Gayet F, et al. Surgical treatment of subaortic stenosis after biventricular repair of double-outlet right ventricle. J Thorac Cardiovasc Surg. 1996;112:1570–1578; discussion 1578–80. Yoo SJ, Kim YM, Choe YH. Magnetic resonance imaging of complex congenital heart disease. Int J Card Imaging. 1999;15:151–160. Beekman RP, Roest AA, Helbing WA, et al. Spin echo MRI in the evaluation of hearts with a double outlet right ventricle: usefulness and limitations. Magn Reson Imaging. 2000;18:245–253. Beekman RP, Beek FJ, Meijboom EJ, Wenink AC. MRI appearance of a double inlet and double outlet right ventricle with supero-inferior ventricular relationship. Magn Reson Imaging. 1996;14:1107–1112. Igarashi H, Kuramatsu T, Shiraishi H, Yanagisawa M. Criss-cross heart evaluated by colour Doppler echocardiography and magnetic resonance imaging. Eur J Pediatr. 1990;149:523–525. Niezen RA, Beekman RP, Helbing WA, van der Wall EE, de Roos A. Double outlet right ventricle assessed with magnetic resonance imaging. Int J Card Imaging. 1999;15:323–329. Van Praagh R, Van Praagh S. The anatomy of common aorticopulmonary trunk (truncus arteriosus communis) and its embryologic implications: a study of 57 necropsy cases. Am J Cardiol. 1965; 16:406–425. Collett RW, Edwards JE. Persistent truncus arteriosus: a classification according to anatomic types. Surg Clin North Am. 1949;29: 1245–1270. Litovsky SH, Ostfeld I, Bjornstad PG, Van Praagh R, Geva T. Truncus arteriosus with anomalous pulmonary venous connection. Am J Cardiol. 1999;83:801–804, A10. Goldmuntz E, Clark BJ, Mitchell LE, et al. Frequency of 22q11 deletions in patients with conotruncal defects. J Am Coll Cardiol. 1998;32:492–498. Tworetzky W, McElhinney DB, Brook MM, Reddy VM, Hanley FL, Silverman NH. Echocardiographic diagnosis alone for the complete repair of major congenital heart defects. J Am Coll Cardiol. 1999; 33:228–233. Fyler DC, Buckley LP, Hellenbrand WE, Cohn HE. Report of the New England Regional Infant Cardiac Program. Pediatrics. 1980; 65:377–461. Celoria GC, Patton RB. Congenital absence of the aortic arch. Am Heart J. 1959;58:407–413. Geva T, Gajarski RJ. Echocardiographic diagnosis of type B interruption of a right aortic arch. Am Heart J. 1995;129:1042–1045.

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complex congenital heart disease. Ann Thorac Surg. 2003;76: 542–553; discussion 553–4. Kreutzer J, Keane JF, Lock JE, et al. Conversion of modified Fontan procedure to lateral atrial tunnel cavopulmonary anastomosis. J Thorac Cardiovasc Surg. 1996;111:1169–1176. Muthurangu V, Taylor AM, Hegde SR, et al. Cardiac magnetic resonance imaging after stage I Norwood operation for hypoplastic left heart syndrome. Circulation. 2005;112:3256–3263. Lim DS, Peeler BB, Matherne GP, Kramer CM. Cardiovascular magnetic resonance of pulmonary artery growth and ventricular function after Norwood procedure with Sano modification. J Cardiovasc Magn Resonan. 2008;10:34. Brown DW, Gauvreau K, Powell AJ, et al. Cardiac magnetic resonance versus routine cardiac catheterization before bidirectional glenn anastomosis in infants with functional single ventricle: a prospective randomized trial. Circulation. 2007;116:2718–2725. Tworetzky W, del Nido PJ, Powell AJ, Marshall AC, Lock JE, Geva T. Usefulness of magnetic resonance imaging of left ventricular endocardial fibroelastosis in infants after fetal intervention for aortic valve stenosis. Am J Cardiol. 2005;96:1568–1570. Hjortdal VE, Emmertsen K, Stenbog E, et al. Effects of exercise and respiration on blood flow in total cavopulmonary connection: a real-time magnetic resonance flow study. Circulation. 2003;108: 1227–1231. Be’eri E, Maier SE, Landzberg MJ, Chung T, Geva T. In vivo evaluation of Fontan pathway flow dynamics by multidimensional phase-velocity magnetic resonance imaging. Circulation. 1998;98: 2873–2882. Fogel MA, Weinberg PM, Rychik J, et al. Caval contribution to flow in the branch pulmonary arteries of Fontan patients with a novel application of magnetic resonance presaturation pulse. Circulation. 1999;99:1215–1221. Fratz S, Hess J, Schwaiger M, Martinoff S, Stern HC. More accurate quantification of pulmonary blood flow by magnetic resonance imaging than by lung perfusion scintigraphy in patients with fontan circulation. Circulation. 2002;106:1510–1513. Fogel MA, Gupta KB, Weinberg PM, Hoffman EA. Regional wall motion and strain analysis across stages of Fontan reconstruction by magnetic resonance tagging. Am J Physiol. 1995;269: H1132–H1152. Garg R, Powell AJ, Sena L, Marshall AC, Geva T. Effects of metallic implants on magnetic resonance imaging evaluation of Fontan palliation. Am J Cardiol. 2005;95:688–691.

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79. Kutsche LM, Van Mierop LH. Cervical origin of the right subclavian artery in aortic arch interruption: pathogenesis and significance. Am J Cardiol. 1984;53:892–895. 80. Kaulitz R, Jonas RA, van der Velde ME. Echocardiographic assessment of interrupted aortic arch. Cardiol Young. 1999;9:562–571. 81. Varghese A, Gatzoulis M, Mohiaddin RH. Images in cardiovascular medicine: magnetic resonance angiography of a congenitally interrupted aortic arch. Circulation. 2002;106:E9–E10. 82. Tsai-Goodman B, Geva T, Odegard KC, Sena LM, Powell AJ. Clinical role, accuracy, and technical aspects of cardiovascular magnetic resonance imaging in infants. Am J Cardiol. 2004;94:69–74. 83. Nielsen J, Powell AJ, Gauvreau K, et al. Magnetic resonance imaging predictors of coarctation severity. Circulation. 2005;111:622–628. 84. Fontan F, Baudet E. Surgical repair of tricuspid atresia. Thorax. 1971;26:240–248. 85. Kreutzer GO, Vargas FJ, Schlichter AJ, et al. Atriopulmonary anastomosis. J Thorac Cardiovasc Surg. 1982;83:427–436. 86. Jonas RA, Castaneda AR. Modified Fontan procedure: atrial baffle and systemic venous to pulmonary artery anastomotic techniques. J Card Surg. 1988;3:91–96. 87. Bridges ND, Mayer Jr JE, Lock JE, et al. Effect of baffle fenestration on outcome of the modified Fontan operation. Circulation. 1992;86:1762–1769. 88. Tireli E. Extracardiac Fontan operation without cardiopulmonary bypass: how to perform the anastomosis between inferior vena cava and conduit. Cardiovasc Surg. 2003;11:225–227. 89. Gentles TL, Mayer Jr JE, Gauvreau K, et al. Fontan operation in five hundred consecutive patients: factors influencing early and late outcome. J Thorac Cardiovasc Surg. 1997;114:376–391. 90. Wilson WR, Greer GE, Tobias JD. Cerebral venous thrombosis after the Fontan procedure. J Thorac Cardiovasc Surg. 1998;116:661–663. 91. Day RW, Boyer RS, Tait VF, Ruttenberg HD. Factors associated with stroke following the Fontan procedure. Pediatr Cardiol. 1995; 16:270–275. 92. Jacobs ML. Complications associated with heterotaxy syndrome in Fontan patients. Semin Thorac Cardiovasc Surg Pediatr Card Surg Annu. 2002;5:25–35. 93. Lam J, Neirotti R, Becker AE, Planche C. Thrombosis after the Fontan procedure: transesophageal echocardiography may replace angiocardiography. J Thorac Cardiovasc Surg. 1994;108:194–195. 94. Deal BJ, Mavroudis C, Backer CL. Beyond Fontan conversion: surgical therapy of arrhythmias including patients with associated

Pulmonary Vein Imaging Thomas H. Hauser and Dana C. Peters

With the advent of radiofrequency (RF) ablation for the treatment of atrial fibrillation, there has been increased interest in the accurate determination of pulmonary vein anatomy to help plan the procedure and to monitor for postablation pulmonary vein stenosis. Contrast-enhanced magnetic resonance angiography (CE-MRA) readily demonstrates the pulmonary veins and is the method of choice for these required serial imaging studies. Cardiovascular magnetic resonance (CMR) offers a preferred alternative to computed tomography angiography (CTA) because of favorable contrast agent toxicity and absence of radiation exposure. The latter aspect is especially important for this group, which often has repeated studies. In this chapter, we review the normal and anomalous pulmonary venous anatomy and the utility of imaging before and after pulmonary vein isolation.

IMAGING METHOD The pulmonary veins can be identified by using standard anatomic and functional CMR imaging sequences. While these methods are usually sufficient to identify the anatomic relationship of the pulmonary veins to the heart and the other major vascular structures, the pulmonary veins are usually imaged by using CE-MRA. A three-dimensional (3D) spoiled gradient echo sequence is acquired during the first pass of gadolinium contrast.1 Clinical protocols vary but have common elements.2–10 The technique uses short repetition times (TR ¼ 3 to 6 msec), a high flip angle (30 to 60 ), and fractional echoes to provide T1 weighting and minimal flow artifacts. The spatial resolution varies from 1 to 2 1 to 2 mm in-plane with 2- to 4-mm slices, before interpolation. A single 3D volume requires a 10- to 20-second breath hold to suppress ventilatory motion, but scan time can be shortened by using smaller fields of view, shorter repetition times, partial Fourier, lower spatial resolution, or parallel imaging. Electrocardiographic triggering is not employed, although it is recognized that the position and shape of the pulmonary veins change throughout the cardiac cycle.7,11 Images obtained with this method reflect maximal pulmonary vein anatomy.12 Axial slabs are usually acquired, using either sequential or centric k-space filling. For the pulmonary vasculature, the arterialvenous transit time is very short (4 to 7 seconds);13 therefore, artery-vein separation is highly challenging and generally not targeted. Contrast is injected with a dose of 0.1 to 0.2 mmol/kg at a rate of 1 to 2 mL/sec, followed by a 10- to 20-mL saline flush. A precontrast mask can be acquired, although mask subtraction is not essential for pulmonary venography, since the background signal in the lungs is

very low. Often, a second time frame is acquired immediately after the first pass image to ensure acquisition during peak contrast. Timing of the acquisition to the first pass of contrast through the pulmonary veins is critical and is achieved either by using a bolus timing scan14 or with fluoroscopic triggering.15 For either method, imaging is timed to begin with the appearance of contrast in the left atrium (LA).

IMAGE DISPLAY Once the 3D CE-MRA dataset has been obtained, the images can be transferred to a workstation for further manipulation and analysis (Fig. 32-1). The simplest and often most informative method to display the images is to dynamically view two-dimensional (2D) slices within the 3D dataset in the axial, coronal, and sagittal planes. The axial images usually provide a good overview of the pulmonary veins and their relationship to the LA, but the coronal and sagittal images are frequently required to determine specific anatomic findings, such as a left common or anomalous pulmonary vein. Although 2D slices are very useful for viewing the individual pulmonary veins, it is difficult to produce a single summary image of the anatomy. Maximal intensity projection (MIP) and 3D reconstructions displayed as shaded surface or volume-rendered images take full advantage of the 3D dataset and provide very good summary images. These are most useful when the displayed volume is limited to the LA and pulmonary veins. Because the aorta is directly posterior to the left-sided pulmonary vein, it frequently obscures them from view in the MIP images. Three-dimensional reconstructed images are frequently preferred because the aorta can be excluded from the displayed volume. Moving images can also be generated to better display the anatomy. By convention, the LA and pulmonary veins are viewed in the posterior-anterior orientation. Direct anatomic measurements should not be obtained from these postprocessed images but rather should be obtained from the 2D slices.

PULMONARY VEIN EMBRYOLOGY A clear understanding of pulmonary vein embryology is important for understanding both normal pulmonary vein anatomy, nonpathologic variations from the normal anatomy, and congenital anomalies. The pulmonary veins and Cardiovascular Magnetic Resonance 441

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CHAPTER 32

VASCULATURE AND PERICARDIUM

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Figure 32-1 Normal anatomy and quantification of pulmonary vein size. These images show the normal complement of four pulmonary veins along with the left atrium (LA) and descending aorta (Ao). The right inferior (RI), right superior (RS) and left inferior (LI) pulmonary veins are shown in the axial plane (A and B). The left superior (LS) and right superior (RS) pulmonary veins are shown in the coronal plane from the posterior-anterior orientation (C). The left-sided (D) and right-sided (E) pulmonary veins are shown in the sagittal plane (anterior to the left). All of the pulmonary veins are shown in the axial maximal intensity projection (F) and posterioranterior volume rendered (G) images. The aorta has been removed from the volumerendered image to show all of the pulmonary veins. All of these images were derived from the same three-dimensional magnetic resonance angiography dataset. The gray lines in the maximal intensity projection image (panel F) correspond to the location in the sagittal plane in which the pulmonary veins separate from the LA and from each other (panels D and E). The maximal diameter, perimeter, and cross-sectional area can be easily measured in the sagittal images.

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LA are derived from the primitive common pulmonary vein. The primitive pulmonary venous system initially has no connection with the heart and drains into the cardinal veins and the umbilicovitelline system. At approximately the fourth week of gestation, the pulmonary venous drainage coalesces into a single vessel.16 At the same time, an outgrowth of the primitive LA extends toward the pulmonary venous system to meet this vessel to form the primitive common pulmonary vein and the venous connections to the cardinal veins and the umbilicovitelline system degenerate. The common pulmonary vein then expands to form the smooth-walled body of the LA, while the primitive LA forms the trabeculated left atrial appendage.17 The branches of the primitive common pulmonary vein form the adult pulmonary veins. The development of the 442 Cardiovascular Magnetic Resonance

LA and pulmonary veins is asymmetrical, the two rightsided pulmonary veins developing first while the left-sided pulmonary venous drainage enters the LA through a single trunk that eventually bifurcates to form two veins.18

NORMAL AND VARIANT PULMONARY VENOUS ANATOMY Most commonly, there are four pulmonary veins that enter the LA: right superior, right inferior, left superior, and left inferior (see Fig. 32-1). Each of the veins is directed laterally, the inferior veins being directed posteriorly and

Figure 32-2 Variant pulmonary venous anatomy. These images were obtained from a patient with right middle and left common pulmonary veins. A, The right middle (RM) pulmonary vein is shown in the axial plane. B, The left common (LC) pulmonary vein is shown in the coronal plane from the posterior-anterior orientation. The single left common (C) and all three right pulmonary veins (D) are shown in the sagittal plane (anterior to the left) along with the pulmonary artery (PA) immediately adjacent to the right superior pulmonary vein. All of the pulmonary veins are shown in the axial MIP (E) and posterior-anterior volume rendered (F) images. The aorta has been removed from the volume rendered image. The right middle pulmonary vein is obscured by the right inferior pulmonary vein and is best seen with cranial angulation (G). It is frequently necessary to manipulate the point of view to see all of the pulmonary veins.

LA, while more incorporation results in additional pulmonary veins (Fig. 32-3). These variations in pulmonary venous anatomy have not yet been identified as a cause of pathology.

CONGENITAL PULMONARY VENOUS ANOMALIES Congenital pulmonary venous anomalies account for up to 3% of all congenital heart disease and approximately 2% of all deaths from congenital heart disease in the first year of

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the superior veins being directed anteriorly. The left superior pulmonary vein frequently has a cranial angulation and may appear to arise from the superior portion of the LA. Variant, nonpathologic pulmonary vein anatomy is very common, present in approximately 40% of patients.2,19 Although numerous variations have been described, the most common variations in the usual anatomy are the presence of a single left common pulmonary vein or an additional right middle pulmonary vein (Fig. 32-2).3 These variations occur because of more or less incorporation of the primitive common pulmonary vein into the LA. Less incorporation leads to apparent fusion of pulmonary veins prior to entering the

VASCULATURE AND PERICARDIUM

C B

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Figure 32-3 Incorporation of the primitive common pulmonary vein into the LA. The incorporation of the primitive common pulmonary vein is variable and results in nonpathologic variations in the normal anatomy. This figure shows the results of variable incorporation of the left-sided pulmonary veins. The most common pattern is two left-sided pulmonary veins (plane B). With less incorporation of the common pulmonary vein into the LA, there is only a single left common pulmonary vein (plane C). With more incorporation, there are additional pulmonary veins (plane A). Source: From Ghaye B, Szapiro D, Dacher JN, et al. Percutaneous ablation for atrial fibrillation: the role of cross-sectional imaging. Radiographics 2003;23:S19–33.

life.20 The congenital anomalies that affect pulmonary veins are atresia, stenosis, and anomalous connections, which can be total or partial. These conditions occur when the normal connections of the primitive pulmonary venous system form abnormally or if embryologic connections to the cardinal vein or umbilicovitelline systems persist, and they are frequently associated with other major congenital cardiac anomalies.21 Anomalous pulmonary venous connections are the most common congenital pulmonary venous anomaly.20 In total anomalous pulmonary venous connection, there is no connection of the pulmonary veins to the LA such that all of the pulmonary venous drainage enters the right atrium (RA) directly or via a systemic vein. This anomaly is necessarily associated with an atrial right-to-left shunt. Pulmonary venous hypertension may result from twists in the veins or compression from adjacent vascular structures,22 while small atrial septal defects restrict systemic blood flow.23 Either of these conditions results in cyanosis and congestive heart failure. Although the mortality rate for symptomatic uncorrected infants is 80% at 1 year,20 surgical repair is usually feasible and reduces the mortality rate to less than 25%.24 444 Cardiovascular Magnetic Resonance

In partial anomalous pulmonary venous return, one or more pulmonary veins, but not all, enter the RA or a systemic vein. There is usually an associated atrial septal defect, frequently of the sinus venosus type with the right superior or right middle pulmonary veins draining into the superior vena cava.20 The physiology of this anomaly is similar to that of an atrial septal defect and depends on the magnitude of the leftto-right shunt and the presence of increased pulmonary vascular resistance.21 Patients are often asymptomatic if the shunt is relatively small and the pulmonary vascular resistance is normal, and the diagnosis might not be made until adulthood. The scimitar syndrome, named after the characteristic chest radiograph finding, is a specific form of partial anomalous pulmonary venous connection in which all of the venous drainage from the right lung enters the inferior vena cava (Fig. 32-4). This rare syndrome is also associated with anomalous arterial supply of the right lower lobe from the aorta, dextroposition of the heart, and hypoplasia of the right lung.25 Congenital pulmonary vein atresia is defined as the absence of any connection of the pulmonary veins to either the LA or any other vascular structure. This is a very rare condition that is not compatible with life, although the infant may survive for a short period of time owing to small connections between the pulmonary veins and esophageal or brachial veins.21 Congenital pulmonary vein stenosis can involve a focal segment of one or more pulmonary veins or may more diffusely involve an entire pulmonary vein and is usually associated with other congenital cardiac malformations. Severe stenosis frequently results in cyanosis, congestive heart failure, and death, although surgical repair is possible if only focal stenosis is present.21 Imaging patients with congenital anomalies by using CE-MRA is a valuable method for determining the

Figure 32-4 Scimitar syndrome. This coronal MIP image in the anterior-posterior orientation shows the typical findings of the scimitar syndrome. The single right pulmonary vein (RPV) enters the inferior vena cava (IVC). The right atrium (RA) and descending aorta (AoD) are also shown. Source: From Greil GF, Powell AJ, Gildein HP, Geva T. Gadolinium-enhanced three-dimensional magnetic resonance angiography of pulmonary systemic venous anomalies. J Am Coll Cardiol 2002;39:335–341.

PULMONARY VEINS AND THE PATHOPHYSIOLOGY OF ATRIAL FIBRILLATION Atrial fibrillation is the most common sustained cardiac arrhythmia, affecting more than 2 million people in the United States,27 and is a major cause of morbidity and mortality, accounting for more than 400,000 hospitalizations each year28 and increasing the risk of death by 50%.29 It quintuples the risk of stroke30 and is the attributed cause for 15% of all strokes, totaling more than 100,000 per year.28 The associated hospital costs alone are estimated at over $1 billion.28 Although several antiarrhythmic drugs are available for the treatment of atrial fibrillation, maintenance of sinus rhythm is frequently suboptimal31–33 and all these drugs are associated with significant side effects or adverse events.34 Data suggest that the pulmonary veins play a critical role in the pathophysiology of atrial fibrillation. As was noted above, the pulmonary veins and LA are both derived from the primitive common pulmonary vein17 and have many anatomic similarities. Both are smooth-walled structures that have electrically active myocardium. Approximately 90% of pulmonary veins contain atrial myocardium.35 Although the myocardium in the atrium is uniform, myocardium in the pulmonary veins is frequently discontinuous and fibrotic. Patients with a history of atrial fibrillation uniformly have myocardium in the pulmonary veins, with an increased rate of structural abnormalities. These structural abnormalities result in abnormal electrical activation. Conduction within pulmonary veins is slow and anisotropic. Proarrhythmic reentrant beats and sustained focal activity can be easily induced.36 A landmark study demonstrated that the proarrhythmic electrical activity in pulmonary veins is directly responsible for the generation of atrial fibrillation in many patients.37 Among those with paroxysmal atrial fibrillation, 94% were found to have ectopic foci in the pulmonary veins that were responsible for the induction of atrial fibrillation. RF ablation of these foci resulted in complete suppression of atrial fibrillation in a majority of patients. After the report of these findings, several related procedures were developed for the treatment of atrial fibrillation.37–41 Each of these procedures uses RF ablation to electrically isolate the pulmonary veins from the LA, with or without additional ablation within the LA. Short-term success rates range from 65% to 85% in patients with paroxysmal atrial fibrillation, with a reduction in morbidity and improved quality of life.42

anatomy and after the procedure to screen for pulmonary vein stenosis. The accurate determination of pulmonary vein anatomy is critical for the planning and execution of atrial fibrillation ablation. To achieve success, the operator must place a series of RF lesions that encircle the pulmonary veins and electrically isolate them from the LA.43 This necessarily requires that the pulmonary vein anatomy be determined prior to the procedure. In the initial development of the procedure, the pulmonary veins were identified by using invasive contrast venography.44 Although this can be done successfully, it greatly increases the procedure time and provides only projection images of the pulmonary veins. Most centers now use CE-MRA or CTA to determine the pulmonary vein anatomy prior to the procedure. Either technique provides high-resolution 3D tomographic images of the pulmonary veins and other mediastinal structures. Advantages of CE-MRA are the use of lower-morbidity gadolinium-based contrast and the lack of ionizing radiation. Long-term toxicity related to nephrogenic systemic fibrosis does need to be considered for patients with impaired renal function (see Chapter 6). These images can also be imported into the 3D electrophysiologic mapping systems that are an integral part of the procedure to combine anatomic and functional information during the procedure.45 It is also important to identify the relationship of the pulmonary veins to other mediastinal structures to avoid complications during the procedure. The formation of an atrial-esophageal fistula is a rare but catastrophic complication that is caused by excessive heating of the posterior LA wall and the adjacent esophagus.46–49 The esophagus and its relationship to the LA can be readily identified on standard anatomic MR sequences (Fig. 32-5). The esophagus almost always directly abuts the LA and is usually closer to the left-sided pulmonary veins, but the location is highly variable.50–52 The esophagus is frequently within 5 mm of the pulmonary veins at a location that probably increases the risk for the formation of an atrial-esophageal

LA RI

Eso

IMAGING BEFORE AND AFTER ATRIAL FIBRILLATION ABLATION Pulmonary vein imaging is usually performed before atrial fibrillation ablation to determine the pulmonary vein

Figure 32-5 Anatomic relationship of the esophagus to the pulmonary veins and the left atrium. This spin echo T1-weighted image shows the esophagus (Eso) immediately posterior to the left atrium and adjacent to the right inferior pulmonary vein. The esophagus is compressed between the enlarged LA and the spinal column. Eso, esophagus; LA, left atrium; RI, right inferior pulmonary vein. Cardiovascular Magnetic Resonance 445

32 PULMONARY VEIN IMAGING

pulmonary venous anatomy. CE-MRA is generally able to identify all pulmonary venous anomalies, providing new information in 75% of cases and identifying previously unsuspected anomalies in 30%.26

VASCULATURE AND PERICARDIUM

pulmonary vein angioplasty is usually successful in restoring normal flow and alleviating symptoms.66 Newer techniques that have emphasized placing ablation lesions closer to the LA under intracardiac echocardiographic guidance have reduced the rate of pulmonary vein stenosis,57 but screening is still recommended for all patients.

RS

LA LI

Figure 32-6 Pulmonary vein stenosis. There is severe stenosis of the left inferior pulmonary vein and moderate stenosis of the right superior pulmonary vein, indicated by the dashed arrows. There is prestenotic dilation of the left inferior pulmonary vein. LA, left atrium; LI, left inferior pulmonary vein; RS, right superior pulmonary vein.

fistula.53,54 The risk of causing an atrial-esophageal fistula may be reduced by avoiding ablation in the region of the LA closest to the esophagus, but this may be difficult, as the esophagus is mobile and may move during the course of the procedure.55 Some practitioners advocate radiopaque nasogastric tube “markers” in the esophagus during the ablation procedure to avoid this complication. The pulmonary veins are imaged after the procedure to screen for pulmonary vein stenosis. Pulmonary vein stenosis is an uncommon but severe complication of atrial fibrillation ablation (Fig. 32-6).2,19,44,56–62 The application of RF energy to the pulmonary veins causes intimal proliferation and myocardial necrosis that can result in stenosis or occlusion.63 Severe stenosis occurs in up to 5% of patients after the procedure and results in pulmonary hypertension and decreased perfusion of the affected lung segments.64,65 Patients frequently present with cough or dyspnea, but a significant proportion are asymptomatic.58 Stenosis is most likely to occur in smaller pulmonary veins in which the ablation lesions were placed further into the pulmonary vein trunk and with greater extent of ablation.61,62 If stenosis does occur,

QUANTIFICATION OF PULMONARY VEIN SIZE The accurate measurement of pulmonary vein size is essential for serial assessment of pulmonary vein stenosis and to further investigate the role of pulmonary veins in the generation and maintenance of atrial fibrillation. Most investigators have measured pulmonary vein diameters in a specified plane, usually at the ostia.2,44,62 These measurements tend to have poor reproducibility for several reasons (Fig. 32-7). Identification of the true ostia is very difficult because the pulmonary veins and LA are embryologically related with no clear anatomic border between them. The pulmonary vein ostia are not round, such that measurements taken at the same location vary significantly with the plane of measurement.2,3 A further complication is that most measurements are derived from nongated images, while the pulmonary vein size varies significantly over the cardiac cycle.7,67 These difficulties were highlighted in a study comparing pulmonary vein diameter measurements performed by using CTA, intracardiac echocardiography, transesophageal echocardiography, and X-ray venography in the same patients.68 Each of these methods identified different numbers and positions of pulmonary veins with a poor correlation between diameter measurements obtained with each imaging modality. Tomographic imaging of the pulmonary veins using CECMR (or CTA) has several advantages. All of the anatomic information is obtained in a single 3D dataset that can be manipulated in numerous ways. This allows for anatomic measurements in any desired plane, including determination of the perimeter and cross-sectional area that may be more meaningful measures of pulmonary vein size.

LI

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A

B

Figure 32-7 Difficulty in measurement of pulmonary vein diameters. These images show a right superior pulmonary vein in the axial plane (A) and a left inferior pulmonary vein in the sagittal plane (B). There is no clear anatomic border between the right superior pulmomary vein and the left atrium (A) such that several potential diameter measurements are possible (solid lines). The left inferior pulmonary vein is oval (B). The diameters measured in the axial plane (horizontal line) and coronal plane (vertical line) differ from each other and from the true maximal diameter (dashed line). LA, left atrium; LI, left inferior pulmonary vein; RS, right superior pulmonary vein. 446 Cardiovascular Magnetic Resonance

RS

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RM LA

LI

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Ao

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A RI

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Figure 32-8 Late gadolinium enhancement CMR. These images show the pulmonary vein anatomy (A) in a patient who underwent LGE CMR before (B) and 6 weeks after (C ) pulmonary vein ablation. He has variant anatomy with an additional right middle pulmonary vein (panel A). Prior to ablation (panel B), there is no hyperenhancement of the pulmonary veins. There is increased signal in the right superior and right middle pulmonary veins due to artifact from the ventilatory compensation technique. After ablation (panel C ), there is evidence of pulmonary vein scar (dashed arrows). The inset in panel C shows a reformatted image of the left inferior pulmonary vein that shows circumferential scar. Ao, aorta; LA, left atrium; LI, left inferior pulmonary vein; RI, right inferior pulmonary vein; RM, right middle pulmonary vein; RS, right superior pulmonary vein.

A simple method for determining pulmonary vein size in the sagittal plane is highly reproducible and provides these additional measures.3 The maximal diameter and cross-sectional area are measured at the location in the sagittal plane at which the pulmonary veins separate from the LA and from each other (see Fig. 32-1). This is easily determined by scrolling through a reconstruction of the 3D dataset in the sagittal plane. Because the measurements are made in a standard plane and location, reproducibility is greatly improved in comparison to standard diameter measurements.3 This allows for more accurate determination of interstudy differences in pulmonary vein size and increased statistical power in research studies. Even in the absence of severe stenosis, this method can identify small changes in pulmonary vein size after atrial fibrillation ablation that may be due to hemodynamic changes related to the restoration of sinus rhythm.4 The determination of the perimeter and cross-sectional area is also advantageous. Patients with larger summed total pulmonary vein cross-sectional area are more likely to have recurrent atrial fibrillation after ablation independent of the type of atrial fibrillation or left atrial size.69 Diameter measurements do not have predictive value.42

map of the ablation lines. This could be useful either after the procedure to assess patient outcomes or as a real-time tool to guide ablation during interventional CMR. Ventricular scar can be detected by using the LGE technique,70 in which a strongly T1-weighted CMR image is acquired 10 to 20 minutes after the injection of gadolinium contrast.71 Gadolinium contrast remains concentrated in the regions of scar, compared to muscle or blood, owing to reduced clearance and the large contrast distribution volume in fibrotic regions.72 To detect scar in the LA, the standard LGE method is modified to achieve higher spatial resolution (1.3 1.3 5 mm) by acquiring a 3D volume during free breathing with ventilatory motion compensated imaging (Fig. 32-8).73 The clinical utility of this technique is currently being assessed. Preliminary data suggest that the circumferential extent of hyperenhancement is predictive for clinical success (freedom from atrial fibrillation recurrence).74 Though of uncertain clinical impact, esophageal hyperenhancement is quite common.75

LATE GADOLINIUM ENHANCEMENT PULMONARY VEIN IMAGING

CE-MRA easily images the pulmonary veins and is very useful to define normal, variant, and anomalous pulmonary vein anatomy. Imaging is usually performed before atrial fibrillation ablation to plan for the procedure and afterward to screen for pulmonary vein stenosis. Quantification of pulmonary vein size is important for the comparison of serial studies and is easily accomplished by examining images in the sagittal plane. LGE imaging of the pulmonary veins is a new technique that may be useful to define the extent of ablation and to guide follow-up procedures. In the future, interventional CMR with RF ablation in the CMR laboratory may also be available (see Chapter 43).

An added advantage of CE-MRA (versus CTA) is the opportunity to also image LA scar that results from the ablation procedure. RF ablation for the treatment of atrial fibrillation results in scarring of the pulmonary vein and LA.63 Scar imaging has the potential to noninvasively assess the completeness of ablation by providing a precise anatomic

CONCLUSION

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63. Taylor GW, Kay GN, Zheng X, Bishop S, Ideker RE. Pathological effects of extensive radiofrequency energy applications in the pulmonary veins in dogs. Circulation. 2000;101:1736–1742. 64. Kluge A, Dill T, Ekinci O, et al. Decreased pulmonary perfusion in pulmonary vein stenosis after radiofrequency ablation: assessment with dynamic magnetic resonance perfusion imaging. Chest. 2004; 126:428–437. 65. Arentz T, Weber R, Jander N, et al. Pulmonary haemodynamics at rest and during exercise in patients with significant pulmonary vein stenosis after radiofrequency catheter ablation for drug resistant atrial fibrillation. Eur Heart J. 2005;26:1410–1414. 66. Qureshi AM, Prieto LR, Latson LA, et al. Transcatheter angioplasty for acquired pulmonary vein stenosis after radiofrequency ablation. Circulation. 2003;108:1336–1342. 67. Bowman AW, Kovacs SJ. Prediction and assessment of the timevarying effective pulmonary vein area via cardiac MRI and Doppler echocardiography. Am J Physiol Heart Circ Physiol. 2005;288: H280–H286. 68. Wood MA, Wittkamp M, Henry D, et al. A comparison of pulmonary vein ostial anatomy by computerized tomography, echocardiography, and venography in patients with atrial fibrillation having radiofrequency catheter ablation. Am J Cardiol. 2004;93:49–53. 69. Hauser TH, Essebag V, Baldessin F, et al. Larger pulmonary vein crosssectional area is associated with recurrent atrial fibrillation after pulmonary vein isolation. Circulation. 2005;112(S):II–555. 70. Kim RJ, Wu E, Rafael A, et al. The use of contrast-enhanced magnetic resonance imaging to identify reversible myocardial dysfunction. N Engl J Med. 2000;343:1445–1453. 71. Simonetti OP, Kim RJ, Fieno DS, et al. An improved MR imaging technique for the visualization of myocardial infarction. Radiology. 2001;218:215–223. 72. Judd RM, Lugo-Olivieri CH, Arai M, et al. Physiological basis of myocardial contrast enhancement in fast magnetic resonance images of 2-day-old reperfused canine infarcts. Circulation. 1995;92:1902–1910. 73. Peters DC, Wylie J, Hauser TH, et al. Detection of pulmonary vein and left atrial scar after catheter ablation using 3D navigator-gated delayed enhancement magnetic resonance imaging: initial experience. Radiology. 2007;243:690–695. 74. Peters DC, Wylie JV, Hauser TH, et al. Recurrence of atrial fibrillation correlates with extent of post-procedural late gadolinium enhancement: a pilot study. JACC Cardiovasc Imaging. 2009;2:308–316. 75. Meng J, Peters DC, Hsing JM, et al. Late gadolinium enhancement of the esophagus is common on cardiac MR several months after pulmonary vein isolation: preliminary observations. Pacing Clin Electrophysiol. January 2010 (online).

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50. Lemola K, Sneider M, Desjardins B, et al. Computed tomographic analysis of the anatomy of the left atrium and the esophagus: implications for left atrial catheter ablation. Circulation. 2004;110: 3655–3660. 51. Tsao HM, Wu MH, Higa S, et al. Anatomic relationship of the esophagus and left atrium: implication for catheter ablation of atrial fibrillation. Chest. 2005;128:2581–2587. 52. Cury RC, Abbara S, Schmidt S, et al. Relationship of the esophagus and aorta to the left atrium and pulmonary veins: implications for catheter ablation of atrial fibrillation. Heart Rhythm. 2005;2: 1317–1323. 53. Sanchez-Quintana D, Cabrera JA, Climent V, Farre J, Mendonca MC, Ho SY. Anatomic relations between the esophagus and left atrium and relevance for ablation of atrial fibrillation. Circulation. 2005;112:1400–1405. 54. Monnig G, Wessling J, Juergens KU, et al. Further evidence of a close anatomical relation between the oesophagus and pulmonary veins. Europace. 2005;7:540–545. 55. Good E, Oral H, Lemola K, et al. Movement of the esophagus during left atrial catheter ablation for atrial fibrillation. J Am Coll Cardiol. 2005;46:2107–2110. 56. Moak J, Moore H, Lee S, et al. Case report: pulmonary vein stenosis following RF ablation of paroxysmal atrial fibrillation: successful treatment with balloon dilation. J Interv Card Electrophysiol. 2000;4: 621–631. 57. Saad EB, Rossillo A, Saad CP, et al. Pulmonary vein stenosis after radiofrequency ablation of atrial fibrillation: functional characterization, evolution, and influence of the ablation strategy. Circulation. 2003;108:3102–3107. 58. Saad EB, Marrouche NF, Saad CP, et al. Pulmonary vein stenosis after catheter ablation of atrial fibrillation: emergence of a new clinical syndrome. Ann Intern Med. 2003;138:634–638. 59. Scanvacca M, Kajita L, Vieira M, Sosa E. Pulmonary vein stenosis complicating catheter ablation of focal atrial fibrillation. J Cardiovasc Electrophysiol. 2000;1:677–681. 60. Yang M, Akbari H, Reddy GP, Higgins CB. Identification of pulmonary vein stenosis after radiofrequency ablation for atrial fibrillation using MRI. J Comput Assist Tomogr. 2001;25:34–35. 61. Arentz T, Jander N, von Rosenthal J, et al. Incidence of pulmonary vein stenosis 2 years after radiofrequency catheter ablation of refractory atrial fibrillation. Eur Heart J. 2003;24:963–969. 62. Dill T, Neumann T, Ekinci O, et al. Pulmonary vein diameter reduction after radiofrequency catheter ablation for paroxysmal atrial fibrillation evaluated by contrast-enhanced three-dimensional magnetic resonance imaging. Circulation. 2003;107:845–850.

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CHAPTER 33

Thoracic Aortic Disease Stephan Kische, Hu¨seyin Ince, and Christoph A. Nienaber

The anatomic and functional characteristics of the aorta, which may at first glance appear relatively straightforward, are now recognized to be complex. Recent insights from modern imaging technology and from a better understanding of the hydraulic principles associated with the variety of diseases affecting the aorta have helped the medical community to realize the multiple facets of in vivo aortic pathology as well as its varied clinical presentation. Diagnostic modalities such as transesophageal echocardiography (TEE), cardiovascular magnetic resonance imaging (CMR), and spiral computed tomography (CT) have all been shown to be useful to interrogate the aorta, both in chronic disease and in acute aortic syndromes. X-ray contrast angiography, the former gold standard in acute and chronic aortic syndromes, has been relegated to a secondary role after the emergence of the noninvasive techniques, most importantly CMR, with their high sensitivity, specificity, and practical advantages.1–7 However, none of the diagnostic modalities listed above is ideal for all patients, and for a given individual, knowledge of both accuracy and limitations in the presenting clinical scenario are required.6–10 While the information content of CMR may greatly overlap with established methods such as echocardiography, CT, or angiography, the technique is more accurate and comprehensive. Although the costeffectiveness of CMR has not been proven in all areas,11 CMR is the preferred modality in selected areas of aortic disease, such as aneurysm, dissection, and its precursors, congenital and inherited heart diseases, and in particular for postoperative follow-up of aortic repair and cardiac malformations. This chapter focuses on the possibilities and emerging advantages of CMR with respect to a spectrum of thoracic aortic pathologies.

PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE IN AORTIC IMAGING Spin Echo Cardiovascular Magnetic Resonance Imaging Spin echo T1-weighted imaging provides the best anatomic detail of the aortic wall and pathologic conditions such as atheromatous plaques, intimal flaps, or intramural hemorrhage and is still the basis of any aortic study12,13 while T2-weighted images (repetition time: 2/3 R-R; echo time: 450 Cardiovascular Magnetic Resonance

80 to 100 msec) can be used in tissue characterization of the aortic wall or blood components. Electrocardiographic (ECG) triggering is essential in minimizing motion and pulsatility artifacts. Slice thickness of 3 to 8 mm and an echo time (TE) of 20 to 30 msec are standard, while repetition time (TR) is determined from the R-R interval of the ECG. A shorter acquisition time can be achieved with fast spin echo pulse sequences whereby a long train of echoes is acquired by using a series of 180 radiofrequency (RF) pulses; washout effects are even more substantial than in conventional spin echo techniques. A superior black-blood effect is achieved by using preparatory pulses14 (such as presaturation, dephasing gradients, and preinversion) with one or more additional RF pulses outside the plane to suppress the signal intensity of in-flowing blood and nullify the blood signal (Fig. 33-1). Therefore, “black-blood” fast T1- and T2-weighted spin echo sequences have improved image quality, with substantial saving of time, and constitute, at present, the method of choice for morphologic assessment of the thoracic aorta. Usually, a conventional study of the thoracic aorta is first acquired in the axial plane, displaying the orientation of the great arteries and optimal visualization of mural lesions perpendicular to their long axes. Images in additional planes, the oblique sagittal or coronal view, are then acquired, depending on the anatomy and diagnostic problems, to define the longitudinal extent of the disease.

Gradient Echo Cardiovascular Magnetic Resonance Imaging and Flow Mapping Gradient echo techniques provide dynamic and functional information, although with fewer details of the vessel wall. The bright signal of the blood pool on gradient echo images results from flow-related enhancement obtained by applying RF pulses to saturate a volume of tissue. With a short TR (4 to 8 msec) and low flip angle (20 to 30 ), maximal signal is emitted by blood flowing in the voxel. An ECG signal is acquired with the imaging data so that the images, acquired with a high degree of temporal resolution throughout the cardiac cycle (up to 20 to 25 frames), are reconstructed in the different phases of the cardiac cycle and can be displayed in cine format. Flow-related enhancement is produced by inflow of unsaturated blood exposed to only one RF pulse. As result, the laminar moving blood displays a bright signal in contrast to stationary tissues. The signal can be reduced if the flow is low, as in aortic aneurysms. Mural thrombi can be identified by

AAo

RPA

DAo

Magnetic Resonance Angiography Figure 33-1 Black-blood vascular imaging of the aorta obtained with cardiac gating and breath holding. Axial image shows the ascending (AAo) and descending (DAo) aorta at the level of the right pulmonary artery (RPA). Note the excellent suppression of the luminal blood signal and demonstration of the vessel wall.

persistent low-signal intensity in different phases of the cardiac cycle. Turbulent flow produces rapid spin dephasing and results in a signal void, providing additional information in many pathologic conditions such as coarctation, aortic valve insufficiency, aortic aneurysm, and dissection.15 Particularly in aortic dissection, the detection of entry and reentry sites is a special capability of functional CMR that can be helpful in planning both surgical and endovascular therapy. Accurate quantitative information on blood flow is obtained from modified gradient echo sequences with parameter reconstruction from the phase rather than the amplitude of the magnetic resonance (MR) signal; this is also known as flow mapping or phase contrast or velocity-encoded cine CMR16,17 (Fig. 33-2). In each pixel of velocity images, the phase of the signal is related to the velocity component in the direction of a bipolar velocity phase-encoding gradient. In the phase image, the velocity of blood flow can be determined for any site of the vascular Figure 33-2 Phase-contrast imaging of the aorta. Magnitude (A) and phase (B) axial images show the ascending (AAo) and descending (DAo) aorta at the level of the main pulmonary artery (MPA in panel A). Flow encoding was superior to inferior. On the phase image (B), the ascending aorta and main pulmonary artery appear black, and the descending aorta appears white, owing to the opposite directions of flow in these arteries.

AAo

A variety of magnetic resonance angiography (MRA) techniques, including various pulse sequences, methods of data acquisition, and postprocessing, have been developed in the evaluation of vascular structures.20 However, after the introduction of faster gradient systems, three-dimensional (3D) contrast-enhanced MRA constitutes the method of choice for the evaluation of the aorta.21–26 The technique relies on the contrast-induced T1-shortening effects of the contrast medium whereby saturation problems with slow flow or turbulence-induced signal voids are avoided. During the short intravascular phase, the paramagnetic contrast agent provides a signal in the arterial or venous system, enhancing the vesselto-background contrast-to-noise ratio irrespective of flow patterns and velocity. Pulsatility artifacts are minimized, even in the ascending aorta and without ECG gating. The paramagnetic contrast agent (e.g., Gadolinium-DTPA) is generally administered intravenously (antecubital vein). Bolus timing is necessary to ensure peak enhancement during the middle of CMR acquisition. The flow rate should be adjusted to guarantee that the contrast volume is injected in a period not exceeding the acquisition time, and the start delay can be easily monitored by a real-time fluoroscopic triggering when a power injector is employed.27 Improved gradient systems allow a considerable reduction of the minimum TRs and TEs and the acquisition of complex 3D datasets within a breath

MPA

DAo

A

B

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system. Flow velocity is calculated by using a formula in which velocity is proportional to change in the phase angle of protons in motion. MR maps of flow velocity are obtained two-dimensionally, which is particularly important in profiles of nonuniform flow such as that in the great vessels. On phase images, the gray value of a pixel depends on the velocity and direction with respect to the imaging plane. Signals below a defined range are considered noise and are eliminated by a subtraction process. Quantitative data on flow velocity and flow volume are obtained from the velocity maps through a region of interest. The mean blood flow is estimated by multiplying the spatial mean velocity and the cross-sectional area of the vessel. Vector mapping has been used to describe flow patterns in different aortic diseases (e.g., hypertension, aneurysms, dissection Marfan syndrome, coarctation).18,19

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Figure 33-3 Gadolinium-enhanced 3D MRA of the thoracic aorta (surface-shaded display algorithm).

hold interval of under 30 seconds. With the support of maximum intensity projection (MIP) images and the 3D multiplanar reformation, this technique delineates all the morphologic details of the aorta and its side branches in any plane in a 3D format (Fig. 33-3).

DISSECTION OF THE THORACIC AORTA Acute aortic dissection is a life-threatening medical emergency requiring prompt diagnosis and treatment.28 The 14-day period after onset has been designated as an acute

Stanford Type A

phase because the rates of morbidity and mortality are highest during this period: 1% to 2% per hour in the first 24 hours after onset and 80% within 2 weeks.29,30 Classic aortic dissection is characterized by a laceration of the aortic intima that allows blood to course through a false lumen in the medial layer. Dissection can occur throughout the length of the aorta. The two most frequently used classifications (DeBakey and Stanford) are based on the anatomic location and extension of intimal flap (Fig. 33-4). DeBakey’s nomenclature is based on the anatomic site of the intimal tear and the extent of the resulting dissection.31 In a type I dissection, the intimal tear originates in the ascending aorta, and the dissecting hematoma extends past the origin of the left subclavian artery. Type II dissections are confined to the ascending aorta. Type III dissections begin after the origin of the left subclavian artery and extend distally. The Stanford classification is conceptually founded on prognostic grounds. Type A dissections involve the ascending aorta (Fig. 33-5), regardless of the site of the entry tear, and type B dissections spare the ascending aorta and often imply a better prognosis.32 In general, acute dissections involving the ascending aorta require emergency surgery, while descending aortic dissections may be successfully managed with medical therapy alone.33,34 Thus, rapid diagnosis of the dissecting process and a delineation of its anatomic details are critical for successful management.35 The primary diagnostic goal is not only a clear anatomic delineation of the intimal flap and its extension, but also the detection of the entry and reentry sites and the presence and degree of aortic insufficiency and flow in the aortic branches.7 The relationship between true and false lumina and the visceral vessels, with any involvement of the iliac arteries and the identification of the entry and reentry sites, is crucial in patient selection for transcatheter endovascular repair of type B acute and chronic aortic dissection as an alternative to open surgery.36,37 In a suspected case of aortic dissection, the standard CMR examination should begin with spin echo “black-blood” sequences. In the axial plane, the intimal flap is detected as a straight linear image inside the aortic lumen. The true lumen can be differentiated from the false lumen by the anatomic Figure 33-4 Commonly used classification systems for aortic dissection. While the DeBakey classification (I, II, IIIa, IIIb) focuses on the anatomic extent, the Stanford classification (A, B) highlights the involvement of the ascending aorta and the prognostic aspects of dissection.

Stanford Type B

Type II

Type III a

De Bakey Type I Type III b

452 Cardiovascular Magnetic Resonance

B Figure 33-5 Stanford type A aortic dissection. A, Axial T1-weighted black-blood image shows nearly circumferential compression of the true aortic lumen by a false lumen (arrow). High signal intensity in the false lumen makes it difficult to differentiate thrombosis from flowing blood. B, Axial reformatted image from contrast-enhanced MRA shows an intimal flap (black arrow) with flow in the false lumen (white arrow).

features. In addition, the visualization of remnants of the dissected media as cobwebs adjacent to the outer wall of the lumen may help to identify the false lumen. The leakage of blood from the descending aorta into the periaortic space, which can appear with high signal intensity and can result in a left-sided pleural effusion, is usually better visualized on axial images. A high signal intensity of a pericardial effusion indicates a bloody component and is considered to be a sign of impending rupture of the ascending aorta into the pericardial space. A detailed anatomic map of aortic dissection must indicate the type and extension of dissection and distinguish the origin and perfusion of branch vessels from the true or false channels. Therefore, a further spin echo sequence on the sagittal plane should be performed, and in stable patients, adjunctive gradient echo sequences or phase contrast images can be instrumental in identifying aortic insufficiency and entry or reentry sites as well as in differentiating slow flow from thrombus in the false lumen.38,39 The third step in the diagnosis of aortic dissection and definition of its anatomic detail relies on the use of gadoliniumenhanced 3D MRA. Because 3D MRA is rapidly acquired without any need of ECG triggering and gadolinium has minimal toxicity in patients with good renal function, this technique may even be used with severely ill patients.40 With spin echo sequences, artifacts caused by imperfect ECG gating, respiratory motion, or slow blood pool can result in intraluminal signal, simulating or obscuring an intimal flap. In gadolinium-enhanced 3D MRA, the

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A

intimal flap is easily detected, and the relationship with aortic vessels is clearly depicted. Entry and reentry sites appear as a segmental interruption of the linear intimal flap on axial or sagittal images. The analysis of MRA images should not be limited to viewing MIP images or surface-shaded display; it should also include a complete evaluation of reformatted images in all three planes to confirm or improve spin echo information and exclude artifacts. Combining the spin echo with MRA images completes the diagnosis and anatomic definition.41 At present, CMR is one of the most accurate tools in the detection of aortic dissection. A high degree of spatial resolution and contrast and the capability for multiplanar acquisition provide excellent sensitivity and specificity rating at approximately 100% in the published series.6,41–43 Each of the current noninvasive imaging modalities has advantages and limitations in the evaluation of suspected aortic dissection. In selecting the study of choice, one must begin by considering what diagnostic information needs to be obtained. First and foremost, the study must confirm or refute the diagnosis of dissection. Second, the study must determine whether or not the dissection involves the ascending aorta (Type A) or is confined to the descending aorta (Type B). Third, a number of anatomic features of the dissection should be identified, including its extent, the sites of entry and reentry, the presence of thrombus in the false lumen, the extent of branch vessel involvement, the presence and severity of aortic insufficiency, the presence of a pericardial effusion, and the presence of coronary artery involvement. It is also important to consider the accuracy of the diagnostic information obtained, as a false negative diagnosis may result in avoidable death, while a false positive diagnosis might lead to unnecessary surgery. According to the present information, CMR and TEE are the most sensitive modalities, with both performing better than aortography. The sensitivities of aortography, CT, and CMR are all quite high, while the specificity of TEE may be comparable only when a strict definition of a positive study is applied. Finally, availability, speed, safety, and cost should be taken into consideration in comparing various modalities. Aortography is rarely immediately available, requires transport of the patient, is a lengthy study, has the associated risks of both an invasive study and intravenous contrast, and is the most expensive. Yet it may be necessary in selected patients who are being considered for combined surgical treatment, especially those with a high likelihood of coronary artery disease or evidence of involvement of major arterial trunks arising from the aorta. CT scanning has the advantage that it is more easily obtained in less time and is noninvasive, but it is overall less accurate than the other techniques. CMR is usually less available in most hospitals, requires transportation of the patient, and is considered undesirable for unstable patients or those requiring very close monitoring. Meanwhile, TEE is readily obtained, quick to complete at the bedside and thus ideal for unstable patients, and the least costly of the four imaging techniques. Therefore, TEE with its accuracy, safety, speed, and convenience is often considered the study of first choice in cases of suspected dissection. In some institutions, TEE has assumed this role, with many surgeons taking patients to the operating room on the basis of the diagnostic findings of TEE alone, especially when the

VASCULATURE AND PERICARDIUM

A

B

Figure 33-6 Contrast-enhanced MRA of chronic Type B dissection originating from the aortic arch region. A, Follow-up MRA at 7 days after stent-graft placement shows a completely sealed proximal entry to the thrombosed false lumen. B, The diameter of the true lumen is normalized, and the descending aorta is reconstructed.

likelihood of coronary disease and any compromise of major arterial trunks is low. These options may render coronary angiography obsolete in the routine workup of aortic dissection. The optional approach to detecting dissection of the thoracic aorta should be a noninvasive strategy using CMR in all hemodynamically stable patients and TEE in patients who are too unstable for transportation. Comprehensive and detailed evaluation can thus be reduced to a single noninvasive imaging modality in the evaluation of suspected aortic dissection. While CMR may be less practical than TEE for the evaluation of patients presenting with suspected aortic dissection, it is well suited for studying patients with stable or chronic dissections. The extraordinary accuracy of CMR with its high-quality images may make it the gold standard for defining aortic anatomy in such patients. It appears an accepted policy to follow patients after successful evaluation and surgical treatment by CMR to identify any subsequent aneurysm formation or extension of dissection. Moreover, CMR and MRA have proven to be of particular importance for the individual morphometry of aortic lesions considered for endovascular covered stent-graft treatment. The MRA appears the optimal diagnostic tool to custom-design an individual stent-graft for lesions such as aneurysms, aortic dissections, and localized aortic ulcers likely to penetrate the aortic wall. With a customized aortic endoprosthesis, such pathomorphologic entities are becoming more frequently an ideal target for interventional treatment rather than surgical approaches that have high mortality and morbidity44 (Fig. 33-6).

AORTIC INTRAMURAL HEMATOMA An important differential diagnosis of aortic dissection is intramural hematoma (IMH), which usually presents with the same classic clinical picture and risk profile as overt 454 Cardiovascular Magnetic Resonance

aortic dissection.45 The noninvasive tomographic modalities such as CMR may be used to identify IMH with no luminal component, which is considered an imminent precursor of aortic dissection.46,47 IMH was first described by Krukenberg in 1920 as “dissection without initial tear” and was originally considered a distinct entity at necropsy. However, with high-resolution tomographic imaging, the in vivo diagnosis of IMH is now feasible, and the data suggest that IMH is a precursor of dissection, as there is a high (up to 30%) rate of progression to overt dissection.48–50 Both CMR and CT imaging can identify IMH in 12.8% of patients with acute aortic syndromes; this percentage is almost identical to autopsy results of 13.2% with no identifiable intimal tear.46 Because of these findings in vivo before death, IMH appears more likely to be a variant of dissection than a separate entity.46,51,52 Typical epiphenomena of dissection such as aortic insufficiency, pericardial or pleural effusion may also occur in IMH.52,53 Moreover, arterial hypertension is the most frequent predisposing factor for IMH as well as for overt dissection, with an estimated incidence of 85%. Spontaneous rupture of aortic vasa vasorum, especially of nutrient vessels to the media layer, has been suggested to initiate the process of aortic disintegration without an intimal tear. With a pathogenesis that explains the high rate of progression to overt aortic dissection and a prognosis and survival that are similar to those in aortic dissection,54,55 urgent diagnosis of IMH is very important. The extent of IMH can be considerable, with abnormal thickness of the aortic wall up to 30 mm both asymmetrical or symmetrical in circumference with an extent of 3 to 30 cm. It should be noted that angiography is of no use for IMH, since there is no intraluminal component. The diagnosis of IMH relies on the visualization of intramural blood and/or evidence of localized increased wall thickness.56 The high density of fresh hematoma on CT scans appears specific for IMH. CMR techniques, however, not only visualize the blood in the wall, but also allow an assessment of the age of the hematoma based on signal changes caused by the formation of methemoglobin (Fig. 33-7). Acute IMH (early stage) is well imaged on T2-weighted spin echo images due to high initial signal intensity of blood, whereas blood of 1 to 5 days of age has lower signal intensity on T2 images. High signal intensity within the aortic wall on T1 spin echo images suggests subacute IMH, whereas acute IMH may be determined on T1 images from the isodense appearance of blood and aortic wall.46,53 TEE has also been emphasized as a diagnostic tool; however, the differentiation of IMH from severe atherosclerosis with local wall thickening may be difficult, and IMH may be diagnosed only retrospectively with serial evaluation (resolution or progression to dissection). Moreover, false positive findings of local thickening on tangential scans and around the hemiazygos vein may be more likely with TEE. Both TEE and CT may result in false positive and false negative findings (pathologic wall thickness without hematoma), whereas the segmental extent of IMH is usually correctly defined with CMR. Although TEE has an excellent sensitivity to detect aortic dissection, the definite distinction between IMH and normal findings may require a second tomographic modality such as CT or CMR, since a false negative result (or false exclusion of IMH) is more likely to be avoided with independent

33 THORACIC AORTIC DISEASE

A

B

C

Figure 33-7 T1-weighted spin echo axial image of intramural hematoma of the ascending and descending aorta. The abnormal wall thickening (arrows) present intermediate signal intensity in panel A (oxyhemoglobin, acute phase) and high signal intensity in panel B (methemoglobin, subacute phase). C, T2-weighted spin echo image signal intensity is high in the acute phase (recent hemorrhage).

morphologic information. In conclusion, as a precursor of dissection, IMH requires diagnostic attention by use of high-resolution tomographic imaging; owing to its physical properties, CMR may play a prominent role not only to diagnose IMH, but also to assess its age and to differentiate IMH from mural thrombosis; angiography certainly is not diagnostic. Given the poor results and unfavorable outcome with medical treatment, early surgical repair should be considered for all patients with ascending aortic involvement (type A IMH) and for any patient with recurrent pain. Conversely, surgery may not be required in patients with IMH of the descending aorta. Yet all patients may benefit from serial follow-up CMR imaging to rule out progression regardless of treatment strategy, owing to new lesions or spontaneous relapses even after surgical repair.

PENETRATING AORTIC ULCER Ulcers occur in the presence of aortic atherosclerosis and may mimic subacute dissection or may develop without major symptoms. In contrast to IMH, aortic ulcers are characterized on angiography by focal contrast enhancement beyond the confines of the aortic lumen but communicating with the lumen (Fig. 33-8).57 At present, there is no consensus on the prognosis and outcome of aortic ulcers.58,59 Persistent pain, hemodynamic instability, and signs of expansion should trigger surgical or endovascular treatment, whereas asymptomatic patients can be managed medically. An incidence of transmural rupture ranging from 8% to 42% has been reported in the medical literature.60,61 Both IMH and ulcers are unrelated to intimal lacerations as in acute aortic dissection.62 Lacerations and IMH usually occur at points of greatest hydraulic stress (right lateral ascending aorta or adjacent to the ligamentum arteriosum), whereas penetrating ulcers are typically found in the descending or abdominal aorta. Conversely, discrete penetrating atheromatous ulcers (“giant ulcers”) have been suspected as one cause of intramural bleeding.63 In such a chronic setting, the hematoma is confined to the rim

Figure 33-8 Penetrating atherosclerotic ulcer in a 61-year-old man with high blood pressure. Sagittal oblique contrast-enhanced MRA shows a diffusely aneurysmal descending aorta with a large outpouching in the medial anterior region (arrow), an appearance consistent with penetrating atherosclerotic ulcer. Prior studies showed development of a focal aneurysm 2 years earlier.

adjacent to the ulcer. Both MRI and TEE have helped to elucidate the pathogenic background and the complex anatomic peculiarities of these important features, but accurate diagnosis of penetrating ulcers is sometimes difficult, and no test is ideal.64 CT may demonstrate the surrounding hematoma and displaced calcifications in most cases; in addition to this, CMR can differentiate subacute intramural hematoma from chronic intraluminal thrombus.65 No large-scale studies of CMR and aortic ulcers are currently available, but owing to its versatility and imaging features, CMR appears best suited to characterize this form of aortic Cardiovascular Magnetic Resonance 455

VASCULATURE AND PERICARDIUM

pathology as a constellation of diffuse aortic atherosclerosis, an ulceration with focal wall thickening and missing evidence of aortic dissection.

THORACIC AORTIC ANEURYSM Limited dilatation of a blood vessel can be either a true aneurysm or a false aneurysm. True aneurysms involve all layers of the aortic wall and result from the degeneration of the elastin fibers within the media. A false aneurysm, or pseudoaneurysm, is not a true aneurysm but rather a contained perforation of the vessel wall with penetration of the intima and media. Pseudoaneurysms form after focal penetration of the intima and media by trauma (20% to 25% of thoracic aneurysms) or by infection (about 5% of thoracic aneurysms). Atherosclerosis may result in the formation of pseudoaneurysms if a true aneurysm ruptures but is contained by the adventitia and periaortic tissues. Pseudoaneurysms, therefore, have a narrow “neck” leading to the “aneurysm.” True aneurysms usually result from the atherosclerotic process of infiltration and damage to the aortic media. Seventy percent of the aneurysms located in the thoracic aorta are associated with severe atherosclerosis, with up to 25% of patients having concomitant abdominal aortic aneurysms. Thoracic aortic aneurysms are common, with a recently reported incidence of 10.9 cases per 100,000 persons per year.66 They may involve all thoracic aortic segments, requiring different therapeutic strategies with a surgical, interventional, or hybrid approach (Fig. 33-9).67–71 In modern series, aneurysms of the ascending aorta occur most commonly (60%), followed by aneurysms of the descending

A

aorta (40%), whereas aneurysms involving the arch or thoracoabdominal aorta occur less often.68 Atherosclerotic aneurysms are usually fusiform, involving long segments of the aorta. Saccular aneurysms are less common, but like fusiform aneurysms, they are most often the result of atherosclerosis; infection, trauma, and degenerative disease may also result in saccular aneurysms. Aneurysms involving solely the ascending aorta tend to be saccular in shape. Often, they are caused by annuloaortic ectasia, Marfan syndrome, syphilis, or poststenotic dilatation resulting from aortic stenosis; the sinotubular junction is preserved in aortic stenosis but markedly dilated in the other entities. Pseudoaneurysms can be seen superimposed on poststenotic true aneurysms as a result of infective endocarditis of the aortic valve. If isolated to the descending aorta, saccular aneurysms may be traumatic or infective, but atherosclerosis remains the leading cause. Saccular and sinus Valsalva aneurysms in the aortic root very often have associated aortic insufficiency. The clear delineation of the location, extent, and shape of an aneurysm; its relation to branch vessels and adjacent structures; and its associated complicating factors, such as rupture, periaortic hematoma or infection, hemopericardium, or aortic valve insufficiency, are all of importance. CMR is effective in identification and characterization of thoracic aortic aneurysms as well as in evaluation of their pathophysiologic consequences.72–74 Standard spin echo sequences are helpful in evaluation of alterations of the aortic wall and periaortic space. Periaortic hematoma and areas of high signal intensity within the thrombus may indicate instability of the aneurysm and are well depicted on spin echo images. Atherosclerotic lesions are visualized as areas of increased thickness with high signal intensity and irregular profiles. In contrast to transaxial imaging, oblique planes allow precise determination of lumen

B

Figure 33-9 A, MRA images of an aortic arch aneurysm in parasagittal-oblique and axial orientation. B, Follow-up study after a hybridprocedure with initial head-vessel-debranching and staged endovascular stent-graft repair demonstrates perfect reconstruction of the proximal aorta. 456 Cardiovascular Magnetic Resonance

TRAUMA TO THE AORTA

Figure 33-10 3D shaded-surface-rendered contrast-enhanced MRA depicts a large true aneurysm of the thoracic descending aorta in a 38-year-old man.

diameter. With fat suppression techniques, the outer wall of the aneurysm can easily be distinguished from periadventitial fat tissue. Contrast-enhanced 3D MRA can provide precise topographic information about the extent of an aneurysm and its relationship to the aortic branches (Fig. 33-10). The capability of contrast MRA to visualize the Adamkiewicz artery represents an important advance in avoiding postoperative neurologic deficit secondary to spinal cord ischemia.75,76 Optimal treatment of thoracic aneurysms requires accurate definition of size, anatomy, rate of growth, presence of dissection or thrombus, and involvement of adjacent structures, including the aortic valve. Only CMR is comprehensive and consistently provides all this information. Aortography may underestimate the size of aneurysms if significant intra-aortic thrombus is present and may fail to fully define saccular structures because of stagnant intraaneurysmal blood flow. CT scanning is noninvasive and suitable for serial examinations but requires potentially nephrotoxic contrast. Transthoracic echocardiography usually visualizes the ascending aorta adequately but may not image the arch and is useless for the descending aorta, as is TEE for the abdominal aorta. Both TEE and CMR seem well suited for serially evaluating any thoracic aortic aneurysm; in addition to measuring aneurysm dimensions, both TEE and CMR can detect fistulas, false channels, intraluminal thrombosis, and aortic valvular insufficiency. At present, no large trials exist to compare the value of CMR and TEE for thoracic aneurysms; there is, however, consensus that CMR is accurate and versatile but most likely not as cost-effective in serial studies. Although CMR promises to emerge as the standard in the diagnosis and management

Aortic trauma is usually generated by deceleration of the body in serious motor vehicle crashes, pedestrian injuries, and falls from height.77 The aortic segment that is subjected to the greatest strain by rapid deceleration forces is located just beyond the area of the isthmus aortae. In clinical series, aortic rupture occurs at this location in 90% of cases.78,79 The traumatic lesion is a transverse tear variably extending from the intimal to adventitial layers.80,81 Intimal hemorrhage without any laceration has been described in pathologic series but was not recognized in the clinical setting before the advent of high-resolution tomographic imaging modalities. Recent reports indicate that intimal hemorrhage with and without partial intimal laceration tends to heal spontaneously. When the lesion involves intimal and medial layers, false aneurysm formation occurs. The aneurysm is fusiform in the case of a circumferential lesion, while in a partial laceration, it appears as localized diverticulum (Fig. 33-11). Periaortic hemorrhage may occur irrespective of the type of lesion. Complete rupture of the aorta, including the adventitial layer and periadventitial connective tissue, leads to immediate exsanguination. However, false aneurysm or occlusion of the side of rupture may permit temporary survival. A long examination time and difficult access to the polytraumatized patient have been considered to be the main limitations of CMR in acute aortic pathology. The development of fast CMR techniques has shortened the examination time to a few minutes; therefore, MRI can be used even in critically ill patients.82

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Figure 33-11 Contrast-enhanced MRA of chronic traumatic aortic lesion. A, Partial laceration of the aortic wall results in a diverticular aneurysm (shaded surface rendering). B, Discharge imaging after endovascular stent-graft repair reveals complete restitution of thoracic aortic integrity. Cardiovascular Magnetic Resonance 457

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of thoracic aortic aneurysms, cost considerations have to be evaluated before recommendation of widespread use of MRI scanning to follow stable patients with subcritical aneurysms.

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The value of CMR in detecting traumatic aortic rupture in comparison to angiography and conventional CT was reported in a series of 24 consecutive patients.83 The potential for CMR to detect the hemorrhagic component of a lesion by its high signal intensity is beneficial in traumatized patients. On spin echo images in the sagittal plane, a longitudinal visualization of the thoracic aorta makes it possible to distinguish a partial lesion from a lesion encompassing the entire aortic circumference. This discrimination is of prognostic significance because a circumferential lesion may be more likely to rupture.83 The presence of periadventitial hematoma and/or pleural and mediastinal hemorrhagic effusion may also be considered a sign of instability. In the same sequence used to evaluate the aortic lesion, without the need of any additional time, the wide field of view of CMR provides a comprehensive evaluation of chest trauma such as lung contusion and edema and pleural effusion and rib fractures. Furthermore, if delayed surgery is considered, CMR may be used to monitor thoracic and aortic lesions because it is noninvasive and repeatable. MRA provides an excellent display of the aortic lesion and its relationship to supra-aortic vessels. However, it does not add any diagnostic value to spin echo CMR, and it cannot supply information on parietal lesions and hemorrhagic fluids outside the aortic vessel. Recently, the development of endovascular technique has provided additional opportunities in the treatment of acute onset and chronic traumatic aortic disease (see Fig. 33-11).84–88

Aortic Coarctation Coarctation is a common congenital anomaly, with an incidence of 20 to 60 per 100,000 live births, and represents 5% to 8% of all congenital cardiovascular disorders. The obstructive lesion results from an abnormality in the aortic media and refers to an enfolding of the posterolateral aortic wall in the region of the ligamentum or ductus arteriosus. This is usually a discrete phenomenon that occurs just distal to the ductus and is also labeled postductal coarctation. Since it is usually asymptomatic in the neonatal period, it is also referred to as adult coarctation. Coarctation may occur proximal to the ductus and may present itself shortly after birth; this variety has been termed preductal or infantile coarctation. It is less common than adult coarctation and is usually associated with hypoplasia of the arch between the left subclavian artery and the ductus of the aortic isthmus. There is usually a dilation of the descending aorta distal to the coarctation. As a result of the obstruction caused by the coarctation, collateral vessels develop to increase flow into the descending aorta. The intercostal arteries serve as a major source of collateral flow. The increased flow through these vessels results in their dilation. This, in turn, can result in notching along the inferior aspect of the ribs, which usually takes 8 to 10 years to become significant enough to be observed on a chest radiograph. With its multiplanar image acquisition, large field of view, and dynamic quantitative flow imaging capacity, CMR appears the modality of choice for evaluation of coarctation (Fig. 33-12). The first step in quantification of the disease is to obtain an anatomic display of the extent and severity of the stenotic segment, and spin echo images 458 Cardiovascular Magnetic Resonance

A

B

Figure 33-12 A, MRA of aortic coarctation depicts focal stenotic segment (arrow) at the isthmic zone and multiple intercostal collateral arteries joining the descending aorta. B, Medium-sized Dacron patch aneurysm (arrow) diagnosed 30 years after coarctation surgery.

well depict the morphologic features of the coarctation as reported first by Amparo and colleagues.89 The left oblique sagittal view centered on the middle of the ascending and descending aortas is an ideal orientation that may also demonstrate associated aortic stenosis, left ventricular hypertrophy, and ventricular septal defects. This is important because there is a high association of bicuspid aortic valve and ventricular septal defects with coarctation. The severity of the stenosis can be expressed as the ratio of the diameters or cross-sectional areas measured at the coarcted segment and above the diaphragm.90 However, although the anatomic narrowing of the aorta establishes the diagnosis of coarctation, an assessment of its clinical significance depends on determining its hemodynamic effects. Cine CMR has been applied to evaluate flow turbulence across the coarctation; the severity of coarctation is quantified on the basis of the length of flow void.91 Further functional information can be provided by CMR flow mapping, which can define the severity of the stenosis by measuring velocity jets at the level of coarctation and mean flow deceleration in the descending aorta.92,93 With this technique, it is possible to predict the coarctation severity with good sensitivity and specificity (95% and 82%, respectively) compared with catheter angiography.94–96 Flow mapping is also able to quantify the flow pattern and volume of collateral flow in the descending aorta,97 which are other important parameters of the severity of coarctation, and this information may be crucial in the choice of surgical strategy. The severity of aortic coarctation can be measured with flow mapping MRI using two different methods. In the first method, flow mapping CMR is used to quantify collateral blood flow.96 Normally, flow volume in the proximal descending aorta is slightly (about 7%) higher than in the distal portion, but in coarctation, there may be greater blood flow distally, owing to retrograde collateral blood flow into the aorta arising from the intercostal arteries and other aortic branches. Flow volume is estimated at two sites in the aorta: one just below the

A

B

endovascular stent-grafting, have currently been applied to the treatment of postsurgical patch aneurysms with excellent results, avoiding the need for further surgical intervention.113,114

Aortitis Although there are many causes of aortitis, Takayasu arteritis is the type most often studied with CMR because of the diffuse stenotic nature of the disease, which often makes vascular access impossible by catheterization.115 In Takayasu disease, the aortic arch vessels are primarily affected, but thoracic and abdominal aorta may also be involved (Fig. 33-13). Active inflammatory disease demonstrates diffuse thickening of the aortic wall, typically enhanced after gadolinium administration in T1-weighted images. The chronic stage, however, is characterized by extensive perivascular fibrosis without postcontrast enhancement.116 MRA is the preferred imaging modality for the study of branch vessels stenosis and has replaced invasive angiography, which carries a risk of pseudoaneurysm formation at the site of arterial puncture. Recently, the capability of F-18 FDG hybrid camera positron emission tomography combined with CMR to detect early stages of Takayasu aortitis has been demonstrated.117

INTERVENTIONAL CARDIOVASCULAR MAGNETIC RESONANCE IMAGING CMR guidance of vascular interventional procedures offers several potential advantages over fluoroscopy-guided techniques, including image acquisition in any desired orientation, superior 3D soft tissue contrast with simultaneous

C

Figure 33-13 Takayasu arteritis. A, Coronal MIP 3D MRA shows significant stenosis of right common carotid artery at its origin (small arrow). The left subclavian artery has two stenotic segments (large arrows) with a small area of poststenotic dilatation in between. There are also some luminal irregularities of the left common carotid artery. Axial unenhanced (B) and gadolinium-enhanced (C) T1-weighted CMR shows wall thickening of the ascending aorta (arrow), which is enhanced on the gadolinium-enhanced image. Cardiovascular Magnetic Resonance 459

33 THORACIC AORTIC DISEASE

coarctation site and the other at the level of the diaphragm. Collateral flow is present if distal aortic flow is greater than proximal flow and it is possible to quantify collateral circulation subtracting flow volume in the proximal descending aorta from that in the distal portion. The presence of collateral flow indicates a hemodynamically significant coarctation. In the second method, flow mapping CMR is used to estimate peak flow velocity through the most severely narrowed segment; the pressure gradient across this site can be calculated with the modified Bernoulli equation (P ¼ 4V2).95 The pressure gradient calculated with flow mapping CMR can be used to determine the need for surgical/endovascular repair (a pressure gradient greater than 15 mm Hg is considered an indication for intervention). However, this threshold is arbitrary; therefore, it may be more appropriate to use the measurement of collateral flow. Several therapeutic strategies are available for the treatment of aortic coarctation, depending on the morphology of the affected aorta as well as the age and clinical condition of the patient.98,99 Surgery for aortic coarctation is recommended at an early age because long-term results seem to be better. Recently, interventional procedures and balloon angioplasty have come into wide use and provide good results, especially in mild or moderate cases.100,101 An accurate selection of favorable anatomy by high-resolution imaging modalities is particularly important in interventional procedure to ensure a low rate of complications and restenosis.102 An increased risk for aneurysm formation at the site of repair has been reported after both synthetic patch aortoplasty and subclavian-flap arterioplasty (see Fig. 33-12).103–105 Moreover, restenosis, aortic dissection, and pseudoaneurysms have been reported after surgery or balloon angioplasty in up to 42% of patients.106–109 Therefore, routine follow-up is recommended for patients who underwent repair of an aortic coarctation, independently of surgical technique used and timing of the repair.110–112 New interventional techniques, such as

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visualization of the interventional device, absence of ionizing radiation, and avoidance of nephrotoxic contrast media. The feasibility of real-time CMR-guided interventions has been demonstrated for a wide range of vascular interventional procedures in animals.118–125 For peripheral stent placement, the feasibility of CMR guidance has also been shown in patients.126 In terms of endovascular aortic stent-graft placement, MRI appears to be particularly useful, as it can provide preinterventional evaluation of aortic pathology, real-time interventional image guidance, and immediate evaluation of treatment success or procedurerelated complications as well as follow-up examinations.127,128 Passive device tracking using material-induced susceptibility artifacts of the interventional instrument for CMR visualization requires no hardware or instrument modifications and appears to be promising in terms of potential clinical applications. However, MR-compatible instruments with satisfactory susceptibility artifacts are required, as well as CMR-compatible guide wires with adequate mechanical support. To date, the limitations of interventional CMR are long procedure times, lack of true realtime monitoring, and stent artifacts, which necessitate further modifications before they can be recommended for

clinical use.129–133 However, MRI-guided interventions are feasible and may become an important advance in the near future.

CONCLUSION With recent advances in the understanding of aortic diseases, both power and versatility have put magnetic resonance imaging in the focus of diagnostic workup in the entire spectrum of clinical aortic pathology. Technical refinements, from classic anatomic imaging to 3D gadoliniumenhanced MRA and tissue characterization, have rendered CMR ideal for assessment of acquired disease such as aortic dissection, intramural hematoma, and aneurysm, along with postoperative follow-up evaluation, with better reliability and safety than other imaging modalities. Moreover, congenital pathology of the aorta, including aortic arch anomalies and coarctation, can be noninvasively evaluated by MRI. The guidance of vascular interventional procedures by CMR offers potential advantages over fluoroscopy-guided techniques and may become an important advance in the near future.

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109. Serfontein SJ, Kron IL. Complications of coarctation repair. Semin Thorac Cardiovasc Surg Pediatr Card Surg Annu. 2002;5:206–211. 110. Bogaert J, Kuzo R, Dymarkowski S, et al. Follow-up of patients with previous treatment for coarctation of the thoracic aorta: comparison between contrast-enhanced MR angiography and fast spin-echo MR imaging. Eur Radiol. 2000;10(12):1847–1854. 111. Celermajer DS, Greaves K. Survivors of coarctation repair: fixed but not cured. Heart. 2002;88(2):113–114. 112. Therrien J, Thorne SA, Wright A, et al. Repaired coarctation: a “costeffective” approach to identify complications in adults. J Am Coll Cardiol. 2000;35(4):997–1002. 113. Bell RE, Taylor PR, Aukett M, et al. Endoluminal repair of aneurysms associated with coarctation. Ann Thorac Surg. 2003;75(2):530–533. 114. Ince H, Petzsch M, Rehders T, et al. Percutaneous endovascular repair of aneurysm after previous coarctation surgery. Circulation. 2003;108(24):2967–2970. 115. Nastri MV, Baptista LP, Baroni RH, et al. Gadolinium-enhanced threedimensional MR angiography of Takayasu arteritis. Radiographics. 2004;24(3):773–786. 116. Choe YH, Kim DK, Koh EM, et al. Takayasu arteritis: diagnosis with MR imaging and MR angiography in acute and chronic active stages. J Magn Reson Imaging. 1999;10(5):751–757. 117. Meller J, Grabbe E, Becker W, et al. Value of F-18 FDG hybrid camera PET and MRI in early takayasu aortitis. Eur Radiol. 2003; 13(2):400–405. 118. Buecker A, Spuentrup E, Grabitz R, et al. Real-time-MR guidance for placement of a self-made fully MR-compatible atrial septal occluder: in vitro test. Rofo. 2002;174(3):283–285. 119. Buecker A, Adam GB, Neuerburg JM, et al. Simultaneous real-time visualization of the catheter tip and vascular anatomy for MR-guided PTA of iliac arteries in an animal model. J Magn Reson Imaging. 2002;16(2):201–208. 120. Buecker A, Neuerburg JM, Adam GB, et al. Real-time MR fluoroscopy for MR-guided iliac artery stent placement. J Magn Reson Imaging. 2000;12(4):616–622. 121. Fink C, Bock M, Umathum R, et al. Renal embolization: feasibility of magnetic resonance-guidance using active catheter tracking and intraarterial magnetic resonance angiography. Invest Radiol. 2004;39 (2):111–119. 122. Raval AN, Karmarkar PV, Guttman MA, et al. Real-time magnetic resonance imaging-guided endovascular recanalization of chronic total arterial occlusion in a swine model. Circulation. 2006;113 (8):1101–1107. 123. Raval AN, Telep JD, Guttman MA, et al. Real-time magnetic resonance imaging-guided stenting of aortic coarctation with commercially available catheter devices in swine. Circulation. 2005;112 (5):699–706. 124. Raman VK, Karmarkar PV, Guttman MA, et al. Real-time magnetic resonance-guided endovascular repair of experimental abdominal aortic aneurysm in swine. J Am Coll Cardiol. 2005;45(12):2069–2077. 125. Spuentrup E, Ruebben A, Schaeffter T, et al. Magnetic resonance– guided coronary artery stent placement in a swine model. Circulation. 2002;105(7):874–879. 126. Manke C, Nitz WR, Djavidani B, et al. MR imaging-guided stent placement in iliac arterial stenoses: a feasibility study. Radiology. 2001;219(2):527–534. 127. Mahnken AH, Chalabi K, Jalali F, et al. Magnetic resonance-guided placement of aortic stents grafts: feasibility with real-time magnetic resonance fluoroscopy. J Vasc Interv Radiol. 2004;15(2 Pt 1):189–195. 128. Eggebrecht H, Kuhl H, Kaiser GM, et al. Feasibility of real-time magnetic resonance-guided stent-graft placement in a swine model of descending aortic dissection. Eur Heart J. 2006;27(5):613–620. 129. Eggebrecht H, Zenge M, Ladd ME, et al. In vitro evaluation of current thoracic aortic stent-grafts for real-time MR-guided placement. J Endovasc Ther. 2006;13(1):62–71. 130. Konings MK, Bartels LW, Smits HF, et al. Heating around intravascular guidewires by resonating RF waves. J Magn Reson Imaging. 2000;12(1):79–85. 131. Ladd ME, Quick HH, Debatin JF. Interventional MRA and intravascular imaging. J Magn Reson Imaging. 2000;12(4):534–546. 132. Nitz WR, Oppelt A, Renz W, et al. On the heating of linear conductive structures as guide wires and catheters in interventional MRI. J Magn Reson Imaging. 2001;13(1):105–114. 133. Quick HH, Zenge MO, Kuehl H, et al. Interventional magnetic resonance angiography with no strings attached: wireless active catheter visualization. Magn Reson Med. 2005;53(2):446–455.

Cardiovascular Magnetic Resonance Angiography: Carotids, Aorta, and Peripheral Vessels Robert R. Edelman, James W. Goldfarb, and Agnes E. Holland

Cardiovascular magnetic resonance angiography (MRA) has become widely accepted for the evaluation of abnormalities throughout multiple vascular territories. Advances have allowed MRA to progress rapidly from a developmental technique to a noninvasive clinical imaging tool providing vital information in the clinical care of patients every day. Blood that flows through magnetic field gradients and radiofrequency (RF) fields produces signal changes that can be used to distinguish blood vessels from stationary surrounding tissue. Pulse sequences that exploit the effect of blood motion to directly visualize blood vessels, without the use of a contrast agent, include the time-of-flight (TOF)1 and phase contrast (PC)2 MRA techniques. These noncontrast MRA approaches are excellent for imaging of vessels in healthy subjects with normal, continuous, laminar flow. However, in patients with vascular disease that disturbs the laminar pattern and direction of flow, the quality of TOF and PC images is degraded. By administering a paramagnetic contrast agent intravenously, the depiction of vessels is no longer dependent on blood inflow and motion. The use of contrast agents alleviates many of the problems encountered with flow-dependent MRA techniques.3,4 The paramagnetic contrast agents are injected intravenously, and the image data are collected during first passage of the contrast agent through the vascular territory of interest, depicting the blood vessels with a brief high signal intensity. This allows imaging of a large field of view (FOV) that encompasses an extensive region of vascular anatomy. The development of high-powered gradient technology has resulted in significantly shorter acquisition times, making it possible to acquire an entire three-dimensional (3D) highresolution volume dataset in a single breath hold using fast gradient echo techniques. Contrast-enhanced MRA (CE-MRA) also has many advantages over conventional X-ray angiography. It is noninvasive, providing high-resolution images without the need for arterial access. This may account for the popularity and widespread use of CE-MRA throughout the world. Another advantage is that the paramagnetic contrast agents used for CE-MRA have a much more favorable safety profile than do the conventional iodinated contrast agents used in X-ray angiography,5,6 though nephrogenic systemic fibrosis is a concern for patients with impaired renal function.7–9 (See Chapter 6.)

This chapter provides a brief review of the underlying basic principles of MRA and an overview of the various MRA techniques and their application, with examples in clinical practice.

BASIC PRINCIPLES AND TECHNIQUES The essence of MRA is creating high contrast between flowing blood and surrounding stationary tissue. To distinguish vessels from surrounding stationary tissue, either the intrinsic mechanisms of blood flow can be used, or a contrast agent can be added intravenously. Techniques depending on the intrinsic mechanisms of blood flow include TOF and PC angiography. The gadolinium-enhanced 3D techniques involve the addition of contrast material to visualize blood vessels.

Time-of-Flight Cardiovascular Magnetic Resonance Angiography TOF methods depend on the inflow enhancement of flowing blood and the relative saturation of stationary tissue to produce high contrast. The signal from the background tissue is suppressed by multiple RF pulses, such that the magnetic spins associated with background tissue do not have enough time to regain their longitudinal magnetization. This saturation of spins from the stationary tissue makes stationary tissue appear dark in TOF images. The bright signal intensity of flowing blood is the consequence of the continuous inflow of fresh, unsaturated blood into the imaging slice, which produces more signal than the surrounding stationary tissue, as the latter was repeatedly exposed to RF pulses. This effect is known as flow-related enhancement.10,11 The flow-related enhancement can be maximized by using sequences with a short repetition time (TR). Ideally, the imaging slice should be thin and oriented perpendicular to the direction of flow. Electrocardiographic (ECG) gating can be used to optimize the flow and consequently the blood signal in arteries.

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CHAPTER 34

VASCULATURE AND PERICARDIUM

Phase Contrast Magnetic Resonance Angiography In PC angiography, the contrast between flowing blood and the surrounding stationary tissue is generated by the flowinduced phase shift of the moving blood. When a magnetic spin in flowing blood is subjected to a gradient magnetic field, it acquires a phase shift that is proportional to its velocity.12–14 In the presence of a pair of bipolar gradient lobes, the stationary spins do not develop a net phase shift, whereas spins moving along the direction of the magnetic field gradient develop a net phase shift. Because the flow-induced phase shifts are proportional to the signal intensity, flow direction and velocity can be derived directly from PC data.

Gadolinium-Enhanced ThreeDimensional Cardiovascular Magnetic Resonance Angiography Instead of relying on blood flow, as with TOF and PC techniques, the gadolinium-enhanced MRA techniques utilize the addition of a gadolinium chelate to create intravascular signal. The gadolinium chelates are paramagnetic contrast agents that preferentially shorten the T1 relaxation time of blood in proportion to their local concentration. With this strategy, blood can be imaged irrespective of flow and image contrast is based on T1 relaxation rather than on flow effects. Therefore, flow artifacts and in-plane saturation effects, which occur with the TOF and PC techniques, are largely eliminated.3 In addition, these gadolinium-enhanced techniques allow in-plane imaging of vessels, so the number of image sections required to image an extensive region of vascular anatomy is strongly reduced.15

Gadolinium Chelates Gadolinium is a paramagnetic ion that shortens the longitudinal or spin lattice (T1) relaxation times of nearby protons.16 (See Chapter 6.) When added to blood, it shortens the T1 relaxation time of blood according to Equation 1: 1=T1 ¼ 1=1200 msec þ R1 ½Gd

(Equation 1)

1200 msec ¼ T1 of blood without gadolinium @ 1:5 T R1 ¼ relaxivity of gadolinium ½Gd ¼ gadolinium concentration in the blood Gadolinium itself is toxic and must be bound to a chelator, such as diethylenetriamine pentaacetic acid (DTPA), to be used safely in humans. Most of the currently available U.S. Food and Drug Administration (FDA) -approved gadolinium chelates are extracellular agents. They readily diffuse through the capillary walls into the extravascular space. Consequently, imaging of arteries has to be done during the initial arterial phase of the contrast injection to minimize signal from gadolinium that crossed into surrounding stationary tissues and/or passed through the capillary system and into veins. 464 Cardiovascular Magnetic Resonance

Gadolinium chelates can be transiently bound to larger molecules, such as albumin, that do not readily pass through the capillary membranes. This binding reduces the tumbling rate, with the consequence that such agents (e.g., gadobenate dimeglumine) demonstrate higher relaxivity than standard gadolinium chelates and can improve the quality of the MRA. Another kind of contrast agent is the so-called blood pool agent, of which one agent, gadofosveset trisodium (Lantheus Medical Imaging, N. Billerica, MA), is approved for CE-MRA in the United States and Europe. Ultra-small iron oxide particles also have potential as blood pool contrast agents (Fig. 34-1). The safety profile of the gadolinium chelates is much more favorable than that of the conventional iodinated contrast agents.5,6,17 With gadolinium chelates, the incidence of acute adverse events is extremely low, and idiosyncratic reactions are very low.4 In addition, gadolinium contrast agents have no nephrotoxicity, even when administered at high doses.18–20 This is particularly relevant for the evaluation of patients with renal failure. However, long-term toxicity due to nephrogenic systemic fibrosis must be considered in patients with impaired renal function who are exposed to gadolinium chelates.

Three-Dimensional Cardiovascular Magnetic Resonance Angiography Imaging Pulse Sequences A fast 3D spoiled gradient echo sequence is used, with TR ¼ 3 to 8 msec, TE ¼ 1 to 3 msec, and flip angle ¼ 20 to 40 . The imaging volume can be acquired in any desired orientation and the acquisition time historically ranged from 20 to 40 seconds. As will be considered later in the chapter, scan times as short as 2 seconds can now be obtained by using parallel imaging techniques, so breath hold capability is seldom a limitation any more. Scan times on this order allow imaging to be done during a prolonged breath hold. At postprocessing, the 3D nature of the dataset allows viewing of the data from any desired angle, yielding multiple images in various anatomic orientations. This assumes that thin (i.e., <3 mm thick) slices are acquired. In some situations, it may be useful to use thicker slices (e.g., 10 mm) to reduce the number of slices and allow faster scanning. Although the ability to rotate the maximum intensity projection (MIP) is lost, scan times can be as short as 1 second.21 The high temporal resolution enables the depiction of organ perfusion and time-resolved imaging of high-flow vascular lesions.

Vessel Brightness (T1) Versus Gadolinium Dose An adequate amount of paramagnetic contrast agent should be administered to ensure that the T1 of blood is well below the T1 of the brightest surrounding tissue. This ensures that blood will appear bright when compared with background tissue. Because the background tissue with the shortest T1 relaxation time is usually fat (T1 ¼ 280 msec at 1.5 T), a sufficient dose of gadolinium chelate should be injected at a sufficient rate to ensure that the T1 relaxation time of the arterial blood is well under 280 msec.15 The blood T1 relaxivity for various doses of gadolinium chelates can be calculated from Equation 1 and plotted. (The currently available gadolinium chelates have a relaxivity of approximately 4.5 mM1sec1 at 1.5 T.)

½Gd ¼

ðrate of Gd injection in mol=secÞ ðcardiac output in L=secÞ

(Equation 2)

The blood T1 for a given cardiac output and injection rate can be calculated by using Equations 1 and 2.

Bolus Timing

Figure 34-1 Whole body MRA obtained over 5-minute acquisition using ferumoxytol (Advanced Magnetics, Inc., Cambridge, MA), an ultrasmall iron particle that behaves as a blood pool agent (intravascular half–life: 14 hours).

Currently available FDA-approved gadolinium contrast agents can readily cross the capillary membranes. A few minutes after the injection, there is already a redistribution of the contrast agent into the extracellular space. Therefore, gadolinium-enhanced MRA should be performed during the early arterial phase (first pass) of the contrast agent. The timing of the bolus must be very precise such that imaging coincides with the arrival of the contrast agent bolus in the vascular territory of interest.

Proper delay between the start of the contrast agent injection and the start of the cardiovascular magnetic resonance (CMR) scan is crucial. For maximum arterial enhancement, the contrast injection has to be timed such that the acquisition of the central lines of k-space coincides with the peak arterial gadolinium concentration. If the central lines of k-space are inadvertently acquired before the peak arterial enhancement, a coarse “ringing” artifact combined with “widening” of the vessel margins occurs.23 When the central lines of k-space are acquired too late (i.e., after the peak arterial enhancement), there is reduced arterial signal intensity and increased venous signal intensity, resulting in a sometimes confusing overlap between arterial and venous structures. This reduces the accuracy of the imaging modality. There are different types of CMR pulses sequences, with the central lines of k-space acquired in the very beginning of the scan, after 25% of the scan, and after 50% of the scan. To time the gadolinium injection bolus correctly, the exact transit time of the contrast agent from the injection site (typically in the arm) to the blood vessels of interest has to be determined. The transit time of the contrast agent depends on several factors, including stroke volume, heart rate, valvular function, injection rate, and catheter site. Currently, there are several different methods to determine this transit time. The test bolus technique involves the injection of a small dose (1 to 2 mL) of gadolinium, at the same rate as the actual contrast injection, followed by a 10- to 15-mL saline flush.24,25 A thick two-dimensional (2D) gradient echo section is acquired every 1 to 2 seconds through the vascular territory of interest for approximately 1 minute. The contrast transit time can be determined visually or by drawing a region of interest and measuring the signal intensity on sequential images. Another way of timing the arrival of the contrast agent bolus is through the use of automatic triggering.26,27 This technique involves automatic synchronization of central k-space image data with the arterial phase of contrast material bolus infusion (MR Smartprep, GE Medical Systems, Milwaukee, WI; MobiTrak, Phillips Medical Systems, The Netherlands). A spin echo (SE) sequence with orthogonal 90 and 180 is used to monitor the signal intensity in a single large, 4 4 12 cm voxel, which is placed in the vascular territory of interest. The arrival of the gadolinium bolus corresponds with an increase of signal intensity in the large voxel. Either a technician or a computer program senses the arrival of the gadolinium and triggers the 3D gradient echo sequence. Cardiovascular Magnetic Resonance 465

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Assuming adequate mixing of the contrast agent and blood, the arterial blood gadolinium concentration is related to the injection rate and cardiac output as follows:3,22

VASCULATURE AND PERICARDIUM

The time-resolved 3D CE-MRA is a technique that is virtually independent of the exact contrast transit time.28 It involves the rapid collection of successive 3D datasets beginning immediately after contrast injection so that at least one of the 3D datasets aligns with the arterial phase of the contrast injection. With this technique, acquisition times for acquiring each 3D dataset are on the order of 2 to 10 seconds. The initiation of the injection and initiation of the first image acquisition are simultaneous, and multiple vascular phases are collected (i.e. arterial, tissue perfusion, venous) as the contrast agent flows through the vascular bed of interest. In addition to simplifying bolus timing, this technique allows evaluation of the temporal enhancement of structures, such as the kidneys, that are caught within the 3D volume. Since organ function depends on its level of perfusion, regional blood flow information might aid in the assessment of hemodynamically significant stenoses.

Parallel Imaging CE-MRA techniques require strong, fast magnetic field gradients to achieve the short repetition and echo times. However, the engineering complexity and cost of the gradients are reaching practical limits. In addition, neuromuscular stimulation may occur if the gradients are switched at too rapid a rate. An alternative means to shorten scan time is to use parallel imaging.29,30 (See Chapter 3.) Parallel imaging uses a property of phased array RF coils, specifically that the sensitivity to tissue signals varies with distance from each element of the coil, to accelerate the scan process. For each k-space line that is encoded by the magnetic field gradients, one or more additional lines is encoded using the phased array coil. Acceleration factors ranging from 2 to 32 have been obtained by this approach. There is a signal-to-noise (SNR) penalty, which worsens with increasing acceleration factor, and image artifacts may also be introduced, depending on the type of RF coil. Nonetheless, the method has proved robust for a variety of CE-MRA applications.31–33

Cardiovascular Magnetic Resonance Angiography at 3 Tesla Imaging at 3 T provides a near twofold boost in SNR for MRA compared with 1.5 T and results in improved image quality.34,35 The extra SNR can be used to improve spatial resolution by a factor of two or scan time by a factor of four. Benefits are seen for both TOF and CE-MRA. Imaging at 3 T also entails certain challenges, such as a fourfold higher rate of power deposition and increased artifact from magnetic susceptibility. Nonetheless, unlike the situation for CMR, in which imaging is problematic for certain imaging techniques, the challenges for MRA are relatively minor and easily overcome for clinical studies. For instance, we have used a combination of a spatially nonselective RF excitation, which uses much less power than a standard slab-selective excitation, in combination with parallel imaging to obtain high-resolution MRA of the abdomen at 3 T with striking image quality (Fig. 34-2). 466 Cardiovascular Magnetic Resonance

Figure 34-2 Breath hold 3D MRA at 3 T provides striking detail of the aorta and branch vessels. A spatially nonselective RF excitation was used to minimize power deposition at 3 T, and parallel imaging was used to shorten scan time.

THREE-DIMENSIONAL CONTRAST-ENHANCED MAGNETIC RESONANCE ANGIOGRAPHY COMPARED TO THREE-DIMENSIONAL CONTRAST-ENHANCED COMPUTED TOMOGRAPHY ANGIOGRAPHY MRA and computed tomography angiography (CTA) are in some respects quite similar because each acquires image data in a tomographic manner. Three-dimensional CEMRA does have several advantages over contrast-enhanced CTA (CE-CTA). Both methods use an intravenous injection of a contrast agent and make images during a short period of time, when the agent is in the vessels of interest. The major difference is how the image is made and the side effects of the injected contrast agent. MRA and CTA images are made by using a magnetic field and X-ray beams, respectively. The MRA dataset may be acquired in any orientation, while the CTA is acquired in the transverse plane. The paramagnetic agents that are used for contrastenhanced MRA have proven to be much safer than the iodinated agents. The incidence of anaphylaxis with paramagnetic agents is rare.5,6 The paramagnetic contrast agents are not nephrotoxic and can be administered safely to patients with mild to moderately impaired renal function. As will be discussed later in the chapter in relation to renal artery stenosis, CE-MRA allows angiographic investigation without grave concern for inducing renal failure in this

CARDIOVASCULAR MAGNETIC RESONANCE ANGIOGRAPHY OF THE EXTRACRANIAL CAROTID ARTERIES Surgical intervention (and, more recently, endovascular stenting) is recognized to be beneficial in certain patient groups with extracranial carotid disease.39,40 Whereas Xray angiography was the gold standard for the NASCET (North American Symptomatic Carotid Endarterectomy Trial) study, it involves a small stroke risk on the order of 0.5% to 4%.41 Consequently, there is a strong need for an accurate noninvasive test for extracranial carotid disease. Although duplex sonography is an accurate test in experienced hands and less costly than DSA or MRA, the accuracy is highly operator-dependent. MRA is an accurate test in most centers and is now widely used to evaluate both intracranial and extracranial carotid disease. TOF MRA can be divided into 2D and 3D approaches (Fig. 34-3). The 2D technique involves the acquisition of a series of contiguous or overlapping thin (e.g., 2 mm thick) slices in an axial plane. Superior tracking presaturation eliminates venous inflow signal. Scan times are on the order of 5 minutes (or half that with twofold parallel

Figure 34-3 Comparison of digital subtraction angiography (left) with maximum intensity projection from 2D TOF (middle) and 3D TOF (right) CMR in a patient with stenosis of the proximal internal and external carotid arteries. The stenoses are exaggerated in the 2D TOF acquisition in comparison with 3D TOF and DSA.

acceleration). 2D TOF is sensitive to slow flow but tends to exaggerate the severity of a stenosis. The 3D TOF technique involves data acquisition sequentially over several thick volumes, which are then phase-encoded in the slice direction to produce a series of contiguous thin slices. Three-dimensional TOF provides greater specificity but suffers from signal loss when flow is slow or in-plane. For contrast-enhanced imaging, there are also two general approaches. The first approach involves a relatively lengthy acquisition (e.g., 30 to 60 seconds) and uses ellipticocentric phase encoding to minimize venous enhancement. The second approach involves a shorter time-resolved acquisition (e.g., 5 to 15 seconds). Time-resolved images may be acquired by using parallel imaging or the time-resolved imaging of contrast kinetics (TRICKS) technique. Each approach has certain benefits. The ellipticocentric approach allows scans to be acquired with high spatial resolution. However, the great vessel origins are not always well shown, owing to breathing motion. The time-resolved acquisition is short enough to allow breath holding so the vessel origins are well shown. However, spatial resolution tends to be lower. Our own preference is for the time-resolved technique; especially since high parallel acceleration factors now enable spatial resolution to be boosted nearly to the levels offered by 3D TOF. Although not very useful for body applications, TOF MRA is sensitive and specific for carotid bifurcation disease and often provides superior spatial resolution in comparison with CE-MRA.42 Some investigators report better specificity for TOF in comparison with CE-MRA (97% for TOF versus 81% for CE-MRA).43 Nonetheless, CE-MRA is also highly accurate (Fig. 34-4). Willinek and colleagues reported excellent sensitivity (100%) and specificity (99%) for 70% or greater diameter stenosis by CE-MRA.44 Clearly, technical factors play a substantial role in the accuracy of the two approaches. From a practical standpoint, Cardiovascular Magnetic Resonance 467

34 CARDIOVASCULAR MAGNETIC RESONANCE ANGIOGRAPHY: CAROTIDS, AORTA, AND PERIPHERAL VESSELS

important patient population.18–20 However, for patients with eGFR<30 mL/min/1.73m2, nephrogenic systemic fibrosis may emerge weeks to months after gadolinium injection.36–38 CTA data are acquired exclusively in the transverse plane, while MRA allows imaging to be performed in the sagittal and coronal and flexible oblique imaging planes. This provides greater volume coverage. Conventional CMR units allow maximal FOVs from 400 to 500 mm. This feature is particularly important for the evaluation of the extent of aortic disease, both proximally into the thorax and distally into the iliac arteries. MRA can depict vessels in their full extent at high resolution with fewer slices than can CTA, which requires many thin sections for adequate distinction of the vascular anatomy. This illustrates another disadvantage of CTA: the relatively large radiation dose to which the patient is exposed when these multiple thin sections are being obtained. In contrast to CTA, MRA can also provide functional information, such as flow velocity and flow volume. These measurements are not made by using contrast-enhanced techniques. Rather, they can be calculated by using PC sequences. With the gated cine sequences, the blood flow can be displayed dynamically during the cardiac cycle. MRA data can be easily reconstructed and displayed using the multiplanar reconstruction and MIP algorithms. Postprocessing CE-CTA images is considerably more complicated. It involves segmenting out structures on the basis of their density, such as bone, that obscure the vascular anatomy from each individual slice. This can be a tedious and lengthy procedure. One major clinical advantage of CTA over MRA is that the easier setup in a nonmagnetic environment allows clinically unstable patients to be more comfortably studied with CTA than with MRA.

VASCULATURE AND PERICARDIUM

Figure 34-4 Carotid CE-MRA showing a right carotid artery stenosis at the level of the carotid bifurcation. A targeted maximum intensity projection (MIP) reconstruction shows the severity of the stenosis from multiple angles. In the lower right corner, a coronal MIP of the 3D dataset shows the vasculature of the neck and head.

CE-MRA is much faster than TOF and less motion sensitive, and the vessel origins are better evaluated. Improvements in parallel imaging methodology should enable CEMRA to match or exceed the spatial resolution of TOF methods and thereby further improve specificity.45 Imaging at 3 T will likely further improve the accuracy of both TOF and CE-MRA techniques, owing to the near twofold boost in SNR.46,47 In addition to its utility for carotid stenosis, CE-MRA is a robust diagnostic method for other extracranial carotid pathology, such as dissection.48 In cases of suspected dissection, it is important to acquire a set of fat-suppressed T1-weighted images in addition to CE-MRA so as to detect blood products in the vessel wall. CMR offers capabilities for evaluating carotid stenosis beyond simply depicting the cross-sectional area of the vessel lumen. Although the eventual clinical impact is not yet known, magnetic resonance (MR) has already proven capable of imaging and characterizing various components of carotid plaque, including hemorrhage, lipid, and fibrous cap.49 (See Chapter 25.) For instance, ipsilateral cerebral infarcts were found to be more common in patients having a lipid core (decreased T2 signal) within plaque at the carotid bifurcation than in those without.50

MAGNETIC RESONANCE ANGIOGRAPHY OF THE BODY Aorta MRA of the aorta is among the most common studies performed. Various radiologic imaging modalities have been employed for the assessment of disease of the thoracic and abdominal aorta (See Chapter 33.) For the visualization of the thoracic aorta, plain film radiography, CT, transesophageal echocardiography, and conventional X-ray aortography are all established techniques. For imaging of the abdominal aorta, conventional angiography, ultrasound and CT are widely used. Of these three techniques, CT is 468 Cardiovascular Magnetic Resonance

currently the most widely used, because of its widespread availability, high throughput, low cost and diagnostic accuracy. Recently, CE-MRA has emerged as a robust and reliable technique for imaging both the thoracic and abdominal aorta.51–54 Indications for evaluating the aorta include occlusive .disease in patients with generalized atherosclerosis,55,56 suspected aneurysm,57 dissection,58 follow-up surgical repair, and congenital malformations such as coarctation. Both noncontrast and CE-MRA are used to image the aorta. The non-contrast-enhanced techniques include the black-blood SE sequences and the bright-blood gradient recalled echo (GRE) sequences. The development of turbo spin echo or fast spin echo (FSE) techniques has significantly reduced the scan time and is quickly replacing SE imaging. For FSE and GRE imaging of the thoracic aorta, ECG gating is necessary to reduce artifacts associated with the motion of the heart during the cardiac cycle.54 In contrast, imaging of the abdominal aorta is not directly affected by cardiac motion, and ECG gating is usually not required. However, studies of the abdominal aorta may be subject to pulsation artifacts, which can be reduced by the use of spatial saturation bands for SE techniques and flow compensation algorithms for GRE techniques. The black-blood sequences are used to provide anatomic information (aneurysm size, anatomic location) and to evaluate any diseases involving the aortic wall, such as atherosclerosis and aortitis. The bright-blood sequences comprise TOF techniques. Especially useful for evaluation of the thoracic aorta are the ECG-gated bright-blood cine MR sequences, which provide functional information displaying the blood flow through the aorta dynamically during the cardiac cycle. The PC techniques can be used to calculate the flow velocity and flow volume. These parameters may be particularly useful for the hemodynamic evaluation of focal stenoses, such as in coarctation of the aorta.59 The black-blood (SE) and bright-blood (GRE) sequences are excellent techniques for the imaging of vessels in healthy subjects with steady laminar flow, which is perpendicular to the imaging plane. Owing to vessel wall irregularities, such as stenoses, many patients with vascular disease

in the coronal approach. The abdominal aorta is typically imaged in the coronal plane, ensuring the smallest number of slices and the shortest acquisition time. The renal and iliac vessels should be included in the 3D volume. For the abdominal aorta, the imaging FOV typically varies between 30 and 34 cm. Imaging of the thoracic aorta typically requires a larger FOV, ranging from 36 to 40 cm. These dimensions depend on the extent of the aortic disease and the patient’s body habitus. The slice thicknesses for imaging of both the thoracic and abdominal aorta vary between 1.5 to 3.5 mm. First a test 3D dataset is acquired precontrast to ensure accurate positioning and to search for possible disturbing aliasing artifacts. This precontrast 3D volume can also be used as baseline data for subtraction to enhance vessel visualization. Typically, two postcontrast datasets are obtained, with an interval of approximately 10 seconds between the two. During this time, the patient can breathe freely. The probability of obtaining a diagnostic imaging study is increased by acquiring two postcontrast datasets. The second dataset predominantly shows the enhancement of veins because it is acquired later. Subtraction of the latter from the first contrast-enhanced 3D volume ideally results in a dataset showing the arteries only. Another advantage of obtaining two postcontrast 3D datasets is that structures with slow flow, such as large aneurysms, may require additional time to fully enhance and are therefore visualized optimally only on the second 3D volume. Because CE-MRA predominantly depicts the vascular lumen, every aorta protocol should also include ECG-gated T1-weighted SE postcontrast scans to visualize vessel walls. These sequences are acquired in the transverse/axial plane, typically with two to three signal averages for the reduction of respiratory motion artifacts. In addition, depending on the nature of the vessel disease, GRE sequences, showing the pulsatile blood flow and also PC sequences quantifying flow dynamics, may be desirable.62 The majority of thoracic and abdominal aortic aneurysms occur secondary to atherosclerosis. Other causes include infection, inflammation, trauma, syphilis, cystic medial necrosis, and valvular disease (thoracic aorta). There seems to be a strong association of aneurysmatic changes of the aortic wall with smoking63 and hypertension. Aneurysms can be further subdivided by their gross appearance into fusiform (circumferential enlargement of the involved segment) and saccular shaped (asymmetric or focal outpouching of the involved segment). Aneurysms can be categorized as being either true, with all three aortic wall layers involved in the aneurysm, or false (pseudoaneurysms). In the latter case, there is a break in the intima and media such that the wall of the aneurysm is formed by the adventitia only. Most abdominal aortic aneurysms are localized infrarenally and are fusiform in shape. Typically, the patient is asymptomatic, and a pulsatile mass is palpated on physical examination. If the aneurysm is untreated, life-threatening complications, such as rupture, thrombosis, and embolism, frequently occur. Mycotic aneurysms are a rare entity resulting from weakening of the vessel wall by a bacterial infection. This can be either a primary infection or a secondary infection in a preexisting aneurysm. Mycotic aneurysms are usually saccular and are most often found in a suprarenal location. They may demonstrate paravertebral fluid or soft tissue mass, para-aortic Cardiovascular Magnetic Resonance 469

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have turbulent flow, which is neither laminar nor perpendicular to the imaging plane. As a result, image quality is seriously degraded by local artifacts. Decreased flow velocities, caused by impaired cardiac function or aneurysmatic changes of the vessel wall, provide less contrast between the vessel and the surrounding background tissue. The differentiation of slow flow from thrombosis and from artifacts can be quite difficult. This is another limitation that makes these techniques, which are strongly dependent on flow, less accurate for clinical use in many patients. The shortcomings of these methods illustrate the need for techniques that are not dependent on blood flow, such as 3D CE-MRA. With 3D CE-MRA techniques, many of these shortcomings are overcome, as these techniques are not dependent on blood flow and provide excellent vessel-to-background contrast. The image quality is not affected by slow flow; consequently, differentiation of thrombus from slow flow is less likely to be a problem. 3D CE-MRA gradient echo sequences are fast techniques. For the thoracic and abdominal aorta, ECG triggering is generally not required.51,60,61 Acquisition times are approximately 20 seconds, with a sagittal or oblique sagittal volume, allowing a complete study of the thoracic aorta and branches to be performed in a single breath hold. For imaging of both the thoracic and the abdominal aorta, a body phased array surface coil should be used to optimize SNR. In large patients, the body phased array coil may not provide enough coverage of the volume of interest or not allow the patient to be comfortably moved into the machine’s bore. In these patients, the body coil is often the best choice. For studies of the thoracic aorta, special attention should be given to the positioning of the body phased array coil to ensure that the arch vessels are included in the coil FOV. Patients are routinely imaged in the supine position. Our current imaging protocol for the thoracic aorta consists of a localizer sequence, a contrast bolus timing sequence, a T1-weighted precontrast sequence, the 3D pre- and post-CE-MRA sequence, and a T1-weighted postcontrast sequence. As a localizer sequence, an ECG-gated 2D GRE or FSE sequence is appropriate. It is advisable to acquire the localizer during breath holding in the same breath holding position, that is, end inspiration or end expiration, as is planned for the 3D CE-MRA. The anatomy during breath holding is shown on the localizer, and positioning the 3D-imaging slab on this localizer ensures that this anatomy will be included in the important contrast acquisition. A single-slice GRE sequence can be performed in many orientations (axial, sagittal, coronal, or parasagittal plane) for proper timing of the contrast bolus. For evaluation of the thoracic aorta, the gradient echo 3D CE-MRA sequence may be obtained in either the sagittal, coronal, or parasagittal plane. The 3D volumes that are acquired in the coronal plane do provide more anatomic coverage (inclusion of a greater portion of the arch vessels), but may require a longer acquisition time than is feasible in a breath holding period for many patients (This may be a problem in older scanners that are not equipped with high-powered gradients.) For this reason, the sagittal or parasagittal plane may be preferred for imaging of the thoracic aorta. This approach provides less volume coverage but has resolution and slice thickness equal to those

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gas, or osteomyelitic involvement of an adjacent vertebral body. Mycotic aneurysms are more common in women and may present either insidiously or with sepsis and rupture. Although a high percentage of mycotic aneurysms yield negative cultures, the most commonly associated organisms are staphylococcus and salmonella. MRA can be used to visualize both intraluminal and extraluminal anatomy. The intraluminal anatomy is shown by the 3D CE-MRA GRE technique, enabling detailed visualization of the aneurysm morphology providing cross-sectional and, if desired, reformatted images in any orientation. Advantageously, the 3D technique clearly shows the relationship of the aneurysm to the origin of the brachiocephalic or renal arteries. It is chiefly this property that has led to an increased use of this technique for imaging of aortic aneurysms. The precise dimensions of the aneurysm can be evaluated by subvolume MIP and MPR (multi-planar reconstruction) postprocessing techniques. The extraluminal anatomy of the aneurysm is visualized on the T1-weighted SE sequences. These sequences should be performed after the 3D CE-MRA scans but while a noticeable concentration of the contrast agent is still present in the circulation. The presence of contrast aids in the differentiation of flowing blood from mural thrombus. Finally, late gadolinium enhancement (LGE) of the vessel wall, indicating inflammation as is seen in aortitis and mycotic aneurysms when these T1-weighted sequences are used post contrast. T2-weighted SE images should be obtained for mycotic aneurysms. The T2 images show perivascular edema or fluid collections and GRE cine sequences can also be used to demonstrate flow within the aneurysm.64 Several clinical studies have reported high accuracy of 3D CE-MRA for the detection of aneurysms involving the thoracic and abdominal aorta.3,51,53 A sensitivity of 100% and specificity of 100% have been reported for the morphologic analysis of aneurysms. This involves assessment of maximum aneurysm diameter, luminal patency, as well as the proximal and distal aneurysm extent.52,65 CE-MRA has an 88% to 100% sensitivity and a 97% to 100% specificity for the detection of stenoses or occlusions of the aorta.3,66 Aortic dissection occurs when blood dissects through the endothelial lining and into the media of the aortic wall through an intimal tear. In most instances, it is related to degeneration of an aging aorta and may be accelerated by hypertension. In younger patients, there is often an underlying process such as a bicuspid or unicuspid aortic valve, coarctation, pregnancy, connective tissue disorders such as Marfan or Ehlers-Danlos syndrome, relapsing polychondritis, Turner or Noonan syndrome, thoracic cage deformities, and, rarely, systemic lupus erythematosus or giant cell arteritis. Dissection may also occur at sites of iatrogenic trauma, such as at the location of a prior aortic incision, cross-clamping, or traumatic catheterization. The two most common sites for the initiation of aortic dissection are the proximal ascending aorta and the descending aorta just distal to the origin of the left subclavian artery. These reflect sites of maximal mechanical strain on the aorta caused by flexion. The dissection channel generally spirals. In the ascending aorta, the false lumen is usually anterior and to the right. In the arch, the false lumen is superior and slightly posterior. In the descending aorta, it is posterior and to the left. The false channel may 470 Cardiovascular Magnetic Resonance

compress the true lumen. Branch arteries may be perfused by either the false lumen, the true lumen, or both. Arch and iliac vessels are supplied by the false lumen in approximately one half of patients. Renal arteries are fed by the false lumen in approximately 25% of cases. Therapeutic management depends on the proximal extent of the dissection. This is reflected in the Stanford classification system, which describes all aneurysms involving the ascending aorta as type A (60%) and all other dissections as type B. Type A dissections represent a surgical emergency with surgical resection of the dissected segment and replacement with a synthetic graft. If the aortic valve is also compromised, the valve is also replaced. Patients with type B dissection are usually treated medically, surgical treatment being reserved for cases in which visceral or lower extremity circulation is threatened or medical therapy has failed. 3D CE-MRA is an excellent technique for the diagnosis and follow-up of chronic aortic dissections. The intimal flaps are typically clearly visualized on both the source images and MPRs. On the projection images, the intimal flaps might be obscured and not visible. Therefore, the source images remain the most important tools in the diagnosis of aortic dissection. It is not advisable to make a diagnosis based solely on the projection images. The involvement of branch vessels, if any, can be determined by using reformations of the original 3D images. It can be evaluated whether the branch vessels are perfused by the true lumen, the false lumen, or both. The extension of the dissecting membrane into branch vessels should also be evaluated. The axial reformations are particularly helpful for the assessment of the communications between the true and false lumen, that is, entry and reentry tears. Recent studies report high sensitivity (92% to 96%) and high specificity (100%) of the 3D CE-MRA technique for the diagnosis of thoracic aortic dissection.51,53

Renal Arteries Two to five percent of all hypertension cases are directly caused by stenosis of the renal artery and/or its branches.67 Renovascular hypertension, defined as high blood pressure caused by renal artery stenosis, is the most common cause of secondary hypertension. There are two main lesion groups that cause renal artery stenosis. The first main group comprises atherosclerotic lesions, causing 90% to 95% of all cases of renovascular disease.68 These lesions have twice the prevalence in men than in women. Atherosclerotic lesions are more typical for older age groups, with an average age of 55 years. The stenosis can be caused by either a plaque in the renal artery itself or an aortal atherosclerotic plaque impinging on the origin of the renal artery and narrowing its proximal portion. The second main group of lesions causing renal artery stenosis is a heterogeneous group of intrinsic structural abnormalities of the arterial renal artery wall termed fibromuscular dysplasia. This type of stenosis is most frequently seen in women and typically manifests at a younger age (average: 35 years) than atherosclerotic lesions. These lesions consist of fibrous and fibromuscular dysplasia of the renal artery wall and typically have no atherosclerotic components.68

The current gold standard for determining the anatomic presence of renal artery stenosis is invasive X-ray angiography. X-ray angiography provides excellent resolution of the presence and extent of a stenosis, usually identifying the specific cause.89 Digital subtraction technology has reduced the amount of iodinated contrast volume needed and the risk of contrast nephropathy.90 The use of smaller catheters has decreased morbidity associated with the arteriotomy, allowing renal angiography to be performed on an outpatient basis.91 Risks associated with angiography include allergic reaction, reduction in renal function after administration of iodinated contrast medium, and arteriotomy and catheter manipulation complications. The presence of an angiographic stenosis at X-ray angiography does not necessarily imply that the hypertension is related to the lesion (or will reverse after treatment). Several studies have reported high sensitivity (88% to 96%) and high specificity (87% to 100%) of helical CT angiography for the grading of renal artery stenosis greater than 50% of the luminal diameter.92–95 This technique involves capturing the arterial phase of a properly timed bolus of intravenous contrast agent over a region of interest. Currently, CTA has lower spatial and temporal resolution than conventional angiography or digital subtraction angiography does. Image postprocessing is more time-consuming than that for MRA. The most serious drawback of CTA is the large load of iodinated contrast material, which can be contraindicated in patients with impaired renal function and radiation exposure. Doppler ultrasonography enables the indirect detection of an anatomic lesion by measurement of changes in renal artery blood flow patterns. It has also been used for the detection of renal artery stenosis, and either the main renal artery or intrarenal branch vessels can be interrogated. Unfortunately, technically inadequate studies have been reported in up to 40% of cases with ultrasonographic examination of the main renal arteries. The failures are caused predominantly by obesity, bowel gas, aortic aneurysm, and/or recent surgery.96 Consequently, several studies have demonstrated high sensitivity (84% to 93%) and high specificity (95% to 98%) for detecting renal artery stenosis,97–99 while other studies have reported a sensitivity as low as 0%100,101 and a specificity of 37%.100 Several Doppler ultrasound techniques have been employed for the detection of renal artery stenosis. The renal aortic ratio involves measuring the peak velocity in the renal artery and comparing it with that from the adjacent abdominal aorta. Handa and colleagues102 demonstrated that the acceleration index in segmental arteries may be used to quantify the severity of the damping of the Doppler waveform and predict significant renal artery stenosis. They reported a sensitivity of 100% and specificity of 93% for renal artery stenosis greater than 50%. The examination was diagnostic in 98% of the 149 vessels. Patriquin and colleagues103 showed that stenosis greater than 75% was detected with a sensitivity and specificity of 100%, using the acceleration index in segmental and interlobar arteries of children. Stavros and colleagues104 also evaluated segmental renal artery branches. They reported a sensitivity of 89% and a specificity of 83% for detecting stenosis of more than 60%. Currently, the use of Doppler sonography as a screening tool for renal artery stenosis is controversial. The diagnostic value of the examination is highly dependent Cardiovascular Magnetic Resonance 471

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Renal artery stenosis is a progressive disease that may result in gradual and silent (almost one half of patients are not hypertensive69) loss of functional renal tissue. The incidence increases with age.70 It has been estimated that renal artery stenosis accounts for 5% to 15% of all patients developing end-stage renal disease each year.69 For patients with hypertension due to renal artery stenosis, interventional therapy is far superior to pharmacologic therapy. Renovascular hypertension tends to be resistant to medical therapy, and regulation of the blood pressure does not reduce the ischemic injury to the kidney. Interventional therapy with surgery or angioplasty restores blood flow to the ischemic kidney and also reduces the need for or complexity of lifelong antihypertensive therapy, thereby reducing the number of patients who require dialysis and decreasing the morbidity and mortality associated with long-term hypertension.71 The challenge to the clinician is to identify patients with renal artery stenosis in a safe and cost-effective manner. The diagnosis should be made sufficiently early in the natural history that the patient may benefit maximally from intervention (prior to development of end-stage renal dysfunction). The tests that have been used to diagnose renal artery stenosis can be subdivided into tests that detect physiologic abnormalities (functional) and tests that detect an anatomic lesion of the renal artery (morphologic). No single test that can provide both types of information.72 Tests that detect physiologic abnormalities associated with renal artery stenosis have included the rapid sequence intravenous pyelogram (no longer used), measurement of peripheral and renal vein renin activity,73 the captopril test with sampling of plasma renin, and captopril renography. Tests that directly detect an anatomic lesion of the renal artery include intravenous digital subtraction angiography, CTA,74 and renal artery MRA. With duplex sonography, anatomic lesions can be detected indirectly by virtue of changes in flow patterns. While each of these tests has shown promise, none has been universally accepted. The captopril test is a relatively safe and inexpensive test. It is based on the increase in plasma renin activity after captopril administration. However, its accuracy is affected by age,75 race,76 renal function,77 intake of antihypertensive medications,77–80 renal arteriolar changes,77 food intake before the test,81 and dilution of plasma. Consequently, there is a great variability in reported sensitivities (38% to 100%) and specificities (72% to 100%) among different studies. The administration of captopril can worsen renal function in the setting of renal artery stenosis, although the effect is transient. Radionuclide renography is accurate in demonstrating differential glomerular filtration rates between the two kidneys.82,83 When captopril is used in association with 99m TcDTPA or 99mMag3 renography, the difference is magnified by resultant decrease in the glomerular filtration rate of the kidney affected by renal artery stenosis. Captopril acts by reducing the angiotensin-II-mediated constriction of the efferent arteriole, thus lowering glomerular pressure.84 In contrast, the glomerular filtration rate in patients with primary hypertension does not change after captopril administration.85 Consequently, the use of these drugs enhances the sensitivity and specificity of the routine renal scintigram, with reported sensitivities ranging from 91% to 94% and specificities ranging from 93% to 97%.86–88

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on the expertise of the operator, the accessory vessels are often poorly visualized,100,105,106 and as was mentioned previously, although several studies report high sensitivities and specificities, it is not clear that high technical success is widely reproducible. The above-mentioned shortcomings of both conventional angiography, digital subtraction angiography, and other techniques, added to the recently available costeffective treatment options for renal artery stenosis, illustrate the need for a novel noninvasive screening modality for this disease. Many investigators have regarded renal MRA as this promising technique for noninvasive screening of patients with suspected renal artery disease. Thus, various MRA imaging strategies for visualizing the renal arteries have been proposed, including TOF and PC techniques with 2D107,108 or 3D acquisitions.109–112 Another method that has been applied to image the renal arteries is signal targeting with alternating radiofrequency. This method involves tagging blood in a feeding vessel (suprarenal abdominal aorta) and then imaging it, after it has moved into a target vessel (renal artery).113,114 Though many studies initially reported promising results, the widespread use of these methods has been limited by the inability of these techniques to fully and reliably visualize the renal arteries, owing to motion artifacts caused by respiration, vessel tortuosity, in-plane flow saturation effects, and limited spatial resolution. The renal arteries are typically 4 to 5 mm in diameter, similar in size to the left main coronary artery. Added to these problems, the accessory renal arteries are often smaller and typically not well depicted. The use of 3D CE-MRA has overcome many of these limitations, being a fast technique that enables imaging of a large 3D volume, providing accurate and reliable visualization of the major renal arteries and accessory renal arteries in their full length. A typical imaging protocol for CE-MRA of the renal arteries is quite similar to investigations of the aorta. It consists of a localizer sequence, a contrast bolus timing sequence, and the 3D pre- and post-CE-MRA sequence. A body phased array coil should be used for optimal signal. As a localizer sequence, a 2D gradient-refocused sequence or FSE sequence is appropriate, prescribed in the sagittal and/or coronal plane. It should be acquired during breath holding in the same position, that is, end expiration or end inspiration, as that in which the 3D CE-MRA will be performed. The 3D CEMRA volume for evaluating the renal arteries is prescribed in a fashion similar to that in imaging of the aorta, although for most studies, the FOV can typically be reduced to 32 cm or less. The imaging volume should be positioned in the coronal plane, preferably slightly tilted parallel to the abdominal aorta (28 to 40 slices with thickness < 2.5 mm). These parameters provide high-resolution images within a breath holding period of less than 30 seconds. Parallel imaging techniques should be used if available to shorten scan time and improve spatial resolution. The 3D PC techniques depict blood flow through the renal arteries and depict turbulence and slow flow, both of which are indicative of stenosis, as signal voids.112 Therefore, while 3D CE-MRA provides anatomic images of the renal arteries, PC sequences provide additional functional information. Consequently, some studies suggest the use of 3D PC techniques as an adjunct to the renal 472 Cardiovascular Magnetic Resonance

CE-MRA.115,116 In addition, postcontrast T1-weighted images should be obtained to look for the renal excretion of gadolinium-DTPA, which typically occurs within minutes after the contrast injection. This last study is especially important for the evaluation of transplanted kidneys. For the detection of renal artery stenosis greater than 50% with 3D CE-MRA, reported sensitivities and specificities vary from 93% to 100% and from 71% to 92%, respectively.66,115,117–119 These techniques are also superior to noncontrast MRA techniques in the detection of accessory renal arteries115,118 and have found use in the preoperative detection of accessory renal arteries in living renal allograft donors.120 3D CE-MRA has also been successfully applied in the postoperative evaluation of transplanted renal arteries.121 In these patients, it is recommended to add an axial T2-weighted sequence with fat saturation, prior to giving gadolinium contrast, to assess for perinephric fluid collections and hydronephrosis. Owing to the pelvic location of transplanted kidneys, these studies are generally less susceptible to respiratory motion and can be carried out without breath holding.

Mesenteric Arteries Chronic intestinal ischemia is caused in the majority of cases by atherosclerotic narrowing or obstruction of the major splanchnic arteries. These narrowings or obstructions compromise the blood flow to the intestine, resulting in postprandial abdominal pain, called intestinal (abdominal) ischemia. The pain typically starts within a half-hour of eating and persists for hours. The significant weight loss in these patients is primarily due to a decreased intake of food. Mucosal damage caused by the chronic ischemia adds to the weight loss. Although patients with generalized atherosclerotic disease often have involvement of the splanchnic circulation, only few of these patients have the symptoms of mesenteric ischemia. This is due to the extensive collateral circulation between the celiac artery and superior mesenteric arteries and, to a lesser extent, the inferior mesenteric arteries. Therefore, chronic stenosis or even occlusion of all three major mesenteric arteries frequently occurs without abdominal symptoms.122,123 The high therapeutic success rates of various surgical techniques for reestablishing blood flow have illustrated the need for a reliable imaging technique to diagnose stenosis and preoperatively map the mesenteric arteries.124–128 Although conventional invasive X-ray angiography remains the gold standard, it does have many disadvantages, including invasiveness, cost, and nephrotoxic iodinated contrast. More recently, Doppler sonography has been suggested for imaging of the mesenteric vessels.129–132 However, sonography has several severe limitations, including a 25% rate of nonvisualization of both vessels and failure to obtain adequate signal owing to overlying bowel gas, excess adipose tissue or vessel wall calcification. Possibly, it also overestimates the prevalence of double-vessel disease.132 Noncontrast MRA techniques such as PC MRA have failed to reliably depict the mesenteric arteries. The disappointing results achieved with these techniques (even when

Peripheral Vessels The leading cause of occlusive arterial disease of the extremities in patients over 40 years of age is atherosclerosis. The highest incidence of peripheral vascular disease occurs in the sixth and seventh decades of life and is slightly more common among men than women.141 The same risk factors that have been identified for other types of cardiovascular disease, that is, coronary artery disease and cerebrovascular disease, also have been found to correlate with peripheral vascular disease. These risk factors include hypertension, low levels of high-density lipoprotein cholesterol, and high levels of triglycerides. Risk factors that show a particularly strong correlation with peripheral vascular disease are cigarette smoking and diabetes mellitus/impaired glucose intolerance.142 Atherosclerotic lesions have a predilection for sites of increased turbulence or complex flow, arterial branching points, and areas with increased mechanical wall stress. Several noninvasive techniques have been developed for the detection of peripheral vascular disease and for the evaluation of the severity of the stenosis. These tests include Doppler ultrasonography, pulse volume recording, segmental blood pressure measurement, postocclusive reactive hyperemia testing, transcutaneous oximetry, and color-assisted ultrasound imaging.143 These noninvasive tests are safe, can be performed on an outpatient basis, and are readily repeatable. It should be noted, however, that the ultrasound techniques are operator dependent and that the accuracy of these tests does depend on the skill and experience of the examiner.144 Imaging of the peripheral arteries with ultrasound has several drawbacks. Obese extremities are difficult to evaluate, and high-quality ultrasound images distal to the midpopliteal level are difficult to obtain consistently in many patients. The estimation of the degree of luminal narrowing by peak velocity measurements can be inaccurate. These noninvasive tests are used to evaluate and determine the extent of vascular stenosis and/or occlusion. Once it has been determined that the patient’s symptoms are due to hemodynamically significant vascular disease, it should be determined whether the patient is eligible for a revascularization procedure. Owing to the risks and toxicities associated with this technique, only patients who are considered for an intervention should undergo conventional angiography. These complications are well described145,146 and have decreased somewhat with the widespread use of digital subtraction technology and smaller catheter size.147 The systemic complications associated with conventional angiography include nausea, vasovagal attack, angina, allergic reaction, renal dysfunction, and death. Local complications include hematoma, dissection, thrombosis, embolus, and pseudoaneurysm, sometimes requiring additional therapy such as blood transfusion or surgery. The estimated prevalence of local complications varies from 4.1% to 23.2%.148 Conventional peripheral X-ray angiography can be uncomfortable, as large volumes of iodinated contrast are required, with subsequent filling of many muscular arterial branches.149 Interpretation of X-ray angiograms may be inaccurate, owing to poor filling (slow flow) into distal vessels. Conventional angiography can fail to demonstrate distal vessels suitable for reconstructive surgery in up to 70% of patients Cardiovascular Magnetic Resonance 473

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systolically gated133) are likely to be due to the triphasic flow pattern in these arteries. There are several other shortcomings that preclude 3D PC MRA from being used as a screening tool for mesenteric artery stenosis. These limitations include phase ghosting, motion artifact related to long scan times, and uncertainty regarding choice of appropriate velocity-encoding gradient value. The 3D CE-MRA techniques have been successfully applied to image the mesenteric circulation.134–136 Two main factors have added to this success: (1) Most stenoses occur within 1 cm from the orifice, where the diameter of the superior mesenteric artery and celiac artery is the greatest and well within the resolution capacity of the 3D techniques, and (2) these techniques are flow independent. With the 3D CE-MRA of the inferior mesenteric artery, a smaller-caliber vessel is less reliably visualized and is therefore prone to be more frequently graded incorrectly.135 A 3D MRA study of the mesenteric vessels is carried out in the following manner. Prior to the examination, the patients can be administered glucagon intravenously to reduce bowel motion and may be given a high-calorie meal. This results in an increased blood flow to the intestine, allowing visualization of the smaller branch vessels. Patients should be positioned supine, preferentially with their arms placed above their heads to minimize aliasing artifacts, and for optimal signal, the body phased array coil should be used. As with all 3D CE-MRA, the protocol comprises a localizer sequence, contrast bolus sequence, and 3D gradient echo sequence pre- and post-contrast. The 3D slab is positioned in the coronal plane and should include the aorta and the portal vein in addition to the splanchnic vessels. If the primary clinical question is to determine disease involving the proximal mesenteric arteries, the slab should be positioned in the sagittal plane. The number of slices should be 30 to 32, with a slice thickness of 2.0 to 3.0 mm. Overall, these parameters should be tailored for the individual patient so that the entire 3D slab can be imaged within a single breath holding period. As was discussed previously for other vascular territories, a 3D volume should be acquired before injection of the contrast agent to check for correct positioning. This precontrast image dataset can later be subtracted from the postcontrast images to increase vessel conspicuity. A delayed 3D MRA dataset should be acquired to image the venous phase of the contrast bolus injection, which shows the portal vein and mesenteric veins. In addition, the severity of the mesenteric ischemia can be functionally assessed by using other noncontrast CMR techniques that provide quantitative information, such as blood flow and blood oxygen saturation. Postprandial PC cine CMR of the superior mesenteric artery, superior mesenteric vein, and portal vein showed that in patients with mesenteric ischemia, the percentage change in postprandial blood flow in both the superior mesenteric artery and vein is significantly less than that seen in healthy subjects.137,138 CMR evaluation of blood oxygen saturation139 in the superior mesenteric vein can be used as a measure of the degree of acute flow reduction in the superior mesenteric artery, which suggests mesenteric ischemia in a canine model.140 These functional CMR techniques are not suitable for screening purposes. However, when added to a 3D CE-MRA protocol, they may aid in the diagnosis of mesenteric ischemia.

VASCULATURE AND PERICARDIUM

with severe disease.150,151 In addition, background structures such as cortical bone can make interpretation of crossing arteries difficult. Asymmetrical stenoses can be underestimated by conventional angiography, and film magnification can affect the interpretation of lesions. The dose of iodinated contrast and the timing of its administration affect the quality of the study. X-ray angiography is expensive. The examination often requires several hours to perform, with a minimum of 4 hours required for recovery of the patient after uncomplicated procedures. Because of its several advantages over conventional angiography, many investigators have attempted to develop an MRA technique to reliably image the peripheral arteries. Peripheral MRA using the 2D TOF technique has been evaluated and compared to conventional angiography with mixed results.152–155 Currently, it is widely believed that the 2D TOF MRA techniques do not have the diagnostic accuracy to replace conventional angiography for several reasons. In the peripheral arteries, the nature of the flow is highly pulsatile, with possible retrograde flow during diastole, yielding a decreased blood signal on the TOF images. This pulsatile flow also causes ghost artifacts, which can be reduced by synchronizing the data acquisition to a single phase of the cardiac cycle, though at the expense of increased imaging time.156,157 Vessels with horizontal flow, such as collateral vessels, tend to have less signal intensity on 2D TOF images and may appear stenotic, and retrograde flow and collateral flow are often missed. Retrograde flow can also complicate the use of presaturation slabs. Turbulent flow causes dephasing, which results in signal loss in areas of stenosis and/or aneurysms, causing an overestimation of the degree of luminal narrowing. Motion of the patient during the acquisition of 2D data results in an apparent offset of the vessel that can simulate a vascular lesion. Background suppression can be imperfect, particularly if there is abundant fat.158 In addition, 2D TOF should be prescribed in a plane perpendicular to the blood flow direction. In the peripheral arteries, this is the transaxial plane. Thus, multiple 2D slabs have to be acquired in the transaxial direction to cover the total length of the vessels in the extremities. The blood flow in the peripheral arteries is typically slow and is visualized best when using thin slice acquisitions (1.5 to 3 mm). It can therefore take up to 2 hours to complete a study visualizing the peripheral arteries in their total length, that is, from the iliac bifurcation of the aorta to the dorsalis pedis artery. These long imaging times have also contributed to the limited use of the 2D TOF techniques in routine clinical practice. Noncontrast 3D TOF studies have similar limitations as well as a long examination time. Despite the above-mentioned limitations of the noncontrast 2D TOF techniques, Owen and colleagues159 demonstrated that these MRA techniques can detect runoff vessels distal to occlusions in patients with peripheral arterial occlusive disease, with greater sensitivity than X-ray angiography. This may have consequences for the surgical management and may alter the treatment plan between revascularization and amputation. The introduction of the 3D peripheral CE-MRA techniques has altered the imaging strategies for the peripheral arteries. The 3D contrast techniques and conventional X-ray angiography have similar approaches to imaging 474 Cardiovascular Magnetic Resonance

vessels, as both follow the contrast agent as it travels through the vascular region of interest. This approach makes the 3D contrast-enhanced techniques relatively independent of flow in comparison to the 2D noncontrast TOF techniques. With the 3D CE-MRA techniques, the imaging data can be acquired in the coronal plane, so the vessels in the extremities can be imaged in their total length in under 4 minutes. The adult legs are long, and the imaging volume encompassing the peripheral arteries is too large to be covered in a single acquisition. Consequently, either the imaging range should be restricted or multiple acquisitions should be performed. There are two main approaches for acquiring multiple acquisitions with the 3D contrast techniques. One approach is to give multiple low doses of gadolinium chelate while covering a different part of the vasculature each time. With this technique, higher SNR and higher spatial resolution can be achieved. Imaging can be completed in 2 to 4 minutes, a period short enough that the legs can easily be held still. Subtraction techniques can correct for background tissue enhancement and venous signal resulting from previous injections.160 Another approach is quite similar to the technique applied in X-ray angiography, that is, the bolus chase technique. With this MRA technique, a single contrast medium bolus is tracked by fast multiple 3D slabs as it flows down through the leg vasculature. To follow the contrast agent, the table or the patient needs to be moved, typically two or three times between sets of 3D volumes, to ensure total coverage of the leg vessels.161–163 With this technique, the peripheral arteries can be visualized in their full length in 4 minutes.164 Care should be taken that the patient is positioned such that the arteries remain within the imaging volume over the full range of table motion. Currently, our preferred approach for peripheral MRA represents a hybrid of time-resolved imaging and the bolus chase technique.165 A bolus infusion of 0.1 mmol/kg gadolinium chelate is given at 2 cc/sec while a time-resolved MRA of the calf and pedal vessels is performed (e.g., using TRICKS or parallel imaging). This station is typically acquired by using a peripheral phased array coil or torso phased array coil to obtain the best SNR. Slices are 2 mm thick, and in-plane spatial resolution is on the order of 1 mm or better. Next, a bolus chase, moving table acquisition is done, using a bolus of 0.1 mmol/kg at 2 cc/sec followed by a slow infusion of 0.1 mmol/kg at 0.5 cc/sec. The respective mask images are subtracted from the contrast-enhanced images for each table station, and MIPs are created. The bolus chase study can be done with a peripheral array coil, if available, but the body coil is often sufficient, since spatial resolution need not be as high as that for the calf and foot vessels. Use of a tourniquet or blood pressure cuff inflated above venous pressure can help to reduce venous filling if only the moving table study is done and time-resolved imaging is not performed (Fig. 34-5). With 3D CE-MRA, the reported sensitivities and specificities for detecting and grading hemodynamically significant stenoses of the peripheral arteries of the lower extremities are 89% to 97% and a specificity of 95% to 98%,160,161,164 respectively. The 3D CE-MRA techniques are also useful for the assessment of peripheral bypass graft patency and for the detection of arteriovenous fistulas of the peripheral arteries.15

Historically, MRA of the hand and wrist has received considerably less attention in the literature than has MRA of the lower extremities. While some groups have employed the 2D TOF and PC MRA techniques for imaging of the hand,166–168 the use of these techniques is complicated by the necessity of imaging perpendicular to the flow. This limitation requires separate acquisitions for the superficial and deep palmar arches and digital arteries. Because of the relatively lengthy acquisition times, the TOF and PC MRA of the hand and wrist are often degraded by patient motion.

34 CARDIOVASCULAR MAGNETIC RESONANCE ANGIOGRAPHY: CAROTIDS, AORTA, AND PERIPHERAL VESSELS

Figure 34-5 Bolus chase CE-MRA in the same fast-flow patient with (right) and without (left) thigh compression. Dedicated blood pressure cuffs inflated to 60 mm Hg eliminated venous enhancement in the calves, especially on the right side. Source: Courtesy of Drs. Honglei Zhang and Martin Prince.

Although there has been considerable interest in the adaptation of CE-MRA for imaging of the lower extremities, optimization of techniques for CE-MRA of the hand has received relatively less attention to date. Rofsky169 described the use of a 3D-gradient echo sequence with a gadolinium chelate to image the hand and wrist. A double dose (0.2 mmol/kg) of gadolinium-DTPA was administered at an even rate throughout the entire 2- to 3.5-minute acquisition, and a single 3D slab was imaged. In his experience, the 3D CE-MRA sequences are very suitable for obtaining detailed anatomy of the hand and wrist. Cardiovascular Magnetic Resonance 475

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SUMMARY Though noncontrast TOF and PC MRA are occasionally suitable for normal, healthy subjects, they have largely been supplanted by 3D CE-MRA techniques. CE-MRA is currently the most reliable and diagnostically accurate technique for evaluation of many vascular territories throughout the body. It is safe, cost-efficient, and fast, though issues related to nephrogenic systemic fibrosis must be

remembered. Consequently, CE-MRA already rivals conventional X-ray angiography in many cases as an initial angiographic exam. Advantageous recent developments include parallel imaging, 3 T imaging, and introduction of new contrast agents with higher relaxivity or blood pool properties. Further refinement, with improvements in hardware, contrast agents, and pulse sequences will undoubtedly broaden the clinical utility of CE-MRA even more in the near future.

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artery stenosis] Leistungsfahigkeit der CT-angiographie beim nachweis von nierenarterienstenosen. Rofo Fortschr Geb Rontgenstr Neuen Bildgeb Verfahr. 1994;161(6):519–525. Johnson PT, Halpern EJ, Kuszyk BS, et al. Renal artery stenosis: CT angiography-comparison of real-time volume-rendering and maximum intensity projection algorithms. Radiology. 1999;211:337–343. Kaatee R, Beek FJ, de Lange EE, et al. Renal artery stenosis: detection and quantification with spiral CT angiography versus optimized digital subtraction angiography. Radiology. 1997;205(1):121–127. Lewis BD, James EM. Current applications of duplex and color Doppler ultrasound imaging: abdomen. Mayo Clin Proc. 1989;64 (9):1158–1169. Hansen KJ, Tribble RW, Reavis SW, et al. Renal duplex sonography: evaluation of clinical utility. J Vasc Surg. 1990;12(3):227–236. Kohler TR, Zierler RE, Martin RL, et al. Noninvasive diagnosis of renal artery stenosis by ultrasonic duplex scanning. J Vasc Surg. 1986;4(5):450–456. Taylor DC, Kettler MD, Moneta GL, et al. Duplex ultrasound scanning in the diagnosis of renal artery stenosis: a prospective evaluation. J Vasc Surg. 1988;7(2):363–369. Berland LL, Koslin DB, Routh WD, Keller FS. Renal artery stenosis: prospective evaluation of diagnosis with color duplex US compared with angiography. Work in progress. Radiology. 1990;174 (2):421–423. Desberg AL, Paushter DM, Lammert GK, et al. Renal artery stenosis: evaluation with color Doppler flow imaging. Radiology. 1990;177 (3):749–753. Handa N, Fukanaga R, Ogawa S, Matsumoto M, Kimura K, Kamada T. A new accurate and non-invasive screening method for renovascular hypertension: the renal artery Doppler technique. J Hypertens Suppl. 1988;6(4):S458–S460. Patriquin HB, Lafortune M, Jequier JC, et al. Stenosis of the renal artery: assessment of slowed systole in the downstream circulation with Doppler sonography. Radiology. 1992;184(2):479–485. Stavros AT, Parker SH, Yakes WF, et al. Segmental stenosis of the renal artery: pattern recognition of tardus and parvus abnormalities with duplex sonography. Radiology. 1992;184(2):487–492. Mollo M, Pelet V, Mouawad J, Mathieu JP, Branchereau A. Evaluation of colour duplex ultrasound scanning in diagnosis of renal artery stenosis, compared to angiography: a prospective study on 53 patients. Eur J Vasc Endovasc Surg. 1997;14(4):305–309. Strotzer M, Fellner CM, Geissler A, et al. Noninvasive assessment of renal artery stenosis: a comparison of MR angiography, color Doppler sonography, and intraarterial angiography. Acta Radiol. 1995;36 (3):243–247. Kent KC, Edelman RR, Kim D, Steinman TI, Porter DH, Skillman JJ. Magnetic resonance imaging: a reliable test for the evaluation of proximal atherosclerotic renal arterial stenosis. J Vasc Surg. 1991;13 (2):311–318. Debatin JF, Spritzer CE, Grist TM, et al. Imaging of the renal arteries: value of MR angiography. Am J Roentgenol. 1991;157(5):981–990. Bass JC, Prince MR, Londy FJ, Chenevert TL. Effect of gadolinium on phase-contrast MR angiography of the renal arteries. Am J Roentgenol. 1997;168(1):261–266. Borrello JA, Li D, Vesely TM, Vining EP, Brown JJ, Haacke EM. Renal arteries: clinical comparison of three-dimensional time-of-flight MR angiographic sequences and radiographic angiography. Radiology. 1995;197(3):793–799. Loubeyre P, Trolliet P, Cahen R, Grozel F, Labeeuw M, Minh VA. MR angiography of renal artery stenosis: value of the combination of threedimensional time-of-flight and three-dimensional phase-contrast MR angiography sequences. Am J Roentgenol. 1996;167(2):489–494. Wasser MN, Westenberg J, van der Hulst VP, et al. Hemodynamic significance of renal artery stenosis: digital subtraction angiography versus systolically gated three-dimensional phase-contrast MR angiography. Radiology. 1997;202(2):333–338. Edelman RR, Siewert B, Adamis M, Gaa J, Laub G, Wielopolski P. Signal targeting with alternating radiofrequency (STAR) sequences: application to MR angiography. Magn Reson Med. 1994;31(2):233–238. Wielopolski PA, Adamis M, Prasad P, Gaa J, Edelman R. Breath-hold 3D STAR MR angiography of the renal arteries using segmented echo planar imaging. Magn Reson Med. 1995;33(3):432–438. De Cobelli F, Vanzulli A, Sironi S, et al. Renal artery stenosis: evaluation with breath-hold, three-dimensional, dynamic, gadoliniumenhanced versus three-dimensional, phase-contrast MR angiography. Radiology. 1997;205(3):689–695.

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116. Prince MR, Schoenberg SO, Ward JS, Londy FJ, Wakefield TW, Stanley JC. Hemodynamically significant atherosclerotic renal artery stenosis: MR angiographic features. Radiology. 1997;205(1):128–136. 117. Bakker J, Beek FJ, Beutler JJ, et al. Renal artery stenosis and accessory renal arteries: accuracy of detection and visualization with gadolinium-enhanced breath-hold MR angiography. Radiology. 1998;207 (2):497–504. 118. Hany TF, Debatin JF, Leung DA, Pfammatter T. Evaluation of the aortoiliac and renal arteries: comparison of breath-hold, contrastenhanced, three-dimensional MR angiography with conventional catheter angiography. Radiology. 1997;204(2):357–362. 119. Rieumont MJ, Kaufman JA, Geller SC, et al. Evaluation of renal artery stenosis with dynamic gadolinium-enhanced MR angiography. Am J Roentgenol. 1997;169(1):39–44. 120. Buzzas GR, Shield III CF, Pay NT, Neuman MJ, Smith JL. Use of gadolinium-enhanced, ultrafast, three-dimensional, spoiled gradient-echo magnetic resonance angiography in the preoperative evaluation of living renal allograft donors. Transplantation. 1997;64(12):1734–1737. 121. Johnson DB, Lerner CA, Prince MR, et al. Gadolinium-enhanced magnetic resonance angiography of renal transplants. Magn Reson Imaging. 1997;15(1):13–20. 122. Cunningham CG, Reilly LM, Stoney R. Chronic visceral ischemia. Surg Clin North Am. 1992;72(1):231–244. 123. Kurland B, Brandt LJ, Delany HM. Diagnostic tests for intestinal ischemia. Surg Clin North Am. 1992;72(1):85–105. 124. Calderon M, Reul GJ, Gregoric ID, et al. Long-term results of the surgical management of symptomatic chronic intestinal ischemia. J Cardiovasc Surg Torino. 1992;33(6):723–728. 125. Geelkerken RH, van Bockel JH, de Roos WK, Hermans J, Terpstra JL. Chronic mesenteric vascular syndrome: results of reconstructive surgery. Arch Surg. 1991;126(9):1101–1106. 126. Rheudasil JM, Stewart MT, Schellack JV, Smith RB, Salam AA, Perdue GD. Surgical treatment of chronic mesenteric arterial insufficiency. J Vasc Surg. 1988;8(4):495–500. 127. Roberts L, Wertman DAJ, Mills SR, Moore AVJ, Heaston DK. Transluminal angioplasty of the superior mesenteric artery: an alternative to surgical revascularization. Am J Roentgenol. 1983;141(5):1039–1042. 128. Sniderman K, Strandness DE, Van Breda A, eds. Vascular Diseases: Surgical and Interventional Therapy. New York: Churchill Livingstone; Transluminal angioplasty in the management of chronic intestinal ischaemia. p. 803–819. 1994. 129. Jager KA, Fortner GS, Thiele BL, Strandness DE. Noninvasive diagnosis of intestinal angina. J Clin Ultrasound. 1984;12(9):588–591. 130. Koslin DB, Mulligan SA, Berland LL. Duplex assessment of the splanchnic vasculature. Semin Ultrasound CT MR. 1992;13(1):34–39. 131. Moneta GL, Yeager RA, Dalman R, Antonovic R, Hall LD, Porter JM. Duplex ultrasound criteria for diagnosis of splanchnic artery stenosis or occlusion. J Vasc Surg. 1991;14(4):511–518. 132. Roobottom CA, Dubbins PA. Significant disease of the celiac and superior mesenteric arteries in asymptomatic patients: predictive value of Doppler sonography. Am J Roentgenol. 1993;161(5):985–988. 133. Wasser MN, Geelkerken RH, Kouwenhoven M, et al. Systolically gated 3D phase contrast MRA of mesenteric arteries in suspected mesenteric ischemia. J Comput Assist Tomogr. 1996;20(2):262–268. 134. Hany TF, Schmidt M, Schoenenberger AW, Debatin JF. Contrastenhanced three-dimensional magnetic resonance angiography of the splanchnic vasculature before and after caloric stimulation. Original investigation. Invest Radiol. 1998;33(9):682–686. 135. Meaney JF, Prince MR, Nostrant TT, Stanley JC, Gadoliniumenhanced MR. Angiography of visceral arteries in patients with suspected chronic mesenteric ischemia. J Magn Reson Imaging. 1997;7 (1):171–176. 136. Shirkhoda A, Konez O, Shetty AN, Bis KG, Ellwood RA, Kirsch MJ. Mesenteric circulation: three-dimensional MR angiography with a gadolinium-enhanced multiecho gradient-echo technique. Radiology. 1997;202(1):257–261. 137. Burkart DJ, Johnson CD, Reading CC, Ehman RL. MR measurements of mesenteric venous flow: prospective evaluation in healthy volunteers and patients with suspected chronic mesenteric ischemia. Radiology. 1995;194(3):801–806. 138. Li KC, Whitney WS, McDonnell CH, et al. Chronic mesenteric ischemia: evaluation with phase-contrast c cine MR imaging. Radiology. 1994;190(1):175–179. 139. Wright GA, Hu BS, Macovski A. I.I. Rabi Award. Estimating oxygen saturation of blood in vivo with MR imaging at 1.5 T. J Magn Reson Imaging. 1991;1(3):275–283 .

156. Ho KY, de Haan MW, Oei TK, et al. Angiography of the iliac and upper femoral arteries using four different inflow techniques. Am J Roentgenol. 1997;169(1):45–53. 157. Selby K, Saloner D, Anderson CM, Chien D, Lee RE. MR angiography with a cardiac-phase–specific acquisition window. J Magn Reson Imaging. 1992;2(6):637–643. 158. Edelman RR. MR angiography: present and future. Am J Roentgenol. 1993;161(1):1–11. 159. Owen RS, Carpenter JP, Baum RA, Perloff LJ, Cope C. Magnetic resonance imaging of angiographically occult runoff vessels in peripheral arterial occlusive disease. N Engl J Med. 1992;326(24):1577–1581. 160. Rofsky NM, Johnson G, Adelman MA, Rosen RJ, Krinsky GA, Weinreb JC. Peripheral vascular disease evaluated with reduced-dose gadoliniumenhanced MR angiography. Radiology. 1997;205(1):163–169. 161. Meaney JF, Ridgway JP, Chakraverty S, et al. Stepping-table gadolinium-enhanced digital subtraction MR angiography of the aorta and lower extremity arteries: preliminary experience. Radiology. 1999;211 (1):59–67. 162. Wang Y, Lee HM, Avakian R, Winchester PA, Khilnani NM, Trost D. Timing algorithm for bolus chase MR digital subtraction angiography. Magn Reson Med. 1998;39(5):691–696. 163. Wang Y, Lee HM, Khilnani NM, et al. Bolus-chase MR digital subtraction angiography in the lower extremity. Radiology. 1998;207 (1):263–269. 164. Ho KY, Leiner T, de Haan MW, Kessels AG, Kitslaar PJ, van Engelshoven JM. Peripheral vascular tree stenoses: evaluation with moving-bed infusion-tracking MR angiography. Radiology. 1998;206 (3):683–692. 165. Pereles FS, Collins JD, Carr JC, et al. Accuracy of stepping-table lower extremity MR angiography with dual-level bolus timing and separate calf acquisition: hybrid peripheral MR angiography. Radiology. 2006;240(1):283–290. 166. Disa JJ, Chung KC, Gellad FE, Bickel KD, Wilgis EF. Efficacy of magnetic resonance angiography in the evaluation of vascular malformations of the hand. Plast Reconstr Surg. 1997;99(1):136–144. 167. Dobson MJ, Hartley RW, Ashleigh R, Watson Y, Hawnaur JM. MR angiography and MR imaging of symptomatic vascular malformations. Clin Radiol. 1997;52(8):595–602. 168. Holder LE, Merine DS, Yang A. Nuclear medicine, contrast angiography, and magnetic resonance imaging for evaluating vascular problems in the hand. Hand Clin. 1993;9(1):85–113. 169. Rofsky NM. MR angiography of the hand and wrist. Magn Reson Imaging Clin N Am. 1995;3(2):345–359.

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140. Li KC, Pelc LR, Puvvala S, Wright GA. Mesenteric ischemia due to hemorrhagic shock: MR imaging diagnosis and monitoring in a canine model. Radiology. 1998;206(1):219–225. 141. Creager MA, Dzau VJ, Isselbacher KJ, Braunwald E, et al. eds. Harrison’s Principles of Internal Medicine. 13 ed. New York: McGrawHill; 211, Vascular diseases of the extremities. p. 1135–1143. 1994. 142. Criqui MH, Denenberg JO, Langer RD, Fronek A. The epidemiology of peripheral arterial disease: importance of identifying the population at risk. Vasc Med. 1997;2(3):221–226. 143. Creager MA. Clinical assessment of the patient with claudication: the role of the vascular laboratory. Vasc Med. 1997;2(3):231–237. 144. O’Keeffe ST, Persson AV. Use of noninvasive vascular laboratory in diagnosis of venous and arterial disease. Cardiol Clin. 1991;9(3):429–442. 145. Hessel SJ, Adams DF, Abrams HL. Complications of angiography. Radiology. 1981;138(2):273–281. 146. Shehadi WH, Toniolo G. Adverse reactions to contrast media: a report from the Committee on Safety of Contrast Media of the International Society of Radiology. Radiology. 1980;137(2):299–302. 147. Waugh JR, Sacharias N. Arteriographic complications in the DSA era. Radiology. 1992;182(1):243–246. 148. Olivecrona H. Complications of cerebral angiography. Neuroradiology. 1977;14(4):175–181. 149. Katzen BT. Peripheral, abdominal, and interventional applications of DSA. Radiol Clin North Am. 1985;23(2):227–241. 150. Patel KR, Semel L, Clauss RH. Extended reconstruction rate for limb salvage with intraoperative prereconstruction angiography. J Vasc Surg. 1988;7(4):531–537. 151. Flanigan DP, Williams LR, Keifer T, Schuler JJ, Behrend AJ. Prebypass operative arteriography. Surgery. 1982;92(4):627–633. 152. Baum RA, Rutter CM, Sunshine JH, et al. Multicenter trial to evaluate vascular magnetic resonance angiography of the lower extremity. American College of Radiology Rapid Technology Assessment Group. JAMA. 1995;274(11):875–880. 153. Quinn SF, Demlow TA, Hallin RW, Eidemiller LR, Szumowski J. Femoral MR angiography versus conventional angiography: preliminary results. Radiology. 1993;189(1):181–184. 154. Mulligan SA, Matsuda T, Lanzer P, et al. Peripheral arterial occlusive disease: prospective comparison of MR angiography and color duplex US with conventional angiography. Radiology. 1991;178(3):695–700. 155. Yucel EK, Kaufman JA, Geller SC, Waltman AC. Atherosclerotic occlusive disease of the lower extremity: prospective evaluation with two-dimensional time-of-flight MR angiography. Radiology. 1993;187 (3):637–641.

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CHAPTER 35

Pulmonary Artery Cardiovascular Magnetic Resonance David Grand and David A. Bluemke

Pathologies involving the pulmonary arterial system include pulmonary embolus, pulmonary arterial hypertension, and congenital anomalies. These diseases present a unique clinical diagnostic challenge often requiring multiple diagnostic examinations spanning the entire radiologic armamentarium. Initially, cardiovascular magnetic resonance (CMR) played only a minor role in evaluation of the pulmonary arterial system. However, with the advent of improved, high-performance gradients and bolus gadolinium contrast-enhanced magnetic resonance angiography (CE-MRA), CMR is now uniquely suited for this challenge. It not only provides precise, three-dimensional (3D) anatomic information, but also provides functional data that are critical to a complete evaluation of the pulmonary arterial system. Noninvasive and requiring neither iodinated contrast nor exposure to ionizing radiation, CMR consistently demonstrates anatomic accuracy comparable to that of computed tomography (CT) while providing reproducible functional data previously available only by invasive X-ray catheterization.

PULMONARY EMBOLISM Pulmonary embolus remains a challenging and elusive diagnosis that currently requires a multimodality approach. Because the presenting signs and symptoms are nonspecific and overlap in a wide variety of common ailments, suspicion of pulmonary embolus is an extremely common indication for imaging. To date, no single noninvasive examination has consistently demonstrated sufficient accuracy to exclude the diagnosis. Therefore, the current workup typically includes two or more of the following tests: blood assays (e.g., D-dimer), contrast-enhanced CT angiography (CE-CTA) or ventilation/perfusion radionuclide scanning, Doppler ultrasound of the deep venous system, and occasionally the “gold standard”: invasive X-ray pulmonary angiography. From the resulting array of diagnostic information, the clinician must then assess the patient’s likelihood of having the disorder and initiate treatment. The importance of accurate and early diagnosis of pulmonary embolus is easily lost among stacks of requests for imaging and relatively low positive test results. The annual incidence in the United States has been estimated as high as 630,000 per year.1 Even more concerning is the diagnostic dilemma it presents. In one study of patients with major pulmonary embolus, the diagnosis was made 480 Cardiovascular Magnetic Resonance

ante mortem in only 30%.2 These data underscore both the difficulty clinicians face in establishing the diagnosis and the severity of the consequences if treatment is not initiated promptly. Catheter-based X-ray pulmonary angiography is considered the gold standard for the diagnosis of pulmonary embolism. However, this test is invasive and carries a low but real risk of vascular and renal complications. In addition, direct injection of contrast material into the pulmonary arterial system is dangerous and is contraindicated in patients with pulmonary hypertension. Since the pretest probability of pulmonary embolus is typically low, owing to nonspecific presenting signs and symptoms, an invasive procedure is certainly not ideal. Ventilation/perfusion scanning with nuclear radiotracers provides a noninvasive means of diagnosing pulmonary embolus. While quite specific when the results are normal and 88% sensitive when the scan shows high probability, only 41% of patients with pulmonary embolus have a high-probability scan.3 Doppler ultrasound is quite useful for detection of deep venous thrombosis, thought to be the predisposing condition for development of pulmonary embolus. However, a normal ultrasound does not exclude the diagnosis. Of patients with angiographically proven pulmonary emboli, only 50% have positive ultrasound studies.4 The use of CE-CTA to evaluate for suspected pulmonary embolus has exploded in recent years as scan times have grown shorter, the number of detector rows has increased, collimation has become increasingly thin, and the availability of CT has become widespread in the emergency department. Sensitivities and specificities as high as 94% have been reported.5 Widely regarded as a powerful new tool for diagnosis and very extensively utilized, CECTA is not optimal. The sensitivity decreases dramatically when small, subsegmental emboli are considered. Additionally, CT scanning requires potentially allergic and nephrotoxic iodinated intravenous contrast and exposes the patient to ionizing radiation. The latter is of particular concern in young adults and those who are referred for multiple CE-CTA scans because of a history of prior pulmonary embolism. Thus, although a number of imaging tests are available in the radiologic armamentarium, no single test meets our diagnostic needs for safety, ease, and accuracy. As CEMRA techniques were developed, optimized, and shown to be effective, there has been renewed interest in applying them to the clinical problem of pulmonary embolus. Unfortunately, access to CMR is relatively limited, particularly

A

from the emergency department and during off-hours. However, CMR has already been shown to be a powerful tool for diagnosing pulmonary embolism. As access to CMR increases and physicians become more comfortable with CE-MRA techniques, it has the potential to complement if not supplant currently utilized imaging studies. There are two fundamental and complementary approaches to CMR diagnosis of pulmonary embolus. The earliest approach builds on the model of radionuclide ventilation/perfusion imaging and utilizes contrast-enhanced CMR to evaluate pulmonary perfusion (Fig. 35-1), to identify segmental, wedge-shaped perfusion defects indicative of pulmonary embolism. Unfortunately, CMR lung imaging is fraught with difficulty both because the lungs have little proton density (and therefore generate little signal) and because the air within the lungs creates an extensive susceptibility artifact, further diminishing what little signal there was initially. Increasingly fast gradients as well as bolus gadolinium contrast enhancement have made CMR lung perfusion available for over a decade,6 and studies have demonstrated that CMR lung perfusion and nuclear medicine scintigraphy are equivalent for detection of perfusion defects.7,8 CE-MRA techniques have dramatically advanced with the advent of fast 3D acquisitions with bolus gadolinium administration (Fig. 35-2), displacing time-of-flight imaging. These techniques allow for data acquisition in a single breath hold with excellent spatial resolution and signal-tonoise ratio. Postprocessing with maximum intensity projections provides the user with infinite orientations from which to view the dataset. The CMR findings of acute pulmonary embolism include intravascular filling defects and abrupt cutoff of intravascular signal (Fig. 35-3). Studies to evaluate the efficacy of this technique have uniformly demonstrated excellent results4 that are comparable with those of CE-CTA for larger emboli and with accuracies falling for smaller, more distal emboli. Accuracy for small, subsegmental emboli can likely be improved by combined CE-MRA and CMR lung perfusion images.9 As a result, some have suggested that CMR could become a “one-stop shop” in the evaluation of suspected pulmonary emboli by combining the above CMR techniques with magnetic

35 PULMONARY ARTERY CARDIOVASCULAR MAGNETIC RESONANCE

Figure 35-1 A, Cardiovascular magnetic resonance (CMR) lung perfusion showing a lobar perfusion defect on the right lung apex (arrow). B, Computed tomography scan of the hilum in the same patient shows masslike consolidation of calcified lymph nodes (arrow) due to fibrosing mediastinitis around the right pulmonary artery.

B

Figure 35-2 Normal contrast-enhanced MRA of the pulmonary arteries showing excellent branch vessel detail.

resonance venography, a highly sensitive and specific technique for detection of deep venous thrombosis.10

PULMONARY ARTERY HYPERTENSION Pulmonary artery hypertension is an enigmatic progressive condition characterized by abnormally elevated blood pressure of the pulmonary artery circulation with varied etiologies but with a common, poor prognosis. Normally, the pulmonary arterial system is a low-resistance vascular bed. Pulmonary hypertension is present when the systolic and mean pressures in the pulmonary arteries exceed 30 mm Hg and 20 mm Hg, respectively. As pressures rise, Cardiovascular Magnetic Resonance 481

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LT UPPER PA

Figure 35-3 Acute and chronic pulmonary embolism. A, Left (LT) pulmonary artery (PA) branches are enlarged centrally but show distal pruning consistent with chronic pulmonary embolus with distal vessel pruning. B, Right (RT) pulmonary artery branches show abrupt cutoff (arrow) due to pulmonary embolus.

RT UPPER PA

RT MAIN PA

A

LT MAIN PA

B

patients experience an insidious onset of nonspecific symptoms, such as progressive dyspnea and exercise intolerance, which all too often are mistaken for signs of coronary artery disease, thereby delaying diagnosis and treatment. Early diagnosis is critical; one study demonstrated a 30% survival rate at 5 years among patients with a mean pulmonary artery pressure greater than 40 mm Hg but only 10% when the pressure exceeded 50 mm Hg.11 Establishing the diagnosis currently requires numerous tests interpreted in concert. CMR has the potential not only to accurately diagnose pulmonary hypertension and its impact on right ventricular (RV) mass and volumes (see Chapter 11), but also to differentiate its causes, triage patients with regard to potentially effective treatments, and monitor and predict outcome. The utility of CMR is quite favorable in comparison with echocardiography and right heart catheterization (Table 35-1). Although there are numerous causes of pulmonary artery hypertension, for simplicity they can be thought of as belonging to one of two broad categories based on their etiology: thromboembolic and nonthromboembolic.

Chronic thromboembolic pulmonary artery hypertension is a delayed sequela of acute pulmonary embolism, occurring in 0.1% to 0.5% of those who survive the acute event.12 Nonthrombotic pulmonary arterial hypertension can be the result of primary pulmonary hypertension, an idiopathic condition, or chronic hypoxia, which occurs in many common diseases, including chronic obstructive pulmonary disease, emphysema, obstructive sleep apnea, congestive heart failure, and cystic fibrosis. Treatment options differ significantly between thrombotic and nonthrombotic disease, underscoring the importance of accurate diagnosis and characterization. The techniques described in the previous section are well suited for diagnosis and characterization of pulmonary arterial hypertension. Findings on CE-MRA indicative of the disease include an enlarged (>28 mm) right main pulmonary artery diameter and abnormal proximal to distal tapering (“pruning”) of pulmonary vasculature (Figs. 35-3 and 35-4). In one study, these findings alone allowed the diagnosis to be established with a sensitivity of 89% and a 94% negative predictive value.13 CE-MRA also assists in

Table 35-1 Comparison of Cardiovascular Magnetic Resonance, Echocardiography and Right Heart Catheterization for Characterizing the Right Heart and Pulmonary Artery Parameter RV volumes RV ejection fraction RV stroke volume RV mass RV strain RV pressure RV remodeling (septal curvature) Tricuspid regurgitation Miscellaneous (pericardial effusion, pulmonary embolism) RA volume RA pressure PA size PA distensibility Quantitative lung flow

Cardiovascular Magnetic Resonance

Echocardiography

Right Heart Catheterization

þþþþ þþþþ þþþþ þþþ þþþ þ þþþ þþ þþþ

þþ þþ þþ þ þþ þþþ þþ þþþ þþ

þ þ þ þþþþ þ þ

þþþ þþ þþþ þþþ þþþ

þþ þþ þþ þ þ

þþþ þ þ

PA, pulmonary artery; RA, right atrium; RV, right ventricle. Source: Adapted from Benza R, Biederman R, Murali S, Gupta H. Role of cardiac magnetic resonance imaging in the management of patient with pulmonary arterial hypertension. J Am Coll Cardiol. 2008;52:1683–1692.

482 Cardiovascular Magnetic Resonance

A

C

35 PULMONARY ARTERY CARDIOVASCULAR MAGNETIC RESONANCE

Figure 35-4 Tetralogy of Fallot with situs inversus and pulmonary valve reconstruction. (A, B) Two phases from a time resolved pulmonary artery MRA obtained at 3-second intervals. In panel A, the enlarged left pulmonary artery (arrow) is identified, prior to the lung perfusion phase. In panel B, hypoperfusion of the left lung apex (arrow) is seen. C, Conventional high-resolution pulmonary MRA performed during a 15-second breath hold. There is considerable overlap of the pulmonary arteries by the aorta and pulmonary veins. D, Posterior view of the pulmonary vasculature. An enlarged pulmonary is present on the right side (thin arrow), with pruning of pulmonary vessels on the left side (thick arrow). E, Phase contrast images obtained perpendicular to the left pulmonary artery (arrow); the magnitude image is on the left, and the phase image is on the right. Measurement of blood flow to each lung can be performed and quantified with this method.

B

D

E

establishing the etiology of the hypertension. Findings associated with chronic thrombotic hypertension reflect the underlying pathology of the disease, recanalization, and organization of thrombus. CE-MRA demonstrates thrombotic wall thickening and irregularity as well as intraluminal webs. Surgical treatment, pulmonary thromboendarterectomy, can be performed only if thrombus is limited to proximal vessels. As multidetector CT has essentially supplanted catheter angiography for the clinical diagnosis of pulmonary embolism, it has quickly become the study of choice for evaluation of pulmonary hypertension.

However, studies have shown equal accuracy of CMR when compared with CE-CTA in identifying the critical findings for establishing the diagnosis and determination of treatment options in chronic thromboembolic pulmonary arterial hypertension.14 CMR perfusion imaging can also be used to diagnose and characterize pulmonary arterial hypertension. Patients with pulmonary arterial hypertension typically demonstrate perfusion defects that characteristically are segmental in thrombotic hypertension and patchy and/or diffuse in nonthrombotic hypertension (see Fig. 35-4A and 35-4B). Cardiovascular Magnetic Resonance 483

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As was previously discussed, CMR lung perfusion has been shown to have equivalent accuracy in detecting perfusion defects in the detection of acute pulmonary embolus. For detection of chronic changes, the accuracies are again comparable. As with acute pulmonary embolism, combined CMR lung perfusion and CE-MRA distinguish chronic thrombotic from primary pulmonary arterial hypertension with an accuracy of 90%.15 CMR assessment of the pulmonary arteries not only can accurately diagnose pulmonary arterial hypertension and differentiate its cause, but also can provide functional assessments which can help guide treatment and predict

response. CMR can provide structural and quantitative functional analysis of the RV, including RV ejection fraction (see Chapter 11) and mass, which are not readily obtained by echocardiography. CMR data also provide data regarding pulmonary arterial hypertension by the configuration of the interventricular septum which bows into the left ventricle with increasing right heart pressures.16 The amount of curvature is related linearly to the right heart pressure as described by Laplace’s Law.17 Pulmonary artery hypertension has also been associated with a vortex of blood flow in the main pulmonary artery on velocity encoded CMR (Fig. 35-5).18 Additionally, dilatation of the bronchial

PA

PA

PA

PV PV

PV RV RV

A

RV

B

C

E

F

Vortex

D

Blood moving continuously upward along anterior wall

Vortex

G

H

No blood moving continuously upward along anterior wall

I

Figure 35-5 Typical flow patterns in the right ventricular (RV) outflow tract at difference phases of the cardiac cycle in a patient with pulmonary artery (PA) hypertension. A, D, and G, A patient with manifest pulmonary hypertension. B, E, and H, A patient with latent pulmonary hypertension. C, F, and I, A healthy subject. At maximum outflow (A, B, and C), flow profiles are homogeneously distributed across the main pulmonary artery. In the manifest (D, G) and latent (E, H) pulmonary hypertension setting, a vortex was formed that is not seen in the corresponding normal subject. PV, pulmonary valve. Source: Adapted and reproduced with permission from Reiter G, Reiter U, Kovacs G, et al. Magnetic resonance–derived 3-dimensional blood flow patterns in the main pulmonary artery as a marker of pulmonary hypertension and a measure of elevated mean pulmonary artery pressure. Circ Cardiovasc Imaging. 2008;1:23–30. 484 Cardiovascular Magnetic Resonance

CONGENITAL VASCULAR DISORDERS A wide spectrum of congenital heart disease (CHD) and congenital vascular anomalies affect the pulmonary arteries. Precise anatomic delineation of the abnormality is essential to plan and perform operative therapy. Previously, diagnosis and characterization of complex CHD and vascular diseases has been the realm of cardiac catheterization and echocardiography. However, CMR has demonstrated the ability to accurately assess the anatomic and functional abnormalities critical for surgical planning without ionizing radiation or arterial puncture.27,28 (See Chapters 28, 29, 30, and 31.) Numerous studies have demonstrated the efficacy of CMR in precise anatomic delineation of complex CHD affecting the pulmonary arterial and venous circulation. This includes detecting and planning surgery for a wide range of disorders, including the spectrum of tetralogy of Fallot (see Fig. 35-4), hypoplastic left heart syndrome, hypoplastic pulmonary arteries, partial anomalous pulmonary venous return (see Fig. 35-6), and patent ductus arteriosus (Fig. 35-7). Postoperative complications of surgical repair affecting the aorta and pulmonary artery are also readily evaluated (Fig. 35-8) with near 100% accuracy when compared with invasive X-ray catheterization and have even delineated vascular anomalies that were not seen at catheterization but were later proven at surgery. With current advanced CMR hardware, the imaging time can be under 4 seconds.29–31 In addition to anatomic delineation, CMR can provide functional data that are critical for assessing complex congenital cardiac and vascular anomalies. Previously, the domain of perfusion scintigraphy or invasive dilution oximetry, velocity-encoded CMR has been proven to successfully quantify asymmetric lung perfusion, which is particularly useful in patients with Fontan circulation and in the postoperative assessment of patients who have

PDA

Figure 35-6 Partial anomalous pulmonary venous return. Draining vessels from the right lung apex return to the superior vena cava (long arrow). Owing to high pulmonary flow, the main pulmonary artery is significantly enlarged (short arrow).

Figure 35-7 CE-MRA demonstrating a patent ductus arteriosus (PDA) (arrow). Cardiovascular Magnetic Resonance 485

35 PULMONARY ARTERY CARDIOVASCULAR MAGNETIC RESONANCE

arterial system has been noted in patients with pulmonary arterial hypertension as a collateral pathway for providing blood to the lungs for gas exchange. The degree of collateralization not only is indicative of the degree of hypertension, but also has been shown to correlate with improved outcomes after surgery for thrombotic hypertension.19 Previously, the degree of bronchial arterial shunting could be measured only by an invasive contrast medium dilution technique. However, by using velocity-encoded CMR of the pulmonary arteries (see Fig. 35-4E), increased bronchial arterial shunting is seen in patients with chronic thrombotic hypertension, which resolved after surgery.20 Finally, late gadolinium enhancement at the RV attachment sites was utilized in patients with pulmonary hypertension (Fig. 35-6), which may be related to abnormal focal wall stressors.21,22 These techniques could be applied to aid in establishing the diagnosis and predicting positive surgical outcomes. CMR perfusion and CE-MRA techniques that have demonstrated the ability to detect acute pulmonary embolus are equally applicable to diagnosis of the anatomic changes of pulmonary arterial hypertension, including thrombotic wall thickening, increased pulmonary arterial diameter, abnormal vessel tapering, and segmental or patchy perfusion defects. Additionally, CMR allows the acquisition of quantitative functional data, including right heart volumes and function and systemic-pulmonary arterial shunting. In addition, CMR allows for serial assessment of patients with pulmonary artery hypertension undergoing therapy.23,24 That all of these data can be acquired during a single examination without ionizing radiation exposure or iodinated contrast material will likely propel CMR to the forefront of diagnosis of pulmonary arterial hypertension.25,26

VASCULATURE AND PERICARDIUM

Figure 35-8 Shaded surface reconstructions of contrastenhanced MRA in a patient with transposition of the great vessels. A, Typical anterior origin of aorta in patient prior to surgical correction. B, Follow-up examination (different patient) after surgery, showing focal pulmonary stenosis, a common complication of vascular switch procedure.

A

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undergone repair of tetralogy of Fallot.32–34 These data can be obtained during the same, single examination to provide a comprehensive assessment of the unique anatomy and physiology of CHD. CE-MRA has also been proven useful in the evaluation of vascular anomalies such as arteriovenous malformations (AVM). Pulmonary AVMs cause pulmonary-to-systemic shunting and a resultant decrease in oxygen saturation as well as a potential conduit for paradoxical embolism and systemic embolism/stroke. Embolization of these AVMs is the current mainstay of treatment, which requires precise preprocedural anatomic delineation of the anomaly or anomalies. CE-MRA has demonstrated excellent accuracy for preprocedural detection of pulmonary AVMs,35,36 which can be multiple and bilateral (e.g., when associated with Osler-Weber-Rendu syndrome). The lack of ionizing radiation and iodinated contrast also makes CMR ideally

suited for evaluation of pulmonary AVMs, as the lesions often require multiple preprocedural examinations and postprocedural follow-up examinations.

CONCLUSION CMR provides precise anatomic and functional information regarding the pulmonary arterial system in a single examination without iodinated contrast or exposure to ionizing radiation. Technological advances have overcome prior limitations and continue to achieve faster scan times while obtaining higher-resolution images. With the potential to diagnose a wide variety of pathologies in a single examination with extremely low risk, CMR may well emerge as the examination of first choice for evaluation of the pulmonary arterial system.

References 1. Dalen JE, Alpert JS. Natural history of pulmonary embolism. Prog Cardiovasc Dis. 1975;17(4):259–270. 2. Carson JL, Kelley MA, Duff A, et al. The clinical course of pulmonary embolism. N Engl J Med.. 1992;326(19):1240–1245. 3. Value of the ventilation/perfusion scan in acute pulmonary embolism. Results of the prospective investigation of pulmonary embolism diagnosis (PIOPED). The PIOPED Investigators. JAMA. 1990;263(20):2753–2759. 4. Gupta A, Frazer CK, Ferguson JM, et al. Acute pulmonary embolism: diagnosis with MR angiography. Radiology. 1999;210(2):353–359. 5. Blachere H, Latrabe V, Montaudon M, et al. Pulmonary embolism revealed on helical CT angiography: comparison with ventilation-perfusion radionuclide lung scanning. AJR Am J Roentgenol. 2000;174(4):1041–1047. 6. Amundsen T, Kvaerness J, Jones RA, et al. Pulmonary embolism: detection with MR perfusion imaging of lung: a feasibility study. Radiology. 1997;203(1):181–185. 7. Yilmaz E, Akkoclu A, Degirmenci B, et al. Accuracy and feasibility of dynamic contrast-enhanced 3D MR imaging in the assessment of lung perfusion: comparison with Tc-99 MAA perfusion scintigraphy. Clin Radiol. 2005;60(8):905–913. 8. Amundsen T, Torheim G, Kvistad KA, et al. Perfusion abnormalities in pulmonary embolism studied with perfusion MRI and ventilationperfusion scintigraphy: an intra-modality and inter-modality agreement study. J Magn Reson Imaging. 2002;15(4):386–394. 9. Seo JB, Im JG, Goo JM, et al. Comparison of contrast-enhanced CT angiography and gadolinium-enhanced MR angiography in the 486 Cardiovascular Magnetic Resonance

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detection of subsegmental-sized pulmonary embolism: an experimental study in a pig model. Acta Radiol. 2003;44(4):403–410. Obernosterer A, Aschauer M, Portugaller H, Koppel H, Lipp RW. Three-dimensional gadolinium-enhanced magnetic resonance angiography used as a “one-stop shop” imaging procedure for venous thromboembolism: a pilot study. Angiology. 2005;56(4):423–430. Riedel M, Stanek V, Widimsky J, Prerovsky I. Longterm follow-up of patients with pulmonary thromboembolism: late prognosis and evolution of hemodynamic and respiratory data. Chest. 1982;81(2): 151–158. Moser KM, Auger WR, Fedullo PF. Chronic major-vessel thromboembolic pulmonary hypertension. Circulation. 1990;81(6):1735–1743. Kruger S, Haage P, Hoffmann R, et al. Diagnosis of pulmonary arterial hypertension and pulmonary embolism with magnetic resonance angiography. Chest. 2001;120(5):1556–1561. Ley S, Kauczor HU, Heussel CP, et al. Value of contrast-enhanced MR angiography and helical CT angiography in chronic thromboembolic pulmonary hypertension. Eur Radiol. 2003;13(10):2365–2371. Nikolaou K, Schoenberg SO, Attenberger U, et al. Pulmonary arterial hypertension: diagnosis with fast perfusion MR imaging and highspatial-resolution MR angiography–preliminary experience. Radiology. 2005;236(2):694–703. Dellegrottaglie S, Sanz J, Poon M, et al. Pulmonary hypertension: accuracy of detection with left ventricular septal-to-free wall curvature ratio measured at cardiac MR. Radiology. 2007;243:63–69.

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have undergone a Fontan operation or bidirectional cavopulmonary connection: initial experience. J Magn Reson Imaging. 2007;25 (4):727–736. Prakash A, Torres AJ, Printz BF, Prince MR, Nielsen JC. Usefulness of magnetic resonance angiography in the evaluation of complex congenital heart disease in newborns and infants. Am J Cardiol. 2007;100 (4):715–721. Greil GF, Powell AJ, Gildein HP, Geva T. Gadolinium-enhanced threedimensional magnetic resonance angiography of pulmonary and systemic venous anomalies. J Am Coll Cardiol. 2002;39(2):335–341. Geva T, Greil GF, Marshall AC, Landzberg M, Powell AJ. Gadoliniumenhanced 3-dimensional magnetic resonance angiography of pulmonary blood supply in patients with complex pulmonary stenosis or atresia: comparison with x-ray angiography. Circulation. 2002;106 (4):473–478. Balci NC, Yalcin Y, Tunaci A, Balci Y. Assessment of the anomalous pulmonary circulation by dynamic contrast-enhanced MR angiography in under four seconds. Magn Reson Imaging. 2003;21(1):1–7. Powell AJ, Tsai-Goodman B, Prakash A, Greil GF, Geva T. Comparison between phase-velocity cine magnetic resonance imaging and invasive oximetry for quantification of atrial shunts. Am J Cardiol. 2003;91 (12):1523–1525, A1529. Roman KS, Kellenberger CJ, Farooq S, MacGowan CK, Gilday DL, Yoo SJ. Comparative imaging of differential pulmonary blood flow in patients with congenital heart disease: magnetic resonance imaging versus lung perfusion scintigraphy. Pediatr Radiol. 2005;35(3):295–301. Fratz S, Hess J, Schwaiger M, Martinoff S, Stern HC. More accurate quantification of pulmonary blood flow by magnetic resonance imaging than by lung perfusion scintigraphy in patients with fontan circulation. Circulation. 2002;106(12):1510–1513. Khalil A, Farres MT, Mangiapan G, Tassart M, Bigot JM, Carette MF. Pulmonary arteriovenous malformations. Chest. 2000;117(5): 1399–1403. Ohno Y, Hatabu H, Takenaka D, Adachi S, Hirota S, Sugimura K. Contrast-enhanced MR perfusion imaging and MR angiography: utility for management of pulmonary arteriovenous malformations for embolotherapy. Eur J Radiol. 2002;41(2):136–146.

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17. Roeleveld RJ, Marcus JT, Faes TJ, et al. Interventricular septal configuration at MR imaging and pulmonary arterial pressure in pulmonary hypertension. Radiology. 2005;234(3):710–717. 18. Reiter G, Reiter U, Kovacs G, et al. Magnetic resonance–derived 3dimensional blood flow patterns in the main pulmonary artery as a marker of pulmonary hypertension and a measure of elevated mean pulmonary artery pressure. Circ Cardiovasc Imaging. 2008;1:23–30. 19. Kauczor HU, Schwickert HC, Mayer E, Schweden F, Schild HH, Thelen M. Spiral CT of bronchial arteries in chronic thromboembolism. J Comput Assist Tomogr. 1994;18(6):855–861. 20. Ley S, Kreitner KF, Morgenstern I, Thelen M, Kauczor HU. Bronchopulmonary shunts in patients with chronic thromboembolic pulmonary hypertension: evaluation with helical CT and MR imaging. AJR Am J Roentgenol. 2002;179(5):1209–1215. 21. Blyth KG, Groenning BA, Martin TN, et al. Contrast enhanced-cardiovascular magnetic resonance imaging in patients with pulmonary hypertension. Eur Heart J. 2005;26(19):1993–1999. 22. Sanz J, Dellegrottaglie S, Kariisa M, et al. Prevalence and correlates of septal delayed contrast enhancement in patients with pulmonary hypertension. Am J Cardiol. 2007;100:731–735. 23. Roeleveld RJ, Vonk-Noordegraaf A, Marcus JT, et al. Effects of epoprostenol on right ventricular hypertrophy and dilation in pulmonary hypertension. Chest. 2004;125:572–579. 24. Reesink HJ, Marcus JT, Tulevski II, et al. Reverse right ventricular remodeling after pulmonary endarterectomy in patients with chronic thromboembolic pulmonary hypertension: utility of magnetic resonance imaging to demonstrate restoration of the right ventricle. J Thorac Cardiovasc Surg. 2007;133:58–64. 25. Du J, Bydder M. High-resolution time-resolved contrast-enhanced MR abdominal and pulmonary angiography using a spiral-TRICKS sequence. Magn Reson Med. 2007;58(3):631–635. 26. Nael K, Fenchel M, Krishnam M, Finn JP, Laub G, Ruehm SG. 3.0 Tesla high spatial resolution contrast-enhanced magnetic resonance angiography (CE-MRA) of the pulmonary circulation: initial experience with a 32-channel phased array coil using a high relaxivity contrast agent. Invest Radiol. 2007;42(6):392–398. 27. Goo HW, Yang DH, Park IS, et al. Time-resolved three-dimensional contrast-enhanced magnetic resonance angiography in patients who

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CHAPTER 36

The Pericardium: Normal Anatomy and Spectrum of Disease Susan B. Yeon and Noriko Oyama

Understanding the anatomy of the pericardium as well as the interaction between pericardial pathology and cardiovascular function is key to detecting and assessing pericardial diseases. A comprehensive assessment of pericardial disease may require integration of dynamic structural and functional indicators of cardiac performance. Clinical evaluation after history and physical examination in subjects with suspected pericardial disease may include electrocardiography (ECG); chest X-ray; clinical laboratory tests; invasive hemodynamic assessment via cardiac catheterization; transthoracic echocardiography (TTE), including Doppler assessment of ventricular inflow and tissue velocity; computed tomography (CT); and cardiovascular magnetic resonance (CMR). Structural and functional characterization provided by CMR combined with other clinical assessment may be useful in diagnosing and managing patients with pericardial disease.

IMAGING MODALITIES Several imaging methods are used to detect and assess pericardial disease. CMR may complement other available imaging methods. Chest X-ray is frequently the first imaging test that suggests the presence of pericardial disease and may prompt further studies. An enlarged cardiac silhouette may suggest the presence of pericardial effusion, although this finding has limited sensitivity and poor specificity.1 The presence of pericardial calcification on chest X-ray may suggest constrictive pericarditis, although this finding is not necessarily indicative of clinical constriction and frequently is not present in cases of constriction.2 Calcification was not seen on chest X-ray in more than 70% of patients with constriction in the 1998 to 1995 Mayo Clinic series.3 The imaging technique most frequently used for the initial evaluation of pericardial disease, particularly cases of suspected effusion and tamponade, is TTE. Echocardiography, including Doppler examination, is an excellent portable tool for identification and assessment of pericardial disease, particularly effusion and tamponade. However, restricted acoustic windows and suboptimal views may limit the usefulness of echocardiography for identification of some loculated pericardial effusions and hematomas and may prevent acquisition of adequate images in certain subjects.4,5 Also, pericardial thickness is not reliably assessed by TTE; it can be better assessed by transesophageal echocardiography (TEE), although the field of view is still limited to available acoustic windows.6 488 Cardiovascular Magnetic Resonance

CMR and CT both provide a larger field of view than do TTE and TEE, enabling examination of the entire chest, including the mediastinum and lungs.7,8 They both provide excellent anatomic definition in standardized views, without limitation to patient-specific windows, as in TTE. Detailed imaging of the pericardium with CT or CMR requires the use of ECG gating or fast imaging techniques to minimize blurring as a result of motion. However, ungated CT or CMR studies performed for other indications occasionally uncover some of the more prominent pericardial abnormalities, such as large pericardial effusions.2 CT or CMR may be useful when findings on echocardiography are nondiagnostic, particularly to assess pericardial thickness in cases of suspected constriction and occasionally to look for loculated pericardial effusions or hematomas.4,8 An advantage of CT over other modalities is its ability to detect pericardial calcification, a frequent finding in constrictive pericarditis.7,9 Disadvantages of CT include the use of ionizing radiation and the need for intravenous iodinated contrast for adequate blood-tissue contrast. Also, in some cases, CT images may not adequately discriminate pericardial fluid from thickened pericardial tissue.7 Electron beam CT is less widely available, but offers the advantage of shortened scan times (<50 to 100 msec), limiting blurring as a result of cardiac motion. Using this technique, the upper limit of normal pericardial thickness may be as low as 2 mm compared with the conventional CT upper limit of 3 to 4 mm.9,10 The high spatial and temporal resolution of electron beam CT also facilitates identification of the pericardial sinuses and recesses.11 The development of multidetector CT (MDCT) offers fast scanning times with improved spatial resolution.12 As with electron beam CT, thin-section CT obtained via MDCT technology is advantageous in detecting pericardial sinuses and recesses, particularly smaller ones.13,14 CMR provides comprehensive images of the pericardium without the use of iodinated contrast or X-ray radiation. CMR may provide some advantages over CT and echocardiography with respect to its ability to characterize pericardial effusions and pericardial masses.5,15,16 As noted earlier, the study of pericardial disease often requires assessment of the functional effect of pathologic changes. CMR is well suited for pericardial assessment because it can provide both morphologic and functional information in a single examination. A combination of CMR sequences can be used to characterize the extent and composition of pericardial changes and provide functional assessment, including chamber compression, interventricular septal

NORMAL ANATOMY The pericardium encloses the heart and adjacent portions of the great vessels.17 It consists of fibrous and serous components. The tough, fibrous outer pericardium is loosely attached to the sternum and costal cartilage and more firmly attached to the central tendon of the diaphragm.18,19 The serous pericardium consists of an inner visceral layer adhering directly to the heart and forming its outer covering (the epicardium) and an outer parietal layer that lines the fibrous pericardium. A film of clear pericardial fluid (15 to 50 mL) normally separates the two serosal surfaces.2,19 The serosal layers merge at two complex parietovisceral lines of reflection, one at the base of the aorta and pulmonary trunk and the other at the insertions of the superior and inferior vena cavae and the four pulmonary veins.17 The pericardial complex generally lies between variable amounts of epicardial and pericardial adipose tissue.7 Although the pericardium is not required for normal cardiac function, several roles in promoting homeostasis have been ascribed to it. Among these are: (1) limitation of intrathoracic cardiac displacement; (2) maintenance of normal ventricular compliance; (3) balancing right ventricular (RV) and left ventricular (LV) output over several cardiac cycles through diastolic and systolic interactions; (4) buffering of changes in chamber filling and therefore output; (5) aiding in atrial filling by providing more negative pericardial pressure during ventricular ejection; (6) limitation of acute dilation; (7) minimizing friction between the cardiac chambers and surrounding structures; and (8) serving as an anatomic and immunologic barrier to inflammation and infection from contiguous structures, such as the lungs.15,19 However, as discussed later, the presence of an intact pericardium may be detrimental in situations in which fluid fills the pericardial space rapidly, producing tamponade. The pericardial cavity is a complex space, consisting of the pericardial cavity and interconnecting cul-de-sacs Figure 36-1 Normal pericardium. A, Axial electrocardiographic gated spin echo CMR image with pericardium of normal thickness (arrows). Note the epicardial and anterior mediastinal fat outlining the pericardium. B, Similar image at a higher axial level in another subject. As is often the case, the pericardium (arrows) is not well visualized over much of the left ventricle (LV) in the absence of pathology because of lack of adjacent pericardial fat to provide tissue contrast. DA, descending aorta; RA, right atrium; RV, right ventricle.

known as sinuses (as recognized by the Terminologia Anatomica) that have been further subdivided into recesses.11,20,21 As noted earlier, the reflections of the serous pericardium are arranged as two complex “tubes,” one enclosing the base of the aorta and the main pulmonary artery and the other enclosing the vena cavae and the four pulmonary veins.21 The cul-de-sac delimited by the inferior border of the tube surrounding the veins is the oblique sinus. The oblique sinus lies posterior to the left atrium (LA) and includes the posterior pericardial recess.11 The transverse sinus is the space between the two pericardial tubes. The transverse sinus is divided into four recesses: superior aortic, inferior aortic, left pulmonic, and right pulmonic.11 The pericardial cavity includes the postcaval, left pulmonary venous, and right pulmonary venous recesses.11 The normal pericardium is seen as a very thin linear density surrounding the heart. Discrimination of the pericardium from the myocardium on radiologic images requires the presence of interposed epicardial fat or pericardial fluid, with associated enlargement of the pericardial space.18 The amount of epicardial fat correlates with waist circumference, diastolic blood pressure, and fasting insulin level,22 and is significantly related to obesity-related insulin resistance.23 The pericardium is best visualized over the RV and may not be visualized over much of the LV, where there may be less epicardial and pericardial fat to provide adjacent tissue contrast7 (Fig. 36-1B). In addition, technical factors, such as spatial and temporal resolution, may affect the measurement of pericardial thickness because the normal pericardium is a thin, irregularly shaped sac that moves with cardiac and respiratory motion. Such motion affects the pericardial recesses to varying degrees: the superior aortic recess and left pulmonic recess are not surrounded by mobile structures and so are less affected by cardiac motion, whereas the inferior aortic recess, left pulmonary venous recess, and right pulmonic recess abut the LA or LV and are often blurred because of motion.13 The superior aortic recess of the transverse pericardial sinus, which surrounds the ascending aorta, sometimes may be mistaken for an aortic dissection or lymphadenopathy11,24 (Fig. 36-2). The oblique pericardial sinus, which is situated behind the LA, may simulate abnormalities in the esophagus, descending thoracic aorta, and

RV RA

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36 THE PERICARDIUM: NORMAL ANATOMY AND SPECTRUM OF DISEASE

motion, and filling velocity.16 Administration of conventional extracellular gadolinium-based contrast agents (as an off-label use) may assist in the identification of inflammation or neoplasia.

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AA PA

RPA

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DA

CMR and anatomic measurements is likely caused by a combination of volume averaging, motion artifact, and inclusion of small amounts of pericardial fluid, as well as chemical shift artifact for CMR.16,18,30 However, unlike CT, a combination of imaging sequences (spin echo and gradient echo) permits differentiation of small pericardial effusions from pericardial thickening.7,16,18,30 On gradient echo imaging (including steady-state free precession), the pericardium appears hypointense, in contrast to the hyperintensity of pericardial fluid16,18 thus, even normal, trivial amounts of pericardial effusion and loculated effusions can be detected readily. Pericardial thickness on gradient echo images is greater than on spin echo images. In addition, there is greater overlap in pericardial thickness between control subjects and patients with constrictive pericarditis with gradient echo compared with spin echo imaging.27,31 Therefore, spin echo imaging is the standard CMR method recommended for measurement of pericardial thickness.

PERICARDIAL DISEASES Figure 36-2 Normal aortic recesses. Axial electrocardiographic gated spin echo CMR images show the normal positions of the superior aortic recesses (white arrows) posterior and lateral to the ascending aorta. The left pulmonic recess (black arrow) lies laterally to the main pulmonary artery. AA, ascending aorta; DA, descending aorta; LPA, left pulmonary artery; PA, pulmonary artery; RPA, right pulmonary artery.

subcarinal and bronchopulmonary lymph nodes.14,25 Recognizing the appearance of these normal structures is important to avoid mistaking them for mediastinal disease processes.11,14 The reported upper limit for the thinnest portion of normal pericardium seen on CT scan varies among studies using various slice thicknesses and acquisition times. The upper limit (mean þ 2 SD) for the thinnest portion of the pericardium was reported as 3.4 mm using a 10-mm slice thickness and 9.6-second acquisition time10; it was reported as 1.2 mm (using a 10-mm slice thickness) and 0.7 mm (using a 1-mm slice thickness) with a 0.75-second acquisition time.26 As noted earlier, using electron beam CT, the upper limit of normal pericardial thickness was reported as 2 mm using a 3-mm slice thickness and a 100-msec acquisition time. A potential pitfall in determining pericardial thickness with CT is that trivial pericardial effusions as well as partial volume effects may produce the appearance of localized thickening.26 These imaging effects, along with normal intrinsic inhomogeneities in pericardial thickness, produce significant variations in pericardial width that should not be mistaken for disease. Thus, maximum pericardial thickness seen on CT among normal subjects was found to vary widely (mean 3.3 mm, SD 1.2 for 10-mm slice thickness; mean 2.8, SD 1.0 for 1-mm slice thickness).26 With CMR, the normal pericardium appears as a linear low-intensity signal between high-intensity mediastinal and subepicardial fat (see Fig. 36-1), and it is best visualized during systole.27 The thickness ranges from 1.2 0.5 mm in diastole to 1.7 0.5 mm in systole.27 These measurements are slightly higher than the reported anatomic thickness of normal pericardium of 0.4 to 1.0 mm.28,29 Like CT scan, the discrepancy between 490 Cardiovascular Magnetic Resonance

Pericardial Defects Congenital pericardial defects are uncommon. A spectrum of abnormalities exists, ranging from small defects in various locations to total absence of the pericardium (very rare).32 The most common defect is absence of the entire left side of the pericardium. Pericardial agenesis may be associated with congenital heart disease in a minority of cases, including atrial septal defect, patent ductus arteriosus, and tetralogy of Fallot.29 Although pericardial defects are generally found incidentally, smaller defects of the left pericardium may cause potentially serious herniation and entrapment of portions of the heart and infarction.19,32 Chest X-ray may show evidence of cardiac displacement and interposition of lung between the aorta and the pulmonary artery or between the inferior cardiac border and the diaphragm.19 CT and CMR may be helpful in determining the presence or absence of a segment of pericardium; however, frequently, the pericardium is not well visualized over portions of the heart, so diagnosis may be difficult.7 The diagnosis may be made by CT or CMR by noting absence of the superior aortic pericardial recess8 or interposition of lung tissue between the aorta and the main pulmonary artery.8

Pericardial Cysts Pericardial cysts are rare remnants of defective embryologic pericardial development.19 They are usually unilocular, smooth, thin-walled structures filled with clear transudative fluid, and may be attached, with or without a peduncle, to the pericardium.18,19 They occur most commonly at the right anterior cardiophrenic angle, but can be located throughout the mediastinum.7,18 Most patients are asymptomatic, with lesions detected incidentally on chest X-ray or other imaging. However, they may be associated with chest pain, dyspnea, cough, or arrhythmias, presumably as a result of compression of adjacent structures, and may therefore require aspiration.19

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Imaging of a suspected pericardial cyst generally distinguishes it from a mediastinal tumor, hematoma, or loculated pericardial effusion.19 The CT and CMR appearance of a pericardial cyst is a generally nonenhancing, wellcircumscribed, ovoid structure adjacent to the pericardium. The imaging characteristics of the contents are generally similar to those of water. Therefore, the cyst contents have low signal intensity on T1-weighted spin echo images and high, homogeneous intensity on T2-weighted spin echo images18,33 (Fig. 36-3). A cyst may occasionally contain highly proteinaceous fluid that has high signal intensity on T1-weighted images.33 Intracystic septae may be observed, especially after the administration of contrast.34 Cyst contents do not enhance with administration of conventional extracellular gadolinium-based contrast agents.8 A discriminatory feature of cysts is that they often change size and shape with respiratory motion or body position.18

Pericarditis Pericardial disease is associated with a large variety of clinical disorders that may arise locally or systemically.19,30,34,35 Such disease may present clinically as one of the following clinical syndromes in which the term pericarditis may be used: (1) acute pericarditis without evident effusion on clinical examination (although a small effusion is frequently present on imaging); (2) pericardial effusion without major hemodynamic compromise; (3) cardiac tamponade; and (4) constrictive pericarditis.19,36 Large series of patients undergoing extensive clinical evaluation for acute pericarditis have shown a cause in a minority of cases.37,38 Pericardial disease may occur in association with infection, radiation, autoimmune diseases, drug toxicity, diseases of contiguous structures (e.g., myocarditis, myocardial infarction), metabolic disorders (e.g., uremia, dialysis, myxedema), trauma (including iatrogenic causes), and neoplasms, as well as idiopathic causes.19,39 Signs and symptoms of acute pericarditis include chest pain (which may be pleuritic or positional), pericardial friction rub, and widespread ST elevation on electrocardiogram.19,36 Although pericardial effusion is frequently present in acute pericarditis, the absence of an effusion does not

B

exclude the diagnosis. As noted earlier, CT techniques are useful to identify pericardial thickening as well as pericardial effusion, although differentiation between pericardial thickening and small effusions may require reference to other imaging (e.g., TTE, CMR).7 Contrast administration, which is generally used to help delineate blood vessels and the cardiac chamber cavities from myocardium, may also show pericardial enhancement, which is nonspecific evidence of pericardial inflammation, infection, or neoplastic involvement.2 The use of CMR enables detection of pericardial thickening, effusion, and pericardial contrast enhancement (discussed later).2,39–41 Late (10 to 20 minutes after gadolinium-diethyl triaminepentaacetic acid administration) gadolinium enhancement may be useful for distinguishing the following groups: (1) subjects with normal pericardial histology, with or without pericardial effusion, were found to have no pericardial thickening or pericardial contrast enhancement on CMR; (2) subjects with histologic evidence of acute or chronic inflammation of the pericardium were found to have pericardial thickening and contrast enhancement (similar to Fig. 36-4); and (3) subjects with histologic evidence of chronic fibrosing pericarditis had pericardial thickening on CMR, with no enhancement and minimal pericardial fluid.40,41 However, as with CT, CMR contrast enhancement of the pericardium is a nonspecific finding that may indicate inflammation, infection, or tumor involvement.8 In patients being evaluated for metastatic disease, inflammation of the pericardium as a result of infection or other causes may produce nodularity, thickening, and contrast enhancement of the pericardium that may be difficult to distinguish from tumor involvement.41

Pericardial Effusions Pericardial effusion is defined clinically as accumulation of excess fluid in the pericardial sac. Such accumulation may be caused by a variety of causes, with most accumulations arising from the visceral pericardium.19 Types of pericardial effusion include hydropericardium (transudate associated with systemic fluid retention), inflammatory/ Cardiovascular Magnetic Resonance 491

36 THE PERICARDIUM: NORMAL ANATOMY AND SPECTRUM OF DISEASE

Figure 36-3 Pericardial cyst. A well-circumscribed, ovoid pericardial cyst (*) along the right side of the pericardium shows low signal intensity on T1-weighted CMR imaging (A) and high signal intensity on T2-weighted CMR imaging (B). LV, left ventricle; RV, right ventricle.

VASCULATURE AND PERICARDIUM

A

B

C

Figure 36-4 Pericarditis. A, Axial electrocardiographic gated spin echo T1-weighted CMR image showing diffuse thickening of the pericardium (arrows), with a maximal thickness of 7 mm. B, Midventricular short axis steady-state free precession CMR image showing hypointense pericardium (arrows). C, Corresponding late gadolinium enhancement image showing diffuse pericardial enhancements (arrows).

irritative effusion (the majority of significant effusions), hemopericardium, chylopericardium, cholesterol pericarditis, and rarely, lymphopericardium.19,35 Hemopericardium may consist of either frank blood (e.g., from traumatic wounds, including iatrogenic perforation, aortic dissection, or another bleeding source) or bloody effusion (e.g., uremia and many types of infections, particularly tuberculosis). Pericardial effusion may develop early (within days) in response to transmural or nearly transmural myocardial infarction or may develop late (1 week to several months) after myocardial infarction (Dressler syndrome). Similarly, pericardial effusions can be seen early (within days) or late (1 week to 2 months, known as post-pericardiotomy syndrome) after cardiac surgery. Although infectious pericarditis is most commonly viral or postviral (particularly because many idiopathic cases are attributed to this cause), nearly any infectious disease (e.g., bacterial, fungal, parasitic) may involve the pericardium. Viral or postviral effusions often resolve spontaneously, although they may continue as chronic relapsing pericarditis or progress to constriction. Malignant effusions are much more commonly the result of metastatic disease or local invasion rather than primary disease.42,43 Purulent pericarditis, an infection of the pericardial space associated with purulent pericardial fluid, has become less common since the advent of antiobiotics.44 It may be caused by bacterial or fungal pathogens, with Staphylococcus aureus being the most common.44–46 When purulent pericardial infection is suspected, drainage (preferably by subxiphoid pericardiotomy, given the thickness of the fluid and likelihood of adhesions) is generally required, along with intravenous antibiotic therapy.46 Cardiac tamponade causes life-threatening hemodynamic compromise through compression of the heart by accumulating pericardial contents.35,47,48 Tamponade can result from pericardial disease of nearly any cause. Progression to cardiac tamponade is determined by interactions among the following: (1) the rate of fluid accumulation in the pericardial space; (2) the rate of parietal pericardial stretching; and (3) the rate of systemic venous pressure increase to support filling of the right side of the heart.19 Because the pericardium has limited short-term capacity 492 Cardiovascular Magnetic Resonance

to stretch, rapidly accumulating pericardial fluid collections (as from hemorrhage) can produce tamponade at relatively low volumes (250 mL). On the other hand, more slowly accumulating effusions may be accommodated by pericardial stretching over time. Hence, large, slowly accumulating pericardial effusions (>1 to 2 L) may be tolerated, with little or no hemodynamic compromise.47 The symptoms of tamponade are nonspecific, so diagnosis requires careful rapid clinical assessment, including early imaging (usually by TTE). Prompt treatment is generally required via drainage of the effusion, most commonly via pericardiocentesis, although a surgical approach may be indicated when the effusion is loculated or recurrent.36 Pericardial effusions are often detected in subjects with suspected cardiac or pericardial disease as well as incidentally when imaging is performed for other indications. Because cardiac tamponade may occur with pericardial effusions of varying sizes as well as with loculated accumulations, the diagnosis of tamponade should be considered, even for smaller effusions and in cases in which the pericardial effusion or hematoma is not readily apparent on some standard views. Two-dimensional TTE (with Doppler interrogation) is the imaging method most commonly used to detect pericardial effusions, assess their hemodynamic significance, and guide therapeutic pericardiocentesis.49 Features of tamponade seen on TTE include significant respirophasic variation in mitral (>25%) and tricuspid (>50%) Doppler inflow velocities, diastolic collapse of the right atrium, left atrium, or RV, and dilation of the inferior vena cava.35,47,50–52 However, as noted earlier, suboptimal views may limit the usefulness of echocardiography in certain individuals, particularly those with certain loculated pericardial effusions and hematomas that may be difficult to visualize.4,5 On CT images, common transudative effusions have attenuation similar to that of water. Attenuation greater than that of water suggests a more complex, generally more proteinaceous fluid, as in malignancy, hemopericardium, purulent exudates, or myxedema.8 Using CMR, transudative effusions have longer T1 and T2 relaxation times and tend to appear dark on T1-weighted imaging, bright (high signal intensity) on T2-weighted

Figure 36-5 Pericardial effusion. Transudative pericardial effusion (arrows) shows greater signal intensity than epicardial fat (*) steady-state free precession CMR.

imaging, and bright on gradient echo cine images.2,5,8 Transudative pericardial effusions are often even brighter than epicardial fat on gradient echo images (Fig. 36-5).18 More proteinaceous effusions, such as exudates, have shorter T1 and T2 relaxation times and tend to show intermediate signal intensity on both T1- and T2-weighted images (although signal may be reduced by fluid mobility) and lower signal intensity than blood in the ventricular cavities on cine imaging.8,30,32 In hemorrhagic pericardial effusions, the appearance of blood will depend on the age of the collection.2 Fresh hematomas appear homogenously bright on T1- and T2-weighted images.2,8,53,54 Older collections show the effects of T1 and T2 shortening as a result of methemoglobin formation and generally show heterogeneous signal intensity, with areas of high signal intensity on both T1- and T2-weighted images.2,8,53–56 A chronic organized thrombus may have a dark rim and low signal intensity foci that may represent fibrosis, calcification, or hemosiderin deposition.2,8,54,55 The size of pericardial effusions can be readily assessed quantitatively by CMR using volumetric methods.33

Constrictive Pericarditis Constrictive pericarditis as a result of chronic fibrosis or calcification of the pericardial sac is associated with loss of compliance that impedes diastolic cardiac filling.19,36 Pericardial constriction has a wide variety of potential causes, including cardiac surgery, radiation, trauma, infection (tuberculosis; bacterial, particularly purulent; or viral), neoplasia, uremia, connective diseases, Dressler syndrome, sarcoidosis, and drugs. The majority of cases are of uncertain etiology.3,19 Constriction is typically a chronic process, although it can present within days to months after the initial injury, as has been described after cardiac surgery.19 Constrictive pericarditis presents with

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*

symptoms and signs of right-sided heart failure, including edema or anasarca, and fatigue.3,19 Clinically, it may be difficult to distinguish constrictive pericarditis from other causes of heart failure and edema, such as restrictive cardiomyopathy, although current imaging techniques facilitate differentiation of constriction. In both constrictive pericarditis and restrictive cardiomyopathy, ventricular filling is restricted, leading to an increase in diastolic pressure in the cardiac chambers. In restrictive cardiomyopathy, ventricular filling is limited by abnormal myocardial compliance and relaxation, whereas in pure constriction, myocardial relaxation is preserved.57 Although constrictive pericarditis is occasionally transient, it is generally a chronic, persistent condition that necessitates pericardiectomy for relief. However, pericardiectomy may not be indicated in patients with very mild disease or with severe advanced disease because the operative risk is excessively high.49,58,59 Effusive-constrictive pericarditis is an uncommon syndrome in which individuals with pericardial effusion and tamponade show clinical and hemodynamic evidence of pericardial constriction after normalization of intrapericardial pressure by drainage of the effusion.60,61 Effusiveconstrictive disease may be caused by any of the many causes of pericarditis, particularly radiation therapy.62 It is best shown with right heart catheterization at the time of pericardiocentesis to assess for residual elevation in right atrial and ventricular diastolic pressure after normalization of intrapericardial pressure.60 Although this syndrome frequently leads to persistent pericardial constriction that may require pericardiectomy, spontaneous resolution of idiopathic cases has been reported.62 Chest X-ray may show a small cardiac silhouette. Pericardial calcification may be detected in constriction (27% of cases in the 1985 to 1995 Mayo Clinic series3), although this finding is not diagnostic of clinical constriction. Pericardial calcification is often associated with idiopathic disease (67% of cases in the Mayo Clinic series) and is an independent predictor of increased perioperative mortality with pericardiectomy.3 As noted earlier, pericardial thickness is not reliably assessed by TTE, but can be determined using CT or CMR. However, the presence of pericardial thickening by itself does not indicate constriction.33 Also, pericardial thickness is not increased (by imaging and histologic examination) in a significant minority of patients with surgically proven constrictive pericarditis.63 When pericardial thickening is present in constriction, it may be localized. Nonspecific findings that may suggest constriction include ascites, pleural effusion, and occasionally some pericardial effusion. Often, there is dilation of the atria, coronary sinus, inferior vena cava, and hepatic veins.7 Echocardiography may identify flattening of diastolic LV inferolateral wall motion, abnormal septal motion (septal “bounce”; discussed later), and premature opening of the pulmonic valve.64 In addition, TTE may identify the nonspecific findings suggestive of elevated atrial pressure and venous congestion. Conventional Doppler imaging may show a restrictive LV inflow pattern, although this pattern may also be seen in restrictive cardiomyopathy and other conditions associated with high atrial pressure. Respiratory variation in transvalvular inflow patterns suggests constriction.65

VASCULATURE AND PERICARDIUM

Figure 36-6 Constrictive pericarditis. Axial contrast-enhanced computed tomography image (A) and electrocardiographic gated spin echo T1-weighted cardiovascular magnetic resonance image (B) showing focal pericardial calcification and thickening (arrows). The dilated right atrium (RA) and coronary sinus (CS) are suggestive of pericardial constriction. LV, left ventricle; RV, right ventricle.

RV RA LV CS

CS

A

B

Assessment of pericardial thickening with CT shows good agreement with histopathologic identification of pericardial thickening.63,66 As discussed earlier, CT is a sensitive test for the detection of pericardial calcifications. Because calcium produces high attenuation on CT, but a signal deficit on CMR (Fig. 36-6), CT can detect minute amounts of pericardial calcium, whereas CMR may miss significant deposits.7 As discussed earlier, on CMR, normal pericardium has low signal intensity on T1-weighted spin echo imaging, typically seen as a darker stripe between the bright layers of epicardial fat and fat around the pericardium.2,30 Thickened pericardium may have moderate to high signal intensity on spin echo images.8,30 The appearance on T2-weighted imaging is variable, but the signal intensity is usually lower than that of transudative fluid. Thickened pericardium shows intermediate signal intensity that is lower than that of transudative fluid on gradient echo (including steady-state free precession) sequences.30 Normal pericardial thickness seen with CMR is less than 3 mm.27 Thickness of greater than 4 mm indicates pericardial thickening and is strongly suggestive of constrictive pericarditis in the proper clinical setting. Constriction is frequently localized to the right side of the heart and may even be localized to the region of the right

atrioventricular groove.18,33,67 Larger pericardial calcifications may be visualized as regions of low signal intensity by CMR (see Fig. 36-6). The presence of late gadolinium enhancement CMR may suggest pericardial inflammation in effusive-constrictive pericarditis.40 Early diastolic septal flattening, giving the appearance of a septal “bounce,” is suggestive of constrictive pericarditis. This finding, originally noted on TTE, may also be detected by real-time CMR, and can be helpful in distinguishing constriction from restrictive cardiomyopathy.68 This appearance can be visualized on short or long axis cine CMR sequences.68 CMR tagging methods, such as spatial modulation of magnetization, may also be useful in diagnosing constriction. In normal subjects, when tagging stripes laid down in end-diastole are successively imaged during ventricular systole, the normal slippage between myocardium and pericardium results in the appearance of discontinuities, or breaks, in the stripes at the myocardium-pericardium interface. In patients with constrictive pericarditis with adhesion of the parietal pericardium, this slippage is lost in the affected regions. As a result of this “tethering,” tag lines passing through the myocardiumpericardium interface maintain continuity during systolic deformation (Fig. 36-7).2,69

RV LV

RA

LA

A

B

C

Figure 36-7 Constrictive pericarditis. Axial spin echo T1-weighted image showing a moderate-sized partially organized pericardial effusion (A; arrows), particularly prominent around the lateral wall of the left ventricle (LV). Four-chamber spatial modulation of magnetization images (B, end-diastole; C, end-systole) obtained at a similar level, but slightly oblique to A, show pericardial tethering (a lack of sliding at the myocardium-pericardium interface, as indicated by preservation of continuity in each stripe at arrows) with systolic motion of the lateral wall of the LV. LA, left atrium; RA, right atrium; RV, right ventricle. 494 Cardiovascular Magnetic Resonance

PERICARDIAL TUMORS Primary Pericardial Tumors Primary pericardial tumors are rare. Benign pericardial tumors include lipoma, teratoma, fibroma, neuroma, and hemangioma; malignant tumors include mesothelioma, lymphoma, thymoma (may be benign or malignant), sarcoma, and liposarcoma.8,19 Benign pericardial tumors found in children are often associated with large pericardial effusions.19 Primary malignant mesothelioma of the pericardium may cause pericardial effusion or pericardial plaques70 and may lead to pericardial constriction.19 Lymphoma, sarcoma, and liposarcoma typically appear as large, irregular masses, often associated with pericardial effusion.8 Lipomas can be readily recognized by their typical signal characteristics with CMR or CT.2 On CT, lipomas generally have low attenuation. On CMR T1-weighted spin echo images, they have characteristic high signal intensity8 that is not generally altered by contrast administration. Confirmation of the presence of fat signal is achieved by a decrease in signal intensity after application of a fat presaturation technique.34 Depiction of regions of calcium or fat in a pericardial mass on CT or CMR suggests teratoma.8 Fibromas more commonly arise from the pleura, but may arise from the pericardium. They are usually homogenous in appearance and appear isointense to hypointense compared with myocardium on T1-weighted images and hypointense on T2-weighted images.34,68 Fibromas may or may not show contrast enhancement.34,71–73 Hemangiomas are generally bright on T1- and T2-weighted images because of their content of slow-moving blood, and they show strong enhancement after contrast administration.34

Secondary Malignant Pericardial Tumors Secondary malignant pericardial tumors are much more common than primary pericardial tumors and have been found in 10% to 12% of all patients dying with malignancy.73,74 In 110 patients with cardiac metastases, at autopsy, the pericardium was involved in more than 70% of cases and pericardial effusion was found in approximately one third of those with pericardial involvement.74 Malignant pericardial involvement is often clinically silent and may be found incidentally, although in a significant minority of cases, it causes symptoms of

pericarditis, tamponade, or even constriction. Patients with large pericardial effusions who present with tamponade without signs of pericarditis (e.g., chest pain, rub, fever, electrocardiographic changes) may be more likely to have a malignant effusion than other patients with large pericardial effusions.75 Pericardial effusions caused by either malignancy or treatment of malignancy are more likely than other effusions to require repeat pericardiocentesis or surgical management.76 Malignant involvement of the pericardium may result from local invasion (as for lung, breast, esophageal and gastric cancers; lymphoma; thymoma; and pleural mesothelioma) or metastatic spread (as for breast, lung, melanoma, renal, and others).2,19,71,74 Lung cancer, breast cancer, esophageal cancers; melanoma, leukemia, and lymphoma are the malignancies most likely to metastasize to the pericardium.42,73–77 Patients with symptomatic malignant pericardial effusion generally have a poor prognosis, although those with breast cancer, leukemia, and lymphoma may have a better prognosis than others.77–79 Metastatic involvement of the pericardium may be suggested by echo, CT, or CMR. Findings include pericardial effusion and nodular pericardial thickening or pericardial mass.18 However, pericardial effusion and pericardial thickening in patients with malignancy may not be caused by malignant involvement of the pericardium because radiation, drugs, and idiopathic etiologies may all cause pericardial disease in this population.68,77 Hemorrhagic pericardial effusions are strongly suggestive of metastatic pericardial disease.7,75,77 Acute hemorrhagic effusions may be identified by their high signal intensity on T1- and T2-weighted spin echo images. Extension of local tumor to the pericardium may be confirmed by focal loss of the pericardial line, with or without associated effusion.7,30 An intact pericardial line may be seen if an adjacent tumor extends up to the pericardium, but not through it.8 Cine gradient echo CMR may help determine whether the tumor is adherent to the pericardium.71 Most malignant tumors enhance after gadolinium administration.8,72 On noncontrast imaging, most neoplasms have low signal intensity on T1-weighted images and high signal intensity on T2-weighted images. However, metastatic melanoma may have high signal intensity on T1and T2-weighted images because of the paramagnetic metals bound by melanin.79 Lymphomas appear iso- to hypointense to the myocardium on T1- and T2- weighted images.34 They may show lesser contrast enhancement in central regions that may be necrotic.8,80

CONCLUSION The spectrum of pericardial disease produces a variety of clinical presentations, ranging from asymptomatic disorders to severe hemodynamic compromise. CMR can serve as an important tool to characterize pericardial disorders to assist in their diagnosis and management.

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Ventricular filling patterns can be assessed by CMR using flow velocity encoding (phase contrast) sequences. In constrictive pericarditis, increased mitral level E-wave (early filling) may be observed as a consequence of increased diastolic pressure, whereas A-wave (atrial filling) may be of reduced height because of reduced late diastolic filling.16

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insertion of an epicardial pacemaker. Am Heart J. 1995;130: 1298–1300. Jungehulsing M, Sechtem U, Theissen P, et al. Left ventricular thrombi: evaluation with spin-echo and gradient-echo MR imaging. Radiology. 1992;182:225–229. Ferguson ER, Blackwell GG, Murrah CP, Holman WL. Evaluation of complex mediastinal masses by magnetic resonance imaging. J Cardiovasc Surg (Torino). 1998;39:117–119. Rajagopalan N, Garcia MJ, Rodriguez L, et al. Comparison of new Doppler echocardiographic methods to differentiate constrictive pericardial heart disease and restrictive cardiomyopathy. Am J Cardiol. 2001;87:86–94. Seifert FC, Miller DC, Oesterle SN, et al. Surgical treatment of constrictive pericarditis: analysis of outcome and diagnostic error. Circulation. 1985;72:II264–II273. Ling LH, Oh JK, Schaff HV, et al. Constrictive pericarditis in the modern era: evolving clinical spectrum and impact on outcome after pericardiectomy. Circulation. 1999;100:1380–1386. Hancock EW. A clearer view of effusive-constrictive pericarditis. N Engl J Med. 2004;350:435–437. Zagol B, Minderman D, Munir A, D’Cruz I. Effusive constrictive pericarditis: 2D, 3D echocardiography and MRI imaging. Echocardiography. 2007;24:1110–1114. Sagrista-Sauleda J, Angel J, Sanchez A, et al. Effusive-constrictive pericarditis. N Engl J Med. 2004;350:469–475. Talreja DR, Edwards WD, Danielson GK, et al. Constrictive pericarditis in 26 patients with histologically normal pericardial thickness. Circulation. 2003;108:1852–1857. Engel PJ, Fowler NO, Tei CW, et al. M-mode echocardiography in constrictive pericarditis. J Am Coll Cardiol. 1985;6:471–474. Hatle LK, Appleton CP, Popp RL. Differentiation of constrictive pericarditis and restrictive cardiomyopathy by Doppler echocardiography. Circulation. 1989;79:357–370. Oren RM, Grover-McKay M, Stanford W, Weiss RM. Accurate preoperative diagnosis of pericardial constriction using cine computed tomography. J Am Coll Cardiol. 1993;22:832–838. Masui T, Finck S, Higgins CB. Constrictive pericarditis and restrictive cardiomyopathy: evaluation with MR imaging. Radiology. 1992;182: 369–373.

Valvular Heart Disease Philip J. Kilner and Raad H. Mohiaddin

As an imaging modality, cardiovascular magnetic resonance (CMR) offers unrivaled versatility and freedom of anatomic access. In relation to heart valve disease, its relative strengths include the following: Depiction by cine imaging of valve movements and jet flow in planes, or stacks of planes, of any orientation Measurement of right as well as left ventricular volumes and mass by multislice cine imaging Measurement of volume flow and regurgitant fraction (pulmonary and aortic, at least) by phase contrast velocity mapping Assessment of the context and consequences of heart valve disease using the wide fields of view, multiple image slices, and the versatility of tissue characterization available to magnetic resonance. So although CMR is generally regarded as a secondline imaging modality after echocardiography for the assessment of heart valve disease,1 it can have important contributions to make toward decision making in regard to the timing and nature of surgical intervention, particularly in cases in which there have been inconclusive or conflicting findings, perhaps owing to inadequate echocardiographic access, or in which heart valve dysfunction is one aspect of more complex congenital or acquired pathology.2 And potentially, at least, CMR offers several possible methods for the measurement of valve regurgitation,3–6 although work is still needed to optimize and standardize acquisition protocols and to fully validate the techniques used. If CMR is to become established as a reliable second-line modality, several potential weaknesses or pitfalls need to be recognized and avoided or corrected: The slice thickness (typically 6 mm) and the dimensions of voxels (typically about 6 1.5 1.2 mm) of cine and velocity map acquisitions need to be borne in mind. The images are not usually acquired in real time but are reconstructed from data gathered over several heart cycles during a breath hold. For these reasons, valve leaflets are not always well seen, especially if there is arrhythmia. Nor is CMR effective for visualizing the smaller, more mobile vegetations of endocarditis, although it can be useful for identifying an abscess or false aneurysm. It is important to attempt to depict valve and jet structure by cine imaging in several planes and orientations, not just one; appearances and, potentially, interpretation can differ considerably between images of a particular valve acquired in different planes (Fig. 37-1).

Contiguous stacks of cine images are valuable for covering all parts of the mitral or tricuspid valves. The accuracy of phase velocity mapping cannot be taken for granted.7 For jet velocity mapping, the dimensions of the voxels can be large in relation to the size and shape of a narrow or fragmented jet, leading to possible inaccuracies due to signal loss, partial volume averaging, and other artifacts. For the measurement of volume flow, particularly for the calculation of regurgitant fractions, surprisingly large inaccuracies have been found to occur as CMR systems have been “improved” to allow more rapid acquisition in the last 10 years, due to eddy currents and concomitant gradients.7,8 The severity of the problem can vary considerably according to the hardware and software that are used and can change between the upgrades of a system. The uncertainties do not end there, however. The sequence parameters and imaging plane that are needed to be chosen appropriately relative to the characteristics of flow under investigation if artifacts are to be minimized. Measurements of biventricular volume and function are not necessarily as reliable and straightforward as has often been implied, particularly if there is arrhythmia or the right ventricle is structurally abnormal owing to congenital heart disease or previous surgery, and volume analysis remains time consuming and to some extent subjective. The methods of acquisition and analysis need to be specified appropriately for comparisons over time or between patients. In general, the unrivaled versatility of CMR is a strength, but it is also a challenge. There is an ongoing dilemma between continuing development on the one hand and standardization on the other. With these cautionary remarks in mind, the aim of this chapter is to describe some important underlying principles and to guide users to the CMR techniques that we and others have found the most informative. Stenotic and regurgitant lesions of each heart valve are considered below, and an overview of measures of the severity of valve dysfunction is given in Table 37-1, which is adapted from the 2006 ACC/AHA Guidelines for the Management of Patients with Valvular Heart Disease.1 It is important to realize that right-sided valve lesions, particularly pulmonary or tricuspid regurgitation, differ from their counterparts on the left. While echocardiography has become well established in the assessment of valvular lesions of the left heart, CMR, with phase velocity mapping, has particular strengths in relation to the assessment of valvular and perivalvular lesions of the right heart.

Cardiovascular Magnetic Resonance 501

37 VALVULAR HEART DISEASE

CHAPTER 37

FUNCTIONAL CARDIOVASCULAR DISEASE

Figure 37-1 Top left, A systolic frame of an SSFP cine acquisition shows the bright core and dark edges of the jet through the slitlike orifice of a stenosed bicuspid aortic valve. Top right, A phase contrast velocity map, encoded through-plane, shows the dark jet, recording a peak velocity of 4.2 m/sec. Bottom left, Steady-state free precession (SSFP) cine imaging aligned with the jet direction and orientated at right angles to the slit shows the narrower dimension of the jet, which appears bright and clearly delineated. Bottom right, The jet appears broad, diffuse, and dark in the SSFP cine, which is aligned with the length of the slit, its voxels spanning the width of the narrow jet.

Table 37-1 Classification of the Severity of Heart Valve Disease in Adults, Based Partly on the 2006 ACC/AHA Guidelines,1 Which Refer Mainly to Echocardiographic Indices Aortic Stenosis

Peak jet velocity (m/sec) Orifice area (cm2) Orifice area index (cm2/m2) Additional features

Mild

Moderate

Severe

<3 >1.5

3–4 1.0–1.5

>4 <1.0 <0.6 LV hypertrophy, post stenotic dilatation of the ascending aorta

Subaortic stenosis and coarctation should be considered and excluded. Pulmonary Stenosis

Peak jet velocity (m/sec) Valve orifice area (cm2) Additional features

Mild

Moderate

Severe

<3

3–4

>4.0 <1.0 RV hypertrophy, poststenotic dilatation of the MPA and LPA

Subpulmonary and pulmonary artery stenosis should be excluded. Mitral Stenosis

Peak jet velocity Valve orifice area (cm2) Additional features

Mild

Moderate

Severe

<1.2 >1.5

1.2–2.2 1–1.5

>2.2 <1.0 Possible evidence of pulmonary hypertension

Tricuspid Stenosis Valve orifice area (cm2)

<1.0 Continued

502 Cardiovascular Magnetic Resonance

Aortic Regurgitation Mild

Moderate

Severe

Regurgitant volume (mL/beat) Regurgitant fraction

<30 <30%

30–60 30–50%

Regurgitant orifice area (cm2)

<0.1

0.1–0.3

>60 >50% (CMR flow measurements tend to underestimate AR, unless corrected for aortic root motion) >0.3

Pulmonary Regurgitation (assuming near normal pulmonary resistance)

Regurgitant jet or stream width Valve leaflet appearances Regurgitant volume (mL/beat) Regurgitant fraction

“Free” or “Almost Free” (But May Be Well Tolerated)

Mild

Moderate

Narrow, <2 mm Mobile, coapting <30 <25%

moderate 2–5 mm Partly coapting 30–40 20–35%

Unobstructed reversed stream, >6 mm across Ineffective leaflets with wide failure of coaptation. >40 >35%, but modified by up- and down-stream factors. Free PR occurs mainly in the first half of diastole, typically followed by late diastolic forward flow if the right ventricle is full and conduit-like when the atrium contracts.

Mild

Moderate

Severe

Narrow, <1.0 mm Narrow core

1.0–2 mm Bright jet core >2 mm width 30–60 30–50% 0.2–0.4

>2 mm, with extensive jet or swirling LA flow

Additional features

Mitral Regurgitation

Regurgitant jet width No visible jet core

Regurgitant volume (mL/beat) <30 Regurgitant fraction <30% 2 <0.2 Regurgitant orifice area (cm ) Additional features Systolic flow reversal in pulmonary veins

>60 >50% >0.4 Dilated left atrium and pulmonary veins

Tricuspid Regurgitation (assuming near normal pulmonary resistance)

Regurgitant jet width

Mild

Moderate

Severe

Narrow, <2 mm

2–6 mm

>6 6 mm, measured on a through-plane velocity map Dilated right atrium and caval veins

Additional features

Measurements by CMR should generally be comparable to those by echo. They can be more accurate than echo where measurements of flow volume and/ or ventricular volume are used to quantify regurgitation, particularly pulmonary regurgitation. But CMR might not measure the peak velocity of a narrow or fragmented jet as accurately as Doppler echo. These jet velocities depend on the rate of flow as well as the amount of area reduction. Direct planimetry may be feasible by CMR, but only in cases in which the jet core, immediately downstream of an orifice, is coherent and of suitable size, shape, and location relative to the image voxels for clear delineation.

BASIC PRINCIPLES Slice Thickness and the Visualization of Thin Structures The visualization by CMR of heart valves and flow through them is still done mainly by imaging in slices, and the thickness of a CMR slice, for example 6 mm, is generally considerably greater than that of pixel dimensions, which might be 1.0 to 1.6 mm. This means that the image is effectively made up of a layer of relatively long, narrow voxels, as shown in Figure 37-2. A thin structure such as a valve leaflet or narrow jet will be visualized more clearly if it is

aligned parallel with the voxels rather than across the length of the voxels and will therefore be seen most clearly when the structure passes perpendicularly through plane of the image. This is one reason why valve leaflets are sometimes, but not always, visible in cine acquisitions. Awareness of this can help with the choice of imaging plane; for example, the leaflets of a closed mitral valve, particularly if it prolapses to the plane of the annulus or beyond, are likely to be seen clearly in certain long axis but not the short axis slices. And a flat jet through a slitlike orifice will be seen most clearly if the image plane cuts perpendicular to the line of the slit, although the jet will appear narrower in this orientation than if depicted, less clearly, in a plane aligned with the length of the slit (see Fig. 37-1). Cardiovascular Magnetic Resonance 503

37 VALVULAR HEART DISEASE

Table 37-1 Classification of the Severity of Heart Valve Disease in Adults, Based Partly on the 2006 ACC/AHA Guidelines, Which Refer Mainly to Echocardiographic Indices—Cont’d

FUNCTIONAL CARDIOVASCULAR DISEASE

6 mm

The Visualization and Planimetry of Jets by Cine Imaging Stenosis or regurgitation of a diseased heart valve is associated with jet formation and consequent dissipation of energy through shear and turbulence. Appreciation of the structure of jet flow is fundamental to the interpretation of jet appearances on cine images. Jet flow is associated with an upstream region of convergence of streamlines where flow accelerates into the orifice; then, in most situations, there is a relatively stable, high-velocity jet core, typically extending two to three orifice diameters beyond the orifice and, lateral and distal to the core, a parajet region of shear and turbulence (Fig, 37-3). The shear layer is the region with the most extreme spatial gradients of velocity, and this is where much of the turbulence of a jet originates and where there is likely to be most loss of signal in cine acquisitions and potential inaccuracies in phase velocity maps. It is the friction associated with shear and turbulence that dissipates fluid kinetic energy as heat, the fluid becoming less intensely turbulent as it loses kinetic energy and is swept downstream. Each of the characteristic zones of a jet may be identifiable on cine images, but the appearances depend not only on jet shape and structure, but also on the nature of the acquisition: the slice thickness, pixel dimensions and orientation, sequence design, and echo time9 and, 504 Cardiovascular Magnetic Resonance

Figure 37-2 Illustration of the importance of the orientation of a cardiac magnetic resonance image slice, consisting of long, narrow voxels, with respect to a thin structure. The boundaries of the structure will be depicted clearly only if it passes nearly perpendicular to the image slice. Source: Courtesy of Robert Merrifield.

importantly, on the location of the image slice relative to the jet. The relationships between flow features and the recovery of magnetic resonance signal are not straightforward. On the whole, coherent flow of the jet core is associated with local brightening of the signal on cine images. Acquisitions with longer echo times will be more subject to signal loss from regions of high shear and turbulence, while sequences with shorter echo times tend to be less susceptible to signal loss due to turbulence. Steady-state free precession (SSFP) imaging has become widely used in CMR because of the good contrast it gives between blood and adjacent tissues. Its properties also make it suitable for the delineation of jets, but appearances differ somewhat from those of gradient echo imaging. The echo times used for SSFP imaging are usually very short (e.g., 2 msec), thus minimizing signal loss due to the higher orders of motion associated with turbulence. But without velocity compensation, there is marked dephasing and consequent loss of signal in voxels that contain a wide range of velocities simultaneously, which means those that lie in the shear layer at the edge of a jet. SSFP sequences are also susceptible to slight inhomogeneities of the magnetic field, and there is more marked loss of signal in regions of flow disturbance if the magnetic field is slightly off shim. This can affect the appearance of flow in the vicinity of ferromagnetic prosthetic material such as that in the rings of mechanical heart valves. Under optimal conditions, however, typical appearances of a jet on SSFP imaging include a relatively

Coherent jet core Parajet shear layer

Turbulence swept downstream

Figure 37-3 The structure of jet flow as visualized using a suspension of metal pigment in water flowing continuously, left to right, in an open channel indented by a “stenosis.” In this case, the flow upstream and in the core of the jet is laminar. Turbulence generated in the parajet shear layer is swept downstream, dissipating the kinetic energy of the jet. Jet flow through heart valves, particularly through severely stenosed or regurgitant orifices, is likely to be more complex in structure than this, with possible splaying and fragmentation of the jet core and possible asymmetric attachment to one wall or valve leaflet. Source: Photograph: P. Kilner.

bright jet core, clearly delineated on either side by dark lines that represent the shear layers (Figs. 37-1 and 37-4). But this is the case only if there is a coherent core and the image plane is aligned with it. The same jet may appear dark, with no bright core, if the imaging plane is displaced slightly to one side, tangential to the shear layer. For this reason, careful alignment of planes, sequentially cross-cutting one cine acquisition with another and then another, is important in the assessment of heart valve disease by CMR. Each jet should be imaged in at least three planes (see Fig. 37-1), preferably six or more, before the operator can be reasonably confident that the size, shape, and functional significance of a jet have been adequately depicted. Gradient recalled echo (GRE) cine imaging using segmented k-space, also known as fast low-angle shot (FLASH)10 imaging, is generally velocity compensated and less subject to signal loss in the shear layer than SSFP

imaging is. For this reason, GRE cine imaging may be more suitable than SSFP for the direct measurements of orifice area by planimetry,11–15 as long as the echo time is sufficiently short for the high-velocity core of the jet, directed through the image plane, to appear bright, but more work is needed to determine optimal imaging parameters and, if SSFP imaging is used, exactly where the jet should be outlined relative to the dark shear layer. If planimetry is attempted, it is advisable to acquire a stack of slightly overlapping slices, shifting gradually from upstream to immediately downstream of the orifice in, say, 3-mm steps, to ensure that the minimum orifice/jet area is visualized. This approach is likely to be accurate only when the orifice is relatively simple in shape and the jet core is coherent. It is not likely to be accurate if the orifice is irregularly deformed or is a narrow slit, giving rise to a fragmented, splayed, or fanlike jet.7

RVOT LV RV RV

Figure 37-4 Double-chambered right ventricle, or subinfundibular stenosis, shown by steady-state free precession cine imaging in oblique coronal (left) and short axis (right) planes, with the systolic jet from the hypertrophied lower part of the right ventricle into the outflow tract arrowed. The pulmonary valve was not stenosed in this patient. LV, left ventricle; RV, right ventricle; RVOT, right ventricular outflow tract. Cardiovascular Magnetic Resonance 505

37 VALVULAR HEART DISEASE

Upstream convergence zone

FUNCTIONAL CARDIOVASCULAR DISEASE

Phase Contrast Velocity Mapping The principles of phase contrast velocity mapping are explained in Chapter 7. In brief, one or more of the directional components of velocity are encoded in the phase of the signal in each voxel of the image for each phase of the cardiac cycle.16 Velocity can be encoded horizontally or vertically (in the X- or Y-direction) in the image plane, but more often, and generally more reliably, the through-plane Z-component is encoded, this being in the direction of the slice select gradient. As with Doppler flow measurement, phase contrast velocity mapping is represented by shifts of phase and so is subject to aliasing. To avoid this, an appropriate velocityencoding range (Venc) should be chosen before acquiring velocity data in a particular cavity or vessel. An appropriate Venc exceeds the peak expected velocity by 20% to 50%, meaning that 180 of phase contrast corresponds to 120% to 150% of the expected peak velocity. If the Venc is too low, set below the level of the peak velocity, the result is aliasing, in which case the acquisition should be repeated with the Venc increased appropriately. Too high a Venc, however, reduces the sensitivity of velocity mapping by reducing the amount of velocity-related phase shifts relative to background noise or other artifact-related phase shifts. Velocity mapping is used clinically for jet velocity measurements and for volume flow measurements. Each can be subject to artifacts, generally of distinct kinds in each case, and these can give misleading clinical results if they are not recognized and, if possible, corrected.

Jet Velocity Mapping for the Assessment of Stenoses In jet flow, measurement of the peak velocity of the jet and the rates of change of the peak velocity through time or use of the velocity time integral17 may contribute to the assessment of the severity of stenosis or possibly regurgitation. Recovery of signal is a prerequisite for calculation of velocity, so it is necessary for voxels to lie sufficiently within the coherent core of a jet to measure its velocity accurately. If a jet is narrow and fragmented, this condition may be difficult to fulfil, as there may not be a core of sufficient dimensions. Initial breath hold cine acquisitions should be used, as described above, to decide whether it is worth trying to measure jet velocity by CMR. These, possibly followed by a preliminary in-plane breath hold velocity map, can then be used to locate a through-plane breath hold velocity map to measure the velocities of the core of the jet. Breath hold sequences are recommended, as their echo time will be short and so will minimize several potential artifacts,8 and any respiratory motion of the jet will also be avoided. What may be harder to avoid is movement of the valve related to cardiac motion. The plane chosen should therefore be located at the phase of peak flow, or, better still, a moving plane is located by using a motion-tracking method of acquisition, if available.18 In spite of these precautions, velocity maps, particularly of a narrow jet, may give misleading values. Of particular concern, voxels in shear layer 506 Cardiovascular Magnetic Resonance

at the edge of a jet are subject to signal loss and possible phase errors due to abrupt deceleration. Apparent velocities of edge pixels should therefore be treated with suspicion if they differ from those in the core of the jet, and they may degrade attempts to measure volume flow at or downstream of a narrow orifice. These problems explain why CMR velocity mapping is generally less suitable than Doppler echocardiography for the assessment of right ventricular (RV) or pulmonary artery pressures by measurement of the velocities of relatively narrow tricuspid or pulmonary regurgitant jets. In echocardiography, it became routine for jet velocity measurements to be converted to estimated differences of pressure, often referred to as gradients, in units (mmHg) that were familiar from the use of catheterization. This conversion relies on assumptions that may not always be applicable, for example, that all of the jet or stream reaches a uniform peak velocity and that virtually all of its kinetic energy is dissipated in turbulence. We recommend that peak velocities measured at specific locations, by either CMR or Doppler echocardiography, are quoted as such without conversion to presumed differences of pressure. The velocity of a jet through an orifice is, of course, affected by the rate of flow as well as by the degree of narrowing, so a low rate of ejection, for example, due to ventricular dysfunction, will reduce the degree of elevation of velocity through a stenosed outflow valve.

Regurgitant Flow Measurements Phase contrast velocity mapping is able to measure velocities in pixels throughout the plane of acquisition, which can be acquired in any chosen orientation. An advantage of this is that the volume of flow passing through an image plane can be calculated. This is done by multiplying the cross-sectional area of a vessel, for example, the aorta or pulmonary trunk, by the mean through-plane component of velocity in that area for each phase of the cardiac cycle, after delineating the area of the vessel in every frame of the acquisition. In this way, CMR is able to measure the volumes of forward or reversed flow in a large vessel. Furthermore, CMR velocity mapping is the only imaging technique that has the potential to acquire comprehensive information (i.e., three-dimensional, three directionally encoded, time-resolved acquisition) and is well suited for studying spatial and temporal patterns of flow in the human cardiovascular system.19 This can be of interest where flow is multidirectional in cavities or in vessels that are curved or dilated,20 but measurement of one directional component of velocity is adequate for most clinical applications. Through-plane flow measurements are subject to particular kinds of inaccuracy, however, especially when used for regurgitant or shunt flow rather than total flow measurements.8 Background phase offset errors can cause slight inaccuracies of velocity measurement across the whole field of view and through the whole cardiac cycle, typically increasing with distance from the center of the magnet. These errors have been identified as being due mainly to concomitant Maxwell gradients, which can be calculated and corrected by using appropriate software, and to eddy currents in hardware components, which

Fourier CMR Velocity Traces Another type of velocity acquisition that is worth mentioning, although not widely implemented, is CMR Fourier velocity mapping. In this technique, the spatial phase-encoding gradient pulses are replaced by velocity phase-encoding bipolar gradient waveforms. The resulting dataset has only a single spatial dimension (in the readout direction) and measures one component of velocity in the direction of velocity-encoding gradient.21 Like M-mode Doppler, this approach could be adapted to real-time acquisition and display of the time course of flow through a particular valve or orifice. Arguably, it should be made available for routine CMR investigation, for example, to assist in diagnosis of constrictive pericardial disease and diastolic ventricular dysfunction as well as in heart valve disease, although the technique has yet to be optimized and proven in clinical use.

STENOTIC HEART VALVE AND OUTFLOW TRACT LESIONS Aortic Valve Stenosis Stenosis at the level of the aortic valve can be due to congenital malformation of the valve, rheumatic heart disease or degenerative disease in older patients. On the whole, echocardiographic assessment is reliable, but CMR may be called upon when there is limited echocardiographic access, when there are conflicting findings, in complex disease with stenosis at more than one level, or for the assessment of left ventricular (LV) volume and mass. Ranges of measurements reflecting different severities of aortic stenosis and other valve lesions1 are shown in Table 37-1. Several publications on CMR quantification of aortic stenosis have used planimetry of the orifice, or rather the jet,

to make direct estimates of orifice area.12–15 The results seem to have been encouraging, although the spatial resolution, meaning slice thickness as well as pixel dimensions, of cine acquisitions may not be adequate when it comes to the delineation of narrow, splayed, or fragmented jets. In our experience, there is a proportion of orifices and jets that cannot be adequately defined by CMR planimetry. By echocardiography, direct planimetry on two-dimensional images, peak stenotic jet velocity and mean jet velocity are possible approaches to the assessment of the severity of stenosis, although planimetry may be unsatisfactory in calcified or deformed valves, and simple jet velocities depend on flow rate as well as stenotic severity. Probably the most widely accepted echocardiographic approach uses the continuity equation to calculate the area of the valve orifice.22,23 Using this, the volume flow through the left ventricular outflow tract (LVOT) is calculated by multiplying the LVOT velocity time integral (VTI) by the LVOT cross-sectional area (typically calculated echocardiographically from its diameter) to give outflow volume, then dividing this volume by the VTI of the jet at the stenosis, which gives a measurement representing the stenotic valve area. Comparable approaches using CMR measurements have been described.17,24 A stenosed bicuspid aortic valve generally has a slitlike orifice (see Fig. 37-1), whereas degenerate or calcified valves can have irregular or three-pointed orifices, and the considerations described earlier regarding jet shape and the locations of image slices are important. Cine imaging should aim to depict the jet in several orientations, which also help to determine the best strategy for jet velocity mapping and planimetry of the valve.11–15 Both of these should be attempted; unfortunately, each is less likely to be accurate if the orifice is small and irregular and the jet is fragmented. For direct planimetry of the valve orifice, highresolution breath hold FLASH or SSFP imaging in a stack of about five slices, shifting 3 mm at a time from immediately upstream to just downstream of the orifice, should be acquired. The best cine, which is probably the one that shows the smallest, most clearly delineated cross section of the jet, should then be used for direct planimetry. The jet velocity and orifice area might not be accurately measurable in severe stenosis or when the jet is fragmented. Multiple in-plane cine views together with assessment of the impact of the lesion of ventricular mass and function are particularly important in such cases and, in combination, should make the severity of the stenosis apparent. Cine imaging should show the location and direction of the jet core, and the orientation of the slit, if applicable. Inplane breath hold velocity mapping with the direction of velocity encoding aligned with that of the jet can be helpful to visualize the jet, but it is likely to underestimate peak velocity if the jet is narrow. The above acquisitions should then be used to align a velocity-mapping slice to transect the jet core, with velocity of an appropriate range (Venc) encoded through the image plane. The velocity of the jet should be analyzed with caution in the resulting image, particularly if the stenosis is severe. Signal loss can allow background noise or other artifacts to contribute spurious phase shifts, particularly in the shear region at the edge of a jet, so peak jet velocity should be measured in a region of good signal recovery within the core of the jet. Cardiovascular Magnetic Resonance 507

37 VALVULAR HEART DISEASE

are harder to predict and eliminate.7 The result can be an upward or downward shift of the baseline of the flow curve, which can significantly affect the calculation of a regurgitant fraction. The amount of the shift of the flow curve depends not only on the underlying cause of the artifact, but also on the Venc that is chosen (the higher the Venc, the greater the error), and on the cross-sectional area of the vessel measured (the larger the area, the greater the error). A further source of inaccuracy can be movement of the plane of a regurgitant valve itself relative to the plane of acquisition. Because such movements, for example, of the aortic valve in diastole or of the mitral valve in systole, displace a volume of blood in the direction opposite that of the regurgitant flow, they cause underestimation of the regurgitation measured, the error increasing with the amount of valve movement and the cross section of the flow area measured. It is an error that can be overcome by appropriately implemented motion tracking,18 although this is not yet available on most systems. Both manufacturers and users need to be aware of these potential sources of inaccuracy and work together toward the optimization of systems and sequences.

FUNCTIONAL CARDIOVASCULAR DISEASE

Subaortic Stenosis There are important anatomic differences between the left and the right ventricular outflow tracts that result in distinct types of subvalvular stenotic lesion. Two types of subaortic stenosis need to be distinguished. One, which is typical of hypertrophic cardiomyopathy is caused by narrowing in the proximal part of the outflow tract, between the hypertrophied basal part of the septum and the septal leaflet of the mitral valve. As the valve leaflet is mobile, its tip can be drawn still closer to the septum by the low pressure of the jet, a fluid dynamic phenomenon known as the Venturi effect, causing systolic anterior motion of the mitral valve, with probable vibration of the leaflet. Although vibration is rapid and unlikely to be visible directly, it may explain why velocity mapping tends to be unsuccessful when this type of stenosis is significant, whereas Doppler echocardiography is probably more reliable. Such stenoses tend to become very much more severe on exertion, owing to enhancement of the Venturi effect with rapid ejection, and may be associated with induced mitral regurgitation. Because administration of pharmacologic stress can be risky in such cases, with the patient lying in the bore of a magnet, echocardiography with exercise stress is a more appropriate approach to the assessment of severity. CMR is nevertheless valuable, giving good views of all parts of the ventricle and of any associated pathology of the LVOT. It is important to realize that LVOT and right ventricular outflow tract (RVOT) stenosis can be at more than one level and possibly at three distinct levels, including the valve, in a single patient. The second type of subaortic stenosis is caused by a subvalvular ridge or membrane, usually only a few millimeters below the level of the valve. It is advisable here, as elsewhere, to acquire cine images of jet flow in several orientations, especially as it can be hard to distinguish subvalvular from valvular stenosis. In-plane velocity mapping can help to establish the level of flow acceleration, followed by velocity mapping through two planes, each clearly labeled, immediately below and above the level of the valve. If subvalvular stenosis is present, the aortic valve leaflets, even if normal, cannot be expected to open fully, owing to the Venturi effect, so assessment of the aortic valve has to remain cautious when subvalvular stenosis is present unless it is clear that there is additional, more severe stenosis at the level of the valve itself.

Pulmonary and Other Stenoses of the Right Ventricular Outflow Tract Stenosis of the pulmonary valve generally presents as a congenital condition early in life, either isolated or associated with other abnormalities of the RVOT and pulmonary arteries, as in tetralogy of Fallot. Transcatheter or surgical relief of pulmonary stenosis is typically recommended when there are peak-to-peak transpulmonary catheter pressure differences of 40 mm Hg, or 30 mm Hg in symptomatic patients. This probably translates into 508 Cardiovascular Magnetic Resonance

peak velocity measurements of the order of 3 to 4 m/sec by CMR jet velocity mapping, but as was outlined above, jet velocity mapping can be subject to errors, particularly if acquisitions are not perfectly located. The strength of CMR in this setting is in the evaluation of associated pathophysiology, particularly of RV hypertrophy and dysfunction, and possible additional levels of stenosis, including the branch pulmonary arteries. RV to pulmonary artery conduits are also readily evaluated by CMR, although ferromagnetic valve rings or sternal wires may cause localized artifacts. Isolated pulmonary valve stenosis can also present later in life, although a more common progressive lesion in the adult population is double-chambered right ventricle, which can also be called subinfundibular stenosis25 (see Fig. 37-4). Double-chambered right ventricle is a diagnosis that can be misinterpreted on echocardiography as a ventricular septal defect, which frequently coexists, or as stenosis at infundibular or pulmonary valve level. The distinction is important and is usually clear on CMR imaging, as subinfundibular stenosis is treatable, if necessary, by transatrial surgical resection of the obstructing muscle bands without the need for pulmonary valve replacement. It is important to use the excellent access provided by CMR to visualize the level and nature of any RVOT stenosis as clearly as possible. RV to pulmonary artery conduits and the function of their valves can also be evaluated very effectively by CMR cine imaging and velocity mapping.

Mitral and Tricuspid Valve Stenosis Methods of assessment of stenosis are fairly similar for the mitral and tricuspid valves. Both can usually be adequately assessed by two-dimensional and Doppler echocardiography, and more recently by three-dimensional echocardiography, which can contribute useful morphologic information. The cross-sectional shape and area of the stenotic orifice can generally be seen and measured on CMR short axis cine acquisitions and phase velocity maps, oriented to transect the jet at right angles at or very close to the orifice on the downstream side. Motion tracking should be used if available. Planimetry of the stenotic mitral valves has been reported by using SSFP cine imaging26 and has been found to yield orifice areas slightly larger than those measured echocardiographically (e.g., 1.65 versus 1.5 cm2). Not only the area and velocity of the jet but also the time course of jet flow are important, with loss of the normal biphasic mitral flow pattern and prolonged elevation of velocity as stenosis increases in severity. The slope of the decline in the maximum velocity has been used to estimate the pressure half-time (the time it takes until the initial pressure difference across the orifice is halved).27 The time curve of pulmonary venous flow, with marked atrial systolic reversal, can also be informative.28 An important contribution of CMR is in the assessment of associated pathophysiology such as the degree of atrial dilatation, and possible changes of the pulmonary veins, pulmonary arteries, and RV if pulmonary hypertension has developed secondary to mitral stenosis.

General Principles In regurgitant heart valve disease, the optimal time for surgical intervention depends on a balanced assessment of when valve dysfunction is sufficiently severe to warrant the shortand long-term risks of surgery but before the heart muscle sustains irreversible damage. The form, structure, and mobility of regurgitant valve leaflets, particularly mitral and tricuspid, should be characterized with a view to surgical repair. Owing to the slice thickness of acquisitions, this may not be easy by CMR, and two-dimensional echocardiography and three-dimensional echocardiography are probably preferable in this respect.2 The orientations of CMR slices need to be optimized in relation to individual valve structures. The need for intervention is determined partly by the severity of the symptoms and objective clinical indices such as oxygen consumption and exercise capacity, but the measurements of regurgitant volume or regurgitant fraction and ventricular function are important. Measurement of valvular regurgitation is clinically important but is difficult to quantify with current echocardiographic and angiographic techniques. Several CMR methods are available to assess the severity of valvular regurgitation. Regurgitation of any single heart valve, if isolated and uncomplicated by shunting, may be calculated by comparison of the LV and RV stroke volumes measured by planimetry from a standard multislice short axis volume acquisition, preferably using software that can combine short axis data with information on the locations and movements of the mitral and tricuspid valve planes in long axis acquisitions.29 The ventricle with the leaking valve will have excess stroke volume that represents the volume of leakage. The accuracy of this approach is dependent on the image quality and accuracy of planimetry, which, if done manually, takes considerable postprocessing time. Automated methods have still to be developed and validated. Hypertrophy and deformity of the right ventricle, for example, in congenital heart disease or after surgery for it, present particular problems. Standardized approaches need to be taken to the inclusion or exclusion of myocardial trabeculations and any aneurysmal regions of the RVOT.30 If appropriately performed, CMR measurements of right and LV volumes should have greater accuracy and reproducibility than those attempted by any other technique.31 Where more than one valve is regurgitant, ventricular volume measurement can be combined with aortic or pulmonary phase contrast flow measurements to calculate regurgitation of the inflow valve. The systolic flow measured in the great vessel by velocity mapping is subtracted from the stroke volume of the associated ventricle (for example, mitral regurgitant volume equals the LV stroke volume minus systolic aortic flow).4 But this depends on the accuracy of techniques used, which may need to be validated for a particular CMR and image-processing system.

Aortic Regurgitation Uncomplicated aortic regurgitation can be relatively easily measured by through-plane velocity mapping, the velocitymapping plane being located to transect the aortic root just

above valve level, at about the sinotubular junction (Fig. 37-5). If the valve is unstenosed, a Venc of 2 m/sec should be adequate, and it is important not to set the Venc too high to maintain the sensitivity of the measurement. For this reason, mixed stenotic and regurgitant valve disease presents a problem, requiring a higher Venc in the systolic phase at least. If possible, the higher Venc should be used for the systolic frames only, and a lower, more sensitive Venc should be used for diastole, when the reversed flow back toward the regurgitant valve will be of low velocity. Flow measurement through a plane located higher across the ascending aorta is also possible but will be more subject to underestimation of regurgitant flow for the reason explained below. An important issue is the volume change of the section of aortic root beneath the velocity mapping plane through the course of the cycle. This volume change is related to ventricular contraction and the compliance of the aorta, the aortic root being pulled downward and expanding in systole and returning, piston-like, through the plane of acquisition in the diastolic phase when the regurgitation is being measured. This upward movement of the root accounts for apparent slight diastolic forward flow by velocity mapping in a healthy individual with a competent valve, and it causes underestimation of the regurgitant volume or fraction in almost all cases with aortic regurgitation. The error is small if the root is small and immobile, for example, after previous surgery, but can be significant if the root is dilated and ventricular function is good. The appropriate way to avoid this underestimation is to use motion tracking for the velocity acquisition.18 If this is not available, approximate calculation of the change of volume of the root from its cross-sectional area times the distance of valve plane movement may be attempted to give an idea of how much the underestimation of regurgitant flow is likely to be.

Pulmonary Regurgitation Although significant pulmonary regurgitation is relatively uncommon as an isolated or acquired lesion, it is commonly found as a residual lesion following repair of tetralogy of Fallot32,33 or valvotomy for congenital pulmonary stenosis. Pulmonary regurgitation, even if free, meaning that there is no effective valve function, is usually well tolerated for decades, but late morbidity and mortality can ensue, related to eventual deterioration of RV function and the onset of potentially fatal arrhythmias.34,35 Although uncontained by effective valve action, the pulmonary regurgitant volumes and regurgitant fractions can remain surprisingly low and considerably less than would be expected on the systemic side of the heart. Regurgitant fractions of 35% to 45% are, in our experience, typical of free pulmonary regurgitation. Higher values may be attributable to additional factors such as an unusually capacitant right ventricle on the upstream side35 and, on the downstream side, large and compliant pulmonary arteries, raised pulmonary vascular resistance at branch arterial or microvascular level, or combinations of these.36 In other words, the pulmonary regurgitant fraction, above a level of about 30%, reflects other aspects of right-sided pathophysiology, which should be assessed, as much as dysfunction of the valve itself. We suggest use of the terms free or almost free rather than severe pulmonary stenosis when there is not an effective valve and a regurgitant fraction below 50%. If the pulmonary regurgitant fraction exceeds 50%, however, it should be called Cardiovascular Magnetic Resonance 509

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500 400 300 200 100 0 0

200

400

600

800

1000

−100 −200

Figure 37-5 Moderate aortic regurgitation shown by steady-state free precession cine imaging (top) and phase contrast velocity mapping (bottom left and right). With respect to the oblique coronal LVOT image (top left), the black lines mark the plane chosen for imaging the valve leaflets, and the white lines show the plane chosen for through-plane flow measurement, close to the level of the sinotubular junction. The flow curve shows systolic forward and diastolic reversed flow through the velocity mapping plane, recording a regurgitant fraction of 32% in this case. This is likely to be an underestimate, probably by at least 5%, owing to upward diastolic movement of the dilated aortic root relative to the plane of velocity acquisition.

severe, and the contributory factors should be considered, for example, unusually compliant pulmonary arteries and/or elevated pulmonary vascular resistance. For pulmonary flow measurement, a suitable velocitymapping plane transects the pulmonary trunk, or the ventriculopulmonary conduit, above expected valve level and proximal to the bifurcation (Fig. 37-6). Viewing of RVOT cines in sagittal and oblique cross-cut orientations should help to determine the most appropriate level in a particular patient. A Venc of 2 m/sec is usually sufficient to measure systolic and diastolic velocities, but a higher Venc will be needed if there is a degree of stenosis.

Mitral Regurgitation Mitral regurgitation is probably the most difficult valve lesion to assess by CMR. If the regurgitant orifice is discrete, rounded, and in the severe range, it may be possible to measure the cross-sectional area of the jet by planimetry 510 Cardiovascular Magnetic Resonance

(Fig. 37-7). But the orifice is often extended and narrow, perhaps with more than one defect, and prolapse of a leaflet can cause asymmetric attachment of a fanlike jet to one leaflet or wall of the atrium, which can be very difficult to evaluate visually (Fig. 37-8). The thickness of CMR slices also makes the assessment of valve structure problematic in some cases, and the comments above about cine acquisition in multiple slices are important. Visualization of the dynamic morphology of the valve and its suspensory apparatus is needed with a view to repair.37 To achieve this, a contiguous stack of 5- or 6-mm-thick cine slices is recommended to give comprehensive coverage of all scallops of the mitral valve, starting at the anterolateral commissure and continuing down to the inferior commissure. The stack is oriented parallel to an oblique LVOT long axis plane, orthogonal to the central part of the valve, progressing down from the superior commissure through each scallop of the valve (A1 to P1, A2 to P2, A3 to P3) so that any regurgitant jet, prolapse, or tethering of part of the valve can be visualized and located with a view to possible surgical repair.

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PA

RV

30 Late diastolic forward flow

Pulmonary flow

20 Systolic forward flow

10 0 0 -10

100 200 300 400 500 600 700 800 900 1000 Diastolic Time (msec) reversal

-20 -30

Figure 37-6 Free pulmonary regurgitation, late after repair of tetralogy of Fallot, shown by using steady-state free precession cine imaging (top left) and through-plane velocity mapping (top right and below). There was no evidence of effective valve action. The regurgitant fraction was 38%, which is typical of such cases. Movement of the outflow tract is not a problem after this type of surgery. Notice the early diastolic reversal of flow, when blood pours back to the right ventricle (RV) from the compliant pulmonary artery (PA), and the late diastolic forward flow, when the full ventricle acts as an open conduit for flow returning from the systemic veins, boosted by atrial systole.

Figure 37-7 A central jet of mitral regurgitation depicted passing in-plane (left) and through-plane (right, white arrows) relative to steadystate free precession cine acquisitions. In this case, the through-plane jet (black arrow) shows a bright, coherent core that can be fairly clearly delineated, making planimetry feasible. But this is not always the case; regurgitant jets may be narrow and fanlike or fragmented. Cardiovascular Magnetic Resonance 511

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Figure 37-8 A mitral regurgitant jet attaching asymmetrically to the posterolateral leaflet of the valve from the dominant left ventricle in a patient with a ventricular septal defect and subaortic outlet right ventricle. The patient, who had undergone a Fontan procedure, was fluid loaded at the time, with a pleural effusion, which can be seen as bright layers adjacent to the lung. The mitral regurgitation was less severe after diuresis.

Several approaches to the quantification of MR have been described,4–6 and each can be subject to errors. If the other three heart valves are demonstrably competent, then: Mitral regurgitant volume ¼ the stroke volume difference ðLVSV RVSVÞ Mitral regurgitant fraction ¼ mitral regurgitant volume LVSV But accurate calculation of RV as well as LV stroke volume is based on four sets of relatively time-consuming multislice area measurements. An alternative, more widely applicable approach is to calculate: Mitral regurgitant volume ¼ LV stroke volume LV outflow volume ðmeasured as systolic aortic forward flow; by mapping velocities through a plane transecting the aortic root above the aortic valveÞ A third approach is to measure diastolic mitral inflow and systolic aortic outflow: Mitral regurgitant volume ¼ diastolic mitral inflow systolic aortic outflow But owing to systolic displacement of the mitral valve, a moving slice acquisition should be used for these flow measurements, particularly the diastolic mitral inflow volume,18 and this has yet to be implemented on most commercial systems. Through-plane velocity mapping is also recommended for further visualization (rather than quantification) of the regurgitant jet(s) on the atrial side of the line of coaptation. 512 Cardiovascular Magnetic Resonance

Mild tricuspid regurgitation is common in healthy individuals. More severe tricuspid regurgitation can occur owing to congenital malformation of the valve, notably in Ebstein’s anomaly, or to acquired heart disease, including any cause of dilation of the RV with secondary failure of coaptation. In Ebstein’s anomaly, parts of the tricuspid valve, usually the septal and mural leaflets, are displaced toward the apex with “atrialization” of the basal and inferior parts of the RV. This is commonly associated with severe tricuspid regurgitation and, in some cases, an atrial septal defect or pulmonary valve dysfunction. If the tricuspid regurgitant jet is not visible in a standard four-chamber view, a contiguous stack of transaxial cines can be acquired to visualize all parts of the tricuspid valve and right ventricle and to locate the regurgitant stream, which should then be cross-cut vertically and again horizontally before velocity mapping of the jet. Transaxial cines are also usually more suitable than short axis cines for measuring RV volumes in most cases of Ebstein’s anomaly. If pulmonary resistance is normal or nearly normal, the pressure difference propelling the jet of tricuspid regurgitation is low in comparison with regurgitant lesions of the left heart, and through-plane velocity mapping, encoded through a plane transecting the tricuspid regurgitant jet, is a very useful way of determining the size and shape of the regurgitant orifice (Fig. 37-9). The Venc that is chosen is likely to be in the range of 2 to 3 m/sec, depending on relative RV and atrial pressures. The relationships between orifice area and severity are not simple, but tricuspid regurgitant orifice diameters of 6 mm are associated with severe regurgitation. The driving difference of pressure is clearly relevant, however. In the presence of pulmonary hypertension, the assessment of tricuspid regurgitation becomes more like that of mitral regurgitation.

CMR IN PATIENTS WITH MECHANICAL HEART VALVES The metal rings of prosthetic heart valves generally have only slight ferromagnetic properties, so torque on the ring in the magnetic field is small in comparison with the strength of anchorage necessary for the valve to stay in place. It is therefore regarded as safe to image patients with prosthetic heart valves,38 although variable amounts of image artifact are caused by metallic components, depending upon the nature of the alloy and the imaging sequence used. Lack of signal in blood passing metallic components can exaggerate apparent disturbances of flow, so caution is required in the assessment of possible stenosis, regurgitation, or paravalvar leaks in this situation. A narrow regurgitant “wash jet” is normal through most prosthetic valves in the closed position. Evaluation of prosthetic valve function by catheter and echocardiography is also more difficult than with native valves, so CMR may be called upon. Although velocity mapping is unlikely to be accurate in the immediate vicinity of a valve, volume flow, regurgitant fraction, and ventricular function may be studied as for a native valve as

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Figure 37-9 Severe tricuspid regurgitation. The regurgitant jet or stream may be visible on steady-state free precession cine imaging, but it is much more clearly delineated by phase contrast velocity mapping, encoded through a plane that transects the jet, as indicated by the white lines, with a Venc of 250 cm/sec. The velocity map shows the triangular cross section of the regurgitant jet, 10 to 20 mm wide in this case.

long as appropriate planes of imaging can be chosen outside the magnetic field distortion caused by metal of the ring.

CONCLUSION For investigation of valvular heart disease, the important strengths of CMR lie in visualization and quantification of regurgitation; the measurement of ventricular volumes, function, and mass; and the investigation of associated congenital or acquired pathology. Freedom of access and free orientation of planes, together with ability to measure

poststenotic jet velocities, also give CMR a role in cases such as calcified aortic stenosis, in valvular dysfunction of the right heart including ventriculopulmonary conduits, and in patients with chest deformity or lung abnormality where ultrasonic access may be limited.

ACKNOWLEDGMENTS We wish to thank our colleagues at the CMR Unit of the Royal Brompton Hospital for their help and support and British Heart Foundation and CORDA the Heart Charity for support and grants.

References 1. Bonow RO, Carabello BA, Chatterjee K, et al. ACC/AHA Guidelines for the management of patients with valvular heart disease. Circulation. 2006;114:e84–e231. 2. Schwitter J. Valvular heart disease: assessment of valve morphology and quantification using MR. Herz. 2000;25(4):342–355. 3. Wyttenbach R, Bremerich J, Saeed M, Higgins CB. Integrated MR imaging approach to valvular heart disease. Cardiol Clin. 1998;16 (2):277–294. 4. Kon MW, Myerson SG, Moat NE, Pennell DJ. Quantification of regurgitant fraction in mitral regurgitation by cardiovascular magnetic resonance: comparison of techniques. J Heart Valve Dis. 2004;13(4):600–607. 5. Westenberg JJ, Doornbos J, Versteegh MI, et al. Accurate quantitation of regurgitant volume with MRI in patients selected for mitral valve repair. Eur J Cardiothorac Surg. 2005;27(3):462–466. 6. Gelfand EV, Hughes S, Hauser TH, Yeon SB, Goepfert L, Kissinger KV, et al. Severity of mitral and aortic regurgitation as assessed by cardiovascular magnetic resonance: optimizing correlation with Doppler echocardiography. J Cardiovasc Magn Reson. 2006;8(3):503–507. 7. Kilner P, Gatehouse P, Firmin D. Flow measurement by magnetic resonance: a unique asset worth optimising. J Cardiovasc Magn Reson. 2007;9:723–728.

8. Chernobelsky A, Shubayev O, Comeau CR, Wolff SD. Baseline correction of phase contrast images improves quantification of blood flow in the great vessels. J Cardiovasc Magn Res. 2007;9: 681–685. 9. Kilner PJ, Firmin DN, Rees RSO, Martinez JE, Pennell DJ, Mohiaddin RH, et al. Valve and great vessel stenosis: Assessment with CMR jet velocity mapping. Radiology. 1991;178:229–235. 10. Edelman RR, Wallner B, Singer A, Atkinson DJ, Saini S. Segmented TurboFLASH: method for breath-hold CMR imaging of the liver with flexible contrast. Radiology. 1990;177:515. 11. Friedrich MG, Schulz-Menger J, Poetsch T, Pilz B, Uhlich F, Dietz R. Quantification of valvular aortic stenosis by magnetic resonance imaging. Am Heart J. 2002;144(2):329–334. 12. John AS, Dill T, Brandt RR, Rau M, Ricken W, Bachmann G, et al. Magnetic resonance to assess the aortic valve area in aortic stenosis: how does it compare to current diagnostic standards? J Am Coll Cardiol. 2003;42(3):519–526. 13. Kupfahl C, Honold M, Meinhardt G, Vogelsberg H, Wagner A, Mahrholdt H, et al. Evaluation of aortic stenosis by cardiovascular magnetic resonance imaging: comparison with established routine clinical techniques. Heart. 2004;90(8):893–901.

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14. Debl K, Djavidani B, Seitz J, et al. Planimetry of aortic valve area in aortic stenosis by magnetic resonance imaging. Invest Radiol. 2005; 40(10):631–636. 15. Reant P, Lederlin M, Lafitte S, Serri K, Montaudon M, Corneloup O, et al. Absolute assessment of aortic valve stenosis by planimetry using cardiovascular magnetic resonance imaging: comparison with transoesophageal echocardiography, transthoracic echocardiography, and cardiac catheterisation. Eur J Radiol. 2006;59(2):276–283. 16. Mohiaddin RH, Longmore DB. Functional aspects of cardiovascular nuclear magnetic resonance imaging: techniques and application. Circulation. 1993;88:264–281. 17. Yap SC, van Geuns RJ, Meijboom FJ, Kirschbaum SW, McGhie JS, Simoons ML, et al. A simplified continuity equation approach to the quantification of stenotic bicuspid aortic valves using velocityencoded cardiovascular magnetic resonance. J Cardiovasc Magn Reson. 2007;9:899–906. 18. Kozerke S, Schwitter J, Pedersen EM, Boesiger P. Aortic and mitral regurgitation: quantification using moving slice velocity mapping. J Magn Reson Imaging. 2001;14(2):106–112. 19. Weigang E, Kari FA, Beyersdorf F, Luehr M, Etz CD, Frydrychowicz A, et al. Flow-sensitive four-dimensional magnetic resonance imaging: flow patterns in ascending aortic aneurysms. Eur J Cardiothorac Surg. 2008;34:11–16. 20. Frydrychowicz A, Arnold R, Hirtler D, Schlensak C, Stalder AF, Hennig J, et al. Multidirectional flow analysis by cardiovascular magnetic resonance in aneurysm development following repair of aortic coarctation. J Cardiovasc Magn Reson. 2008;10:30. 21. Mohiaddin RH, Gatehouse PD, Henien M, Firmin DN. Cine magnetic resonance Fourier velocimetry of blood flow through cardiac valves: comparison with Doppler echocardiography. J Magn Reson Imag. 1997;7:657–663. 22. Otto CM, Pearlman AS, Comess KA, et al. Determination of the stenotic aortic valve area in adults using Doppler echocardiography. J Am Coll Cardiol. 1986;7(3):509–517. 23. Otto CM. Valvular aortic stenosis: disease severity and timing of intervention. J Am Coll Cardiol. 2006;47(11):2141–2151. 24. Caruthers SD, Lin SJ, Brown P, Watkins MP, Williams TA, Lehr KA, et al. Practical value of cardiac magnetic resonance imaging for clinical quantification of aortic valve stenosis: comparison with echocardiography. Circulation. 2003;108(18):2236–2243. 25. Kilner PJ, Sievers B, Meyer GP, Ho SY. Double-chambered right ventricle or sub-infundibular stenosis assessed by cardiovascular magnetic resonance. J Cardiovasc Magn Reson. 2002;4(2):373–379.

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26. Djavidani B, Debl K, Lenhart M, et al. Planimetry of mitral valve stenosis by magnetic resonance imaging. J Am Coll Cardiol. 2005;45 (12):2048–2053. 27. Lin SJ, Brown PA, Watkins MP, et al. Quantification of stenotic mitral valve area with magnetic resonance imaging and comparison with Doppler ultrasound. J Am Coll Cardiol. 2004;44(1):133–137. 28. Mohiaddin RH, Amanuma M, Kilner PJ, Pennell DJ, Manzara C, Longmore DB. MR phase-shift velocity mapping of mitral and pulmonary venous flow. J Comput Assist Tomogr. 1991;15(2):237–243. 29. Maceira AM, Prasad SK, Khan M, Pennell DJ. Normalized left ventricular systolic and diastolic function by steady state free precession cardiovascular magnetic resonance. J Cardiovasc Magn Reson. 2006;8(3):417–426. 30. Davlouros PA, Kilner PJ, Hornung TS, Li W, Francis JM, Moon JC, et al. Right ventricular function in adults with repaired tetralogy of Fallot assessed with cardiovascular magnetic resonance imaging: detrimental role of right ventricular outflow aneurysms or akinesia and adverse right-to-left ventricular interaction. J Am Coll Cardiol. 2002;40(11):2044–2052. 31. Grothues F, Moon JC, Bellenger NG, Smith GS, Klein HU, Pennell DJ. Interstudy reproducibility of right ventricular volumes, function, and mass with cardiovascular magnetic resonance. Am Heart J. 2004;147 (2):218–223. 32. Rebergen SA, Chin JGL, Ottenkamp J, Vander wall EE, De Roos A. Pulmonary regurgitation in the late post-operative follow-up of tetralogy of Fallot: volumetric quantification by nuclear magnetic resonance velocity mapping. Circulation. 1993;88:2257–2266. 33. Chaturvedi RR, Redington AN. Pulmonary regurgitation in congenital heart disease. Heart. 2007;93:880–889. 34. Gatzoulis MA, Balaji S, Webber SA, Siu SC, Hokanson JS, Poile C, et al. Risk factors for arrhythmia and sudden cardiac death late after repair of tetralogy of Fallot: a multicentre study. Lancet. 2000;356 (9234):975–981. 35. Shinebourne EA, Babu-Narayan SV, Carvalho JS. Tetralogy of Fallot: from fetus to adult. Heart. 2006;92(9):1353–1359. 36. Kilner PJ, Balossino R, Dubini G, Babu-Narayan SV, Taylor AM, Pennati G, et al. Pulmonary regurgitation: the effects of varying pulmonary artery compliance, and of increased resistance proximal or distal to the compliance. Int J Cardiol. 2009;133:157–166. 37. Chan KM, Wage R, Symmonds K, et al. Towards comprehensive assessment of mitral regurgitation using cardiovascular magnetic resonance. J Cardiovasc Magn Reson. 2008;10:61. 38. Shellock FG. Prosthetic heart valves and annuloplasty rings: assessment of magnetic field interactions, heating, and artifacts at 1.5 Tesla. J Cardiovasc Magn Reson. 2001;3(4):317–324.

Cardiomyopathies Matthias G. Friedrich

Cardiomyopathies (CMP) usually present as chronic, progressive myocardial diseases, often with distinct patterns of morphologic, functional, structural, and electrophysiologic changes. On clinical, morphologic, and histologic grounds, they have been classified into five categories: Dilated (DCM), hypertrophic (HCM), restrictive (RCM), arrhythmogenic right ventricular (ARVC), and unclassified.1 More recently, an updated classification proposed the use of the terms primary (predominantly cardiac) and secondary cardiomyopathies, which would refer to infiltrative systemic diseases such as amyloidosis.2 In numerous forms of CMP, there are diffuse or local tissue characteristics that are related to the underlying pathophysiology. Therefore, visualization of myocardial inflammation, scarring, or infiltration is of special importance. The diagnosis of the CMP often is established by exclusion of other cardiovascular etiologies and an accurate characterization of the phenotype. Therapy is guided by knowledge of the individual stage and hemodynamic relevance of the disease; long-term monitoring is needed in most patients. Thus, imaging techniques are important for both the diagnosis and therapy of the specific CMP. During recent years, cardiovascular magnetic resonance (CMR) has become the gold standard of in vivo imaging of cardiomyopathies. A multisociety consensus deemed CMR a highly appropriate test for the structural and functional assessment of cardiomyopathies.3

CLINICAL QUESTIONS AND THE CONTRIBUTION OF CARDIOVASCULAR MAGNETIC RESONANCE Patients with CMP may present with symptoms related to cardiac dysfunction or arrhythmia or on the occasion of a diagnostic workup. Table 38-1 shows specific information typically requested by referring physicians. In routine clinical practice, transthoracic echocardiography is frequently used to assess left ventricular (LV) parameters. It is widely available, noninvasive, fast, and straightforward to perform in most patients. However, standard M-mode and two-dimensional (2D) echocardiography results suffer from substantial interstudy and interobserver variability,4,5 and the reliability of follow-up measurements is limited.6 Other problems include the poor ultrasound transmission of lungs and bones, angular errors of the

imaging plane, control for plane localization, difficulties in identifying the endocardial border, and parallel shift of the ultrasonic plane out of the targeted center. This may result in reduced accuracy of (especially systolic) intraventricular dimensions and subsequently ejection fraction,7 even with the use of newer techniques such as acoustic quantification,8 automated border detection,9 or threedimensional postprocessing. But probably the most important limitation of echocardiography is the lack of techniques to characterize tissue pathology itself. Since comprehensive information on all relevant parameters can be obtained very efficiently in a single CMR session, many diagnostic centers use CMR as the diagnostic tool of choice. CMR is considered the gold standard not only for noninvasively quantifying biventricular mass, volume, and function, but also for providing in vivo tissue pathology. With substantial prognostic data now available, the detection of structural changes of the myocardium has gained more and more attention, making the accurate definition of the myocardial phenotype in a patient with CMP probably the most important diagnostic contribution of CMR.

CARDIOVASCULAR MAGNETIC RESONANCE APPROACH TO THE PATIENT WITH CARDIOMYOPATHY Morphology and Function CMPs are often characterized by specific alterations of ventricular and myocardial geometry and/or function. CMR allows for a noninvasive assessment of related parameters with the highest accuracy and reproducibility currently achievable.10,11 It is superior to 2D echocardiography in determination of LV mass12 and volumes.13 Its reproducibility is superior to that of standard echo14,15 and thus allows for a reduction of the sample size of clinical trials.13,16 To assess volumes and mass, generally white-blood gradient echo sequences should be applied with 20 to 30 phases per heartbeat. Today, steady state free precession (SSFP) techniques offer the best contrast between blood and endocardium and therefore are preferred. Breath hold techniques with acquisition times of 15 to 20 seconds reduce blurring of the endocardium-blood border, although a small shift of the heart’s position between the breath

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Table 38-1 CMR-Derived Information in Cardiomyopathies LV and/or RV volume and morphology, including regional changes of wall thickness or ventricular shaping Global LV and/or RV function, including stroke volume and ejection fraction Regional LV and/or RV function (arrhythmogenic right ventricular cardiomyopathy/dysplasia) LV and/or RV myocardial mass Intraventricular obstruction (suspected hypertrophic obstructive cardiomyopathy) Global and regional structural changes of the myocardium, including findings suggestive for inflammation, fibrosis, iron content, or other abnormal tissue composition

hold studies may occur. Multislice techniques covering both ventricles within one to three breath holds are available. To cover the whole diastolic phase, techniques have been developed with continuous data acquisition and retrospective gating.17 Under routine clinical circumstances, a biplanar approach (long axis and short axis view) may be sufficient,18–20 but most published data are available for the short axis approach, covering the entire LV from the mitral plane to the apex.21,22 However, it is important to notice that there is a substantial shortening of the ventricular long axis, leading to a smaller number of slices covering the heart in systole than in diastole.23 This may induce errors of up to 20% in quantifying volumes and function. Recently, it has been shown that multiple long axis views not only provide accuracy similar to that of the multiple short axis approach, but also have better reproducibility.24 Accurate definition of the anatomic axis of the heart is crucial, and at least three angulated scouts should be used.25 The slice thickness should be less than or equal to 10 mm; in case of circumscribed or subtle global changes, it should be reduced adequately. To maximize reproducibility of follow-up measurements, trabecular tissue should be included into the cavity,26 although quantitative mass data may be less accurate.

Tissue Characterization Several contrast and noncontrast approaches are available to assess the myocardial tissue and will be discussed in this chapter. Gadolinium-enhanced T1-weighted CMR has been successfully applied to visualize myocardial inflammation and infiltration. Whereas images obtained early (1 to 3 minutes) after gadolinium injection can be used to visualize inflammation,27,28 images acquired later (10 or more minutes) after injection indicate irreversible myocardial injury (late gadolinium enhancement, or LGE).29 T2-weighted imaging using short TI inversion recovery techniques is a robust technique to visualize fluid accumulation such as edema and effusion in inflammatory diseases.28 More recently, the quantification of T1 relaxation times has been proven as an additional option to visualize tissue pathology.30

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Metabolism CMR allows the semiquantitative assessment of several nuclei, and numerous studies have been published on 1H and 31P CMR spectroscopy (CMRS) in cardiomyopathies. Changes of high-energy phosphates as studied by 31P CMRS in cardiomyopathies have been reported for DCM31–33 and HCM.34,35 However, CMRS is still an experimental approach. Whereas 1H CMRS is characterized by a strong signal from water-bound protons and difficulties of spectral interpretation, 31P CMRS is limited by the weakness of the phosphorus signal. Thus, voxels have to be too large to cover circumscribed myocardial regions, and spectra are often altered by blood or adjacent tissue (e.g., skeletal muscle). Furthermore, CMRS requires extensive experience of the investigator, strong physicist support, and sophisticated hardware and software. Therefore, the number of centers with access to this promising tool is currently limited.

DILATED CARDIOMYOPATHY DCM is characterized by progressive dilation of the heart with loss of contractile function. Its etiology is unclear in about half of cases,36 but the typical pattern of DCM may be the end stage of a disease process, initiated by myocardial inflammation, toxic agents such as alcohol and drugs, or a genetic disorder.37 The histologic hallmark of DCM is a progressive interstitial fibrosis with a numeric decrease of contractile myocytes.

Function and Morphology For phenotyping patients with DCM, CMR can be used not only to quantify global LV function, but also to analyze wall thickening,38 visualize impaired fiber shortening,39 and calculate global40 and regional41 end-systolic wall stress. Right ventricular (RV) morphology and function also are accurately assessed by CMR42 and can be referred to published normal values.43,43a A key finding in DCM is LGE of the LV. Patients with the clinical diagnosis of DCM from heart failure clinics (normal coronary angiography with an electrocardiogram (ECG) showing no infarction) show three patterns of LGE: (1) no LGE (59%); (2) midwall enhancement of the circumferential fibers particularly in the septum (28%); and (3) typical LGE of myocardial infarction (13%).44 These findings suggest that clinical diagnosis of DCM may sometimes be wrong when the coronaries are used as the key diagnostic criterion. The reason for this is that some 25% of infarctions do not present to hospital, and in some, there is eventual lysis of the occlusive thrombus, leaving a healed but minor coronary plaque that is not deemed significant with lumenography. In some patients, dual pathology may exist. Further studies have confirmed the excellent discrimination of DCM from myocardial infarction as a cause of heart failure.45 The patients with midwall fibrosis have been shown to have significantly more future cardiac events in prospective follow-up including arrhythmias and hospitalizations.46

Metabolic Cardiovascular Magnetic Resonance As CMRS studies have shown, high-energy phosphate metabolism is altered in DCM.54 Moreover, a low ratio of phosphocreatine (PCr) to adenosine triphosphate (ATP) as assessed by CMRS was shown to be of prognostic value

in DCM.33 A study with a similar technique related these changes to a reduction of creatine kinase activity.55 A more recent approach has proven the utility of 1H CMRS for detecting myocardial steatosis as an early marker in patients with glucose tolerance abnormalities.56 This allows for assessing the role of lipotoxicity in several cardiomyopathies. New techniques as well as simplification of existing protocols will strengthen the clinical role of metabolic CMR.

MYOCARDITIS It is known that myocarditis can be the underlying cause of DCM,57 but recently, myocarditis has attracted increasing attention as a frequent cause of less severe disease58 associated with symptoms such as dyspnea, fatigue, chest discomfort, and palpitations. Although the disease is benign in most cases, it may have a fulminant course, and patients may present with features of acute myocardial infarction. It is, however, important to emphasize that myocardial inflammation is not confined to viral myocarditis but also occurs in other myocardial diseases such as HCM and DCM.59,60 Often, atypical symptoms and findings in younger patients raise the concern about myocarditis. CMR provides very useful information on function and on reversible and irreversible injury. Of note, tissue characterization adds significant information beyond LV function, which often is found to be normal in myocarditis. Figure 38-1 shows the CMR findings in a patient with myocarditis.

Figure 38-1 CMR in a patient with acute myocarditis. Top left, Systolic steady-state free precession frame in a mid short axis view with mild hypokinesis of the anteroseptal segment. Bottom left, Postcontrast inversion recovery late gadolinium enhancement gradient echo image with several areas of high signal intensity, likely reflecting irreversible injury (arrows). Note that in contrast to myocardial infarction, the subendocardial layer is spared. Middle, Non-breathhold T1-weighted fast spin echo images before (upper panel) and after (lower panel) contrast. In the global quantitative analysis, the relative signal increase normalized to skeletal muscle was increased. Top right, T2-weighted triple inversion recovery fast spin echo image in a mid short axis, with increased signal intensity in the anterior wall. A global quantitative analysis normalized to skeletal muscle revealed an increased ratio. Bottom right, Postcontrast inversion recovery gradient echo image (late gadolinium enhancement) in a mid short axis view with midwall high signal intensity (arrow).

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The presence and extent of LGE were more predictive of events than was the ejection fraction, and this suggests a possible role for LGE in DCM for the improved choice of patients requiring implantable cardioverter-defibrillators. Later studies have confirmed the association of anatomic variables,46a as well as LGE with events.47 By using CMR, atrial volumes and function have been assessed in sinus rhythm48,49 and atrial fibrillation.50 CMR may be the method of choice for the longitudinal follow-up of patients with DCM after pharmacologic51,52 or surgical53 intervention, allowing for a reduction of the sample size of clinical trials.13,16 Costs could be reduced markedly, and time could be saved in clinical research. Its high sensitivity to detect subtle changes also allows the physician to fine-tune therapy, thus improving the patient’s quality of life and prognosis. Moreover, it may be possible to avoid repeated studies using less precise methods, which are inconsistent. More precise adjustment of therapy and reduction of admissions for repeat studies are likely to overcome the additional costs of a CMR study. However, an analysis of cost effectiveness of CMR when replacing other modalities is still awaited.

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Function and Morphology As in other myocardial diseases, CMR allows for the most accurate quantification of LV size, mass, and function. This includes quantitative data on stroke volume and cardiac output.

Tissue Characterization: Early Enhancement The inflammatory activity of acute myocarditis can be detected by using contrast-enhanced T1-weighted fast spin echo techniques with image acquisition during the first minutes after gadolinium injection and signal normalization to skeletal muscle.27,28,61,62 Increased uptake is most likely caused by a combination of increased inflow (inflammatory hyperemia), slow interstitial wash-in/washout-kinetics (capillary leakage and edema), diffusion into cells through leaky membranes (necrosis), and an increased volume of distribution (scar tissue). Long-term follow-up revealed persistence of these changes in patients with clinical and functional evidence for ongoing inflammation.63 An increased gadolinium accumulation has also been described in patients with myocardial inflammation due to Chagas myocarditis,64 sarcoidosis,65–67 eosinophilic myocarditis,68 and polymyositis-related myocardial inflammation.69 In suspected chronic myocarditis, early enhancement and edema ratio were found to be the most sensitive CMR markers for immunohistologically defined inflammatory activity.70

Edema Increased signal intensity in T2-weighted imaging reflects inflammation-related myocardial edema.28,62,71,72 Patients with myocarditis typically show regional and/or global high signal intensity, which usually equals or exceeds that of skeletal muscle as an internal reference by a factor of 2. It is important to note that the low signal-to-noise ratio of T2-weighted sequences has to be compensated for by using thicker slices than usual (15 mm or even 20 mm). Since edema often is not a focal phenomenon but a more diffuse phenomenon, the sensitivity to detect very small lesions is not an issue. Furthermore, the use of surface coils should be avoided for T2weighted imaging, because the low signal in the regions more distant from the coil reduces the sensitivity and makes a quantitative comparison to the skeletal muscle impossible. Triple inversion recovery spin echo techniques should be used to suppress slow-flow artifacts and fat.28 In a recent biopsycontrolled study, the edema ratio showed a sensitivity of 67% to detect chronic myocarditis.70 Novel T2-weighted sequences have further improved the robustness of CMR of myocardial edema.73

Late Gadolinium Enhancement Following the success in visualizing myocardial scarring in coronary artery disease, LGE has been used to assess myocarditis. Presumably visualizing irreversible myocardial microinjury,29 the images show regions of increased 518 Cardiovascular Magnetic Resonance

signal,28,62,74,75 with a very good agreement between CMR findings and histopathology.75,76 Of importance, the regional distribution of myocarditis lesions differs from that of ischemic scars, since—unlike myocardial infarcts—they generally do not involve the subendocardial layers,28,75,77 findings that allow for differentiating acute myocarditis from myocardial infarction.78 It is important, however, to note that the sensitivity of LGE to detect myocarditis may be lower than that of T1/early enhancement or T2weighted CMR.28,70,79,80

Combined Protocols To overcome the limitations of single sequences, a combination of T2-weighted, T1-weighted early enhancement, and T1-weighted LGE techniques has been successfully used to rule out or establish the diagnosis of myocarditis.28,62 Such a combined approach provides a higher diagnostic accuracy than each of the individual sequences.28,70

Follow-up In uncomplicated myocarditis, the signal abnormalities in T1/early enhancement as well as functional abnormalities (if present) typically disappear within the first weeks,27 whereas their persistence may indicate a less favorable prognosis.63 Thus, a follow-up, including assessment of LV volumes and function, is advisable. CMR is regarded as the most useful noninvasive tool to assess myocarditis,81,82 and its diagnostic accuracy may exceed that of other available invasive and noninvasive approaches.83 Therefore, using proven protocols, it should be a first-line diagnostic tool in centers with access to CMR. Protocol standardization and the proof of its applicability in multicenter trials are being undertaken.

HYPERTROPHIC CARDIOMYOPATHY HCM is characterized by disarray of cardiomyocytes and most often by regional or global hypertrophy with impaired diastolic function and, in its obstructive form, systolic narrowing of the left ventricular outflow tract (LVOT). It is also a common cause of sudden death in the young.84 Histologically, areas of hypertrophy show a pattern of myofibrillar disarray and patchy areas of necrotic tissue. Since the diagnostic value of endocardial biopsy is limited,85–87 the diagnosis is generally made by noninvasive imaging. Echocardiography is used to visualize wall thickening and, in obstructive HCM, accelerated flow in the LVOT.88 However, flow quantification is hampered by a high interstudy variability89 and limited accuracy of pressure gradient measurements.90 Thus, although frequently used, echocardiography may provide a limited basis for the characterization of morphology and hemodynamic relevance of HCM in the individual. In many centers, CMR is now considered the primary diagnostic tool to assess various aspects of the disease. Figure 38-2 shows the CMR findings in a patient with HCM.

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Figure 38-2 Cardiovascular magnetic resonance in a patient with hypertrophic cardiomyopathy. Left, Steady-state free precession images in a two-chamber view (upper panel) and a mid short axis view (lower panel). There is a regional hypertrophy of the basal and midanteroseptal segments (asterisks). Right, Postcontrast inversion recovery gradient echo image (late gadolinium enhancement) in the same orientations, showing regional signal increase in the affected segments as well as in the inferoseptal region (arrows).

Function and Morphology

Tissue Characterization

Today, CMR is widely accepted as delivering accurate data on morphology, mass, function, tissue characterization, and hemodynamic relevance of obstruction. Owing to its high sensitivity for detecting regional morphologic changes and its noninvasiveness, CMR is especially useful for screening families of index patients. SSFP gradient echo sequences are most useful for visualizing morphology and quantifying regional and/or global left ventricular mass. A stack of short axis views should be used, since long axis views may overestimate mass, stroke volume, and ejection fraction, even if they deviate only slightly from the true anatomic long axis. Individual angulations of imaging planes may be needed to assess regional wall morphology. Besides its accuracy to quantify LV mass,12,91 CMR is more accurate than echocardiography for investigating regional hypertrophy92,93,93a,93b and is suitable to determine different phenotypes of this disease such as apical forms,94,95 and coexistent apical aneurysms,95a which may be overlooked by echocardiography.96,97 It is important to emphasize that total LV mass may be normal, even in cases with marked regional hypertrophy.97a CMR screening is also cost-effective.97b Analysis of the early untwisting motion of the myocardium using myocardial tagging may be helpful for assessing early diastolic function in hypertrophic heart diseases.98 Other functional changes that can be detected by the use of myocardial tagging are a reduction of posterior rotation, reduced radial displacement of the inferior septal myocardium,99 heterogeneity of regional function, 100 and reduced threedimensional myocardial shortening.101 The use of geometric indices allows for differentiating pathologic hypertrophy from physiologic forms, such as athlete’s heart.102 CMRderived data on atrial volume have been shown to correlate well with other markers of cardiac performance in patients with HCM and indicate that there may be a prognostic significance to atrial size.103 Recently, Germans and colleagues showed that CMR detects crypts in the basal and midinferoseptal myocardium in HCM gene carriers without relevant LV hypertrophy.104

CMR visualizes focal areas of LGE,105–108 which are caused by regional fibrosis.109,110 Typical locations are the insertion areas of the right ventricle and regions with marked hypertrophy. Its extent correlates with Q waves in the ECG.111 The identification of these lesions has come into focus, since mere measurement of septal wall thickness is obviously insufficient to serve as the major risk stratification tool.112 Retrospective data indicate that the extent of LGE correlates with known markers for an increased risk of malignant arrhythmias,113,113a,113b even in patients with HCM phenotypes that are not characterized by LV hypertrophy.114 Prospective studies on the value of these findings are underway, but it is easy to speculate that the visualization of intramyocardial fibrosis will become a cornerstone for phenotyping and risk-stratifying patients with HCM.115 For visualizing myocardial fibrosis, standard inversion recovery gradient echo techniques are used. Because regions may be small, a contiguous stack with slices of 8 mm or less can improve the sensitivity of the test. CMRS has also been established to investigate phosphate metabolism in HCM. Myocardial PCr/ATP ratio and the signal of phosphomonoesters were found to be changed in patients with HCM.34,116 Spindler and colleagues were able to correlate diastolic dysfunction to a decrease of myocardial energy reserve related to high-energy phosphate metabolism.117 Interestingly, metabolic alterations as assessed by CMR spectroscopy are not related to the severity of LV hypertrophy.35 Using T2-weighted sequences, focal edema can be visualized by CMR118 and appears to be colocalized with LV hypertrophy and irreversible injury.119

Left Ventricular Outflow Tract Obstruction LVOT obstruction is an important determinant of the clinical features and outcome of HCM.120 Thus, its quantitative assessment requires special attention. The turbulent Cardiovascular Magnetic Resonance 519

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systolic jet of LVOT obstruction is easily detected in gradient echo sequences, including SSFP, and can be used to quantify flow velocity and to calculate the pressure gradient according to standard approaches used for valvular stenosis.121 However, since pressure gradients vary greatly in HCM, direct planimetry of the LVOT has been proposed and successfully applied in HCM.122 A cross-sectional systolic image is used to measure the minimal systolic area between the anterior mitral valve leaflet and the interventicular septum. Besides a high temporal resolution, careful selection of the cross-sectional imaging plane in a systolic long axis frame is necessary, as are thin slices, to achieve accurate and reproducible results. Data show LVOT area planimetry may be much less prone to preload variations than is pressure gradient quantification.123 Systolic anterior motion of the anterior mitral valve leaflet, which is a typical feature of obstructive HCM, is detectable by CMR124 and is best visualized in the fourchamber view, an LVOT view, or a short axis view through the valvular plane perpendicular to the outflow tract. CMR can also be used to assess related abnormalities such as alterations of coronary blood flow125 or perfusion deficits.126 Microvascular dysfunction as a known feature of HCM can be assessed by using multiparametric CMR protocols.127 The clinical role of these approaches, however, remains to be evaluated.

Follow-up Serial CMR studies may reveal morphologic changes during the natural course of the disease,95 but the impact of an interventional therapy in HCM also can be reliably monitored and quantified. This has been shown for myectomy124,128 and for interventional septal artery ablation.122,129,130 Besides LV mass, the degree of LV outflow tract obstruction is a very important diagnostic target in this setting. Planimetry of the LV outflow tract has been introduced as a useful tool in the follow-up of HCM patients after interventional therapy.122 Furthermore, the

regional septal injury can be characterized by LGE.131 Mitral valve regurgitation is frequent in HCM and is probably due to a pathologic change of leaflet geometry; therefore, assessment of this should be included in a CMR workup.

ARRHYTHMOGENIC RIGHT VENTRICULAR CARDIOMYOPATHY Figure 38-3 shows CMR findings in a patient with ARVC. CMR is helpful in the exclusion or verification of hypertrophy due to other causes, such as amyloidosis.30,132 ARVC is an inherited myocardial disease with progressive degeneration of the right and, not infrequently, LV myocardium, most often with localized dysfunction. Its pathogenesis is still not fully understood, but there are clear genetic links to proteins contained in the desmosome, which provides structural links between myocytes, and protein dysfunction presumably results in chronic myocyte injury and fibrofatty repair.133,134 Morphologic features include fibrous and/or fatty replacement of myocardial tissue, extensive wall thinning, and atypical arrangement of trabecular tissue, with an associated morphologic spectrum ranging from subtle changes to extensive fibrofatty dysplasia of the right ventricle,135,136 leading to RV congestive heart failure only in very rare cases. The clinical course is generally determined by the occurrence of severe ventricular arrhythmias with a substantial risk of sudden cardiac death.137 The morphologic substrate to ventricular arrhythmias probably are fibromuscular bundles isolated from each other by fatty tissue, leading to reentry phenomena.138 Of importance, pathology studies have revealed LV involvement in up to 76% of cases.135 In 1994, a set of criteria was defined to establish the diagnosis,139 and currently, the most important contribution of CMR to the diagnostic workup of ARVC is detailed visualization of morphologic and functional abnormalities according to these criteria (Table 38-2). Figure 38-3 CMR in a patient with arrhythmogenic right ventricular cardiomyopathy. Left, Diastolic steady-state free precession frame in a fourchamber view (upper panel) and a basal short axis view (lower panel), showing an irregular shape of the right ventricle. There is a regional dilation of the basal right ventricle with marked trabeculation (asterisks). Top right, Postcontrast inversion recovery gradient echo image (late gadolinium enhancement) in a four-chamber view, showing focal late enhancement, consistent with left ventricular involvement. Bottom right, Early diastolic axial steadystate free precession frame showing a focal aneurysm with loss of trabecular structure, wall thinning and diastolic bulging (arrow).

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Major Severe RV dilation with RV dysfunction with no or only mild LV involvement Localized aneurysms as defined by akinetic or dyskinetic areas with diastolic bulging Severe segmental RV dilatation

Minor Mild global RV dilatation and/or dysfunction with normal left ventricular function Regional RV hypokinesia Mild segmental RV dilatation

Adapted from McKenna WJ, Thiene G, Nava A, et al. Diagnosis of arrhythmogenic right ventricular dysplasia/cardiomyopathy. Task Force of the Working Group Myocardial and Pericardial Disease of the European Society of Cardiology and of the Scientific Council on Cardiomyopathies of the International Society and Federation of Cardiology. Br Heart J. 1994;71:215–218.

Since a total of at least two major criteria, one major plus two minor criteria, or four minor criteria have to be present to diagnose ARVC, most often additional criteria (by family history, ECG, and sometimes biopsy) are required to make the diagnosis. In combination with these other criteria, CMR generally allows for a noninvasive, safe diagnosis to decide whether or not to implant a defibrillator device. The ESC Task Force criteria applicable to imaging techniques do not include tissue pathology. Thus, although CMR can provide additional information on trabecular abnormalities, apparent fatty infiltration, and wall thinning, only regional or global dilation or dysfunction can be considered diagnostic criteria. The diagnostic potential of CMR has not been fully exploited yet, owing to the lack of standardized protocols and a subsequently higher interobserver variability.140,141 Protocol standardization, the improving ability of CMR to detect small changes, and validation studies will allow us to incorporate tissue parameters as an integral part of the diagnostic criteria. This will also overcome the current limitation of the criteria to request a normal left ventricle. A review of ARVC and recommended CMR techniques to assess ARVC142 and reference values for the RV43 have been published.

Function and Morphology CMR visualizes ventricular cavities and walls with an excellent depiction of myocardial anatomy. To obtain the high spatial resolution needed to evaluate this entity, a prone position often yields better results than a supine position does. A phased array coil is mandatory. SSFP gradient echo sequences detect the regional wall motion changes, such as global or local hypokinesis localized early diastolic bulging or circumscribed saccular outpouchings, that are described in ARVC.143–148 A slice thickness of 6 mm or less is suggested. Orthogonal image planes (axial and sagittal) and additional short axis views reveal the best results, although a second plane may not be needed in a perfectly normal single-plane study. Although data on the diagnostic accuracy are still sparse, focal abnormalities are likely to provide the highest specificity to differentiate ARVC from other diseases affecting the RV. The RV free wall “accordion sign” has been described in ARVC.149 In patients with known ARVC as defined by the ESC Task Force criteria, RV dilation and RV dysfunction,150

trabecular abnormalities, RV free wall hypertrophy, and wall thinning are frequent findings.151 It is important to emphasize that minor regional variations of RV wall motion can be seen in healthy subjects, limiting the specificity of these findings. Furthermore, former CMR studies have shown morphologic and functional abnormalities in up to 76% of patients with idiopathic RVOT tachycardia.152–155 Although a recent study found no abnormalities in these patients,156 this underscores the importance of using a set of criteria instead of single observations to make a diagnosis. In a study in 100 ARVC patients, Dalal and colleagues described increased RV volumes and a lower RV ejection fraction in patients with inducible ventricular tachycardia, compared to those without ventricular tachycardia. Additional findings included wall thinning and regional functional abnormalities.157

Tissue Characterization Because the sensitivity of endomyocardial biopsy may be limited by the regional nature of the changes,139 its invasiveness, and the low incidence of septal involvement,158 CMR very often has an important role in the workup of a patient with suspected ARVC. T1-weighted spin echo studies should be used to visualize fatty infiltration and wall thinning as well as dysplastic trabecular structures.142 The addition of a saturation band over the atria in T1-weighted images may reduce slow flow artifacts of inflowing blood and thereby enhance the image quality at the endocardial border of the RV apex.144 Additional studies with fat saturation allow for differentiating fat from fibrous tissue. Fatand water-selective gradient echo sequences may help to differentiate the different tissue components.159 Inversion recovery prepared sequences can be used to visualize fibrous tissue. The most frequently observed tissue abnormality in patients with ARVC is a high intramyocardial T1 signal consistent with fat.151 Although it is tempting to use the visualization of fat in the RV wall as a diagnostic parameter, it is important to notice that fatty replacement can also be found in other cardiac diseases such as dilated cardiomyopathy, myocarditis, or alcoholic myocardial injury160,161 and its specificity has been questioned.134 LGE was observed in more than half of patients with ARVC in the RV162,163 and LV,164 possibly reflecting fibrous degeneration, one of the main pathologic features of this disease. These results, however, remain to be reproduced. A comprehensive CMR protocol has been shown to be more sensitive than published Task Force criteria for detecting abnormalities in patients with an ARVC genotype.165

RESTRICTIVE CARDIOMYOPATHY RCM, most often caused by infiltration of the myocardium, is characterized by a severe disturbance of relaxation and compliance, biatrial dilation, normal LV size, and normal systolic function. Atrial thrombi also occur. The main differential diagnostic consideration is constrictive pericarditis, which needs to be excluded in patients with suspected RCM. Cardiovascular Magnetic Resonance 521

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Table 38-2 CMR-Derived Task Force Criteria for Arrhythmogenic Right Ventricular Cardiomyopathy

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CMR studies in RCM focus on myocardial morphology and function as well as on the exclusion of constrictive pericardial disease.

Morphology and Function LV size and wall thickness are quantified in short axis SSFP image sets, but long axis views are also needed. Biatrial dilation is easily visualized with the four-chamber view. CMR volumetry of the enlarged atria may be useful.166,167 Concomitant mitral regurgitation should be visualized. The exclusion of relevant pericardial thickening makes constrictive pericarditis less likely.168 This is possible by T1-weighted spin echo techniques.169,170 It should be noted, however, that constrictive pericarditis may occur without detectable pericardial thickening.171 Recently, real-time gradient echo CMR was shown to visualize a leftward inversion during the onset of inspiration and flattening of the septum during early ventricular filling in constrictive pericarditis but not in RCM.172

Tissue Characterization Aside from cardiac amyloidosis, CMR has not been extensively used to characterize tissue in RCM, mainly because of the rarity of its occurrence. Spin echo sequences with T1-weighted and triple-inversion recovery T2-weighted sequences can be used to detect atrial thrombi,173 although slow-flowing blood in the atria may lead to false positive results.174

VENTRICULAR NONCOMPACTION Ventricular noncompaction (VNC) or noncompaction cardiomyopathy is a recently described inherited myocardial disease, and our knowledge is still limited to small

series.175,176 It is characterized by the combination of left ventricular dilation, regional or global hypokinesis, and an abnormal relationship between a thin compact wall and a thick layer of noncompacted, sometimes spongyappearing myocardium.177 Clinical events include ischemia, arrhythmia, and sudden cardiac death, even in early childhood.178,179 The echocardiographic phenotype has been described by Jenni and colleagues.177 Figure 38-4 shows the CMR of a patient with VNC.

Function and Morphology With SSFP cine imaging, the abnormalities are easily detectable. To visualize the trabecular meshwork without having partial volume effects induced by intertrabecular blood, a high temporal (>25 phases/sec) and spatial (matrix > 256 256, field of view < 350 mm, slice thickness of 8 mm or less) resolution should be used. Since some trabeculation of the left ventricle is to be considered normal, it is important to assess systolic wall motion and to quantify the diameter ratio of noncompacted to compacted myocardium. A perpendicular orientation of the imaging plane to the wall should be carefully ensured, since tangential orientation may lead to overestimation of the relative extent of the noncompacted wall. In a recent study, Petersen and colleagues found that a cutoff value of 2.3 for the ratio of noncompacted to compacted wall accurately detected VNC.180

Tissue Characterization There are only few data on other findings in VNC. These include LV thrombi181 and, in a single case with pathology, myocardial fibrosis.182 We have observed LGE consistent with known ischemic necrosis described in this entity.177 The consistent image quality of SSFP sequences and the use of sequences visualizing tissue pathology will result in the detection of more cases and a better understanding of this disease. Figure 38-4 Cardiovascular magnetic resonance in a patient with ventricular noncompaction. Steady-state free precession images in a two-chamber view (top left panel), three-chamber view (bottom left panel), four-chamber view (top right panel), and mid short axis (bottom right panel). All mid and apical segments except the septum reveal a significant amount of noncompacted wall with a thin compact wall. The left ventricular volume is increased, and its systolic function is substantially impaired. Right, Postcontrast inversion recovery gradient echo image (late gadolinium enhancement) in the same orientations, showing regional signal increase in the affected segments as well as in the inferoseptal region (arrows).

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INFILTRATIVE SECONDARY CARDIOMYOPATHIES AND ENDMYOCARDIAL DISEASES The infiltrative cardiomyopathies include sarcoidosis, amyloidosis, and hemochromatosis. In these systemic diseases, the infiltration leads to impairment of function and/ or conduction abnormalities. Since infiltration of the tissue is accompanied by changes of myocardial signal properties, CMR may become a powerful diagnostic tool, although the specificity may be limited.

Sarcoidosis The incidence of myocardial involvement in systemic sarcoidosis may exceed 50%,183 and up to 50% of deaths in sarcoidosis may be related to cardiac involvement,184 primarily due to sudden death and congestive heart failure. Histologic findings are granulomas, which often are transmural. Probably because of the patchy distribution, however, the sensitivity of endomyocardial biopsy is low.185

Morphology and Function Sarcoid heart disease may be accompanied by LV dilation as well as regional and global wall motion abnormalities, which can be more sensitively detected by CMR than with other modalities186 Regional granulomatous inflammation may lead to regional wall thinning and akinesis. Therefore, a careful functional and morphologic assessment by SSFP sequences, using a stack of contiguous slices, is recommended.

Tissue Characterization Sarcoid lesions may lead to a range of signal intensities, possibly because of different stages of disease activity. Muscular sarcoidosis causes high signal intensity areas on T2-weighted CMR.187 In another study of skeletal muscle sarcoidosis, the granulomatous nodules exhibited a central region with low signal intensity in T1- and T2weighted imaging but were surrounded by a high signal ring.188 Gadolinium accumulates in sarcoid lesions of the brain.189 This behavior may be compatible with fibrotic, not active granulomatous nodules with inflammatory response of the surrounding tissue. Similar findings were reported in several cases of sarcoid infiltration of the heart.190–194 T1-weighted imaging with assessment of early enhancement has been reported,65 and recently, LGE was reported in sarcoid heart disease.195 In one study, abnormal early enhancement was present in more than 80% of patients with cardiac sarcoidosis, whereas LGE could be detected in only 25%.196 A combined protocol including a T2-weighted tripleinversion recovery spin echo sequence, a T1-weighted spin echo technique in short and long axis before and after a standard dose of gadolinium (0.1 mmol/kg), and an

Figure 38-5 Cardiovascular magnetic resonance of thickened walls in a patient with cardiac amyloidosis. Diastolic steady-state free precession frame in a four-chamber view showing diffuse myocardial thickening of the ventricular walls and the interatrial septum.

inversion recovery gradient echo sequence is recommended. In addition, the CMR findings could be used to guide the location for endomyocardial biopsy.191 Since the follow-up is very important to guide therapy, serial CMR studies may be very helpful in these patients.

CARDIAC AMYLOIDOSIS Myocardial amyloid protein accumulation can occur in several forms of systemic amyloidosis and may lead to diastolic dysfunction and restriction, accounting for up to 78% of deaths in patients with cardiac amyloidosis and evidence of diastolic dysfunction.197 Cardiac amyloidosis typically causes wall thickening of the ventricular myocardium but also of the interatrial septum. Figures 38-5 and 38-6 show CMR images of a patient with cardiac amyloidosis.

Morphology and Function Cardiac amyloidosis can be visualized by standard (SSFP) sequences, accurate visualization of the atrial septum being of special importance.

Tissue Characterization Amyloid infiltration of the heart is found in almost all cases of primary amyloidosis and in about one fourth of cases of familial amyloidosis. It is associated with a loss of atrial and/or left ventricular function and congestive left ventricular failure. Since the infiltration often is diffuse, endomyocardial biopsy is very sensitive in the detection of amyloid infiltrates.198 CMR provides important additional information in suspected cardiac amyloidosis.30,132,199,199a The CMR approach is mainly directed toward the detection of signal changes after gadolinium administration. Cardiac amyloidosis can be reliably assessed by T1 quantification after gadolinium.30 This study also clearly demonstrated the predominantly subendocardial distribution of the infiltration. The alteration of myocardial T1 may be substantial, typically leading to difficulties in finding a suitable inversion Cardiovascular Magnetic Resonance 523

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Figure 38-6 Cardiovascular magnetic resonance tissue characterization in a patient with cardiac amyloidosis. Left, Non–breath hold T1-weighted fast spin echo images before (upper panel) and after (lower panel) contrast. In the global quantitative analysis, the relative signal increase normalized to skeletal muscle was increased. Top right, T2-weighted triple inversion recovery fast spin echo image in a mid short axis, with increased signal intensity. A global quantitative analysis normalized to skeletal muscle revealed an increased ratio. Bottom right, Postcontrast inversion recovery gradient echo image (late gadolinium enhancement) in a mid short axis view with blood appearing with a low signal intensity and inhomogeneous signal of the myocardium.

time for inversion recovery gradient echo late enhancement images. Recently, gadolinium CMR has been shown to be a powerful predictor of prognosis in amyloidosis, with a reduced transmural T1 difference predicting death.200 These patients have lost the normal endocardial preponderance of amyloid deposition, which indicates later stage disease and a greater cardiac amyloid burden.

MYOCARDIAL SIDEROSIS Myocardium siderosis is characterized by intracellular iron deposits leading to ventricular dilation with progressive loss of function, leading to congestive heart failure and death. It is important to note that the prognosis of many iron-loading conditions depends mainly on cardiac involvement. In thalassemia major, a cardiac death occurs in more than 70% of patients.

Morphology and Function Impairment of LV function is a feature of advanced disease and associated with a poor outcome. The onset of decompensated heart failure may be rapidly fatal. Standard SSFP techniques with complete coverage of the ventricles should be applied.

Tissue Characterization The CMR approach is directed toward the detection of iron deposits as the specific marker for the disease. Iron has strong paramagnetic properties and thus may act as a contrast-generating agent itself, when suitable sequences are used. Myocardial deposits result in signal loss in T1and T2-weighted images,201–206 in the heart.207 A more 524 Cardiovascular Magnetic Resonance

reproducible way to assess myocardial iron deposition uses T2* quantification.208 Myocardial T2* correlates well with tissue iron content and also relates to ventricular function,208 being more sensitive than diastolic function parameters.209 Interestingly, iron deposition in the liver, compared with that in the heart, may deviate substantially.208 T2* imaging of the heart serves as a better guiding parameter for chelation therapy of the heart than does the use of liver iron; this has led to a much better understanding of the treatment of transfusion-induced myocardial siderosis in conditions such as beta-thalassemia major.210 The role of T2* imaging as the tool of choice to assess myocardial iron deposition in a clinical setting, but also in drug trials, has been impressively demonstrated.211,212 Cardiac T2* has been shown to predict the future cardiac events of heart failure and arrhythmia,213 and the introduction of cardiac T2* CMR in the United Kingdom has been associated with a substantial reduction of 71% in cardiac mortality.214

ENDOMYOCARDIAL DISEASES Both forms of endomyocardial fibrosis, the tropical and Loeffler’s endocarditis, lead to mid and apical concentric wall thickening, followed by extensive subendocardial fibrosis and frequent apical thrombus formation. Both ventricles may be affected, often with significant reduction of chamber volumes. The course is determined by progressive diastolic dysfunction and a decrease in stroke volume. Histologic features as detected by endomyocardial biopsy may be myocyte hypertrophy and/or fibrosis, although the specificity is limited by similarity of findings in HCM215 and sometimes in DCM. Figure 38-7 gives an example of a patient with Loeffler’s eosinophilic endocarditis.

38 CARDIOMYOPATHIES

Figure 38-7 Cardiovascular magnetic resonance of a patient with Loeffler’s eosinophilic endocarditis. Left, Diastolic steady-state free precession frame in a threechamber view (upper panel) and an apical short axis view (lower panel), with apparently regionally increased wall thickness (arrows). Right, Postcontrast inversion recovery gradient echo image (late gadolinium enhancement) in similar orientations, showing a thick layer with increased signal and a large intracavitary mass without any contrast accumulation (asterisk), indicating a large intraventricular thrombus (asterisk).

Figure 38-8 Cardiovascular magnetic resonance in a patient with reversible stress-induced cardiomyopathy (Tako-Tsubo cardiomyopathy). Left, Acute stage of the disease (day 2). Top left, Systolic steady-state free precession frame in a two-chamber view with extensive dyskinesis of the mid and apical segments. Middle left, T2-weighted triple inversion recovery fast spin echo image in a two-chamber view. Visible high signal in the mid and apical segments, matching the dysfunctional area. Bottom left, Postcontrast inversion recovery gradient echo image (late gadolinium enhancement) in a two-chamber view with focal areas of high signal intensity (arrows). Note that in contrast to ischemic injury, the subendocardial layer is spared. Right, Follow-up after 8 weeks using the same imaging planes. Top right, Systolic steady-state free precession frame in a two-chamber view showing complete recovery of the mid and apical segments. Middle right, T2-weighted triple inversion recovery fast spin echo image in a two-chamber view. The signal is only very mildly increased in the mid and apical segments. Bottom right, Postcontrast inversion-recovery gradient echo image (late gadolinium enhancement) showing resolution of focal enhancement in the subepicardial myocardium. Cardiovascular Magnetic Resonance 525

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Morphology and Function The morphologic and functional features of endomyocardial fibrosis usually are characterized by severe systolic dysfunction without dilation or even with obliteration. The changes can be reliably visualized and quantified by CMR.216,217

Tissue Characterization Fibrosis or calcification may be visible as a dark rim in brightblood prepared gradient echo sequences but may also reveal intermediate signal intensity.218 Therefore, the use of LGE may substantially improve the diagnostic accuracy of CMR, although our knowledge is so far based on case reports.219–221

STRESS-INDUCED CARDIOMYOPATHY Stress-induced cardiomyopathy, also known as Tako-tsubo cardiomyopathy, is a recently described myocardial disease,

characterized by a stress-induced, reversible extensive wall motion abnormality of the mid and apical LV, mainly occurring in elderly women.222,223 The underlying mechanism is not yet fully understood.

Function and Morphology The typical feature of stress-induced cardiomyopathy is systolic “ballooning” of the LV (Fig. 38-8) and can be easily visualized in long axis views, preferably in the twochamber view.

Tissue Characterization There are only few data on tissue abnormalities in this disease. As was recently reported, myocardial edema and occasionally appearing focal areas of LGE may be observed (see Fig. 38-8), reflecting extensive, typically completely reversible tissue injury.224 It has been suggested that LGE represents a disproportionate increase in extracellular matrix.225

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101. Young AA, Kramer CM, Ferrari VA, Axel L, Reichek N. Threedimensional left ventricular deformation in hypertrophic cardiomyopathy. Circulation. 1994;90:854–867. 102. Petersen SE, Selvanayagam JB, Francis JM, et al. Differentiation of athlete’s heart from pathological forms of cardiac hypertrophy by means of geometric indices derived from cardiovascular magnetic resonance. J Cardiovasc Magn Reson. 2005;7:551–558. 103. Sachdev V, Shizukuda Y, Brenneman CL, et al. Left atrial volumetric remodeling is predictive of functional capacity in nonobstructive hypertrophic cardiomyopathy. Am Heart J. 2005;149:730–736. 104. Germans T, Wilde AA, Dijkmans PA, et al. Structural abnormalities of the inferoseptal left ventricular wall detected by cardiac magnetic resonance imaging in carriers of hypertrophic cardiomyopathy mutations. J Am Coll Cardiol. 2006;48:2518–2523. 105. Wilson JM, Villareal RP, Hariharan R, Massumi A, Muthupillai R, Flamm SD. Magnetic resonance imaging of myocardial fibrosis in hypertrophic cardiomyopathy. Tex Heart Inst J. 2002;29:176–180. 106. Choudhury L, Mahrholdt H, Wagner A, et al. Myocardial scarring in asymptomatic or mildly symptomatic patients with hypertrophic cardiomyopathy. J Am Coll Cardiol. 2002;40:2156–2164. 107. Bogaert J, Goldstein M, Tannouri F, Golzarian J, Dymarkowski S. Original report: late myocardial enhancement in hypertrophic cardiomyopathy with contrast-enhanced MR imaging. AJR Am J Roentgenol. 2003;180:981–985. 108. Amano Y, Takayama M, Takahama K, Kumazaki T. Delayed hyperenhancement of myocardium in hypertrophic cardiomyopathy with asymmetrical septal hypertrophy: comparison with global and regional cardiac MR imaging appearances. J Magn Reson Imaging. 2004;20:595–600. 109. Aso H, Takeda K, Ito T, Shiraishi T, Matsumura K, Nakagawa T. Assessment of myocardial fibrosis in cardiomyopathic hamsters with gadolinium-DTPA enhanced magnetic resonance imaging. Invest Radiol. 1998;33:22–32. 110. Moon JC, Reed E, Sheppard MN, et al. The histologic basis of late gadolinium enhancement cardiovascular magnetic resonance in hypertrophic cardiomyopathy. J Am Coll Cardiol. 2004;43:2260–2264. 111. Papavassiliu T, Fluchter S, Haghi D, et al. Extent of myocardial hyperenhancement on late gadolinium-enhanced cardiovascular magnetic resonance correlates with q waves in hypertrophic cardiomyopathy. J Cardiovasc Magn Reson. 2007;9:595–603. 112. Elliott PM, Gimeno Blanes JR, Mahon NG, Poloniecki JD, McKenna WJ. Relation between severity of left-ventricular hypertrophy and prognosis in patients with hypertrophic cardiomyopathy. Lancet. 2001;357:420–424. 113. Moon JC, McKenna WJ, McCrohon JA, Elliott PM, Smith GC, Pennell DJ. Toward clinical risk assessment in hypertrophic cardiomyopathy with gadolinium cardiovascular magnetic resonance. J Am Coll Cardiol. 2003;41:1561–1567. 113a. Kwon DH, Smedira NG, Rodriguez ER, et al. Cardiac magnetic resonance detection of myocardial scarring in hypertrophic cardiomyopathy: correlation with histopathology and prevalence of ventricular tachycardia. J Am Coll Cardiol. 2009;54:242–249. 113b. Leonardi S, Raineri C, DeFerrari GM, et al. Usefulness of cardiac magnetic resonance in assessing the risk of ventricular arrhythmias and sudden death in patients with hypertrophic cardiomyopathy. Eur Heart J. 2009;30:2003–2010. 114. Moon JC, Mogensen J, Elliott PM, et al. Myocardial late gadolinium enhancement cardiovascular magnetic resonance in hypertrophic cardiomyopathy caused by mutations in troponin I. Heart. 2005; 91:1036–1040. 115. Kim RJ, Judd RM. Gadolinium-enhanced magnetic resonance imaging in hypertrophic cardiomyopathy: in vivo imaging of the pathologic substrate for premature cardiac death? J Am Coll Cardiol. 2003;41:1568–1572. 116. de Roos A, Doornbos J, Luyten PR, Oosterwaal LJ, van der Wall EE, den Hollander JA. Cardiac metabolism in patients with dilated and hypertrophic cardiomyopathy: assessment with proton-decoupled P-31 MR spectroscopy. J Magn Reson Imaging. 1992;2:711–719. 117. Spindler M, Saupe KW, Christe ME, et al. Diastolic dysfunction and altered energetics in the alphaMHC403/þ mouse model of familial hypertrophic cardiomyopathy. J Clin Invest. 1998;101:1775–1783. 118. Melacini P, Corbetti F, Calore C, et al. Cardiovascular magnetic resonance signs of ischemia in hypertrophic cardiomyopathy. Int J Cardiol. 2007. 119. Abdel-Aty H. Abnormalities in T2-weighted cardiovascular magnetic resonance images of hypertrophic cardiomyopathy: regional

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138. Fontaliran F, Arkwright S, Vilde F, Fontaine G. [Arrhythmogenic right ventricular dysplasia and cardiomyopathy: clinical and anatomic-pathologic aspects, nosologic approach]. Arch Anat Cytol Pathol. 1998;46:171–177. 139. McKenna WJ, Thiene G, Nava A, et al. Diagnosis of arrhythmogenic right ventricular dysplasia/cardiomyopathy. Task Force of the Working Group Myocardial and Pericardial Disease of the European Society of Cardiology and of the Scientific Council on Cardiomyopathies of the International Society and Federation of Cardiology. Br Heart J. 1994;71:215–218. 140. Castillo E, Rutberg J, Wu KC, Bluemke DA. Arrhythmogenic right ventricular dysplasia/cardiomyopathy: variability in protocols and image quality between MRI centers. Circulation. 2001;104: II-640. 141. Bluemke DA, Krupinski EA, Ovitt T, et al. MR Imaging of arrhythmogenic right ventricular cardiomyopathy: morphologic findings and interobserver reliability. Cardiology. 2003;99:153–162. 142. Tandri H, Friedrich MG, Calkins H, Bluemke DA. MRI of arrhythmogenic right ventricular cardiomyopathy/dysplasia. J Cardiovasc Magn Reson. 2004;6:557–563. 143. Blake LM, Scheinman MM, Higgins CB. MR features of arrhythmogenic right ventricular dysplasia. AJR Am J Roentgenol. 1994;162: 809–812. 144. Casolo GC, Poggesi L, Boddi M, et al. ECG-gated magnetic resonance imaging in right ventricular dysplasia. Am Heart J. 1987;113: 1245–1248. 145. Frank H, Rimpfl T, Weber H, et al. [Diagnosis of arrhythmogenic right ventricular disease using magnetic resonance tomography]. Z Kardiol. 1991;80:569–573. 146. Ricci C, Longo R, Pagnan L, et al. Magnetic resonance imaging in right ventricular dysplasia. Am J Cardiol. 1992;70:1589–1595. 147. Auffermann W, Wichter T, Breithardt G, Joachimsen K, Peters PE. Arrhythmogenic right ventricular disease: MR imaging vs angiography. AJR Am J Roentgenol. 1993;161:549–555. 148. Menghetti L, Basso C, Nava A, Angelini A, Thiene G. Spin-echo nuclear magnetic resonance for tissue characterisation in arrhythmogenic right ventricular cardiomyopathy. Heart. 1996;76:467–470. 149. Dalal D, Tandri H, Judge DP, et al. Morphologic variants of familial arrhythmogenic right ventricular dysplasia/cardiomyopathy. J Am Coll Cardiol. 2009;53:1289–1299. 150. Bomma C, Dalal D, Tandri H, et al. Regional differences in systolic and diastolic function in arrhythmogenic right ventricular dysplasia/cardiomyopathy using magnetic resonance imaging. Am J Cardiol. 2005;95:1507–1511. 151. Tandri H, Calkins H, Nasir K, et al. Magnetic resonance imaging findings in patients meeting task force criteria for arrhythmogenic right ventricular dysplasia. J Cardiovasc Electrophysiol. 2003;14: 476–482. 152. Carlson MD, White RD, Trohman RG, et al. Right ventricular outflow tract ventricular tachycardia: detection of previously unrecognized anatomic abnormalities using cine magnetic resonance imaging. J Am Coll Cardiol. 1994;24:720–727. 153. Globits S, Kreiner G, Frank H, et al. Significance of morphological abnormalities detected by MRI in patients undergoing successful ablation of right ventricular outflow tract tachycardia. Circulation. 1997;96:2633–2640. 154. Proclemer A, Basadonna PT, Slavich GA, Miani D, Fresco C, Fioretti PM. Cardiac magnetic resonance imaging findings in patients with right ventricular outflow tract premature contractions [see comments]. Eur Heart J. 1997;18:2002–2010. 155. White RD, Trohman RG, Flamm SD, et al. Right ventricular arrhythmia in the absence of arrhythmogenic dysplasia: MR imaging of myocardial abnormalities. Radiology. 1998;207:743–751. 156. Tandri H, Bluemke DA, Ferrari VA, et al. Findings on magnetic resonance imaging of idiopathic right ventricular outflow tachycardia. Am J Cardiol. 2004;94:1441–1445. 157. Dalal D, Nasir K, Bomma C, et al. Arrhythmogenic right ventricular dysplasia: a United States experience. Circulation. 2005;112: 3823–3832. 158. Widmaier S, Jung WI, Sieverding L, et al. Detection of monoester signals in human myocardium by 31P-MRS. MAGMA. 1998;6: 161–163. 159. Schick F, Miller S, Hahn U, et al. Fat- and water-selective MR cine imaging of the human heart: assessment of right ventricular dysplasia. Invest Radiol. 2000;35:311–318. 160. Vikhert AM, Tsiplenkova VG, Cherpachenko NM. Alcoholic cardiomyopathy and sudden cardiac death. J Am Coll Cardiol. 1986;8:3a–11a. 161. Hasumi M, Sekiguchi M, Hiroe M, Kasanuki H, Hirosawa K. Endomyocardial biopsy approach to patients with ventricular tachycardia

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162. 163.

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with special reference to arr Hunold P, Wieneke H, Bruder O, et al. Late enhancement: a new feature in MRI of arrhythmogenic right ventricular cardiomyopathy? J Cardiovasc Magn Reson. 2005;7: 649–655. Hunold P, Wieneke H, Bruder O, et al. Late enhancement: a new feature in MRI of arrhythmogenic right ventricular cardiomyopathy. J Cardiovasc Magn Reson. 2005;649–655. Tandri H, Saranathan M, Rodriguez ER, et al. Noninvasive detection of myocardial fibrosis in arrhythmogenic right ventricular cardiomyopathy using delayed-enhancement magnetic resonance imaging. J Am Coll Cardiol. 2005;45:98–103. Sen-Chowdhry S, Syrris P, Ward D, Asimaki A, Sevdalis E, McKenna WJ. Clinical and genetic characterization of families with arrhythmogenic right ventricular dysplasia/cardiomyopathy provides novel insights into patterns of disease expression. Circulation. 2007;115:1710–1720. Sen-Chowdhry S, Prasad SK, Syrris P, et al. Cardiovascular magnetic resonance in arrhythmogenic right ventricular cardiomyopathy revisited: comparison with task force criteria and genotype. J Am Coll Cardiol. 2006;48:2132–2140. Mohiaddin RH, Hasegawa M. Measurement of atrial volumes by magnetic resonance imaging in healthy volunteers and in patients with myocardial infarction. Eur Heart J. 1995;16:106–111. Jarvinen VM, Kupari MM, Poutanen VP, Hekali PE. A simplified method for the determination of left atrial size and function using cine magnetic resonance imaging. Magn Reson Imaging. 1996;14: 215–226. Restrictive cardiomyopathy or constrictive pericarditis? [editorial]. Lancet. 1987;2:372–374. Sechtem U, Higgins CB, Sommerhoff BA, Lipton MJ, Huycke EC. Magnetic resonance imaging of restrictive cardiomyopathy. Am J Cardiol. 1987;59:480–482. Masui T, Finck S, Higgins CB. Constrictive pericarditis and restrictive cardiomyopathy: evaluation with MR imaging. Radiology. 1992;182:369–373. Talreja DR, Edwards WD, Danielson GK, et al. Constrictive pericarditis in 26 patients with histologically normal pericardial thickness. Circulation. 2003;108:1852–1857. Francone M, Dymarkowski S, Kalantzi M, Bogaert J. Real-time cine MRI of ventricular septal motion: a novel approach to assess ventricular coupling. J Magn Reson Imaging. 2005;21:305–309. Ohyama H, Hosomi N, Takahashi T, et al. Comparison of magnetic resonance imaging and transesophageal echocardiography in detection of thrombus in the left atrial appendage. Stroke. 2003;34: 2436–2439. Raggi P, Daniels M, Shanoudy H, Jarmukli NF. MRI misinterpretation of spontaneous echo-contrast as a large left atrial thrombus. Int J Card Imaging. 1996;12:85–88. Ritter M, Oechslin E, Sutsch G, Attenhofer C, Schneider J, Jenni R. Isolated noncompaction of the myocardium in adults. Mayo Clin Proc. 1997;72:26–31. Oechslin EN, Attenhofer Jost CH, Rojas JR, Kaufmann PA, Jenni R. Long-term follow-up of 34 adults with isolated left ventricular noncompaction: a distinct cardiomyopathy with poor prognosis. J Am Coll Cardiol. 2000;36:493–500. Jenni R, Oechslin E, Schneider J, Jost CA, Kaufmann PA. Echocardiographic and pathoanatomical characteristics of isolated left ventricular non-compaction: a step towards classification as a distinct cardiomyopathy. Heart. 2001;86:666–671. Junga G, Kneifel S, Von Smekal A, Steinert H, Bauersfeld U. Myocardial ischaemia in children with isolated ventricular non-compaction. Eur Heart J. 1999;20:910–916. Yasukawa K, Terai M, Honda A, Kohno Y. Isolated noncompaction of ventricular myocardium associated with fatal ventricular fibrillation. Pediatr Cardiol. 2001;22:512–514. Petersen SE, Selvanayagam JB, Wiesmann F, et al. Left ventricular non-compaction: insights from cardiovascular magnetic resonance imaging. J Am Coll Cardiol. 2005;46:101–105. Petersen SE, Timperley J, Neubauer S. Left ventricular thrombi in a patient with left ventricular non-compaction in visualisation of the rationale for anticoagulation. Heart. 2005;91:e4. Pujadas S, Bordes R, Bayes-Genis A. Ventricular non-compaction cardiomyopathy: CMR and pathology findings. Heart. 2005;91:582. Sugie T, Hashimoto N, Iwai K. [Clinical and autopsy studies on prognosis of sarcoidosis]. Nippon-Rinsho. 1994;52:1567–1570. Perry A, Vuitch F. Causes of death in patients with sarcoidosis: a morphologic study of 38 autopsies with clinicopathologic correlations. Arch Pathol Lab Med. 1995;119:167–172.

206. Sparacia G, Banco A, Midiri M, Iaia A. MR imaging technique for the diagnosis of pituitary iron overload in patients with transfusiondependent beta-thalassemia major. AJNR Am J Neuroradiol. 1998;19:1905–1907. 207. Waxman S, Eustace S, Hartnell GG. Myocardial involvement in primary hemochromatosis demonstrated by magnetic resonance imaging. Am Heart J. 1994;128:1047–1049. 208. Anderson LJ, Holden S, Davis B, et al. Cardiovascular T2-star (T2*) magnetic resonance for the early diagnosis of myocardial iron overload. Eur Heart J. 2001;22:2171–2179. 209. Westwood MA, Wonke B, Maceira AM, et al. Left ventricular diastolic function compared with T2* cardiovascular magnetic resonance for early detection of myocardial iron overload in thalassemia major. J Magn Reson Imaging. 2005;22:229–233. 210. Anderson LJ, Wonke B, Prescott E, Holden S, Walker JM, Pennell DJ. Comparison of effects of oral deferiprone and subcutaneous desferrioxamine on myocardial iron concentrations and ventricular function in beta-thalassaemia. Lancet. 2002;360:516–520. 211. Pennell DJ, Berdoukas V, Karagiorga M, et al. Randomized controlled trial of deferiprone or deferoxamine in beta-thalassemia major patients with asymptomatic myocardial siderosis. Blood. 2006;107:3738–3744. 212. Tanner MA, Galanello R, Dessi C, et al. A randomized, placebocontrolled, double-blind trial of the effect of combined therapy with deferoxamine and deferiprone on myocardial iron in thalassemia major using cardiovascular magnetic resonance. Circulation. 2007;115:1876–1884. 213. Kirk P, Roughton M, Porter JB, Walker JM, Tanner MA, Patel J, et al. Cardiac T2* magnetic resonance for prediction of cardiac complications in thalassemia major (abstract). J Cardiovasc Magn Reson. 2009;11(suppl 1): O2. 214. Modell B, Khan M, Darlison M, Westwood MA, Ingram D, Pennell DJ. Improved survival of thalassaemia major in the UK and relation to T2* cardiovascular magnetic resonance. J Cardiovasc Magn Reson. 2008;10:42. 215. Katritsis D, Wilmshurst PT, Wendon JA, Davies MJ, Webb Peploe MM. Primary restrictive cardiomyopathy: clinical and pathologic characteristics. J Am Coll Cardiol. 1991;18:1230–1235. 216. D’Silva SA, Kohli A, Dalvi BV, Kale PA. MRI in right ventricular endomyocardial fibrosis. Am Heart J. 1992;123:1390–1392. 217. Celletti F, Fattori R, Napoli G, et al. Assessment of restrictive cardiomyopathy of amyloid or idiopathic etiology by magnetic resonance imaging. Am J Cardiol. 1999;83:798–801, a10. 218. Huong DL, Wechsler B, Papo T, et al. Endomyocardial fibrosis in Behcet’s disease. Ann Rheum Dis. 1997;56:205–208. 219. Salanitri GC. Endomyocardial fibrosis and intracardiac thrombus occurring in idiopathic hypereosinophilic syndrome. AJR Am J Roentgenol. 2005;184:1432–1433. 220. Cury RC, Abbara S, Sandoval LJ, Houser S, Brady TJ, Palacios IF. Images in cardiovascular medicine: visualization of endomyocardial fibrosis by delayed-enhancement magnetic resonance imaging. Circulation. 2005;111: e115–7. 221. Cocker M, Abdel-Aty H, Alakija P, Alakija O, Friedrich M. Images in cardiology: cardiac magnetic resonance imaging in Lo¨ffler’s endocarditis. Can J Cardiol. 2008;24:e89–e90. 222. Kurisu S, Sato H, Kawagoe T, et al. Tako-Tsubo-like left ventricular dysfunction with ST-segment elevation: a novel cardiac syndrome mimicking acute myocardial infarction. Am Heart J. 2002;143: 448–455. 223. Sharkey SW, Lesser JR, Zenovich AG, et al. Acute and reversible cardiomyopathy provoked by stress in women from the United States. Circulation. 2005;111:472–479. 224. Abdel-Aty H, Cocker M, Friedrich MG. Myocardial edema is a feature of Tako-Tsubo cardiomyopathy and is related to the severity of systolic dysfunction: insights from T2-weighted cardiovascular magnetic resonance. Int J Cardiol. 2007;132:291–293. 225. Rolf A, Nef HM, Mollmann H, et al. Immunohistological basis of the late gadolinium enhancement phenomenon in tako-tsubo cardiomyopathy. Eur Heart J. 2009;30:1635–1642.

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185. Sekiguchi M, Yazaki Y, Isobe M, Hiroe M. Cardiac sarcoidosis: diagnostic, prognostic, and therapeutic considerations. Cardiovasc Drugs Ther. 1996;10:495–510. 186. Smedema JP, Snoep G, van Kroonenburgh MP, et al. The additional value of gadolinium-enhanced MRI to standard assessment for cardiac involvement in patients with pulmonary sarcoidosis. Chest. 2005;128:1629–1637. 187. Kurashima K, Shimizu H, Ogawa H, et al. MR and CT in the evaluation of sarcoid myopathy. J Comput Assist Tomogr. 1991;15:1004–1007. 188. Otake S, Banno T, Ohba S, Noda M, Yamamoto M. Muscular sarcoidosis: findings at MR imaging. Radiology. 1990;176:145–148. 189. Seltzer S, Mark AS, Atlas SW. CNS sarcoidosis: evaluation with contrast-enhanced MR imaging. AJNR Am J Neuroradiol. 1991;12: 1227–1233. 190. Riedy K, Fisher MR, Belic N, Koenigsberg DI. MR imaging of myocardial sarcoidosis. AJR Am J Roentgenol. 1988;151:915–916. 191. Dupuis JM, Victor J, Furber A, Pezard P, Lejeune LL, Tadei A. [Value of magnetic resonance imaging in cardiac sarcoidosis: apropos of a case]. Arch Mal Coeur Vaiss. 1994;87:105–110. 192. Eliasch H, Juhlin Dannfelt A, Sjogren I, Terent A. Magnetic resonance imaging as an aid to the diagnosis and treatment evaluation of suspected myocardial sarcoidosis in a fighter pilot. Aviat Space Environ Med. 1995;66:1010–1013. 193. Chandra M, Silverman ME, Oshinski J, Pettigrew R. Diagnosis of cardiac sarcoidosis aided by MRI. Chest. 1996;110:562–565. 194. Doherty MJ, Kumar SK, Nicholson AA, McGivern DV. Cardiac sarcoidosis: the value of magnetic resonance imaging in diagnosis and assessment of response to treatment. Respir Med. 1998;92:697–699. 195. Smedema JP, Snoep G, van Kroonenburgh MP, et al. Evaluation of the accuracy of gadolinium-enhanced cardiovascular magnetic resonance in the diagnosis of cardiac sarcoidosis. J Am Coll Cardiol. 2005;45:1683–1690. 196. Schulz-Menger J, Wassmuth R, Abdel-Aty H, et al. Patterns of myocardial inflammation and scarring in sarcoidosis as assessed by cardiovascular magnetic resonance. Heart. 2006;92:399–400. 197. Klein AL, Hatle LK, Taliercio CP, et al. Prognostic significance of Doppler measures of diastolic function in cardiac amyloidosis: a Doppler echocardiography study. Circulation. 1991;83:808–816. 198. Pellikka PA, Holmes Jr DR, Edwards WD, Nishimura RA, Tajik AJ, Kyle RA. Endomyocardial biopsy in 30 patients with primary amyloidosis and suspected cardiac involvement. Arch-Intern-Med. 1988;148:662–666. 199. Matsuoka H, Hamada M, Honda T, et al. Precise assessment of myocardial damage associated with secondary cardiomyopathies by use of Gd-DTPA-enhanced magnetic resonance imaging. Angiology. 1993;44:945–950. 199a. Ruberg FL, Appelbaum E, Davidoff R, et al. Diagnostic and prognostic utility of cardiovascular magnetic resonance imaging in lightchain cardiac amyloidosis. Am J Cardiol. 2009;103:544–549. 200. Maceira AM, Prasad SK, Hawkins PN, Roughton M, Pennell DJ. Cardiovascular magnetic resonance and prognosis in cardiac amyloidosis. J Cardiovasc Magn Reson. 2008;10:54. 201. Siegelman ES, Mitchell DG, Rubin R, et al. Parenchymal versus reticuloendothelial iron overload in the liver: distinction with MR imaging [see comments]. Radiology. 1991;179:361–366. 202. Gandon Y, Guyader D, Heautot JF, et al. Hemochromatosis: diagnosis and quantification of liver iron with gradient-echo MR imaging. Radiology. 1994;193:533–538. 203. Siegelman ES, Mitchell DG, Semelka RC. Abdominal iron deposition: metabolism, MR findings, and clinical importance. Radiology. 1996;199:13–22. 204. Ernst O, Sergent G, Bonvarlet P, Canva Delcambre V, Paris JC, L’Hermine C. Hepatic iron overload: diagnosis and quantification with MR imaging. AJR Am J Roentgenol. 1997;168:1205–1208. 205. Jager HJ, Mehring U, Gotz GF, et al. Radiological features of the visceral and skeletal involvement of hemochromatosis. Eur Radiol. 1997;7:1199–1206.

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CHAPTER 39

Cardiac and Paracardiac Masses Herbert Frank and Francisco Alpendurada

Cardiovascular magnetic resonance (CMR) provides noninvasive and three-dimensional assessment of masses involving the cardiac chambers, pericardium, and extracardiac structures. CMR has become an established method to yield complementary diagnostic information and guide cardiac surgeons in the design of an appropriate therapeutic strategy. Furthermore, to some extent, CMR allows characterization of tumor tissue. The goal of CMR for assessing cardiac and paracardiac masses includes confirming or excluding a mass suspected by X-ray or echocardiography1; assessing its location, mobility, and relationship to surrounding tissues2; imaging the degree of vascularity3; distinguishing solid lesions from fluid4; and determining tissue characteristics and the specific nature of a mass.5

TECHNICAL CONSIDERATIONS For adequate image quality with reduced motion artifacts, CMR is performed with electrocardiographic gating.1 Alternatively, the recently developed navigator technique allows combined electrocardiogram and respiratory triggering.2 Spin echo sequences provide detailed morphologic information on the heart, great vessels, and adjacent structures.3 For T1weighted sequences, echo time (TE) is usually 20 to 30 msec and repetition time (TR) is dependent on the R-R interval. A longer TR is used for T2-weighted sequences, and TE is typically 50 to 90 msec.4 T1-weighted images provide a better signal-to-noise ratio and excellent soft tissue contrast between epicardial fat, myocardium, and rapidly flowing blood.3 T2weighted images have increased image contrast, which may be helpful for tissue characterization. However, T2-weighted images in general have worse signal-to-noise ratios. The turbo or fast spin echo (FSE) technique combines the acquisition of multiple profiles per excitation with the multislice mode, resulting in a marked reduction in imaging time. The contrast of those images is similar to that of spin echo images, with the same TR and equivalent TE. Fat is usually brighter with FSE than with regular spin echo pulse sequences. FSE permits acquisition of a T2-weighted scan in a fraction of time compared with the conventional spin echo sequence. Furthermore, this technique has reduced susceptibility to motion artifacts and is insensitive to field inhomogeneity. The combination of inversion recovery and breath hold techniques led to improved image quality and tissue characterization, which must be assessed in the future.5 When combined with an inversion recovery technique, the signal from fat can be suppressed with a short inversion

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time before the spin echo pulse. Most tissues have a T1 relaxation time longer than that of fat, resulting in a signal increase on T1- and T2-weighted images. This leads to better contrast in these images. Applications of the inversion recovery technique include the distinct recognition of fatty tissues and improved visualization of structures surrounded by fatty tissue, which often makes image analysis complicated. For evaluation of cardiac function and tumor mobility, gradient recalled echo (GRE) techniques (TE 4 to 12 msec) are recommended.6,7 GRE images are characterized by bright signal intensity from rapidly flowing blood, which is useful for the differentiation of thrombus or the assessment of turbulent flow in cases of valvular regurgitation or intracardiac shunt. Because of the low soft tissue contrast, visualization of cardiac masses is not as adequate as with spin echo techniques, unless they are intraluminal. Different pulse sequences used in magnetic resonance for the evaluation of cardiac masses and respective advantages and disadvantages are summarized in Table 39-1.

CONTRAST AGENTS The most commonly used CMR contrast agent uses the paramagnetic element gadolinium in a complex (e.g., chelated with diethylene-triamine-pentaacetic acid)8 with other molecules to reduce toxicity. A number of agents are available. The distribution of nearly all currently available intravenous contrast agents is extracellular, and contrast is enhanced predominantly on T1-weighted images. The normal concentration is 0.1 mmol/kg body weight.9 Gadolinium provides better delineation of the mass by different enhancement of myocardium and tumor tissue as a result of variation in tissue vascularity.10

BENIGN TUMORS OF THE HEART Primary tumors of the heart are very rare, with an incidence of between 0.0017% and 0.19% in unselected patients at autopsy. Three quarters of the tumors are benign, nearly half of them are myxomas, and approximately 10% are lipomas.11–14 Rhabdomyomas, fibromas, hemangiomas, teratomas, and mesotheliomas are found less frequently. Granular cell tumors, neurofibromas, and lymphangiomas are very rare.15

Advantages/ Disadvantages

Sequence

Technical Features

Indication

T1-weighted spin echo

8 12 slices, thickness: 6 10 mm, NSA = 2, TE: 20 25 ms, TR: shortest, time 4 6 mm, axial, sagittal, and coronal orientation. 8 12 slices, thickness: 6 10 mm, NSA = 2, TE: 50, 90 ms, TR: every 2nd or 3rd heartbeat, time: 5 8 min, axial orientation.

Defining anatomic structures, delineation of the mass to the adjacent tissue, visualization of vascular walls.

T2-weighted spin echo, double echo

Detecting the nature of cardiac masses by abnormal T2 values.

T1- and T2weighted fast spin echo

Examination time: 2 4 min, also breath-hold technique

Short inversion recovery T1-weighted spin echo with contrast agent administration Cine gradient echo with flow compensation

Inversion pulse, combination with spin echo or TSE Axial orientation, antecubital Gd-DTPA administration (Dose: 0, 1 mmol/kg)

Suppression of fatty tissue

1 3 slices, thickness 8 10 mm, TE: 9 ms, TR: shortest, 16 phases NSA = 2, time: 1 3 min.

Imaging the hemodynamic effects of a mass, i.e., mobility, transvalvular flow, differentiation between blood flow and thrombus, identify turbulent flow regions.

Cine gradient echo EPI

Breath-hold technique, 1 3 slices, thickness, 8 10 mm, TE: shortest, TR: shortest, time: 10 40 sec

Equal to conventional gradient echo

Signal intensity behavior of suspected masses, assessment of the degree of vascularity.

A: Excellent soft tissue contrast D: Respiratory and flow artifacts A: Better tissue contrast, demonstration of fluid components D: Very long examination time, lower SNR, and increased motion artifacts A: Shorter imaging times, reduced motion artifacts D: Less SNR A: Eliminating artifacts from fat signal To assess invasive and infiltrative components of a tumor A: Functional information D: Low soft tissue contrast A: Shorter acquisition time, reduced motion artifacts D: Less SNR

Modified from Hoffman U, Globits S, Frank H: Cardiac and paracardiac masses: current opinion on diagnostic evaluation by magnetic resonance imaging. Eur Heart J. 1998;19:553–563. EPI, echo-planar imaging; Gd-DTPA, gadolinium-DTPA; NSA, numbers of signal averaged; SNR, signal-to-noise ratio; TE, echo time; TSE, turbo spin-echo; TR, repetition time.

Myxoma Myxomas account for 30% to 50% of all primary cardiac tumors and usually occur sporadically between the third and sixth decades of life.16 Approximately 75% of myxomas originate from the left atrium and 15% to 20% from the right atrium. They usually develop from the interatrial septum close to the fossa ovalis. Only a few myxomas are located in the ventricles.17 Rare aneurysmatic formations of part of the septum secundum can lead to a gallbladderlike cystic structure and can mimic an atrial myxoma on CMR.18 The histologic structure typically shows a myxoid matrix, large blood vessels at the base, and often cysts and areas of hemorrhage and calcification.19 They are generally polypoid, and are often pedunculated, round, or oval, with a smooth surface, often covered with thrombi. Tumors range from 1 to 15 cm in diameter.20 Clinical symptoms appear as a consequence of embolism or intracardiac obstruction and are determined by the size, location, and mobility of the myxoma.21,22 On CMR, myxomas (Fig. 39-1) are mainly diagnosed by the typically pedunculated, jelly-like, and prolapsing appearance as well as by certain signal characteristics.23 Therefore, cine display should be obtained to show the mobility of the tumor.24 Because of their endocardial origin, myxomas are

characterized by an intermediate, but variable signal on spin echo images, similar to that of the myocardium. On GRE images, myxomas often have low signal intensity, which is caused by partial calcification. Therefore, the tumor can always be distinguished from the higher signal of the surrounding, slowly flowing blood.25,26 Intratumorous areas of subacute or chronic hemorrhage are typically characterized by high signal intensity on both short and long echo times.27 After intravenous injection of gadolinium, myxomas show moderately high contrast enhancement, which is caused by high vascularity.9,25

Lipoma Cardiac lipomas (Fig. 39-2) are the second most frequent benign tumors of the heart.15 True lipomas are encapsulated, contain neoplastic fat cells, and occur in young patients.28,29 Approximately 50% arise subendocardially, 25% subepicardially, and 25% from the myocardium.30 Subepicardial lipomas may become quite large and may alter cardiac function, resulting in dyspnea or fatigue,31 and involvement of the coronary arteries has been reported.32 Endocardial lipomas commonly arise from the interatrial septum and are located in the right atrium. Arrhythmias caused by myocardial infiltration have been Cardiovascular Magnetic Resonance 533

39 CARDIAC AND PARACARDIAC MASSES

Table 39-1 Typical Technical Parameters at 1.5 T, for Evaluation of Cardiac or Paracardiac Masses

FUNCTIONAL CARDIOVASCULAR DISEASE

Figure 39-1 Left atrial myxoma. Four-chamber (A) and atrial short axis (B) steady-state freeprecession images showing a left atrial mass attached to the fossa ovalis (arrows). On late gadolinium imaging (C and D), there is typical heterogeneous enhancement within the mass.

A

B

C

D

reported.33 Lipomatous hypertrophy of the atrial septum (Fig. 39-3) is histologically characterized by infiltration of lipomatous cells between atrial muscle fibers. Unlike true lipomas, they are unencapsulated and contain lipoblasts as well as mature fat cells.34 This condition has been described in older, overweight patients, who frequently have atrial fibrillation.35 On CMR, lipomas are characterized by bright signal intensity on T1-weighted images and a slight decrease in signal intensity on T2-weighted images, similar to that of subcutaneous fat.36 Gadolinium injection is not needed because signal intensity will remain unchanged. A decrease in signal intensity using the fat presaturation technique verifies the diagnosis. In lipomatous hypertrophy, bilobed atrial septum thickening with signal intensity comparable to that of subcutaneous fat on T1- and T2-weighted images can be visualized. In contrast to benign lipomatous hypertrophy of the interatrial septum, the very rare liposarcoma usually shows infiltration, inhomogeneities, and fast tumor growth.37

Papillary Fibroelastoma Papillary fibroelastomas are the most common tumors of cardiac valves. These papillary structures consist of elastin 534 Cardiovascular Magnetic Resonance

and collagen layers covered by endothelium, and they predominantly involve the left side of the heart. The aortic valve is most frequently involved, followed by the mitral valve and the left ventricular outflow tract.38 Recognition is important because neurologic events caused by embolization are common complications. Because fibroelastomas are relatively small (usually < 10 mm), mobile structures, echocardiography is the preferred imaging modality. On CMR, they usually present as regular, round lesions attached to the cardiac valves or endothelium. They have low signal on GRE and appear isointense on T1-weighted images and hyperintense on T2-weighted images.39 Although they are relatively avascular (no enhancement with first-pass perfusion imaging), enhancement is seen with late gadolinium imaging (Fig. 39-4).40

Fibroma Fibromas occur primarily in infants and children. They are congenital tumors that are frequently discovered in young adults.41 Typically, fibromas are located intramyocardially, within the ventricular septum. The left ventricle is more often involved than the right ventricle. Surgical excision is recommended, even in asymptomatic patients,42 because of the potential risk of sudden death caused by arrhythmias.

RV LV

A

B

C

A

39 CARDIAC AND PARACARDIAC MASSES

Figure 39-2 Atrial lipoma. A, T1weighted turbo spin echo imaging in a horizontal long axis plane without fat suppression, showing a large intracardiac mass crossing the interatrial septum (arrow). Note the bright appearance of the chest wall (arrowhead) with this imaging technique. B, Imaging in the same plane with the addition of fat suppression. Note the reduced chest wall signal intensity (arrowhead). C, Surgical removal of the mass confirmed the diagnosis of lipoma by histopathology. D, Hematoxylin and eosin stain. LV, left ventricle; RV, right ventricle.

D

B

C

Figure 39-3 Lipomatous hypertrophy of the interatrial septum. Steady-state free precession image (A) showing fatty deposition within the interatrial septum. Note the characteristic bilobed, or hourglass-shaped, appearance, typically sparing the fossa ovalis region (white arrows). The diagnosis of fatty deposition can be confirmed as showing an area of high signal in the interatrial septum on T1-weighted turbo spin echo images (B), with signal drop-out after fat suppression (C). Note the prominent fat deposition around the heart and chest wall (black arrows).

On T1-weighted spin echo images, fibromas are isointense or slightly hyperintense compared with skeletal muscle. Because of the short T2 relaxation time of fibrous tissue, fibromas show a decrease in signal intensity relative to the myocardium on T2-weighted spin echo images.43,44 These masses have low signal on perfusion images and early after gadolinium injection as a result of low tumor vascularity. However,

late after gadolinium injection, they show significantly homogeneous enhancement (Fig. 39-5). A possible problem in diagnosing fibrous tissue might be the presence of fibromuscular elements within the right atrium. Small, nodular soft tissue structures that are isointense to the myocardium and nodules or linear strands in the right atrium are commonly visible and may simulate a Cardiovascular Magnetic Resonance 535

FUNCTIONAL CARDIOVASCULAR DISEASE

A

B

C

D

E

F

Figure 39-4 Papillary fibroelastoma of the aortic valve. Steady-state free precession image (A) showing a solid intraluminal mass (arrow) attached to the commissure of the left coronary and noncoronary aortic valve cusps. T1-weighted (B) and T2-weighted (C) spin echo images showing a solid lesion (arrows) with intermediate signal intensity. The gross specimen (D) appeared as a translucent, gelatinous mass. Hematoxylin and eosin staining of the histologic specimen shows a benign lesion with multiple papillary fronds (E). The papillary fronds consist of three layers, forming a collagenous core with low elastin content, an amorphous intermediate layer, and a delicate coat of singlelayer endothelium (F), as shown with elastica-van Gieson stain.

tumor.45 These structures represent variable degrees of remnants of the crista terminalis and the Chiari network in humans.46

Rhabdomyoma Rhabdomyomas are congenital tumors that are mainly diagnosed in newborn infants. Usually, they arise from the ventricular myocardium at multiple locations. In approximately 50% of patients, tuberous sclerosis can be diagnosed.47 Rhabdomyomas typically have a solid, homogeneous appearance that is hypointense to the myocardium on T1-weighted images and slightly hyperintense on T2-weighted images.48 These masses typically enhance after gadolinium injection (Fig. 39-6).

Cardiac Hemangioma There are few cases reported of arteriovenous hemangiomas of the interventricular septum of the heart. The location of the tumor is predominantly the right or left ventricle, and septal involvement and multiple locations also have been reported.49 The distinction between a hemangioma and a vascular malformation can be difficult.50 On CMR, hemangiomas are characterized as regions of increased signal 536 Cardiovascular Magnetic Resonance

intensity on T2-weighted images compared with the myocardium as a result of slow-flowing blood. After intravenous injection of gadolinium, the vascular nature of the tumor can be easily visualized (Fig. 39-7).51

Leiomyomatosis with Intracardiac Extension Intravenous leiomyomatosis is a rare pathologic entity, and all tumors have been observed in women, most of whom were white and premenopausal.52 The tumor arises either from a uterine myoma or from the vessel wall.53 The tumor generally appears as a large, mobile mass in the right atrium (Fig. 39-8). Because the preoperative evaluation should include assessment of all cardiac chambers and the region of the inferior vena cava, CMR can be considered a primary diagnostic method.54 Because of the myomatous tissue, the signal intensity characteristics are similar to those of muscle.

Other Benign Tumors A number of very unusual conditions that affect the heart cause tumors. These may be infective, such as echinococcus or hydatid cysts (Fig. 39-9), or inflammatory (Behc¸et syndrome mimicking myxoma).

C

E

B

D

F

39 CARDIAC AND PARACARDIAC MASSES

A

Figure 39-5 Fibroma of the lateral wall of the left ventricle. Steady-state free precession four-chamber (A) and short axis (B) images showing a slightly hypointense mass in the lateral wall of the left ventricle (arrows). First-pass perfusion image in the short-axis plane (C) showing hypoperfusion of the mass surrounded by the normally perfused myocardium, which is suggestive of low tumor vascularity (arrow). Delayed myocardial enhancement image (D) showing high signal intensity of the mass compared with the nulled normal myocardium (arrow). Gross pathology of the resected left ventricular mass (E). The cut section (F) shows an off-white whorled surface with foci of calcification. The lesion extended to the inked surgical margin. There was no hemorrhage or necrosis.

MALIGNANT TUMORS OF THE HEART Nearly 25% of all primary cardiac tumors are malignant. Metastases are 20 to 40 times more common than primary malignant tumors and are found at autopsy in 6% of cases of malignant diseases. The most common primary malignant cardiac tumors are various sarcomas and lymphomas.55,56 Because of the small number of cardiac malignancies studied, the differences in tumor age and vascularity, and widespread variability in water content, reliable tissue differentiation of cardiac malignancies is still not possible.57,58 Distinct features of malignant tumors are necrosis, calcification, a high degree of vascularity, infiltration of adjacent tissues, inhomogeneous appearance, and peritumorous edema.59,60

Sarcoma Sarcomas of the heart or great vessels are rare and account for the largest group of malignant primary cardiac tumors in adults. There are various histologic types of sarcoma, such as angiosarcoma, leiomyosarcoma, and liposarcoma.15,56,61 Angiosarcoma (Fig. 39-10) is the most common primary malignant cardiac tumor.15 It is usually located within the right atrium and arises from the

interatrial septum. They appear preferentially between the third and fifth decades of life, with a male-to-female ratio of approximately 2:1.62 In contrast, other types of sarcoma also occur in the left side of the heart, where they are often clinically mistaken for myxoma. Morphologically, angiosarcomas are usually hemorrhagic, often with poorly defined borders. They often invade contiguous structures, such as the venae cavae and tricuspid valve, and are characterized by irregular anastomosing sinusoidal structures with papillary intraluminal tufting.62 On CMR, angiosarcomas have a polymorphic appearance, with a central region of hyperintensity, consistent with necrosis, and moderate signal intensity in peripheral regions on T1- and T2-weighted images.26 Because of the high degree of vascularity, signal enhancement is seen after intravenous injection of gadolinium. Primary leiomyosarcomas are rare, highly aggressive, locally invasive tumors, with a frequency of 0.25%. In 75% of cases, they arise from the inferior vena cava, but also have been reported to originate from the superior vena cava or pulmonary vein.61,63 This neoplasm shows slightly higher signal intensity on T1-weighted spin echo images than in the liver parenchyma, but not as bright as in the adjacent mediastinal fat. Signal intensity is high on T2weighted images. After gadolinium injection, slight contrast enhancement of the tumor can be detected.61 The advantage of CMR is the ability to assess tumor extension into the superior vena cava, pulmonary veins, and heart chambers.64 Cardiovascular Magnetic Resonance 537

FUNCTIONAL CARDIOVASCULAR DISEASE

Figure 39-6 Rhabdomyoma of the interventricular septum. Short axis image from a steady-state free precession image obtained through the proximal septum (A) showing a markedly thickened septum (arrow) that appears to be homogeneous with the rest of the myocardium. Steady-state free precession image of the left ventricular outflow tract (B, arrow). T1-weighted turbo spin echo image (C) showing that the signal from the markedly thickened septum is isointense compared with the rest of the myocardium. T2-weighted image with fat suppression (D) showing hyperintense signal within the thickened septum, suggesting a different structure from the surrounding myocardium. Early after gadolinium injection (E), there is enhancement of the cavities of the left ventricle (LV) and right ventricle (RV), as well as the myocardium, including the thickened septum. Late after gadolinium injection (F), the normal myocardium appears black. There is marked abnormal late enhancement throughout the entire segment of the thickened septum.

LV RV

LV

A

B

C

D

E

F

F

A

B

E

C

D

Figure 39-7 Hemangioma of the right ventricle. Four-chamber steady-state free precession image (A) showing a hypointense, spherical, and pedunculated mass within the mid-right ventricle that is attached to the free wall of the right ventricle (RV) via a short stalk (arrow). Short axis perfusion imaging shows first-pass contrast arrival in the RV (B) and peripheral contrast enhancement of the mass (C, arrow), with subsequent complete filling of the tumor on the second-pass arrival of the contrast in the RV (D, arrow). These findings indicate vascularity of the lesion. Short axis view of 10-minute late gadolinium enhancement (E) shows peripheral enhancement of the RV mass (arrow). This finding favors the diagnosis of cardiac hemangioma. Gross photograph of the resected tumor specimen (F) shows a polypoid, tan, gelatinous, and focally hemorrhagic lesion with a small stalk (arrow). High-power photomicrograph (G) shows marked proliferation of capillary-sized blood vessels (black arrows), within a slightly blue acid mucopolysaccharide ground substance. These features are consistent with capillary-type cardiac hemangioma. LV, left ventricle. 538 Cardiovascular Magnetic Resonance

39 CARDIAC AND PARACARDIAC MASSES

Figure 39-8 Intravenous leiomyomatosis. Atrial short axis (A), right ventricular two-chamber (B), and right ventricular outflow tract (C) steady-state free precession images showing transvenous extension of tumor from the inferior vena cava to the right atrium and right ventricle (arrows). The mass enhances late after gadolinium injection (D, arrow).

A

B

C

D

A

B

C

D

Figure 39-9 Two-chamber (A) and short axis (B) steady-state free precession images showing a large multilocular cystic mass (arrows) involving the inferior wall and producing severe thinning of the adjacent myocardium. Gross image (C) of the excised hydatoid material. Microscopic image of Echinococcus granulosus (D).

Cardiovascular Magnetic Resonance 539

FUNCTIONAL CARDIOVASCULAR DISEASE

A

E

B

F

C

G

D

H

Figure 39-10 Comprehensive evaluation of a right atrial mass. Four-chamber (A) and atrial short axis (B and C) steady-state free precession images showing a large mass in the right atrial free wall invading the right atrial cavity (arrows). The tumor has intermediate signal on T1 (D) and high signal on T2-weighted images (E), which persists after fat suppression (F). The presence of a large, heterogeneous, and invasive mass in the right side of the heart associated with pericardial effusion suggests malignancy. There is significant enhancement early after gadolinium injection, suggesting vascularity (G). The enhancement pattern persists in the late phase (H). Biopsy was performed, and the diagnosis of angiosarcoma was made.

Liposarcomas (Fig. 39-11) are very rare and are not represented in most surgical series of tumors. Grossly, they are bulky tumors as large as 10 cm.62 Liposarcomas often have a pericardial origin, and CMR can detect this pericardial mass with heterogeneous high signal intensity and epicardial infiltration. After injection of gadolinium, liposarcomas may show only slight signal enhancement.65

Lymphoma Primary cardiac lymphomas are rare, and approximately 56 cases have been reported (Fig. 39-12). Postmortem studies have shown cardiac involvement in up to 25% of cases of lymphoma; however, in vivo diagnosis is still rare.59 The incidence of cardiac lymphoma is increasing because of lymphoproliferative disorders related to Epstein-Barr virus in patients with acquired immune deficiency syndrome and in those who have undergone transplantation.62 The mean age at presentation for cardiac lymphoma is 38 years, with a slight predominance in men. On T1- and T2weighted spin echo images, lymphomas appear isointense or hypointense to cardiac muscle. After injection of gadolinium, lymphomas appear heterogeneous, with lessenhancing central regions consisting of necrosis.58

Metastatic Cardiac Tumors The incidence of cardiac metastases has increased to slightly more than 10%. Primary tumors that metastasize to the heart can be divided into three categories of 540 Cardiovascular Magnetic Resonance

incidence: uncommon primary tumors that have a high rate of metastasis to the heart (malignant melanoma, malignant germ cell neoplasm, malignant thymoma); common tumors that have an intermediate rate of cardiac metastasis, but account for the greatest number of cardiac metastases (carcinoma of the stomach, liver, ovary, colon, and rectum); and common tumors with rare metastases to the heart.62 There are four ways in which noncardiac tumors may invade the heart: (1) direct mediastinal infiltration of heart tissue in the case of lung cancer (see Fig. 39-9), breast cancer, or mediastinal lymphoma; (2) hematogenous spread by systemic tumors, such as malignant melanoma, lymphoma, leukemia, or sarcoma; (3) transvenous spreading from the inferior vena cava in the case of renal or hepatic tumors and from the pulmonary vein in the case of lung cancer; and (4) lymphatic spread.15,31 Metastatic cardiac involvement can be localized or diffuse. The nodules can arise at discrete locations in the myocardium or encase the epicardial surface diffusely.

INTRACARDIAC THROMBUS FORMATION Intracardiac thrombus formation often occurs in the left atrium in the case of chronic atrial fibrillation and dilation, or in the left ventricle in the case of wall motion abnormalities seen in cardiomyopathy (Fig. 39-13) or after myocardial infarction.66 Tumor thrombus in the inferior vena cava and right atrium can be seen in renal carcinoma and can mimic a solid cardiac mass.67,68 The diagnosis of

C

39 CARDIAC AND PARACARDIAC MASSES

A

B

D

E

Figure 39-11 Left ventricular liposarcoma. Left ventricular outflow tract (A) and short axis at the papillary muscle level (B) steady-state free precession images showing abnormal tissue enlarging the posteromedial papillary muscle. Some abnormal tissue can also be seen in the mitral valve leaflets (black arrows). T1-weighted turbo spin echo image at the papillary level (C) shows high signal not only in the posteromedial papillary muscle, but also in part of the anterolateral papillary muscle (white arrow). There is signal drop-out on the same plane after fat suppression (D), suggesting fatty infiltration. Increased signal intensity is seen in the late phase after gadolinium injection (E). Surgical resection of the tumoral masses was performed, and the diagnosis of low-grade liposarcoma was made on histologic examination.

Figure 39-12 Involvement of the heart by an aggressive large B-cell lymphoma (arrows). Horizontal long axis steady-state free precession image showing a homogeneous mass involving the lateral aspect of the right ventricle and extending superiorly to the right atrium.

cardiac thrombus is clinically important because patients are at risk for systemic or pulmonary embolization.69,70 However, in most cases, the diagnosis of cardiac thrombus is coincidental and patients are asymptomatic. Although two-dimensional echocardiography is the diagnostic method of choice, false-positive rates of as high as 28% in the detection of left ventricular thrombi and 59% in the detection of left atrial thrombi have been reported.71,72 On CMR images, fresh thrombi on T1-weighted spin echo images often have a higher signal intensity than myocardium, and the contrast is further accentuated on T2weighted spin echo images, consistent with a high amount of hemoglobin.66 However, depending on the age of the thrombus, other signal intensities are possible. After 1 or 2 weeks, paramagnetic compounds in the organizing thrombus, such as deoxyhemoglobin and methemoglobin, cause T1 and T2 shortening, which may result in increased signal intensity on T1-weighted images and decreased signal intensity on T2-weighted images.73 Chronic organized thrombi are of low signal intensity because of loss of water and protons. A problem with differentiation between thrombus and slow-flowing blood occurs, especially in laminated or immobile thrombi on spin echo images.74 Cardiovascular Magnetic Resonance 541

FUNCTIONAL CARDIOVASCULAR DISEASE

Figure 39-13 Four-chamber (A) and two-chamber (B) steady-state free precession images showing two relatively hypointense regular masses adjacent to the apical and anterior walls (arrows). Spoiled gradient echo images with a long inversion time (440 msec) are shown shortly after the administration of gadolinium (C and D). Note the intense contrast between the strongly enhanced blood pool (bright), the myocardium, and the abnormal structures in the left ventricular cavity that return no signal (dark). These findings are consistent with intraventricular thrombi.

A

B

C

D

Compared with thrombus formation, slow-flowing blood shows increased signal intensity on T2-weighted images.43,75 On GRE images, thrombi have the lowest signal intensity compared with other cardiac structures, whereas blood appears brightest.73 If thrombi contain calcification, they appear more heterogeneous.27 To differentiate thrombus from tumor, intravenous injection of contrast agents may be helpful. Thrombi usually do not show signal enhancement after intravenous injection of gadolinium, unless they are already organized.9 Compared with other diagnostic procedures, CMR and CT offer similar sensitivity of approximately 90%, with slightly better specificity compared with two-dimensional echocardiography.74

PERICARDIAL LESIONS Magnetic resonance evaluation of pericardial neoplasms in most cases involves identification of abnormal anatomic structures and boundaries rather than characterization of relative tissue intensities. A few exceptions are included in the differential diagnosis of mediastinal masses, such as fibroma, lipoma, and pericardial cysts. The value of CMR for evaluation of potential neoplasm lies largely in treatment planning, particularly preoperative assessment. The loss of normal anatomic boundaries is an important sign of neoplasm. Neoplastic involvement of the pericardium 542 Cardiovascular Magnetic Resonance

results in focal and diffuse obliteration of the normal pericardial signal. In the case of malignancy adjacent to cardiac structures, visualization of the pericardial line is an indication that pericardial invasion has not occurred.

Tumors The diagnosis of pericardial tumors includes benign lipomas or pericardial cysts and some rare primary malignancies, such as teratomas, mesotheliomas, bronchial pheochromocytomas, lymphomas, fibrosarcomas, and angiosarcomas.76,77 Hemorrhagic effusions result from erosion into intrapericardiac vessels or the myocardial wall, with possible acute or subacute tamponade. Lipoma and the pericardial fat pad are easily distinguished by the high signal of fat tissue.78

Pericardial Metastasis The rate of involvement of the pericardium by metastatic disease at autopsy is much higher than clinically suspected, with an incidence ranging from 1.5% to 22.0%.79 Of malignant pericardial disease, 80% of cases are associated with lung or breast carcinoma, leukemia, and lymphoma.80 Metastatic involvement of the pericardium is characterized

Cysts Pericardial cysts (Fig. 39-14) are rare lesions and are commonly located in the right pericardiophrenic angle.81,82 Pericardial cysts are usually filled with clear fluid. Patients are generally asymptomatic, and the lesion is often discovered on a routine chest film. The appearance is typically stable over a long period. In most cases, no cardiac surgery is necessary.83 On CMR, pericardial cysts appear as paracardiac masses with long T1 and T2 values and flow void, indicating fluid-filled structures. They have low signal intensity on T1-weighted images and increased signal intensity on T2-weighted images.84 After the injection of gadolinium, intracystic septae may be seen. In addition, a line of low signal intensity, representing the pericardial layer, can often be visualized. The significant advantage of CMR is its ability to differentiate these lesions from other mediastinal masses and avoid explorative surgery to determine the diagnosis.

A

B

C

D

TISSUE CHARACTERIZATION Unlike two-dimensional echocardiography, CMR has the potential for tissue characterization by comparing relaxation properties of the mass with a reference tissue (Table 39-2).85 That is based on the observation that significant differences exist in proton density and T1 and T2 relaxation time. Because fat shows very constant T1 and T2 values, it has been used as a reference tissue.86 Previous studies have tried to determine specific T1 and T2 relaxation times of different tissues; however, precise etiologic diagnoses are not possible.87–89 Inversion recovery spin echo sequences are recommended to quantify tissue characteristics because they have the advantage of providing more accurate T1 data (4% dispersion) compared with conventional spin echo sequences (20% dispersion).90 Inhomogeneity, tumor infiltration of adjacent structures, and hemorrhagic peritumorous pericardial effusion suggest malignancy, whereas uniform tissue signal usually indicates benign masses.20 Only lipomas, fibromas, pericardial cysts, and angiomas can be identified by their signal intensity and behavior. Morphologic features seen on CMR are independent predictors of cardiac and paracardiac lesions. Univariate and multivariate analysis of CMR of 55 cardiac and paracardiac lesions identified morphologic features, such as right-sided cardiac location, inhomogeneity of tumor tissue, and presence of effusion, as predictors for malignant cardiac masses.91

E

Figure 39-14 Multimodality imaging in the evaluation of a mediastinal mass. Four-chamber (A) and coronal (B) steady-state free precession images showing a large homogeneous and well-defined mass in the right cardiophrenic angle (arrows). High signal structures on steady-state free precession images represent images either fat or fluid. The mass returns relatively low signal on T1 (C) and relatively high signal on T2-weighted (D) images, which is more evident after fat suppression (E). Therefore, the diagnosis of a fluid-filled pericardial cyst was made. Cardiovascular Magnetic Resonance 543

39 CARDIAC AND PARACARDIAC MASSES

by large effusions out of proportion to the amount of tumor present. This occurs to the extent that neoplasm is the most frequent cause of tamponade. Focal or diffuse plaque-like thickening of the pericardium may occur with signal enhancement after injection of gadolinium.

FUNCTIONAL CARDIOVASCULAR DISEASE

Table 39-2 MRI Features of the Most Common Cardiac Tumors (SI-signal intensity)

Myxoma

Lipoma/lipomatous hypertrophy Papillary fibroelastoma Fibroma

Rhabdomyoma Hemangioma

Intravenous leiomyomatosis Pericardial cysts (simple fluid) (proteinaceous fluid) Angiosarcoma

Lymphoma Liposarcoma

Leiomyocarcoma

Thrombus

Fresh Chronic (older than 2 weeks)

Enhancement after administration of Gd-DTPA

T1-weighted SE

T2-weighted SE

GRE

Intermediate varying SI, calcified areas: hypointense and hemorrhage: increased SI Brightest SI similar to subcutaneous fat, using fat presaturation technique: reduced SI Intermediate SI

Low SI, especially in iron-containing myxomas

Very low SI compared to the surrounding blood pool

Hyperintense

Intermediate SI parallel to subcutaneous fat

Nonspecific

Nonspecific

Intermediate to high SI

Low SI compared to the surrounding blood pool Nonspecific

Hyperintense

Very low compared to myocardium Nonspecific

Nonspecific

Intermediate to slightly hyperintense SI compared to myocardium, when calcification (hypointense) and hemorrhage (hyperintense) are present: heterogeneous. Homogeneous, slightly lower SI than myocardium Intermediate SI

Similar to myocardium

Decrease in SI compared to TI

Strong increased SI Increased SI (due to slow flowing blood in the tumor vessels), higher than myocardium Similar to myocardium

Slight and heterogeneous

Significant increasing SI, heterogeneous.

Nonspecific

Nonspecific

Lowest SI, flow void Low SI, but higher than in normal fluid, no flow void

Highest SI High SI

Nonspecific

Signal enhancement, visualization of intracystic septae

Central hyperintense spot (blood vessels, hemorrhage, or necrosis), surrounded by intermediate SI regions. Isointense to hypointense to cardiac muscle

No change

Nonspecific

Strong

Isointense to myocardium

Nonspecific

Bright SI equal to subcutaneous fat, but heterogeneous: decrease in SI, when fat presaturation is used High SI, slightly higher than liver parenchyma, but not as high as subcutaneous fat homogeneous, commonly connected with proteinaceous pericardial effusion Intermediate, often slightly higher SI than myocardium, slightly higher than blood

Not published

Not published

Heterogeneous with less enhancing central regions Not published

Not published

Not published

Not published

Surrounding slowly flowing blood becomes higher SI than thrombus, contrast between thrombus and myocardium is further accentuated Decreased SI

Thrombus has the lowest SI

No signal enhancement, unless the thrombus is organized.

High SI (oxyhemoglobin) Higher SI (desoxyhemoglobin)

Modified from Hoffman U, Globits S, Frank H: Cardiac and paracardiac masses: current opinion on diagnostic evaluation by magnetic resonance imaging. Eur Heart J. 1998;19:553–563. Gd, gadolinium; GRE, gradient recalled echo; SE, spin-echo; SI, signal intensity.

544 Cardiovascular Magnetic Resonance

Following a step-by-step procedure is suggested (see Table 39-2). (1) Exclusion of the most common benign cardiac tumors by their typical and often pathognomonic appearance. Typical features can be observed in myxoma (septal atrial origin, hypointense, moderate contrast enhancement, no effusion or infiltration); fibromas (left-sided, homogeneous, hypointense, intense contrast enhancement, lipomas (hyperintense on T1-weighted spin echo images, hypointense on fat-suppressed images); and thrombi (typically, enlarged left atria in atrial fibrillation or left ventricular aneurysms, hypointense on GRE, no contrast enhancement). (2) Features specific to malignant tumors, such as an intracardial lesion with pericardial effusion, a paracardial lesion with inhomogeneous tissue composition, or a mass that infiltrates adjacent tissue but is not an inflammatory pseudotumor or a fibroma. (3) With the exception of myxoma, lesions of the right side of the heart suggest malignancy or metastasis.91

PROGNOSIS OF CARDIAC TUMORS Prognosis after surgery is usually excellent in the case of benign tumors. Local recurrence of myxoma is uncommon, but could be related to inadequate resection, multicentricity,

origin in a chamber other than the left atrium, or familial tumor. Nearly all malignant cardiac tumors are rapidly fatal. Sarcomas of the heart have a better prognosis if they arise in the left atrium, have no necrotic regions, and have not metastasized at the time of diagnosis.62 Achievement of remission is unusual and can be achieved with chemotherapy in lymphoma.92

CONCLUSION Cardiovascular magnetic resonance techniques have contributed significantly to the ability to detect cardiac and paracardiac masses and play an important role in diagnostic evaluation. CMR is complementary to echocardiography. Magnetic resonance, because of its larger field of view, adds diagnostic information by assessing the extracardiac components of a mass, such as mediastinal involvement and extension into large pulmonary vessels. CMR allows the exclusion of hiatus hernia, tortuous descending aorta, or a bronchogenic cyst that can mimic cardiac tumors. CMR findings are helpful in characterizing paracardiac masses and guiding therapeutic strategies. Tissue characterization by CMR is limited to the diagnosis of myxomas, fibromas, thrombi, pericardial cysts, and fatty tissue. The signal features of malignant tumors are equivocal and do not permit a tissue diagnosis. However, the inhomogeneous appearance of signal enhancement after injection of gadolinium, the infiltrative components of a tumor, and a hemorrhagic pericardial effusion make the diagnosis of malignancy more likely.

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11. Benjamin HG. Primary fibromyxoma of the heart. Arch Pathol. 1939; 27:950. 12. Heath D. Pathology of cardiac tumors. Am J Cardiol. 1968; 21:315–327. 13. Straus R, Merliiss R. Primary tumor of the heart. Arch Pathol. 1945;39:74–78. 14. Wold LE, Lie JT. Cardiac myxomas: a clinicopathologic profile. Am J Pathol. 1980;101:219–240. 15. McAllister Jr HA, Fenoglio Jr JJ. Tumors of the Cardiovascular system. Atlas of tumor pathology. 2nd series. Fascicle 15. Washington, DC: Armed Forces Institute of Pathology; 1978:1–20. 16. Reynen K. Cardiac myxomas: review article. N Engl J Med. 1995;333 (24):1610–1617. 17. Bicer A, Turhan H, Cagrici G, Yasar AS, Sasmaz H. Asymptomatic left ventricular myxoma diagnosed incidentally by transthoracic echocardiography. Echocardiography. 2005;22:855–856. 18. Seebacher G, Binder T, Frank H, Wolner E, Mohl W. Cystic formation of the foramen ovale mimicking a right atrial myxoma. Ann Thorac Surg. 2006;82:2296–2298. 19. Pritchard RW. Tumors of the heart: review of the subject and report of one hundred and fifty cases. Arch Pathol. 1951;51:98–128. 20. Hall RJ, Cooley DA, MacAllister Jr HR, Frazier OH. Neoplastic heart disease. In: Hurst JW, ed. The Heart, Arteries and Veins. 7th ed. New York: McGraw-Hill; 1990:1382–1403. 21. St John Sutton MG, Mercier L-A, Giuliani ER, Lie JT. Atrial myxomas: a review of clinical experience in 40 cases. Mayo Clin Proc. 1980;55:371–376. 22. Peters MN, Hall RJ, Cooley DA, Leachmann RD, Garcia E. The clinical syndrome of atrial myxoma. JAMA. 1974;230:695–701. 23. Lund JT, Ehman RL, Julsrud PR, Sinak LJ, Tajik AJ. Cardiac masses: assessment by MR. Am J Radiol. 1989;152:469–473. Cardiovascular Magnetic Resonance 545

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24. Go RT, O’Donnell JK, Underwood DA, et al. Comparison of gated cardiac MR and 2D echo of intracardiac neoplasms. Am J Radiol. 1985;145:21–25. 25. Semelka RC, Shoenut JP, Wilson ME, Pellech AE, Patton JN. Cardiac masses. signal intensity features on spin-echo, gradient-echo, gadolinium-enhanced-spin-echo, and TurboFLASH images. J Magn Reson Imaging. 1992;2:415–420. 26. Gomes AS, Lois JF, Child JS, Brown K, Batra P. Cardiac tumors and thrombus: evaluation with MR imaging. Am J Radiol. 1987;149: 895–899. 27. Roos A, Weijers E, Duinen S, Wall EE. Calcified right atrial myxoma demonstrated by MR. Chest. 1989;95:478–479. 28. Reyes CV, Jablokow VR. Lipomatous hypertrophy of the cardiac interatrial septum: a report of 38 cases and review of the literature. Am J Clin Pathol. 1979;5:785–788. 29. Crocker DW. Lipomatous infiltrates of the heart. Arch Pathol Lab Med. 1978;102:69–72. 30. Fine G. Neoplasms of the pericardium and heart. In: Gould SE, ed. Pathology of the Heart and Blood Vessels. 3rd ed. Springfield: Charles C. Thomas; 1968:865. 31. Moulton AL, Jaretzki III A, Bowman OF, Silverstein EF, Bregman D. Massive lipoma of the heart. NY State J Med. 1976;76:1820–1825. 32. Hananouchi GI, Goff WB. Cardiac lipoma: six-year follow-up with MR characteristics, and a review of the literature. Magn Reson Imaging. 1990;8:825–828. 33. Conces DJ, Vix VA, Tarver RD. Diagnosis of a myocardial lipoma by using CT. Am J Roentgenol. 1989;153:725–726. 34. Hutter Jr AM, Page DL. Atrial arrhythmias and lipomatous hypertrophy of the cardiac interatrial septum. Am Heart J. 1971;82:16–21. 35. Kluge WF. Lipomatous hypertrophy of the interatrial septum. Northwest Med. 1969;68:25–30. 36. Levine RA, Weyman AE, Dinsmore RE, et al. Noninvasive tissue characterisation: diagnosis of lipomatous hypertrophy of the atrial septum by nuclear MR. J Am Coll Card. 1986;7:688–692. 37. Applegate PM, Tajik AJ, Ehman RL, Julsrud PR, Miller FA. Twodimensional echocardiographic and MR observations in massive lipomatous hypertrophy of the atrial septum. Am J Card. 1987; 59:489–491. 38. Ngaage DL, Mullany CJ, Daly RC, et al. Surgical treatment of cardiac papillary fibroelastoma: a single center experience with eighty-eight patients. Ann Thorac Surg. 2007;80:1712–1718. 39. de Arenaza DP, Pietrani M, Moon JC, et al. Cardiac fibroelastoma: cardiovascular magnetic resonance characteristics. J Cardiovasc Magn Reson. 2007;9:621. 40. Jahnke C, Hamdan A, Fleck E, Paetsch I. Tissue characterization of a suspected aortic valve fibroelastoma with cardiac magnetic resonance imaging. Circ Cardiovasc Imaging. 2008;1:87–88. 41. Burke AP, Rosado CM, Templeton PA, Virmani R. Cardiac fibroma: clinicopathologic correlates and surgical treatment. J Thorac Cardiovasc Surg. 1994;108:862–870. 42. Oliva PB, Breckinridge JC, Johnson ML, Brantigan CO, O’Meara OP. Left ventricular outflow obstruction produced by a pedunculated fibroma in a newborn: clinical, angiographic, echocardiographic and surgical observations. Chest. 1978;74:590–593. 43. Winkler M, Higgins CB. Suspected intracardiac masses: evaluation with MR imaging. Radiology. 1987;165:117–122. 44. Gamsu G, Stark DD, Webb WR, Moore EH, Sheldon PE. MR of benign mediastinal masses. Radiology. 1984;151:709–713. 45. Meier RA, Hartnell GG. MR of right atrial pseudomass: is it really a diagnostic problem? J Comput Assist Tomogr. 1994;18:398–401. 46. Edwards JE. Congenital malformations of the heart and great vessels. In: Gould SE, ed. Pathology of the Heart. Springfield, IL: Bannerstone House; 1960:61–63. 47. Abushaban L, Denham B, Duff D. 10 year review of cardiac tumors in childhood. Br Heart J. 1993;70:166–169. 48. Hoffmann U, Globits S, Frank H. Cardiac and paracardiac masses. Eur Heart J. 1998;19:553–563. 49. Brizard C, Latremouille C, Jebara VA, et al. Cardiac hemangiomas. Ann Thorac Surg. 1993;56:390–394. 50. Newell II JD, Eckel C, Davis M, Tadros NB. MR Appearance of an arteriovenous hemangioma of the interventricular septum. Cardiovasc Intervent Radiol. 1988;11:319–321. 51. Soberman MS, Plauth WH, Winn KJ, et al. Hemangioma of the right ventricle causing outflow obstruction. J Thorac Cardiovasc Surg. 1988;96:307–309.

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52. Spellacy WN, Maire WJ, Buhi WC. Plasma growth hormone and estradiol levels in women with uterine myomas. Obstet Gynecol. 1972;40:829–834. 53. Bassish MS. Mesenchymal tumors of the uterus. Clin Obstet Gynecol. 1974;17:51–88. 54. Rosenberg JM, Marvasti MA, Obeid A, Johnson LW, Bonaventura M. Intravenous leiomyomatosis: a rare cause of right sided cardiac obstruction. Eur J Cardio-thorac Surg. 1988;2:58–60. 55. Roberts WC, Glancy DL, DeVita VT. Heart in malignant lymphoma: a study of 196 autopsy cases. Am J Cardiol. 1968;22:85–107. 56. Silverman J, Olwin JS, Graettinger JS. Cardiac myxomas with systemic embolization: review of the literature and report of a case. Circulation. 1962;26:99–103. 57. Zeitler .MR of aneurysms and thrombi. Cardiovasc Intervent Radiol. 1986;8:321–328. 58. Dorsay TA, Ho VB, Rovira MJ, Armstrong MA, Brisette MD. Primary cardiac lymphoma: CT and MR findings. J Comp Assist Tomogr. 1993;17(6):978–981. 59. Hendrick RE, Raff U. Image contrast and noise. In: Stark DD, Bradley WB, eds. MR. St. Louis: Mosby-Year Book; 1992:109–144. 60. Tazelaar HD, Locke TJ, McGregor CGA. Pathology of surgically excised primary cardiac tumors. Mayo Clin Proc. 1992;67:957–965. 61. Hoffstetter P, Djavidani B, Feuerbach S, Hofsta¨dter F, Seitz J. Myxoid fibrosarcoma of a pulmonary vein with extension into the left atrium. Am J Roentgenol. 2006;186:365–367. 62. Butany J, Nair V, Naseemuddin A, Nair GM, Catton C, Yau T. Cardiac tumours: diagnosis and management. Lancet Oncol. 2005;6:219–228. 63. Lupetin AR, Dash N, Beckman I. Leiomyosarcoma of the superior vena cava: diagnosis by cardiac gated MR. Cardiovasc Intervent Radiol. 1986;9:103–105. 64. Somers K, Lotte F. Primary lymphosarcoma of the heart: review of the literature and report of 3 cases. Cancer. 1960;3:449–457. 65. Garrigue S, Robert F, Roudaut R, Bonnet J. Assessment of non-invasive imaging techniques in the diagnosis of heart liposarcoma. Eur Heart J. 1995;16:139–141. 66. Dooms GC, Higgins CB. MR Imaging of cardiac thrombi. J Comp Assist Tomogr. 1986;10:415–420. 67. Asensio JMN, Fernandez MRG, Bravo MA, Herrera AM. Kidney tumour mimicking cardiac mass. Intern. J of Cardiol. 2006;106: 401–403. 68. Gomes AS, Lois JF, Child JS, Brown K, Batra P. Cardiac tumors and thrombus: evaluation with MR imaging. Am J Roentgenol. 1987;149:895–899. 69. Visser CA, Kan G, Meltzer RS, et al. Embolic potential of left ventricular thrombus after myocardial infarction: a two-dimensional echocardiographic study of 119 patients. JACC. 1985;5:1276–1280. 70. Hamby RI, Wisoff BG, Davison ET, et al. Coronary artery disease and left ventricular mural thrombi: clinical, hemodynamic, and angiocardiographic aspects. Chest. 1974;66:488–494. 71. Barakos JA, Brown JJ, Higgins CB. MR imaging of secondary cardiac and paracardiac lesions. Am J Roentgenol. 1989;153:47–50. 72. Feigenbaum H. Coronary artery disease. In: Feigenbaum H, ed. Echocardiography. 4th ed. Philadelphia: Lea and Febiger; 1986:489–494. 73. Jungehu¨lsing M, Sechtem U, Theissen P, Hilger HH, Schicha H. Left ventricular thrombi: evaluation with spin-echo and gradient-echo MR imaging. Radiology. 1992;182:225–229. 74. Sechtem U, Theissen P, Heindel W, et al. Diagnosis of left ventricular thrombi by MR and comparison with angiocardiography, computed tomography and echocardiography. Am J Cardiol. 1989;64: 1195–1199. 75. Menegus MA, Greenberg MA, Spindola-Franco H, Fayemi A. MR of suspected atrial tumors. Am Heart J. 1992;123:1260–1268. 76. Stark DD, Higgins CB, Lanzer P, et al. Magnetic resonance imaging of the pericardium: normal and pathologic findings. Radiology. 1984;150: 469–474. 77. Hort W, Braun H. Untersuchungen ueber groesse, wandsta¨rke und mikroskopischen aufbau des herzbeutels unter normalen und pathologischen bedingungen. Arch Kreislaufforsch. 1962;38:1–22. 78. Amparo EG, Higgins CB, Farmer D, et al. Gated MR of cardiac and paracardiac masses: initial experience. Am J Roentgenol. 1984;143: 1151. 79. Mocanda R, Kotler MN, Churchill RJ, et al. Multimodality approach to pericardial imaging. Cardiovasc Clin. 1986;17:409. 80. Theoligides A. Neoplastic cardiac tamponade. Semin Oncol. 1978;5:181. 81. Fraser RG, Pare JAP. Diagnosis of Diseases of the Chest. 2nd ed. Philadelphia, PA: WB Saunders Co; 1977:656. 82. Le Roux BT. Pericardial coelomic cysts. Thorax. 1977;14:27–35.

88. Tscholakoff D. MRT in the diagnosis of cardiovascular system and lung. Acta Med Austriaca. 1986;3:61–65. 89. Schmidt HC, Tscholakoff D, Hricak H, Higgins CB. MR image contrast and relaxation times of solid tumors in the chest, abdomen, and pelvis. J Comput Assist Tomogr. 1985;9:738–748. 90. Walker PM, Marie PY, Danchin N, Bertrand A. Comparison of T1 estimation techniques in cardiac MR. Magn Reson Imag. 1994;12:43–50. 91. Hoffmann U, Globits S, Schima W, et al. Usefulness of magnetic resonance imaging of cardiac and paracardiac masses. Am J Cardiol. 2003;92:890–895. 92. Morillas P, Quiles J, Nunez D, et al. Complete regression of cardiac non-Hodgkin lymphoma. Intern J Cardiol. 2006;106:426–427.

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83. Roberts WC, Spray TL. Pericardial heart disease: a study of its causes, consequences and morphologic features. Cardiovasc Clin. 1976;7: 11–65. 84. Sechtem U, Tscholakoff D, Higgins CB. MR of the abnormal pericardium. Am J Radiol. 1986;147:245–252. 85. Amparo EG, Higgins CB, Farmer D, et al. Gated MR of cardiac and paracardiac masses: initial experience. Am Heart J. 1984;143: 1151–1156. 86. Dooms GC, Hricak H, Margulis AR, Geer G. MR imaging of fat. Radiology. 1986;158:51–54. 87. Schulthess GK, McMurdo K, Tscholakoff D, de Geer G, Gamsu G, Higgins CB. Mediastinal masses: MR imaging. Radiology. 1986;158: 289–296.

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CHAPTER 40

Cardiac Transplantation Nicholas R. Teman, Hee-Won Kim, and Gerald M. Pohost

Cardiac transplantation is the most effective modality for the treatment of end-stage heart failure. Most patients experience tremendous improvement in quality of life following transplantation. Even with the development of new immunosuppressive drugs, however, rejection still remains among the most important causes of morbidity and mortality in patients who have undergone cardiac transplantation. Monitoring these patients requires clinical expertise and a high index of suspicion for acute rejection. The gold standard for diagnosing allograft rejection is endomyocardial biopsy. While this is currently the most effective option, there are many disadvantages associated with this procedure. It is expensive and inconvenient for the patient, requiring a catheterization laboratory. In addition, multiple myocardial samples are necessary to ensure detection of early rejection, which is frequently focal. Moreover, the interpretation of the histopathology of the tissue obtained with the biopsy tends to be subjective, leading to possible errors in diagnosis and discrepancy between the biopsy grade and the actual degree of rejection. Finally, there is an inherent morbidity associated with myocardial biopsy.1 Such morbidity is even more relevant given the increase in cardiac transplantation in young patients. Clearly, a noninvasive approach to monitor rejection of the cardiac allograft would be preferred. Unfortunately, there are no methods that are currently capable of detecting rejection with the sensitivity and specificity needed to avoid biopsy. Measurement of global cardiac function using techniques such as echocardiography and radionuclide methods provides relatively late indicators of rejection. The noninvasive technology that has the most potential to reduce the need for endomyocardial biopsy is cardiovascular magnetic resonance (CMR). While not currently used in the clinical setting, CMR has shown great potential. This section will assess the current state of CMR methods in the noninvasive assessment of cardiac allograft rejection, including angiography (MRA), late gadolinium enhancement (LGE), and cardiovascular magnetic resonance spectroscopy (CMRS), and evaluate the role that CMR may play in the evaluation of the cardiac allograft. The majority of the studies have focused on CMR and CMRS, with minimal discussion of the utility of MRA.

IMAGING EXPERIMENTAL ANIMAL MODELS Early studies with CMR in cardiac transplantation focused on relaxation times and gadolinium enhancement. Aherne and colleagues performed one of the earliest experiments 548 Cardiovascular Magnetic Resonance

using CMR, looking at T1 and T2 values in 15 dogs.2 They found that T2 increased in the dogs that experienced myocardial allograft rejection, and this increase corresponded with the degree of rejection as determined by biopsy. They hypothesized that this increase in T2 was due to an increase in myocardial water content, which correlated with edema due to rejection. In 1988, Konstam and colleagues conducted a study in which they used gadolinium diethylenetriamine pentaacetic acid (Gd-DTPA) to assess the CMR-determined contrast enhancement associated with rejection using a rat model.3 They evaluated the grade of rejection on biopsy and compared it to the degree of T1-weighted enhancement. In their study, nearly every case with moderate to severe rejection was associated with more than one region of myocardial signal enhancement. Subsequent studies have focused on the role of macrophages in acute cardiac rejection. Macrophages are believed to be the major cell type involved in allograft rejection. Investigators have injected particles into a rodent model that are then taken up by macrophages, in order to visualize the rejection process. In 2001, Kanno and colleagues injected ultrasmall superparamagnetic iron oxide (USPIO) particles into rats that had received either syngeneic or allogeneic transplants.4 They reported a difference in signal intensity between these two groups, due to macrophages taking up the USPIO particles. Macrophages that were labeled with these particles decreased in the intensity of the CMR signal, and this effect occurred to a greater degree in rats with increasing grades of rejection, as determined by myocardial biopsy. Macrophages play an important role in rejection; therefore, these findings suggest that the areas observed with CMR were sites of acute rejection. To confirm that rejection was being visualized, the investigators treated the rats with cyclosporin for varying durations. The animals that were treated for a shorter time showed a decrease in the signal intensity, while those that were treated for a longer time did not. Biopsy confirmed that animals that were treated for a shorter time were undergoing moderate rejection, compared to animals that were treated for a longer time, which demonstrated mild rejection. Wu and colleagues reported similar findings in 2004.5 They acknowledged the studies that assessed T1 and T2 relaxation times and gadolinium enhancement but noted that these findings represented edema and necrosis, which are typically found in late rejection. Detecting USPIOlabeled macrophage, on the other hand, would be more clinically useful because the rejection would be identified before necrosis occurred. They also reported that myocardial tagging by spatial modulation of magnetization (SPAMM) or DANTE schemes can be used to identify the

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Figure 40-1 Relative change in CMR myocardial signal intensity (SI) for syngeneic and allogeneic transplants in groups that received one of the ultrasmall superparamagnetic iron oxide varieties. Error bars represent 95% confidence. Source: Reprinted with permission from Penno E, Johnsson C, Johansson L, Ahlstrom H: Comparison of ultrasmall superparamagnetic iron oxide particles and low molecular weight contrast agents to detect rejecting transplanted hearts with magnetic resonance imaging. Invest Radiol 2005;40 (10):648–654.

heterogeneous distribution of myocardial dysfunction in early acute rejection. Penno and colleagues investigated this phenomenon further.6 They compared low molecular weight gadolinium chelate to two varieties of USPIOs by intravenous injection of paramagnetic agents into rodents undergoing acute rejection and a control group of rodents receiving syngeneic transplants. They demonstrated significant differences between rejecting and nonrejecting allografts in each of the two USPIOs (Fig. 40-1). In the group injected with gadolinium chelate, however, a difference in signal intensity could not be appreciated. This suggests that USPIO particles were superior to the gadolinium chelate for visualization of acute rejection. Finally, in 2006, Wu and colleagues reported that micrometer-sized paramagnetic iron oxide (MPIO) particles could be used to visualize individual macrophages with CMR.7 They injected rats with these particles, which were

A

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C

endocytosed by macrophages and carried to the sites of rejection. Areas of high contrast that were visualized in rejecting grafts differed markedly from controls, while grafts without rejection were not distinguishable from controls (Fig. 40-2). Furthermore, the investigators were able to observe individual macrophages for several days. This allowed for an unprecedented study of the time course of acute rejection. This process is advantageous over endomyocardial biopsy because it enables the visualization of the spatial progression of rejection. The investigators observed movement of the macrophages from pericardium to endocardium. Owing to the limitations of myocardial biopsy, this progression had not been previously reported. Accordingly, this technique allows for a greater degree of specificity in determining the stage of rejection, because it evaluates the myocardium in its entirety rather than segmentally. Another important clinical application of CMR imaging of the distribution of MPIO-labeled macrophages is the potential to monitor response to antirejection therapy. This is made possible by the longevity of the label and the ability to observe the cellular response for days at a time.

IMAGING PATIENT STUDIES When CMR-evaluated cardiac morphology is used, little difference is observed in patients with acute rejection compared to those with no evidence of rejection. Therefore, investigations have examined other CMR parameters to aid in the diagnosis of rejection. In 1988, Lund and colleagues found increased T2 values in patients with biopsy-determined acute rejection.8 Lauerma and colleagues performed a study in 1996 using CMR to determine the volumes and cyclic changes in the volumes of the left and right atria.9 When they compared patients with cardiac allografts to controls using conventional gradient echo cine CMR, they found that the transplanted hearts had larger end-systolic volumes and decreased ejection fractions. In addition, the stroke volumes were smaller, and the conduit volumes were larger. The atrial filling and emptying rates were also decreased in the transplant patients. Although the correlation between atrial function

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Figure 40-2 In vivo CMR of allograft heart and lungs, 1 day after intravenous injection of micrometer-sized paramagnetic iron oxide particles. The areas of high contrast likely represent one or more macrophages that have taken up the labeled iron. A, Allograft heart on postoperative day 5. B, C, Allograft heart on postoperative day 6. D, Allograft lung on postoperative day 6. E, Isograft heart on postoperative day 6. Notice the lack of contrast changes in the isograft heart. Shown with 156-mm in-plane resolution at 4.7 T by using a Bruker Biospec AVANCE-DBX magnetic resonance imaging instrument. Source: Reprinted with permission from Wu YL, Ye Q, Foley LM, et al: In situ labeling of immune cells with iron oxide particles: an approach to detect organ rejection by cellular MRI. Proc Natl Acad Sci U S A 2006;103(6):1852–1857. Copyright 2006 National Academy of Sciences, U.S.A. Cardiovascular Magnetic Resonance 549

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and cardiac rejection was not evaluated, CMR was found to be a reliable method for evaluating atrial volume and cyclic atrial function. In 1998, Marie and colleagues using a black-blood CMR pulse sequence (single spin echo/inversion recovery) on 75 cardiac transplant patients, compared T2 values to histologic evidence of significant rejection, and reported that this procedure accurately detected allograft rejection.10 This was associated with 89% sensitivity and 91% specificity. In contrast, echocardiography had a specificity of 91% but a sensitivity of only 44%. In 2001, Marie and colleagues reported a study in 68 post-transplant patients with a high suspicion of acute rejection.11 Again, they used a black-blood CMR pulse sequence and compared the results with endomyocardial biopsy. The results were consistent with their previous study: Increasing grades of acute rejection were associated with increasing T2 relaxation times. They determined that the black-blood sequence was superior to conventional spin echo pulse sequences because it eliminated the confounding effect of the blood signal on T2 measurements. Another interesting finding was that abnormal T2 values are excellent predictors of future rejection. Patients with T2 elevation who did not have moderate rejection on endomyocardial biopsy were likely to develop moderate to severe rejection in the following months. These findings suggest that CMR can predict significant rejection even before it is seen on biopsy. However, echocardiographic parameters and even myocardial biopsy were unable to predict future significant rejection. The distinction between mild and moderate rejection is an important one, because mild rejection is virtually a baseline phenomenon in myocardium from cardiac allografts, requiring no change in antirejection treatment, whereas the presence of moderate rejection suggests a change to a more intense treatment strategy. More recently, Almenar and colleagues used Gd-DTPAenhanced CMR to evaluate 40 patients who had cardiac allografts.12 CMR findings were compared with the results of endomyocardial biopsies. Although most of the comparisons did not achieve statistical significance, there was a significant increase in relative myocardial uptake of GdDTPA in patients who demonstrated significant acute rejection. There was a trend for increased myocardial uptake relative to the degree of rejection seen on biopsy, but this trend did not achieve statistical significance. The authors concluded that CMR was promising as a means to diagnose acute cardiac rejection, but more work is needed.

CORONARY ARTERY CMR PATIENT STUDIES The role of coronary artery CMR in detecting cardiac rejection has not been studied. Investigators have used this technique to detect disturbances in coronary artery flow in cardiac allograft patients. Mohiaddin and colleagues used fast low-angle shot coronary artery CMR to evaluate the proximal coronary arteries in cardiac allograft recipients.13 Sixteen patients were studied for obstruction of coronary artery blood flow. Coronary artery CMR results were compared with X-ray angiography findings to determine 550 Cardiovascular Magnetic Resonance

the sensitivity and specificity of CMR. A CMR finding of right coronary artery stenosis was 100% sensitive and 75% specific when compared to X-ray angiography. These results were superior to those for the left anterior descending (LAD) and left circumflex (LCX) arteries, in which the technique exhibited 86% specificity but poor sensitivity. The right coronary, left main stem, and LAD arteries had a collective specificity of 82% and sensitivity of 56%. The investigators thus reported that while there was some validity to the MRA findings, the usefulness of this approach was limited by low sensitivity and specificity. They suggested the development of new acquisition techniques. They also noted that the accuracy of coronary artery CMR for evaluating the coronary arteries in their study could have been improved with the use of respiratory gating methods. A similar study was performed by Davis and colleagues, who performed coronary artery CMR on 15 adult male allograft recipients.14 They used a breath hold electrocardiogram-gated segmented k-space technique to identify stenotic regions of the coronary arteries. They were able to visualize seven regions of stenosis of the coronary arteries (defined as a compromise of 50% or greater of the luminal diameter). In comparison, standard coronary angiography revealed nine such areas. In the patients with normal coronary anatomy, the investigators were able to identify the left main coronary artery. MRA showed a þ25 anterior (clockwise) RCA ostial rotation as well as a corresponding realignment of the left main coronary artery ostium. They hypothesized that quantifying the rotation of the coronary ostia might enable further exploration of coronary perfusion in the future.

SPECTROSCOPY ANIMAL STUDIES CMRS can demonstrate rejection through changes in the biochemical makeup of the myocardial tissue. 31P spectroscopy can detect allograft rejection before changes occur that are visible with endomyocardial biopsy. One of the first studies to use in vivo 31P CMRS to evaluate rejection in cardiac transplantation was performed in 1986 by Hall and colleagues.15 They induced rejection in dogs that received heterotopic cardiac transplants. The study compared the ability of two-dimensional echocardiography, 31P CMRS, and antimyosin monoclonal antibody (labeled with indium-111) to detect rejection that was confirmed with biopsy. Analysis of both the 31P CMRS and labeled antimyosin monoclonal antibody revealed findings that were consistent with detection of mild to moderate rejection. Evaluation of this progressive rejection by 31P CMRS was demonstrated by decreasing phosphocreatine levels. Another relevant finding was that the changes in the antimyosin antibody and 31P CMRS were detected before any changes on echocardiography, suggesting a decreased clinical utility for echocardiography in detecting acute rejection. In 1987, Canby and colleagues studied changes in cardiac rejection with 31P CMRS using a rat model.16 They analyzed the ratio of phosphocreatine to inorganic phosphate (PCr/Pi) and PCr to beta adenosine triphosphate (PCr/ATP), as well as the intracellular pH. As was noted

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Day 2 (a)

Figure 40-3 Left, CMRS spectra obtained from the control group that received isograft transplants. There are no serial changes in these spectra. Right, Spectra obtained from the allograft transplant group. There is a decrease in PCr signal over time. Source: Reprinted with permission from Canby RC, Evanochko WT, Kirklin JK, et al: Monitoring the bioenergetics of cardiac allograft rejection using in vivo P-31 NMR spectroscopy. J Am Coll Cardiol 1987;9:1067.

in the previous study, myocardial rejection leads to a decrease in PCr. Thus, spectra from the control group show no changes in the 31P metabolites (Fig. 40-3A). In the allogeneic group, however, the PCr signal decreases over time (Fig. 40-3B). This change is due to rejection. The ratios also reflect these changes. In the syngeneic grafts, the PCr/Pi ratio was unchanged or increased. PCr/ Pi steadily decreased in the rats in acute rejection, and this change became statistically significant by day 3 in relation to the syngeneic rats (p < .05) and by day 4 in relation to the baseline allogeneic rats (p <.005). The ratio of PCr to ATP showed a similar trend. It remained unchanged in the syngeneic group and decreased in the allogeneic group. This decrease was significant in comparison to baseline on day 3 (p < .01) and compared to the control group on day 4 (p < .005). The investigators also noted that even though the PCr/Pi ratio is more sensitive than the PCr/ATP ratio, the latter is used more often because of the confounding effects of 2,3-diphosphoglycerate in the blood, which obscures the myocardial Pi. Thus, it is difficult to measure the PCr/Pi ratio. This study also showed that there was no change in intracellular pH among the control rats, but the group with transplant rejection experienced an early alkaline shift, which was followed by acidosis. As was noted earlier, 31P CMRS can detect the early stages of rejection in cardiac allograft patients. In 1998, Fraser and colleagues conducted a study with a dog model to confirm this theory.17 They found that the abnormalities that were evident on 31P CMRS occurred before any other functional or morphologic signs of cardiac rejection. CMRS could therefore be an excellent indicator of early cardiac rejection in the clinical setting. Fraser and colleagues continued their research using the dog model they developed. They observed that the biochemical changes detected by 31P CMRS are reversible.18 After immunosuppressive treatment, the indicators of rejection returned to baseline. Not only can 31P CMRS assess whether a patient is experiencing cardiac rejection, but it can also be used to evaluate the patient’s response to

SPECTROSCOPY PATIENT STUDIES Early studies with 31P CMRS on cardiac transplant patients produced mixed results. In 1990, Herfkens and colleagues demonstrated a trend of decreasing PCr/ATP ratios in heart transplant patients undergoing rejection.20 A report by Bottomley and colleagues was less conclusive.21 They examined the PCr/ATP ratio of patients with cardiac allograft rejection, as determined by endomyocardial biopsy, in comparison with that of healthy control subjects. While there was a significant decrease in the PCr/ATP ratio in the group of rejecting patients (p < .01), this method was unable to distinguish between mild and moderate rejection, a key determinant in the choice of rescue therapy. When the cardiac-averaged ratio was used, the sensitivity of 31P CMRS for detection of patients who needed aggressive antirejection therapy was 50%, and the specificity was 73%. Using the lowest PCr/ATP ratio for this purpose showed a sensitivity of 88% and a specificity of 55%. These findings emphasized the need for further development to make 31P CMRS more clinically useful. One problem with both of these studies is that they evaluated patients at various points in the chronologic progression of rejection. A subset of the patients had early acute rejection, another subset had late acute rejection, and yet another subset was in chronic rejection. Evanochko and colleagues performed a series of studies using either the single-voxel (image-selected in vivo spectroscopy, or ISIS) or the multiple-voxel approach (onedimensional chemical shift imaging, or 1DCSI, with ISIS and gradient phase encoding).22 Figure 40-4 compares the 31 P CMRS findings using ISIS in a patient without rejection as seen on biopsy with those from another patient with transplant rejection. There is a high level of PCr in Figure 40-4A, which represents spectra from the control patient. The spectra depicted in Figures 40-4B through 40-D are from the patient with rejection. Figure 40-4B shows abnormal phosphate metabolism, with a decrease in PCr. The investigators then gave this patient immunosuppressive therapy and monitored the response to therapy with 31P CMRS. Figures 40-4C and 40-4D represent spectra Cardiovascular Magnetic Resonance 551

40 CARDIAC TRANSPLANTATION

Day 7 (d)

treatment. Most important, all of this can be accomplished in a noninvasive manner before there is permanent damage to the myocytes. In 1990, D’Amico and colleagues performed similar research using a dog model and achieved results similar to those of the earlier groups.19 They found a mean decrease in the PCr/Pi ratio of 28% in cardiac grafts with moderate to severe rejection. The investigators next compared 31P CMRS to positron emission tomography (PET). They found that PET was unable to detect allograft rejection in this model, highlighting the greater diagnostic ability of 31P CMRS. Countless other reports have validated the results of these studies. 31P CMRS is clearly capable of assessing acute cardiac rejection in an animal model in the laboratory. This prompted investigators to begin testing this theory on humans in an attempt to gain more clinically relevant insight into this procedure.

FUNCTIONAL CARDIOVASCULAR DISEASE

–25.000

D

C

B PCR

12.500

A

.00000

ATP

–12.500 [ppm]

Figure 40-4 31P CMRS spectra acquired with the single-voxel (ISIS) technique from patients with cardiac allografts. A, A spectrum from a patient with mild rejection. B, C, D, Serial studies from a patient experiencing moderate (B) and mild (C and D) rejection. C and D also correspond to a response to increased immunosuppression. Adapted from Evanochko WT, Bouchard A, Kirklin J, et al. Detection of cardiac transplant rejection in patients by 31-P NMR spectroscopy. In: Proceedings of the Society of Magnetic Resonance in Medicine; 1990:246.

acquired 4 and 6 weeks after the therapy, respectively. These spectra show a correction in PCr levels as the immune response leading to rejection was suppressed. This confirms that the changes seen with 31P CMRS represent the changing nature of allograft rejection and can be used to monitor the progression and response to treatment of these patients. The investigators then acquired spectra using 1DCSI from a normal volunteer and from a patient with decreased PCr/ ATP ratios. Thirty-two slices were taken, each 1 cm thick, and the spectra from each slice displayed adequate signalto-noise ratios for diagnostic purposes. Additionally, the column selection and phase-encoding selection were angulated to maximize the overlap of the selected slices and the left 552 Cardiovascular Magnetic Resonance

ventricle. The investigators used the division of skeletal muscle, cardiac muscle, and fat to study the regional variations in high-energy phosphate metabolism that are often seen in early acute cardiac rejection. They found a decrease in the PCr/ATP ratio in slices that corresponded to myocardial tissue in the patient with rejection. There are several possible explanations for the abnormalities in phosphate metabolites that are evident in acute cardiac rejection. One theory is that there are changes in the myocardial vasculature that produce diffuse capillary endothelial cell swelling, interstitial hemorrhage, and edema. Eventually, this progression results in ischemia.23 Another hypothesis, studied by Evanochko and colleagues, is that cardiac rejection changes the equilibrium constant of the myocardial creatine kinase.24 To test this hypothesis, they theorized that placing a stressor on the patient would potentiate the decrease in the PCr/ATP ratio in the presence of ischemia. In the absence of ischemia, however, the PCr/ATP ratio would not be expected to change. The investigators employed isometric handgrip exercises to place a mild amount of stress on the patient. They then compared spectra acquired before, during, and after the handgrip exercises. The PCr/ATP ratio was corrected for both T1 differences and blood contamination. Figure 40-5 demonstrates the spectrum from one cardiac transplant patient who was studied while exercising. Note that there is a decrease in the PCr/ATP ratio during exercise and a subsequent improvement after the exercise is completed. The handgrip exercise resulted in a small increase in the rate-pressure product in each of the subjects studied. Additionally, the control group usually showed an increase in heart rate, while the transplanted group usually showed an increase in blood pressure. There was no significant change in the PCr/ATP ratio in the control group, while 8 of the 27 transplant patients showed a greater than 2 standard deviation decrease in PCr/ATP ratio (Fig. 40-6). There was no significant change in the remainder of the transplant patients. Two studies by Caus and colleagues used 31P CMRS to detect early allograft dysfunction. In one study, they obtained 31 P spectra from 26 cardiac allografts using one-pulse CMRS sequence to evaluate certain metabolic markers as predictors for future graft failure in 26 cardiac allografts.25 Retrospectively, on the basis of the clinical course and transplantation outcome, they placed the grafts into three categories. Grafts that were transplanted into patients with uneventful outcomes had the highest ratios of PCr/ATP and PCr/Pi. Grafts that were transplanted into patients who experienced early graft failure were associated with lower levels of these ratios. The group with the lowest PCr/ATP and PCr/Pi ratios was composed of patients with grafts that were not suitable for transplantation. The investigators theorized that 31P CMRS could potentially be used to screen for allografts that are more likely to fail after transplantation. Knowledge of this information in advance would greatly reduce the morbidity and mortality associated with cardiac transplantation. Subsequently, Caus and colleagues studied 26 heart transplant recipients and 9 healthy subjects to detect coronary artery vasculopathy with31P CMRS with an acquisitionweighted 3D CSI sequence.26 The transplant patients were grouped on the basis of the presence of coronary artery vasculopathy as defined by X-ray angiography. The investigators reported a significant decrease in the PCr/ATP ratio of the patients with a vasculopathy when compared to those without a vasculopathy. Additionally, transplant patients

30 ATP

N = 11

N = 25

25 20 15

A

10 5 0 –5 –10

B –15 PI –20 –25 –30 Control

Transplant

C Figure 40-6 Comparison of the phosphocreatine/adenosine triphosphate (PCr/ATP) ratio for the control group (left) and the cardiac allograft group (right). The horizontal lines demarcate values that are 2 standard deviations from the control. Source: Reprinted with permission from Evanochko WT, Buchthal SD, den Hollander JA, et al: Cardiac transplant patients response to the 31-P MRS stress test. J Heart Lung Transplant 2002;21:522–529.

20

D

0

–20 [ppm]

Figure 40-5 Serial CMRS spectra from a cardiac transplant patient. A, Baseline spectrum. B, Spectrum acquired during handgrip exercise. C, D, Recovery spectra. There is a significant decrease in phosphocreatine (PCr) resonance in spectrum B, with subsequent return toward baseline in spectrum C. Spectrum C also shows an increase in Pi, which is corrected in spectrum D. Source: Reprinted with permission from Evanochko WT, Buchthal SD, den Hollander JA, et al: Cardiac transplant patients response to the 31-P MRS stress test. J Heart Lung Transplant 2002;21:522–529.

without a coronary artery vasculopathy had similar metabolic profiles to the control group of healthy subjects. Thus, 31P CMRS is potentially useful as a noninvasive method of screening for coronary artery vasculopathy.

CONCLUSION Despite the tremendous amount of progress in the utilization of CMR for the maintenance of patients with cardiac allograft transplantation, there remains much uncertainty

regarding the clinical role for this diagnostic modality. CMR imaging, angiography, and spectroscopy all show a great deal of promise as noninvasive alternatives to endomyocardial biopsy. Table 40-1 compares the various diagnostic modalities that can be used to detect allograft rejection. CMR offers several unique advantages, allowing us to visualize a level of detail that was previously unheard of. Individual macrophages can be followed and analyzed to determine their role in the process of acute rejection. Cellular changes can be detected before any abnormalities are seen on biopsy and, more important, before these changes become irreversible. Other imaging options, such as echocardiography, computed tomography, and PET, do not detect rejection before endomyocardial biopsy. CMRS is also able to detect other early changes of allograft rejection, such as abnormalities in the metabolism of highenergy phosphates. Again, these changes can be seen before other signs of rejection. CMR imaging and CMRS can each be used to monitor the treatment of patients undergoing rejection, because improvements in CMR findings are associated with lesser degrees of rejection. More clinically useful adaptations of these techniques are being developed. Accordingly, CMR is likely to become the most integral diagnostic tool for cardiac allograft rejection.

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40 CARDIAC TRANSPLANTATION

PCR

FUNCTIONAL CARDIOVASCULAR DISEASE

Table 40-1 Comparison of Diagnostic Modalities for the Detection of Cardiac Allograft Rejection Approach

Basis for Diagnostic Evaluation

Advantages

Disadvantages

Endomyocardial biopsy

Direct tissue microscopic evaluation

Current gold standard Allows for grading system with clear progression Reproducible

SPECT

Single-photon emitting radionuclides allow tomographic imaging of the distribution of single photon tracers (99mTc) Positron annihilation with resultant gamma photons in opposite directions (511 keV) Reflected ultrasound tracks motion and evaluates dimensions and wall thickness

Can detect myocyte necrosis and apoptosis Widely available

Invasive Expensive Subjective and prone to error Damages myocardium, causes inherent limitation on viable tissue Small sample size Ionizing radiation exposure Lower spatial resolution

PET Echocardiography

CMR imaging

Radio waves and magnetic fields

MRA

Magnetic and radiofrequency fields image the blood vessels Radio waves and magnetic fields generate spectra indicating the concentration of specific molecules (e.g., ATP)

CMRS

Useful to evaluate myocardial viability

Ionizing radiation Expensive

Widely available Allows assessment of global and regional ventricular function and myocardial thickening Noninvasive, no ionizing radiation, no harmful contrast agents Can detect spatial and temporal progression of rejection Provides reliable assessment of ventricular function Noninvasive No contrast or radiation needed

Usually two-dimensional evaluation Qualitative assessment

Noninvasive Assesses myocardial metabolism

Moderately expensive Problematic with implanted metallic devices Not used for allograft rejection Low spatial resolution Lengthy acquisition time

ATP, adenosine triphosphate; CMR, cardiovascular magnetic resonance; MRA, magnetic resonance angiography; MRS, magnetic resonance spectroscopy; PET, positron emission tomography; SPECT, single-photon emission computed tomography; Tc, technetium.

References 1. Baraldi-Junkins C, Levin HR, Kasper EK, et al. Complications of endomyocardial biopsy in heart transplant patients. J Heart Lung Transplant. 1993;12:63–67. 2. Aherne T, Tscholakoff D, Finkbeiner W, et al. Magnetic resonance imaging of cardiac transplants: the evaluation of rejection of cardiac allografts with and without immunosuppression. Circulation. 1986;74:145–156. 3. Konstam MA, Aronovitz MJ, Runge VM, et al. Magnetic resonance imaging with gadolinium-DTPA for detecting cardiac transplant rejection in rats. Circulation. 1988;78(suppl III):87–94. 4. Kanno S, Wu YJ, Lee PC, et al. Macrophage accumulation associated with rat cardiac allograft rejection detected by magnetic resonance imaging with ultrasmall superparamagnetic iron oxide particles. Circulation. 2001;104(8):934–938. 5. Wu YJ, Sato K, Ye Q, Ho C. MRI investigations of graft rejection following organ transplantation using rodent models. Methods Enzymol. 2004;386:73–105. 6. Penno E, Johnsson C, Johansson L, Ahlstrom H. Comparison of ultrasmall superparamagnetic iron oxide particles and low molecular weight contrast agents to detect rejecting transplanted hearts with magnetic resonance imaging. Invest Radiol. 2005;40(10):648–654. 7. Wu YL, Ye Q, Foley LM, et al. In situ labeling of immune cells with iron oxide particles: an approach to detect organ rejection by cellular MRI. Proc Natl Acad Sci U S A. 2006;103(6):1852–1857. 8. Lund G, Morin RL, Olivari MT, Ring WS. Serial myocardial T2 relaxation time measurements in normal subjects and heart transplant recipients. J Heart Transplant. 1988;7:274–279. 554 Cardiovascular Magnetic Resonance

9. Lauerma K, Harjula A, Jarvinen V, et al. Assessment of right and left atrial function in patients with transplanted hearts with the use of magnetic resonance imaging. J Heart Lung Transplant. 1996;15:360–367. 10. Marie PY, Carteaux JP, Angioi M, et al. Detection and prediction of acute heart transplant rejection: preliminary results on the clinical use of a “black blood” magnetic resonance imaging sequence. Transplant Proc. 1998;30(5):1933–1935. 11. Marie PY, Angioi M, Carteaux JP, et al. Detection and prediction of acute heart transplant rejection with the myocardial T2 determination provided by a black-blood magnetic resonance imaging sequence. J Am Coll Cardiol. 2001;37(3):825–831. 12. Almenar L, Igual B, Martinez-Dolz L, et al. Utility of cardiac magnetic resonance imaging for the diagnosis of heart transplant rejection. Transplant Proc. 2003;35(5):1962–1964. 13. Mohiaddin RH, Bogren HG, Lazim F, et al. Magnetic resonance coronary angiography in heart transplant recipients. Coron Artery Dis. 1996;7(8):591–597. 14. Davis SF, Kannam JP, Wielopolski P, et al. Magnetic resonance coronary angiography in heart transplant recipients. J Heart Lung Transplant. 1997;15(6):580–586. 15. Hall TS, Baumgartner WA, Borkon AM, et al. Diagnosis of acute cardiac rejection with antimyosin monoclonal antibody, phosphorus nuclear magnetic resonance imaging, two dimensional echocardiography and endocardial biopsy. J Heart Transplant. 1986;5:419–424. 16. Canby RC, Evanochko WT, Kirklin JK, et al. Monitoring the bioenergetics of cardiac allograft rejection using in vivo P-31 NMR spectroscopy. J Am Coll Cardiol. 1987;9:1067–1074.

22. Evanochko WT, Bouchard A, Kirklin J, et al. Detection of cardiac transplant rejection in patients by 31-P NMR spectroscopy. In: Proceedings of the Society of Magnetic Resonance in Medicine; 1990:246. 23. Yabe T, Mitsunami K, Okada M, et al. Detection of myocardial ischemia by 31-P magnetic resonance spectroscopy during handgrip exercise. Circulation. 1994;89:1709–1716. 24. Evanochko WT, Buchthal SD, den Hollander JA, et al. Cardiac transplant patients response to the 31-P MRS stress test. J Heart Lung Transplant. 2002;21:522–529. 25. Caus T, Kober F, Mouly-Bandini A, et al. 31P MRS of heart grafts provides metabolic markers of early dysfunction. Eur J Cardiothorac Surg. 2005;28:576–580. 26. Caus T, Kober F, Marin P, et al. Non-invasive diagnostic of cardiac allograft vasculopathy by 31P magnetic resonance chemical shift imaging. Eur J Cardiothorac Surg. 2006;29:45–94.

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17. Fraser CD, Chacko VP, Jacobus WE, et al. Metabolic changes preceding functional and morphologic indices of rejection in heterotopic cardiac allografts. Transplantation. 1988;46:346–351. 18. Fraser CD, Chacko VP, Jacobus WE, et al. Evidence from 31P NMR studies of cardiac allografts that early rejection is characterized by reversible biochemical changes. Transplantation. 1989;48:1068–1700. 19. D’Amico TA, Buchanan SA, Gall SA, et al. Diagnosis of cardiac allograft rejection using PET and MRS. Circulation. 1990;82(suppl III):613. 20. Herfkens RJ, Charles HC, Negro-Vilar R, Van Trigt P. In vivo P-31 NMRS of human heart transplants (abstract). In: Proceedings of the Society of Magnetic Resonance in Medicine. 1988:827. 21. Bottomley PA, Weiss RG, Hardy CJ, Baumgartner WA. Myocardial high energy phosphate metabolism and allograft rejection in patients with heart transplants. Radiology. 1991;181:67–75.

FUNCTIONAL CARDIOVASCULAR DISEASE

CHAPTER 41

Cardiovascular Magnetic Resonance Spectroscopy Stefan Neubauer

Cardiovascular magnetic resonance (CMR) imaging uses the 1H nucleus in water (H2O) and fat (CH2 and CH3 groups) molecules as its only signal source and therefore offers little insight into the biochemical state of cardiac tissue. In contrast, CMR spectroscopy (CMRS) allows the study of many other nuclei with a nuclear spin, that is, with an uneven number of protons, neutrons, or both, and is the only available method for the noninvasive assessment of cardiac metabolism without need for the application of external radioactive tracers. Information on the major nuclei of interest for the metabolic study of cardiac tissue by CMRS is given in Table 41-1, including 1H, 13C, 23 Na, and 31P. While theoretically, many clinical questions can be answered with CMRS, the main reason why MRS has not yet fulfilled its potential in clinical cardiology is related to the fundamental physical limitations of the method: The nuclei that are studied with MRS have a much lower magnetic resonance (MR) sensitivity than 1H has, and they are present in concentrations that are several orders of magnitude lower than those of 1H nuclei of water and fat. Therefore, the temporal and spatial resolution of CMRS has so far remained far behind that of CMR imaging. While this chapter focuses on clinical cardiac CMRS, some explanation of the experimental principles of CMRS is important even for the clinical reader. CMRS has been a widespread method in experimental cardiology, ever since the first 31P-MR spectrum from an isolated heart was obtained by Radda’s group in the 1970s,1 and since then experimental CMRS studies have offered numerous fundamental insights into cardiac metabolism. Furthermore, only with an understanding of the major implications of experimental CMRS studies are we able to fully appreciate the potential of the method and extrapolate to clinical cardiovascular CMRS applications that should become feasible in the future, once the technical challenges that currently limit the clinical utility of CMRS have been overcome. For those who are interested in greater detail on experimental CMRS and methodologic background, comprehensive reviews of the subject are available elsewhere.2–5

PHYSICAL PRINCIPLES The basic principles of CMRS (see Gadian6 for a textbook on the general physical principles of CMRS) are best explained from the most extensively studied nucleus, 31P, and from the most widely used animal model, the isolated buffer-perfused rodent heart. These principles apply to 556 Cardiovascular Magnetic Resonance

MRS of all nuclei. A CMR spectrometer consists of a highfield superconducting magnet (currently up to 18 T) with a bore size ranging between approximately 5 cm and approximately 1 m. The nucleus-specific probe head with the radiofrequency (RF) coils, which are used for MR excitation and signal reception, is seated within the magnet bore. The magnet is interfaced with a computer, a magnetic field gradient system, and an RF transmitter and receiver. CMRS requires much higher magnetic field homogeneity than does CMR imaging. Therefore, before any MRS experiment, the magnetic field must first be homogenized with shim gradients. Spin excitation is achieved by sending an RF impulse into the RF coils. The resulting MR signal, the free induction decay (FID), is then recorded and stored. The FID describes the relationship between time and signal intensity, showing an exponential signal decrease over time. After signal acquisition, the FID is Fourier transformed, resulting in an MR spectrum, which relates resonance frequency and signal intensity. Because of the low sensitivity of CMRS, many FIDs have to be signal-averaged to obtain MR spectra with an adequate signal-to-noise ratio (SNR), that is, the signal amplitude divided by the standard deviation of background noise. The required number of acquisitions is dependent on the concentration of the metabolite under investigation, the filling factor (the mass of the heart relative to the coil size), the natural abundance of the nuclear isotope studied, its relative MR sensitivity (see Table 41-1), the pulse angle, and the pulse repetition time (TR). For a perfused rat heart experiment at greater than 7 T, 100 to 200 FIDs are typically acquired. An important requirement for quantification of spectra is the correction for the effects of partial saturation, which are dependent on the selected pulse angles and TR. A complete MR signal can be obtained only when nuclei are excited from a fully relaxed spin state, that is, when a time of at least 5 T1 has passed since a previous excitation (e.g., T1 of phosphocreatine at 7 T of 3 seconds requires a TR of 15 seconds, while T1 at 1.5 T of approximately 4.4 seconds requires a TR of 22 seconds). Acquisition of fully relaxed MR spectra therefore requires long TRs, leading to prohibitively long acquisition times. In practice, we take advantage of the fact that the initial part of the FID contains most of the signal. Use of shorter TRs therefore yields spectra with higher SNR for a given acquisition time, but some of the signal is lost owing to saturation, that is, spectra are partially saturated. The extent of saturation also varies for different 31 P-resonances, because the T1s of 31P-metabolites such as phosphocreatine (PCr) and ATP are significantly

Natural Abundance

Relative Magnetic Resonance Sensitivity

99.98%

100%

13

1.1%

1.6%

23

100%

9.3%

31

100%

6.6%

Nucleus 1

H

C Na

P

Myocardial Tissue Concentrations H2O 110 M; up to 90 mM (CH3-1H of creatine) labeled compounds, several mM 10 mM (intracellular); 140 mM (extracellular) Up to 18 mM (PCr)

different (T1 of PCr is two to three times longer than T1 of ATP). Therefore, for quantification of partially saturated spectra, saturation factors must be applied for correction. By comparing fully relaxed and saturated spectra, these factors can be determined for each metabolite. In practice, TRs and pulse angles for CMRS are chosen to yield acceptable measurement times at an approximately 20% to 50% degree of saturation. Larger degrees of saturation, resulting from the use of extremely short TRs, would make quantification of spectra unreliable. A 31P-MR spectrum from an isolated, beating rat heart, obtained in 5 minutes at 7 T with a TR of 1.93 seconds and a pulse angle of 45 , is shown in Figure 41-1. A typical cardiac 31P-spectrum shows six resonances, corresponding to the three 31P-atoms of ATP (the resonance at the right shoulder of the a-P-ATP peak represents the P-atom of NADþ), PCr, inorganic phosphate (Pi), and monophosphate PCr

EXPERIMENTAL FOUNDATIONS

ATP

γ–

α–

31

P-Magnetic Resonance Spectroscopy

β–

Pi MPE

10

0

–10

esters (MPE), which mostly represent AMP and glycolytic intermediates. The reason why only a small number of 31Presonances are detectable, in spite of many more 31Pcontaining metabolites present in heart, is that for 31P-nuclei to be visible, they must be present above a certain concentration threshold of approximately 0.6 mM and must be free in solution. Largely immobilized metabolites such as plasma membrane phospholipids do not yield a quantifiable MR signal, owing to very short T2 values; instead, these metabolites contribute to the broad “baseline hump” of 31P-spectra. Owing to the phenomenon termed chemical shift, different metabolites resonate at distinct frequencies, allowing their discrimination from each other. Chemical shift is given in parts per million (ppm) and is quantified in relation to the B0 field. The physical basis of chemical shift is that different positions in the molecule lead to subtle differences in the local magnetic field strength, leading to a spread of 31Pmetabolite resonance frequencies over a range of approximately 30 ppm. After saturation correction, the area under each resonance is directly proportional to the amount of each 31P nucleus in the heart; metabolite resonances can therefore be quantified by integrating peak areas. The recommended method for this is the use of Lorentzian line curve fit algorithms, which removes interpretation bias from quantification of spectra. Relative metabolite levels, such as the PCr/ATP ratio, can then be calculated directly, while absolute metabolite concentrations require comparison of the tissue 31P-resonance areas to those of an external reference standard with a known amount of 31P. Frequently, phenylphosphonate is used for this purpose, as this generates an additional peak in the spectrum that does not overlap with the cardiac 31P-resonances.7,8 The most significant advantage of CMRS over destructive methods such as traditional biochemical assays, in which the tissue has to be frozen and extracted, is that the CMRS measurement is noninvasive. Therefore, spectra can be acquired sequentially, and the response of energy metabolites to ischemia, hypoxia, or inotropic stimulation can be followed on-line. With this approach, each heart can serve as its own control, yielding a more powerful experimental design and substantially reducing the number of required experiments.

–20

Chemical shift (ppm) Figure 41-1 31P-MR spectrum of an isolated, buffer-perfused rat heart obtained within 5 minutes at 7 T. MPE, monophosphate esters; PCr, phosphocreatine; Pi, inorganic phosphate.

With 31P-CMRS, cardiac high-energy phosphate metabolism, that is, the energetic state of the heart, can be monitored noninvasively. ATP is the direct energy source for all energy-consuming reactions in the heart. PCr acts as an energy reservoir and as an energy transport molecule in the creatine kinase/PCr energy shuttle9 (Fig. 41-2). For this, the high-energy phosphate bond is transferred from ATP to creatine at the site of ATP production (the mitochondria), yielding PCr and ADP in a reaction catalyzed by the mitochondrial creatine kinase isoenzyme. PCr diffuses through the cytoplasm to the site of ATP utilization, the myofibrils, where the back reaction (catalyzed by the Cardiovascular Magnetic Resonance 557

41 CARDIOVASCULAR MAGNETIC RESONANCE SPECTROSCOPY

Table 41-1 Atomic Nuclei Most Frequently Studied by MR Spectroscopy

FUNCTIONAL CARDIOVASCULAR DISEASE

Mitochondria

Cytoplasm

Myofibrils

ADP

Phosphocreatine

ADP

Oxidative phosphorylation

ATP-Transfer ATP

Creatine

CKmito

ATP utilization ATP

CK-MM

total, global ischemia, ATP, and PCr resonances have disappeared, and almost all the 31P in heart is present as Pi and monophosphate esters. During reperfusion, PCr and Pi show full recovery, and ATP shows partial recovery. We showed that when hearts are pretreated with endothelin1, a hormone that increases susceptibility to ischemia, recovery of high-energy phosphate metabolism is impaired.13 By summing data from several experiments, Clarke and colleagues14 demonstrated that the decrease of PCr and the increase of Pi were among the very earliest metabolic responses in myocardial ischemia, significant changes that occur within seconds. Therefore, if we were successful in measuring energetics in human myocardium with high temporal and spatial resolution, we could directly image parameters that detect myocardial ischemia within seconds after its onset. No other diagnostic approach currently achieves this. With the magnetization (saturation) transfer method (see, for example, Neubauer and colleagues15), the rate and velocity of the creatine kinase reaction, a measure of ATP transfer from mitochondria to myofibrils, can be measured in vivo. Creatine kinase reaction velocity correlates with cardiac workload16 and with recovery of mechanical function after ischemia.17 Experimental 31P-CMRS studies have contributed substantially to our understanding of the role of energetics in heart failure.15,18–20 Independent of the etiology of heart failure, the failing myocardium shows reduced PCr, unchanged or moderately (by less than 30%) reduced ATP, unchanged or increased Pi, and substantially reduced creatine kinase reaction velocity. These changes are likely to contribute to the impairment of contractile reserve in failing myocardium, due to the failure to maintain appropriate △G values during inotropic stimulation.21 Most recently, genetically manipulated mouse models with selective knockout22 or overexpression23 of the various components of energy metabolism have demonstrated the crucial role of cardiac energetics in normal and failing heart.

myofibrillar-bound MM-creatine kinase isoenzymes) occurs, ATP is reformed and is used for contraction. Free creatine then diffuses back to the mitochondria. The second important function of PCr and creatine kinase is to control the thermodynamic state of the cell, that is, to maintain free cytosolic ADP at low concentration. This is a requirement for normal cardiac function, because ADP determines the free energy change of ATP hydrolysis (△G; kJ/mol), a measure of the amount of energy released from ATP hydrolysis (see Neubauer10 for details on the calculation of free ADP and △G from creatine kinase shuttle metabolites). In the normal heart, △G is approximately 58 kJ/ mol. Many intracellular enzymes such as SR-Ca2þ-ATPase and others will not function properly below a threshold value for △G of about 52 kJ/mol. Pi is formed when ATP is hydrolyzed: ATP ( + ADP þ Pi. Pi increases when ATP utilization exceeds ATP production, for example, during ischemia. Intracellular pH can also be measured with 31P-CMRS, from the chemical shift difference between PCr and Pi, which is pH-sensitive. Cardiac energy metabolism has been investigated by 31 P-CMRS under various experimental conditions. For example, the effect of changes in cardiac workload on energetics has been examined. PCr levels do not change with moderate changes in workload11 but decline with substantial increases in cardiac work.12 ATP content remains constant with varying workload, because the creatine kinase equilibrium favors ATP synthesis over PCr synthesis by a factor of approximately 100. Thus, for any situation of myocardial stress, including ischemia, ATP decreases only when PCr levels are substantially depleted. This is the fundamental reason why the PCr/ATP ratio is a highly sensitive indicator of the energetic state of the heart. Changes of myocardial energy metabolism in experimental models of ischemia and reperfusion highlight the potential of 31PCMRS for the detection of ischemia in the human heart. Figure 41-3 shows an example of 31P-MR spectra during control, ischemia, and reperfusion. After 15 minutes of

Control

15 min ischemia

End of reperfusion

Untreated PCr

Pi

ET-1

10

5

γ−

ATP α−

β−

0 –5 –10 –15 –20

10

558 Cardiovascular Magnetic Resonance

5

0

Figure 41-2 Creatine kinase–PCr energy shuttle. See text for details. CKmito, mitochondrial creatine kinase isoenzyme; CK-MM, MM-creatine kinase isoenzyme.

–5 –10 –15 –20

10

5

0

–5 –10 –15 –20

Figure 41-3 31P-MR spectra from an untreated perfused rat heart and a heart treated with endothelin-1 (ET-1) during control, at the end of 15 minutes of ischemia, and at the end of reperfusion. See text for details. Source: De Groot M et al., J Mol Cell Cardiol. 1998;30:2657–2668. Reproduced with permission from Elsevier.

Other than 31P, the nucleus with the greatest potential for clinical CMRS is 1H. Protons have the highest MR sensitivity of all MR-detectable nuclei as well as high natural abundance (see Table 41-1). 1H is contained in a large number of metabolites, such as creatine, lactate, carnitine, taurine, and —CH3 and —CH2 resonances of lipids.24,25 Measurement of total creatine26,27 in conjunction with 31P-CMRS should allow the noninvasive determination of free ADP and △G of ATP hydrolysis. By means of the oxymyoglobin and deoxymyoglobin resonances, tissue deoxygenation can be measured.28 Technical challenges for 1H-CMRS include the need for suppression of the strong signal from water and the complexity of 1H spectra with overlapping resonances, many of which remain to be characterized. The 13C nucleus has a low natural abundance (1%), and for a 13C-CMRS experiment, the heart has to be supplied with 13C-labeled compounds such as 1-13C-glucose. Cardiac substrate utilization,29 citric acid cycle flux, pyruvate dehydrogenase flux, or beta-oxidation of fatty acids can then be quantified.30,31 Clinical cardiac studies have yet to be reported because of the low sensitivity of 13 C-CMRS and the requirement for high concentrations of exogenous 13C-labeled precursors. 23 Na-CMRS can evaluate changes in total, intracellular and extracellular Naþ during cardiac injury.32 A cardiac 23 Na spectrum shows a single peak representing the total Naþ signal, and to split the intracellular and extracellular Naþ pools into two resonances, paramagnetic shift reagents, such as [TmDOTP]5, are added to the perfusate. This method has been used experimentally to examine the mechanisms of intracellular Naþ accumulation in ischemiareperfusion injury,33 but 23Na-MR shift reagents for clinical use are not yet available. Experimental CMR imaging of total 23Na shows that in acute ischemia, the total myocardial 23Na CMR signal increases owing to the breakdown of ion homeostasis and the formation of both intracellular and extracellular edema (see, for example, Kim and colleagues34 and Horn and colleagues35). Furthermore, 23Na remains significantly elevated during chronic scar formation post coronary ligation35 because of the expansion of the extracellular space in scar, and the area of elevated 23 Na signal correlates closely with histologically determined infarct size. Importantly, 23Na content is not elevated in stunned or hibernating myocardium.35 Therefore, 23 Na CMR may allow detection of myocardial viability without the use of external contrast agents.

CLINICAL CARDIOVASCULAR MAGNETIC RESONANCE SPECTROSCOPY STUDIES Methodologic Considerations Almost all previous human cardiovascular MR spectroscopy studies have interrogated the 31P nucleus. The main reason for the slow progress with clinical CMRS is that the method poses major technical challenges. In practice, total examination time should not be more than 60

minutes; time for signal acquisition is therefore limited. The heart is a rapidly moving organ, currently requiring gating to the electrocardiogram (ECG), and when resolution is further improved, ultimately to respiration as well.36 Most previous clinical CMRS studies have been performed at 1.5 T, that is, at a considerably lower field than experimental studies. The cardiac muscle lies behind the chest wall skeletal muscle, which creates a strong 31P-signal that must be suppressed by spectral localization techniques. Such localization methods include depth-resolved surface coil spectroscopy, rotating frame, 1D-CSI (chemical shift imaging), image-selected in vivo spectroscopy, and 3D-CSI (Fig. 41-4). These methods are reviewed in detail elsewhere.37 Bottomley and colleagues37–40 have pioneered most of the early methodologic development of human CMRS. Most CMRS studies have been performed in the prone position rather than supine, as this reduces motion artifacts and the distance of the heart from the surface coil, thus improving sensitivity. Most spectroscopic techniques require a stack of 1H scout images to be obtained first, which are used to select the spectroscopic volume(s). The low sensitivity of 31 P-CMRS requires relatively large voxel sizes, typically approximately 20 to 70 mL. A 31P-MR spectrum of a healthy subject obtained with 3D-CSI is shown in Figure 41-5. In comparison to the rat heart spectrum, the SNR is lower, and two additional resonances are detected: 2,3-diphosphoglycerate (2,3-DPG), due to the presence of erythrocytes in the interrogated voxel, and phosphodiesters (PDE), a signal due to membrane as well as serum phospholipids. The 2,3-DPG peaks overlap with the Pi resonance, which therefore cannot be detected in blood-contaminated human 31P-MR spectra. Thus, intracellular pH can also not be determined; Pi and pH should become detectable in human myocardium when spatial resolution is increased to minimize blood contamination of 31P-spectra. By calculating the PCr/ATP and PDE/ATP peak area ratios, relative quantification of human 31Pspectra is simple. PCr/ATP is a powerful index of the energetic state of the heart (see the section entitled “Experimental Foundations”), while the meaning of the PDE/ATP ratio is poorly understood, and this ratio probably does not change with cardiac disease. 31P-resonances must be corrected for partial saturation as described for experimental CMRS. At 1.5 T, T1 of PCr is 4.4 0.5 seconds, and T1 of ATP is 2.4 0.4 seconds.37 While experimental studies suggest that 31P-T1s remain constant in the presence of cardiac disease,15,17 this remains to be proven in humans. Furthermore, 31Pspectra require correction for blood contamination: Blood contributes signal to the ATP, 2,3-DPG, and phosphodiester resonances. Because human blood spectra have an ATP/2,3-DPG area ratio of approximately 0.11 and a phosphodiester/2,3-DPG area ratio of approximately 0.19, for blood correction, the ATP resonance area of cardiac spectra is reduced by 11% of the 2,3-DPG resonance area, and the PDE resonance area is reduced by 19% of the 2,3-DPG resonance area.41 31P-metabolite ratios in blood also change in the presence of disease,42 which should be taken into account. A time domain or frequency domain Lorentzian line-fitting algorithm is applied for area integration of resonances in human 31 P-spectra. Cardiovascular Magnetic Resonance 559

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A different approach calibrates the 31P-signal to the tissue water proton content, measured by 1H-CMRS.40 The most advanced technique is spectral localization with optimum point spread function (SLOOP), which allows compartment matching with curved regions of interest that mimic the shape of the heart, and absolute quantification.45 SLOOP requires a 31P reference standard, flip angle calibration, B1 field mapping, and determination of myocardial mass in the interrogated voxel. A long-term goal for human CMRS is to combine 31Pand 1H-spectroscopy to measure free ADP and △G of ATP hydrolysis, as was recently described in canine myocardium.26 Another highly relevant energetic parameter is the rate and extent of ATP transfer (CK-flux) (see the section entitled “Experimental Foundations”). Bottomley and colleagues46 have developed the four-angle saturation transfer (FAST) method, which allows measurement of creatine kinase flux in approximately 30 minutes. The method employs acquisition of four spectra under partially saturated conditions. In two acquisitions, one of the exchanging species is saturated. The other two employ a control saturation. Each pair of saturations is applied with two different flip angles, and the equilibrium magnetization, relaxation times, and reaction rates are calculated from these measurements. Another development is acquisition-weighted 31P-chemical shift imaging,47 which reduces the signal contamination between adjacent voxels, improving spatial resolution (16 mL) and allowing acquisition of spectra from the posterior wall of the heart. In general, a major technical development effort, including advances in coil and sequence design, higher field strength,48,49 and standardization of protocols, will be required to bring cardiovascular MRS into clinical practice.

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Figure 41-4 Basic pulse sequences for localized CMRS with surface coils. A, depth-resolved surface coil spectroscopy. A single section parallel to the plane of the surface coil is selected by applying a magnetic resonance (MR) imaging gradient G in the presence of a modulated radiofrequency excitation pulse of flip angle a. B, The rotating frame MR method uses the gradient inherent in a surface coil to simultaneously spatially encode spectra from multiple sections parallel to the surface coil by means of application of a y flip angle pulse, which is stepped in subsequent applications of the sequence. C, The one-dimensional chemical shift imaging method similarly encodes multiple sections but uses an MR imaging gradient whose amplitude is stepped. D, The image-selected in vivo spectroscopy method localizes to a single volume with selective inversion pulses applied with Gx, Gy, and Gz MR imaging gradients. All eight combinations of the three pulses must be applied, and the resultant signals must be added and subtracted. E, A section-selective threedimensional chemical shift imaging sequence employs MR imaging section selection in one dimension and phase encoding in two dimensions. Source: Bottomley P. Radiology. 1994;191;593–612. Reproduced with permission from the Radiological Society of North America.

Owing to differences in CMRS methods used, the range of “normal” human heart PCr/ATP ratios reported in the literature is considerable, from about 1.1 to 2.4, with an overall average of about 1.8.37 This attests to the need for development of methodologic standards for CMRS. We recently reported absolute PCr levels in normal human myocardium of 9.0 1.2 and ATP levels of 5.3 1.2 mmol/kg wet weight, in agreement with experimental results.50 There is some debate as to whether high-energy phosphate levels decrease with advanced age,51,52 but larger systematic studies on this issue are still outstanding. During stress, PCr/ ATP ratios stay normal for all but extreme levels of inotropic stimulation, when there is a small decrease.53 Most recently, using FAST (see above), high-energy phosphate turnover rates (CK reaction rates) were found to be 0.29 0.06 sec1 in healthy volunteers at rest and did not change during doubling of the cardiac rate-pressure product.54

Absolute quantification of PCr and ATP is technically challenging but is desirable, because the PCr/ATP ratio does not detect simultaneous decreases of both PCr and ATP, which occur in the failing myocardium43 or in the infarcted nonviable myocardium.44 Absolute 31P-metabolite levels can be obtained by acquiring signal from a 31P-standard as well as estimates of myocardial mass from CMR.39

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In patients with hypertension, Lamb and colleagues demonstrated reduced PCr/ATP ratios both at rest and during dobutamine stress,55 and PCr/ATP ratios also

41 CARDIOVASCULAR MAGNETIC RESONANCE SPECTROSCOPY

Figure 41-5 Left, 31P-MR spectrum from a healthy human subject, 3D-CSI technique. Voxel size: 25 27 30 mm (20 mL). Right, Short axis and vertical long axis proton scout images. The entire voxel grid of the 3D-CSI localization is shown in green, and the voxel corresponding to the 31Pspectrum is shown in blue. 2,3-DPG, 2,3-diphosphoglycerate; PCr, phosphocreatine; PDE ¼ phosphodiesters; g-,a-,b-ATP, the three phosphorus atoms of adenosine triphosphate.

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correlated inversely with indices of diastolic function (E deceleration peak). In contrast, two other studies showed no significant changes of cardiac energetics in patients with hypertension.43,51 Differences in patient characteristics, such as severity and duration of hypertension, may be responsible for this discrepancy. Experimental data clearly suggest that cardiac energetics are impaired in longstanding hypertension.56 In contrast, physiologic hypertrophy in the athlete’s heart does not lead to a decrease of the myocardial PCr/ATP ratio, either at rest or during stress;53 experimental studies in exercisetrained rats had predicted this finding.57 In the future, it will be important to unravel the molecular mechanisms responsible for these differences in energy metabolism between physiologic and pathologic hypertrophy.

Diabetes and Obesity Other than via the well-recognized secondary mechanisms (e.g., accelerated coronary disease), diabetes has numerous deleterious effects on cardiac metabolism that may lead to cardiomyopathy. For example, the diabetic heart is insulin-resistant, and glucose utilization is impaired. Plasma free fatty acid levels are elevated, leading to increased expression of mitochondrial uncoupling proteins and reduced expression of glucose transporters (GLUT4). Several studies have examined cardiac energy metabolism in patients with maintained left ventricular ejection fraction and type II58,59 or type I60 diabetes mellitus and have uniformly shown reduced PCr/ATP ratios. PCr/ATP ratios were found to correlate inversely with plasma free fatty acid levels58 and with indices of diastolic function.59,60 Initial evidence also suggests reduced PCr/ATP ratios in patients with uncomplicated obesity and elevated free fatty acid

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levels (author’s own unpublished observations). Together, these results suggest that early cardiac metabolic derangement as observed in diabetes and obesity may contribute to the later development of heart failure, suggesting a possible explanation for the increased incidence of heart failure in diabetes and obesity.

Heart Failure and Cardiac Transplantation Experimental studies have firmly established altered energy metabolism as a hallmark of the chronically failing myocardium (see above). However, initial clinical 31P-CMRS studies, which included mild stages of heart failure, did not find significant reductions of PCr/ATP ratios.61–63 Hardy and colleagues first reported that the myocardial PCr/ATP ratio is significantly reduced (from 1.80 0.06 to 1.46 0.07) in patients with symptomatic heart failure of ischemic or nonischemic etiology.64 Subsequently, we found that the decrease of PCr/ATP ratios in dilated cardiomyopathy correlated with the New York Heart Association (NYHA) class65 and with left ventricular ejection fraction.66 Thus, PCr/ATP ratios decrease in advanced stages of heart failure (Fig. 41-6) but initially remain normal. It is known from experimental work that in heart failure, both PCr and ATP levels decrease in parallel,50 and this cannot be detected by measurement of PCr/ATP ratios. Accordingly, using SLOOP (see above) in patients with heart failure due to dilated cardiomyopathy (ejection fraction: 18%), we have reported that absolute PCr levels were reduced by 51% and ATP levels by 35%, while the PCr/ATP ratio decreased by 25% only.67 Furthermore, significant correlations between LV volumes/ejection fraction and energetics Cardiovascular Magnetic Resonance 561

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Figure 41-6 31P-CMR spectra. From bottom to top: volunteer, dilated cardiomyopathy with normal phosphocreatine/ATP ratio, dilated cardiomyopathy with reduced phosphocreatine/ATP ratio, dilated cardiomyopathy with severely reduced phosphocreatine/ ATP ratio; this patient died 7 days after the CMR examination. DCM, dilated cardiomyopathy; 2,3-DPG, 2,3-diphosphoglycetate; PCr, phosphocreatine; PDE, phosphodiesters; g-, a- and b-P-atom of ATP. Source: Neubauer S. et al. Circulation. 1997; 96:2190–2196. Printed with permission, American Heart Association.

were found, the strongest correlations being observed for PCr and the weakest for the PCr/ATP ratio. Thus, ATP levels are reduced in human heart failure, and the true extent of changes in energy metabolism in heart failure is underestimated when PCr/ATP ratios rather than absolute concentrations are measured. Nakae and colleagues68 used 1 H-CMRS to demonstrate significant reductions of total creatine levels and a correlation of creatine and LV ejection fraction in patients with DCM. Most recently, Weiss and colleagues showed a 50% reduction of CK reaction rates in mild-to-moderate heart failure,54 indicating that energy turnover rates are depressed to a greater extent than steady-state levels of PCr and ATP. PCr/ATP ratios also hold prognostic information on survival of patients with heart failure. We showed that in dilated cardiomyopathy, the myocardial PCr/ATP ratio was a better predictor of longterm survival than was left ventricular ejection fraction or NYHA class69 (Fig. 41-7). Thus, 31P-CMRS may be valuable for prognosis evaluation in heart failure. 562 Cardiovascular Magnetic Resonance

Figure 41-7 Kaplan-Meier life table analysis for total mortality of dilated cardiomyopathy patients divided into two groups split by the myocardial PCr/ATP ratio (<1.60 versus >1.60). Patients with an initially low PCr/ATP ratio showed increased mortality over the study period of, on average, 2.5 years. Source: Neubauer S et al. Circulation. 1997;96:2190–2196. Reproduced with permission from the American Heart Association.

Heart failure trials from the past two decades show that energy-costly treatment, such as beta-receptor mimetics or phosphodiesterase inhibitors, increases mortality, while energy-sparing treatment, such as beta-blockers, ACE inhibitors, or angiotensin-II-receptor blockers, improves survival. Thus, one of the most promising future applications of clinical 31P-CMRS is monitoring of the cardiac energetic response to new forms of heart failure treatment, and it is likely that the PCr/ATP ratio or absolute concentrations of PCr and ATP are powerful surrogate parameters for mortality in this situation. For example, in six patients with dilated cardiomyopathy treated with ACE inhibitors, digitalis, diuretics, and beta-blockers for 3 months, PCr/ATP ratios improved significantly during clinical recompensation, from 1.51 0.32 to 2.15 0.27.41 However, no systematic controlled study has so far used 31P-CMRS to monitor cardiac energetics during heart failure treatment. A treatment trial using 31P-CMRS to monitor cardiac energetics has been reported for Friedreich ataxia. This primarily neurologic disease often is associated with cardiomyopathy, as lack of the mitochondrial protein frataxin leads to deficient mitochondrial respiration and increased free radical damage. In patients with Friedreich ataxia treated with antioxidants (coenzyme Q and vitamin E) for 6 months, Lodi and colleagues70 reported that the myocardial PCr/ATP ratio increased from 1.34 0.59 to 2.02 0.43, demonstrating that cardiac energy metabolism was markedly improved by antioxidative treatment. Energetic derangement measured by 31P-CMRS can detect cardiac transplant rejection in animal models.71 However, while PCr/ATP ratios were reduced in human transplanted hearts with histologic signs of rejection, there was no correlation with biopsy histology scores.72 PCr/ATP ratios show a transient decrease after transplantation as a result of perioperative myocardial ischemia73 but recover after the first few weeks. Thus, PCr/ATP ratios cannot predict transplant rejection. The transplanted

Experimental studies have shown impaired cardiac energy metabolism in advanced LV hypertrophy.75 Similarly, in patients with left ventricular hypertrophy resulting from aortic stenosis or incompetence, Conway and colleagues76 detected reduced PCr/ATP ratios (1.10 0.32 versus 1.50 0.20) when patients were in clinical heart failure, but PCr/ATP ratios were normal (1.56 0.15) for clinically asymptomatic stages. Likewise, in patients with aortic valve disease, we showed reduced PCr/ATP ratios only for NYHA classes III and IV but not for classes I and II.77 When matched for the degree of heart failure, energy metabolism was more affected in aortic stenosis (pressure overload) than in aortic incompetence. We also showed that in aortic stenosis, altered energetics correlated with LV end-diastolic pressures and with end-diastolic wall stress.77 Recently, we reported unchanged absolute ATP concentrations and a 28% decrease of PCr concentrations in aortic stenosis using the SLOOP technique.67 The time course of recovery of cardiac energetics during regression of LV hypertrophy after surgical valve replacement can also be monitored by 31P-CMRS.78 When patients with aortic valve stenosis were studied before and 40 weeks after surgery, the PCr/ATP ratio increased from 1.28 0.17 to 1.47 0.14 (control subjects: 1.43 0.14); that is, energetic impairment was completely reversed 9 months after valve replacement. A long-term prospective clinical study of this subject would demonstrate whether 31 P-CMRS can provide clinical information on the optimum timing of valve replacement.

Ischemic Heart Disease CMRS Stress-Testing for Ischemia Detection Within seconds after reduction of oxygen supply, PCr levels decrease and inorganic phosphate increases, that is, these changes are extremely rapid indicators of myocardial ischemia (see above). If it were feasible to detect these metabolites in human myocardium with high temporal and spatial resolution, a 31P-CMRS-based biochemical stress test would be a powerful diagnostic tool for detecting exercise-induced regional ischemia, requiring only low levels of stress and without the need for intravenous agents or radiation. In selected patients with large anterior wall territories, which become ischemic on exercise, the feasibility of this principle has been demonstrated. Weiss and colleagues (Fig. 41-8) showed that in patients with high-grade LAD stenosis, PCr/ATP ratios were normal at rest, decreased during handgrip exercise (leading to a 30% to 35% increase of cardiac work) from 1.5 0.3 to 0.9 0.2, and returned toward normal during recovery.79 After

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Figure 41-8 PCr/ATP ratios at rest, during handgrip exercise, and during recovery in controls and in patients with stenosis of the left anterior descending coronary artery. PCr/ATP decreased in patients but not in healthy subjects. Source: Weiss RG et al. NEJM 1990;323:1593–1600. Reproduced with permission from the New England Journal of Medicine.

revascularization, PCr/ATP ratios remained constant during exercise. These findings were reproduced by Yabe and colleagues,80 who also demonstrated that a decrease of PCr/ ATP ratios was detected only in patients with reversible defects on thallium scintigraphy (viable myocardium) but not in those with fixed thallium defects (scar), in whom PCr/ATP was already reduced at rest.44 With a 31P-CMRS stress test, we would also be able to test the efficacy of revascularization procedures or of established or new antianginal medication. It is conceivable that a PCr threshold may become a clinically relevant parameter, as the level of exercise achievable without a decrease of myocardial PCr concentrations. This may in the future allow objective fine-tuning of antianginal therapy. The pathophysiologic mechanisms of exercise-induced chest pain in women with a normal coronary angiogram remain unclear, but microvascular dysfunction and subsequent tissue ischemia in the absence of epicardial coronary stenoses has been postulated. Buchthal’s group81 showed that in 7 of 35 women with chest pain and normal coronary arteries, the PCr/ATP ratio decreased by 29 5% during handgrip exercise, providing direct evidence of Cardiovascular Magnetic Resonance 563

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heart develops coronary vasculopathy, and Evanochko and colleagues74 suggested that 31P-CMRS measurements during exercise may be a more sensitive indicator. They studied transplanted patients with 31P-CMRS at rest and stress. Ten of 25 patients showed a decrease of the PCr/ ATP ratio of 26% 4% during exercise; correlations with histologic scores of rejection remain to be established.

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exercise-induced myocardial ischemia. In a subsequent study, they showed that at 3-year follow-up, an initially abnormal 31P-CMRS stress test was a strong predictor of future cardiovascular events.82 31P-CMRS stress testing may facilitate the development and monitoring of treatment for this ubiquitous condition.

Myocardial Viability Assessment Akinetic myocardium, supplied by a stenotic coronary artery, may be nonviable (scarred), or it may be viable, that is, stunned or hibernating. By readjusting the balance between oxygen supply and demand, hibernating myocardium has downregulated its contractility to reduce energy needs. Stunned myocardium shows transient contractile dysfunction following reperfusion after reversible ischemia. 31PCMRS studies of animal models have shown that in stunned and hibernating83 myocardium, ATP levels remain close to normal, while myocardial scar tissue contains negligible amounts of ATP (<1% of normal levels). Thus, a noninvasive diagnostic test that allows measurement of myocardial ATP levels with high spatial (1 to 5 mL) and acceptable temporal resolution (<30 minutes) should be attractive for viability assessment. For example, Kalil-Filho and colleagues84 studied 29 patients with anterior myocardial infarction (MI) 4 and 39 days after MI. All patients showed akinetic anterior myocardium, which had recovered function at the time of the second examination. PCr/ATP ratios were normal in stunned myocardium (1.51 0.17 versus 1.61 0.18 in normal) and did not change during the recovery of contractile function (1.51 0.17 versus 1.53 0.17). These observations were confirmed by Beer and colleagues, who also showed that in patients with nonviable infarcts who failed to recover regional contractile function after 6 months, 31 P-CMRS showed complete absence of PCr.85 The same group recently reported the detection of inferior infarcts by acquisition-weighted 31P-CMRS.86 However, loss of myocardial tissue due to necrosis and scar formation primarily leads to a reduction of both PCr and ATP, and the extent of viable tissue loss can therefore not be detected using the PCr/ATP ratio. Instead, measurement of absolute concentrations of high-energy phosphates is necessary. Only one clinical viability study has previously reported this. Yabe and

colleagues44 showed that absolute myocardial ATP content was significantly reduced in patients with fixed thallium defects (nonviable) but was unchanged in patients with reversible defects (viable myocardium). While these initial results are promising, substantial improvement in spatial resolution is necessary if 31P-CMRS evaluation of viability is to become clinically relevant. Viability detection may also be feasible by measuring total creatine content by localized 1H-CMRS,87 as creatine concentrations in scar tissue are negligible. Owing to the higher MR sensitivity of 1H and because the concentration of CH3creatine protons is approximately tenfold higher than 31 P-concentrations of ATP, resolution of 1H-CMRS is superior to 31P-CMRS, currently at approximately 1 mL. However, several methodologic hurdles (see above) remain to be overcome before clinical 1H-CMRS can be more widely applied. Because the myocardial 23Na signal is elevated in both acute necrosis and in chronic scar (see above), 23Na-CMR imaging may allow detection of myocardial viability without the use of external contrast agents. Owing to 100% natural abundance, relatively high MR sensitivity (9.3% of 1H) and tissue concentration (140 mM extracellular, 10 mM intracellular) of 23Na, and due to its short T1 (30 msec at 1.5 T), allowing for short TR, next to 1H, 23Na yields the highest MRI resolution of all MR-detectable nuclei. Cardiac 23Na-CMR has been demonstrated with a resolution of up to 392 mL in volunteers and patients,88,89 using an ECG-triggered 3D spoiled gradient echo sequence and an acquisition time of approximately 1 hour. In patients studied 8 days after and more than 6 months after acute myocardial infarction, we have obtained the first in vivo 23Na-MR images of infarcted human myocardium90 (Fig. 41-9). All patients after subacute infarction and 12 of 15 patients with chronic infarction showed an area of elevated 23Na signal intensity that significantly correlated with wall motion abnormalities. In a follow-up study, we demonstrated that the total Naþ signal remains significantly elevated in scar over a time period of at least 1 year.91 For 23Na-CMR to become a mainstream clinical viability assessment technique, further improvement of spatial resolution by use of faster gradients, improved coils, and higher field strength, as well as mapping of absolute Naþ content will be required. Figure 41-9 23Na-CMR images of the human heart. Two consecutive short axis 23Na CMR images of a patient with previous anterior myocardial infarction, showing elevated signal intensities (arrows) of the anterior wall. Images obtained with an ECG-triggered three-dimensional spoiled gradient echo sequence. Field of view: 450 450 mm2; matrix: 64 128; flip angle: 70 ; section thickness: 16 mm. Source: Sandstede J et al., Fortschr Ro¨ntgenstr. 2000;172:739–743, reproduced with permission from Georg Thieme Verlag.

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Specific Gene Defects with Cardiac Pathology Clinical cardiac 31P-CMRS has major potential for the noninvasive phenotyping of cardiomyopathies due to specific gene defects, which may eventually be identifiable by a specific metabolic profile. Most of the work in this area has been on hypertrophic cardiomyopathy (HCM), which is, in most cases, due to specific gene mutations, associated with structural disarray of myofibrils and often with substantial increases of left ventricular wall thickness. Energetic derangement has been suggested as the common pathophysiologic mechanism underlying all forms of HCM,92 and experimental studies of transgenic models of HCM confirm this.93 Human studies in HCM63,94–96 have uniformly shown reduced PCr/ATP ratios. For example, Jung and colleagues97 demonstrated that in young, asymptomatic patients with HCM, PCr/ATP ratios were reduced, indicating that energetic imbalance occurs early in the disease process. They also reported that HCM patients with a familial history of the disease showed a more pronounced derangement of energetics than did those patients without a family history.98 Using 1H-CMRS, Nakae and colleagues68 recently reported reduced total creatine content in patients with HCM. In the future, large patient groups with HCM and known specific gene defects will have to be studied to establish whether metabolic phenotyping by 31P- and 1 H-CMRS can identify the underlying genetic mutation. Becker muscular dystrophy, an X-chromosome-linked disease associated with the absence or altered expression of dystrophin in cardiac and skeletal muscle, may lead to the development of cardiomyopathy and heart failure. One study99 showed that both patients (PCr/ATP ratio: 1.55 0.33) and female gene carriers (1.37 0.25) had significantly lower PCr/ATP ratios than did control subjects (2.44 0.33), although all of the carriers and most of the patients showed preserved left ventricular function. Thus, energetic imbalance occurs early in the disease process and may contribute to the development of contractile dysfunction in Becker disease. Recently, altered cardiac energetics were also demonstrated in hereditary hemochromatosis100 and in familial hypercholesterolemia,101 where PCr/ATP ratios returned to normal after treatment with statins.

CARDIOVASCULAR MAGNETIC RESONANCE SPECTROSCOPY AT 3 TESLA The most promising route for the improvement of spatial and temporal resolution of CMRS is the use of magnetic

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field strengths higher than 1.5 T. While early attempts at this49,102 were limited by the technical prototype nature of such magnets, the new generation of 3 T CMR systems that has recently become available is eminently suitable for CMRS studies. Our initial experience shows that a substantial, approximately twofold increase in the SNR (compared to 1.5 T) can be achieved,103 bringing clinical applications substantially closer to reality. An example is shown in Figure 41-10. Also, Boesiger’s group recently reported substantial improvements for 1H-CMRS at 3 T.36

PERSPECTIVE AND GENERAL CONCLUSIONS CMRS allows for the noninvasive assessment of many aspects of cardiac metabolism in normal and diseased heart, providing a wealth of information that should, in theory, be clinically relevant for patient management. While overall, technical progress of CMRS has been much slower than that of CMR imaging, interesting new developments have recently been reported that may pave the way for more mainstream clinical applications of CMRS. These include the assessment of cardiac steatosis by 1HCMRS104,105 and the exciting opportunities of nuclear hyperpolarization methods, which are able to boost the CMRS signal by four to five orders of magnitude, thus, at least experimentally, allowing for a robust and quick assessment of metabolic pathways that has previously been impossible.106 At this point, however, the main obstacle for the widespread implementation of CMRS Cardiovascular Magnetic Resonance 565

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Thus, 31P-CMRS, 1H-CMRS, and 23Na-CMR/CMRS provide an array of methods that might in the future become the preferred approach for viability assessment if substantially higher spatial resolution can be achieved. Unlike currently used techniques, such as dobutamine stress echocardiography, nuclear imaging, or late gadolinium enhancement CMR, the MRS approach provides intrinsic contrast for distinction between viable and nonviable myocardium, is radiation-free, and does not require intravenous agents or stress testing.

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remains its technical complexity and limited resolution, and a major technologic effort is required to substantially improve spatial and temporal resolution. Also, to reduce measurement variability to less than 5%, high signal-tonoise spectra need to be acquired. Finally, standardized acquisition and quantification protocols will have to be developed and agreed upon to provide clinicians with reliable measurements that are also comparable among different centers. These goals should be achievable in the coming years by a major dedicated technical development effort on coil design, sequence design, and hyperpolarization methods and in particular on high-field magnets. High-resolution metabolic imaging would

then finally become a reality in patients with ischemic heart disease, heart failure, valve disease, and genetic cardiomyopathy.

ACKNOWLEDGMENTS The author would like to thank the following colleagues (listed in alphabetical order) for their contributions: Kieran Clarke, Jane Francis, Lucy Hudsmith, Christine Lorenz, Saul Myerson, Matthew Robson, Stefan Roell, Rolf Sauter, Michaela Scheuermann-Freestone, Juergen Schneider, Damian Tyler, Hugh Watkins, and Frank Wiesmann.

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20. Zhang J, Wilke N, Wang Y, et al. Functional and bioenergetic consequences of postinfarction left ventricular remodeling in a new porcine model: MRI and 31 P-MRS study. Circulation. 1996;94: 1089–1100. 21. Ingwall JS. Is cardiac failure a consequence of decreased energy reserve? Circulation. 1993;87(Suppl VII):58–62. 22. ten Hove M, Lygate CA, Fischer A, et al. Reduced inotropic reserve and increased susceptibility to cardiac ischemia/reperfusion injury in phosphocreatine-deficient guanidinoacetate-N-methyltransferaseknockout mice. Circulation. 2005;111:2477–2485. 23. Wallis J, Lygate CA, Fischer A, et al. Supranormal myocardial creatine and phosphocreatine concentrations lead to cardiac hypertrophy and heart failure: insights from creatine transporter-overexpressing transgenic mice. Circulation. 2005;112:3131–3139. 24. Ugurbil K, Petein M, Madian R, Michurski S, Cohn JN, From AH. High resolution proton NMR studies of perfused rat hearts. FEBS Lett. 1984;167:73–78. 25. Balschi JA, Hai JO, Wolkowicz PE, et al. 1H NMR measurement of triacylglycerol accumulation in the post-ischemic canine heart after transient increase of plasma lipids. J Mol Cell Cardiol. 1997;29: 471–480. 26. Bottomley PA, Weiss RG. Noninvasive localized MR quantification of creatine kinase metabolites in normal and infarcted canine myocardium. Radiology. 2001;219:411–418. 27. Schneider J, Fekete E, Weisser A, Neubauer S, von Kienlin M. Reduced (1)H-NMR visibility of creatine in isolated rat hearts. Magn Reson Med. 2000;43:497–502. 28. Kreutzer U, Mekhamer Y, Chung Y, Jue T. Oxygen supply and oxidative phosphorylation limitation in rat myocardium in situ. Am J Physiol Heart Circ Physiol. 2001;280:H2030–H2037. 29. Malloy CR, Jones JG, Jeffrey FM, Jessen ME, Sherry AD. Contribution of various substrates to total citric acid cycle flux and anaplerosis as determined by 13C isotopomer analysis and O2 consumption in the heart. MAGMA. 1996;4:35–46. 30. Lewandowski ED. Cardiac carbon 13 magnetic resonance spectroscopy: on the horizon or over the rainbow? J Nucl Cardiol. 2002;9: 419–428. 31. Weiss RG. 13C-NMR for the study of intermediary metabolism. MAGMA. 1998;6:132. 32. Kohler SJ, Perry SB, Stewart LC, Atkinson DE, Clarke K, Ingwall JS. Analysis of 23Na NMR spectra from isolated perfused hearts. Magn Reson Med. 1991;18:15–27. 33. Van Emous JG, Vleggeert-Lankamp CL, Nederhoff MG, Ruigrok TJ, Van Echteld CJ. Postischemic Na(þ)-K(þ)-ATPase reactivation is delayed in the absence of glycolytic ATP in isolated rat hearts. Am J Physiol Heart Circ Physiol. 2001;280:H2189–H2195. 34. Kim RJ, Lima JAC, Chen EL, et al. Fast 23Na magnetic resonance imaging of acute reperfused myocardial infarction: potential to assess myocardial viability. Circulation. 1997;95:1877–1885. 35. Horn M, Weidensteiner C, Scheffer H, et al. Detection of myocardial viability based on measurement of sodium content: a (23)Na-NMR study. Magn Reson Med. 2001;45:756–764. 36. Schar M, Kozerke S, Boesiger P. Navigator gating and volume tracking for double-triggered cardiac proton spectroscopy at 3 Tesla. Magn Reson Med. 2004;51:1091–1095. 37. Bottomley PA. MR spectroscopy of the human heart: the status and the challenges. Radiology. 1994;191:593–612.

60. Metzler B, Schocke MF, Steinboeck P, et al. Decreased high-energy phosphate ratios in the myocardium of men with diabetes mellitus type I. J Cardiovasc Magn Reson. 2002;4:493–502. 61. Schaefer S, Gober JR, Schwartz GG, Twieg DB, Weiner MW, Massie B. In vivo phosphorus-31 spectroscopic imaging in patients with global myocardial disease. Am J Cardiol. 1990;65:1154–1161. 62. Auffermann W, Chew WM, Wolfe CL, et al. Normal and diffusely abnormal myocardium in humans: functional and metabolic characterization with P-31 MR spectroscopy and cine MR imaging. Radiology. 1991;179:253–259. 63. de Roos A, Doornbos J, Luyten PR, Oosterwaal LJ, van der Wall EE, den Hollander JA. Cardiac metabolism in patients with dilated and hypertrophic cardiomyopathy: assessment with proton-decoupled P-31 MR spectroscopy. J Magn Reson Imaging. 1992;2:711–719. 64. Hardy CJ, Weiss RG, Bottomley PA, Gerstenblith G. Altered myocardial high-energy phosphate metabolites in patients with dilated cardiomyopathy. Am Heart J. 1991;122:795–801. 65. Neubauer S, Krahe T, Schindler R, et al. 31P magnetic resonance spectroscopy in dilated cardiomyopathy and coronary artery disease: altered cardiac high-energy phosphate metabolism in heart failure. Circulation. 1992;86:1810–1818. 66. Neubauer S, Horn M, Pabst T, et al. Contributions of 31P-magnetic resonance spectroscopy to the understanding of dilated heart muscle disease. Eur Heart J. 1995;16(suppl O):115–118. 67. Beer M, Seyfarth T, Sandstede J, et al. Absolute concentrations of highenergy phosphate metabolites in normal, hypertrophied, and failing human myocardium measured noninvasively with (31)P-SLOOP magnetic resonance spectroscopy. J Am Coll Cardiol. 2002;40:1267–1274. 68. Nakae I, Mitsunami K, Omura T, et al. Proton magnetic resonance spectroscopy can detect creatine depletion associated with the progression of heart failure in cardiomyopathy. J Am Coll Cardiol. 2003;42:1587–1593. 69. Neubauer S, Horn M, Cramer M, et al. Myocardial phosphocreatineto-ATP ratio is a predictor of mortality in patients with dilated cardiomyopathy. Circulation. 1997;96:2190–2196. 70. Lodi R, Hart PE, Rajagopalan B, et al. Antioxidant treatment improves in vivo cardiac and skeletal muscle bioenergetics in patients with Friedreich’s ataxia. Ann Neurol. 2001;49(5):590–596. 71. Fraser CD, Chacko VP, Jacobus WE, et al. Early phosphorus 31 nuclear magnetic resonance bioenergetic changes potentially predict rejection in heterotopic cardiac allografts. J Heart Transplant. 1990;9:197–204. 72. Bottomley PA, Weiss RG, Hardy CJ, Baumgartner WA. Myocardial high-energy phosphate metabolism and allograft rejection in patients with heart transplants. Radiology. 1991;181:67–75. 73. Van Dobbenburgh JO, De Groot MC, De Jonge N, et al. Myocardial high-energy phosphate metabolism in heart transplant patients is temporarily altered irrespective of rejection. NMR Biomed. 1999; 12(8):515–524. 74. Evanochko WT, Buchthal SD, den Hollander JA, et al. Cardiac transplant patients response to the (31)P MRS stress test. J Heart Lung Transplant. 2002;21:522–529. 75. Zhang J, Merkle H, Hendrich K, et al. Bioenergetic abnormalities associated with severe left ventricular hypertrophy. J Clin Invest. 1993;92:993–1003. 76. Conway MA, Allis J, Ouwerkerk R, Niioka T, Rajagopalan B, Radda GK. Detection of low phosphocreatine to ATP ratio in failing hypertrophied human myocardium by 31P magnetic resonance spectroscopy. Lancet. 1991;338:973–976. 77. Neubauer S, Horn M, Pabst T, et al. Cardiac high-energy phosphate metabolism in patients with aortic valve disease assessed by 31Pmagnetic resonance spectroscopy. J Investig Med. 1997;45:453–462. 78. Beyerbacht HP, Lamb HJ, van Der Laarse A, et al. Aortic valve replacement in patients with aortic valve stenosis improves myocardial metabolism and diastolic function. Radiology. 2001;219: 637–643. 79. Weiss RG, Bottomley PA, Hardy CJ, Gerstenblith G. Regional myocardial metabolism of high-energy phosphates during isometric exercise in patients with coronary artery disease [see comments]. N Engl J Med. 1990;323:1593–1600. 80. Yabe T, Mitsunami K, Okada M, Morikawa S, Inubushi T, Kinoshita M. Detection of myocardial ischemia by 31P magnetic resonance spectroscopy during handgrip exercise. Circulation. 1994;89:1709–1716. 81. Buchthal SD, den Hollander JA, Merz CN, et al. Abnormal myocardial phosphorus-31 nuclear magnetic resonance spectroscopy in

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38. Bottomley PA. Noninvasive study of high-energy phosphate metabolism in human heart by depth-resolved 31P NMR spectroscopy. Science. 1985;229:769–772. 39. Bottomley PA, Hardy CJ, Roemer PB. Phosphate metabolite imaging and concentration measurements in human heart by nuclear magnetic resonance. Magn Reson Med. 1990;14:425–434. 40. Bottomley PA, Atalar E, Weiss RG. Human cardiac high-energy phosphate metabolite concentrations by 1D-resolved NMR spectroscopy. Magn Reson Med. 1996;35:664–670. 41. Neubauer S, Krahe T, Schindler R, et al. 31P magnetic resonance spectroscopy in dilated cardiomyopathy and coronary artery disease: altered cardiac high-energy phosphate metabolism in heart failure. Circulation. 1992;86:1810–1818. 42. Horn M, Neubauer S, Bomhard M, Kadgien M, Schnackerz K, Ertl G. 31P-NMR spectroscopy of human blood and serum: results from volunteers and patients with congestive heart failure, diabetes mellitus and hyperlipidemia. MAGMA. 1993;1:55–60. 43. Beer M, Sandstede J, Landschu¨tz W, et al. Absolute concentrations of myocardial high-energy phosphate metabolites in normal, hypertrophied and failing human myocardium, measured non-invasively with 31P-SLOOP-magnetic resonance spectroscopy. J Am Coll Cardiol. 2002;40:1267–1274. 44. Yabe T, Mitsunami K, Inubushi T, Kinoshita M. Quantitative measurements of cardiac phosphorus metabolites in coronary artery disease by 31P magnetic resonance spectroscopy [see comments]. Circulation. 1995;92:15–23. 45. Meininger M, Landschutz W, Beer M, et al. Concentrations of human cardiac phosphorus metabolites determined by SLOOP 31P NMR spectroscopy. Magn Reson Med. 1999;41:657–663. 46. Bottomley PA, Ouwerkerk R, Lee RF, Weiss RG. Four-angle saturation transfer (FAST) method for measuring creatine kinase reaction rates in vivo. Magn Reson Med. 2002;47:850–863. 47. Pohmann R, von Kienlin M. Accurate phosphorus metabolite images of the human heart by 3D acquisition-weighted CSI. Magn Reson Med. 2001;45:817–826. 48. Lee RF, Giaquinto R, Constantinides C, Souza S, Weiss RG, Bottomley PA. A broadband phased-array system for direct phosphorus and sodium metabolic MRI on a clinical scanner. Magn Reson Med. 2000;43:269–277. 49. Hetherington HP, Luney DJ, Vaughan JT, et al. 3D 31P spectroscopic imaging of the human heart at 4.1 T. Magn Reson Med. 1995;33: 427–431. 50. Shen W, Asai K, Uechi M, et al. Progressive loss of myocardial ATP due to a loss of total purines during the development of heart failure in dogs: a compensatory role for the parallel loss of creatine. Circulation. 1999;100:2113–2118. 51. Okada M, Mitsunami K, Inubushi T, Kinoshita M. Influence of aging or left ventricular hypertrophy on the human heart: contents of phosphorus metabolites measured by 31P MRS. Magn Reson Med. 1998;39:772–782. 52. Schocke MF, Metzler B, Wolf C, et al. Impact of aging on cardiac high-energy phosphate metabolism determined by phosphorus-31 2-dimensional chemical shift imaging (31P 2D CSI). Magn Reson Imaging. 2003;21:553–559. 53. Pluim BM, Lamb HJ, Kayser HW, et al. Functional and metabolic evaluation of the athlete’s heart by magnetic resonance imaging and dobutamine stress magnetic resonance spectroscopy. Circulation. 1998;97:666–672. 54. Weiss RG, Gerstenblith G, Bottomley PA. ATP flux through creatine kinase in the normal, stressed, and failing human heart. Proc Natl Acad Sci U S A. 2005;102:808–813. 55. Lamb HJ, Beyerbacht HP, van der Laarse A, et al. Diastolic dysfunction in hypertensive heart disease is associated with altered myocardial metabolism. Circulation. 1999;99:2261–2267. 56. Perings SM, Schulze K, Decking U, Kelm M, Strauer BE. Age-related decline of PCr/ATP-ratio in progressively hypertrophied hearts of spontaneously hypertensive rats. Heart Vessels. 2000;15:197–202. 57. Spencer RG, Buttrick PM, Ingwall JS. Function and bioenergetics in isolated perfused trained rat hearts. Am J Physiol. 1997;272: H409–H417. 58. Scheuermann-Freestone M, Madsen PL, Manners D, et al. Abnormal cardiac and skeletal muscle energy metabolism in patients with type 2 diabetes. Circulation. 2003;107:3040–3046. 59. Diamant M, Lamb HJ, Groenevelt Y, et al. Diastolic dysfunction is associated with altered myocardial metabolism in asymptomatic normotensive patients with well-controlled type II diabetes mellitus. J Am Coll Cardiol. 2003;42:328–335.

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women with chest pain but normal coronary angiograms. N Engl J Med. 2000;342:829–835. Johnson BD, Shaw LJ, Buchthal SD, et al. Prognosis in women with myocardial ischemia in the absence of obstructive coronary disease. Circulation. 2004;109:2993–2999. Flameng W, Vanhaecke J, Van Belle H, Borgers M, De Beer L, Minten J. Relation between coronary artery stenosis and myocardial purine metabolism, histology and regional function in humans. J Am Coll Cardiol. 1987;9:1235–1242. Kalil-Filho R, de Albuquerque CP, Weiss RG, et al. Normal highenergy phosphate ratios in stunned human myocardium. J Am Coll Cardiol. 1997;30:1228–1232. Beer M, Sandstede J, Landschutz W, et al. Altered energy metabolism after myocardial infarction assessed by 31P- MR-spectroscopy in humans. Eur Radiol. 2000;10:1323–1328. Beer M, Spindler M, Sandstede JJ, et al. Detection of myocardial infarctions by acquisition-weighted 31P-MR spectroscopy in humans. J Magn Reson Imaging. 2004;20:798–802. Bottomley PA, Weiss RG. Non-invasive magnetic-resonance detection of creatine depletion in non-viable infarcted myocardium. Lancet. 1998;351:714–718. Parrish TB, Fieno DS, Fitzgerald SW, Judd RM. Theoretical basis for sodium and potassium MRI of the human heart at 1.5 T. Magn Reson Med. 1997;38:653–661. Pabst T, Sandstede J, Beer M, et al. Optimization of ECG-triggered 3D (23)Na MRI of the human heart. Magn Reson Med. 2001;45:164–166. Sandstede JJ, Pabst T, Beer M, et al. Assessment of myocardial infarction in humans with (23)Na MR imaging: comparison with cine MR imaging and delayed contrast enhancement. Radiology. 2001;221:222–228. Sandstede JJ, Hillenbrand H, Beer M, et al. Time course of 23Na signal intensity after myocardial infarction in humans. Magn Reson Med. 2004;52:545–551. Ashrafian H, Redwood C, Blair E, Watkins H. Hypertrophic cardiomyopathy: a paradigm for myocardial energy depletion. Trends Genet. 2003;19:263–268. Spindler M, Saupe KW, Christe ME, Seidman CE, Seidman JG, Ingwall JS. A murine model of familial hypertrophic cardiomyopathy shows a markedly impaired response to inotropic stimulation [abstract]. Circulation. 1996;94/8(Suppl I):I–433. Sakuma H, Takeda K, Tagami T, et al. 31P MR spectroscopy in hypertrophic cardiomyopathy: comparison with Tl-201 myocardial perfusion imaging. Am Heart J. 1993;125: 1323–1328.

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Magnetic Resonance Assessment of Myocardial Oxygenation Rohan Dharmakumar and Debiao Li

Under physiologic conditions, myocardial blood flow, oxygen consumption (MVO2), and myocardial mechanical function are intimately related. Therefore, it is not surprising that the most common disease processes involving the heart manifest themselves as imbalances between myocardial oxygen supply and demand. As a consequence, the noninvasive assessment of imbalances in myocardial oxygen supply and demand, particularly on a regional basis, is of critical importance in both cardiovascular research and clinical cardiology. The noninvasive quantification of MVO2 was not possible until it was shown that positron emission tomography (PET), using 11C-acetate, permits accurate quantification of MVO2.1–4 Using this approach, numerous investigators have demonstrated the salutary effects of restoring nutritive perfusion on MVO2 and cardiac function and the importance of preserving MVO2 as a descriptor and probable determinant of myocardial viability in both acute and chronic ischemic disease.5–8 However, PET studies are limited by relatively poor spatial resolution, limited availability, and use of potentially harmful ionizing radiation. Magnetic resonance imaging (MRI) has become the preferred method for many biomedical applications. It is noninvasive, does not require iodinated contrast media or ionizing radiation, and is widely available. MRI can also provide functional and anatomic information in the same setting. Multiple cardiovascular magnetic resonance (CMR) applications have been developed, including anatomic imaging of the heart and great vessels, coronary artery imaging, myocardial infarct imaging with late gadolinium enhancement, myocardial wall motion assessment, and first pass myocardial perfusion measurement. Over the past decade, efforts have also been made to use CMR to determine regional myocardial blood oxygenation levels.9–30 The blood oxygenation state of the myocardium reflects the combined effects of myocardial blood flow and oxygen extraction (which together reflect MVO2). Thus, a change in myocardial blood oxygenation secondary to imbalances between oxygen supply and demand would be useful in assessing disease processes such as coronary artery disease, which leads to impaired coronary vascular reserve. Noninvasive assessment of myocardial venous blood oxygenation may permit the measurement of oxygen extraction. When coupled with flow, these data would allow for measurement of MVO2. Anatomic, functional, and metabolic information can then be obtained in a single CMR study, thereby providing a comprehensive examination for the diagnosis of coronary artery disease and the evaluation of therapies designed to improve the balance between myocardial blood flow and oxygen demand.

For the myocardial system, if we assume that the blood oxygen saturation of hemoglobin in the arterial blood is 100%, whereas that in the venous blood is Y, then the oxygen consumption of the myocardium (MVO2) can be estimated by Fick’s law as MVO2 ¼ F Hct ð1 YÞ where F (in milliliters per gram per minute) is the blood flow to the myocardium and Hct (in percent) is the hematocrit of the blood.31 By assessing changes in F and Y, one can evaluate the changes in MVO2. Another potential application of myocardial venous blood oxygenation assessment is found in the noninvasive evaluation of myocardial perfusion reserve, defined as the ratio between the peak myocardial perfusion rate at maximum vasodilation and basal conditions.32 Under pharmacologic stress induced by vasodilators, such as dipyridamole or adenosine, normal coronary blood flow will increase by severalfold, whereas the oxygen consumption will remain relatively unaltered. As a result, the myocardial venous blood oxygen saturation will be increased. However, if the coronary artery is partially or completely occluded, the increase in blood flow to the myocardial region subtended by the artery will be significantly lower in relation to the healthy myocardium. This observation is illustrated in Figure 42-1. Such perfusion abnormalities lead to regional differences in myocardial venous blood oxygenation, which may allow for the assessment of the functional significance of coronary artery disease. Extensive research has been conducted to evaluate myocardial perfusion using CMR by the administration of exogenous contrast media.33–36 The potential advantage of this method is that multiple examinations can be performed consecutively without the need to wait for the clearance of exogenous contrast materials between baseline and stress studies. In addition, it does not require the temporal resolution required to capture the first passage of the contrast media as in first pass perfusion imaging, permitting increased myocardial coverage and acquisition of images with higher spatial resolution. Perhaps the most significant potential advantage of blood-oxygen-sensitive myocardial imaging over first pass methods is that it can provide a direct knowledge of myocardial oxygenation, the fundamental substrate of cardiac function. In this chapter, we will first provide an overview of the basic biophysical concepts that allow for the assessment of myocardial changes in blood oxygenation. Following this, we will summarize the preclinical and clinical studies Cardiovascular Magnetic Resonance 569

42 MAGNETIC RESONANCE ASSESSMENT OF MYOCARDIAL OXYGENATION

CHAPTER 42

Hypertrophy ↑ HR ↑ Preload Etc.

50% Maximum vasodilation Coronary blood flow

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5.0

70% 3.0 80%

↑ Flow prior to vasodilation

85% 80%

1.0

85%

50% 70%

90%

Coronary pressure Figure 42-1 A schematic showing the relationship between the coronary flow and coronary arterial pressure. The solid curve represents the normal relationship. At a constant level of myocardial metabolic demand, coronary flow is maintained constant over a wide range of coronary pressure, between the bounds of maximum coronary vasodilation and constriction (dashed curves). The solid circle represents the normal operating point under basal conditions; the solid triangle is the flow observed at the same pressure during maximum vasodilation. Myocardial flow reserve is the ratio of flow during vasodilation to that measured before vasodilation. Note that the flow reserve decreases in a nonlinear manner with reduction of coronary pressure (or coronary artery stenosis). Also note that hypertrophy, increased heart rate, and increased preload all decrease the coronary flow reserve. Source: Adopted from Klocke FJ: Measurements of coronary flow reserve: Defining pathophysiology versus making decisions about patient care. Circulation 1987;76(6):1183–1189.

in the assessment of myocardial oxygenation with a focus on recent advances. We will conclude with a brief look at the emerging techniques and an outlook on oxygensensitive myocardial imaging.

MYOCARDIAL BLOOD OXYGEN LEVEL DEPENDENT CONTRAST Blood is a magnetically inhomogeneous medium in which the magnetic susceptibility of red blood cells is strongly dependent on the blood oxygen saturation (%O2), defined as the percentage of hemoglobin that is oxygenated.37–38 Since the susceptibility of blood plasma is generally invariant, the cooperative binding of oxygen to the heme molecules in the red blood cells results in a detectable susceptibility difference 570 Cardiovascular Magnetic Resonance

between plasma and the red blood cells. This susceptibility variation gives rise to local magnetic field inhomogeneities, resulting in local frequency variations that lead to changes in T2* and apparent T2 of whole blood.39–40 This observation has allowed for the acquisition of oxygen-sensitive images permitting the discrimination between arteries and veins.41 Its utility for detecting chronic mesenteric ischemia42 and the identification and quantification of cardiac shunts associated with congenital abnormalities have also been demonstrated.43 An extension of this phenomenon into the microcirculation also provides opportunities for assessing myocardial oxygenation changes. In the myocardium, nearly 90% of the blood volume is within the capillaries;44 accordingly, our discussions will be limited to capillary beds. A change in blood oxygenation in the capillary bed leads to changes in magnetic field variations between the red blood cells and plasma and between the intravascular and extravascular spaces. Following the excitation of the magnetization onto the transverse plane, these field variations cause the spins to lose coherence, leading to a decay of the magnetic resonance (MR) signal.45–50 In particular, the severity of the field variation due to changes in blood oxygen saturation directly determines the rate of loss of the spin coherence (MR signal). This is referred to as the myocardial blood oxygenation level dependent (BOLD) effect. Effectively, this implies that when the capillaries contain deoxygenated blood, all other conditions being the same, the myocardial MR signal associated with the deoxygenated state (resting condition) will be lower than that in the hyperemic state when the capillary oxygenation is substantially greater (30% (resting) versus 80% (hyperemic)). This allows for detecting regional myocardial oxygen deficits as regions of signal loss with imaging sequences that are sensitive to local field inhomogeneities.9 In addition, the sensitivity of the MR signal to the BOLD effect is dependent on the blood volume and choice of pulse sequences used.45–50 Traditionally, gradient and spin echo sequences have been used to probe these effects and are well described by Bauer and colleagues.45,50 The motivation for implementing BOLD effect techniques for evaluation of myocardial venous blood oxygenation9–15 was based on its previous success with brain functional imaging.51–54 However, the heart has a larger blood volume fraction than the brain (10% versus 4%). In addition, the venous blood oxygen saturation in the heart is approximately 30%, compared to approximately 60% in the brain.9 This allows for a wider range of signal change with stimulated flow in the myocardium. In contrast to BOLD brain imaging, the challenge for myocardial BOLD imaging has been image artifacts caused by cardiac motion due to both cardiac and respiratory cycles; specifically, difficulties arise from pulsatile blood flow in the cardiac chambers and the aorta as well as susceptibility variations between tissue and air in the thoracic space.

THE ROLE OF VASODILATORS IN THE ASSESSMENT OF MYOCARDIAL OXYGENATION At the current state of myocardial BOLD imaging, vasodilators are generally accepted to be essential for assessing myocardial blood oxygenation. Their importance has been

oxygen saturation clearly has the dominant effect over increased blood volume, but the apparent R2* change as a function of %O2 is reduced because of the accompanied blood volume effect. In contrast, during hypoxia, both % O2 in coronary sinus and the blood volume fraction increase, and their effects enhance each other. As a result, the apparent change in R2* as a function of %O2 is greater than that if blood volume fraction remains the same. These studies showed that by measuring myocardial blood volume fraction changes using technetium-99m-labeled red blood cells at each interventions and correcting their effects on myocardial R2*, a more linear relationship can be found between R2* and the blood oxygen saturation. Thus, an accurate assessment of myocardial oxygen saturation using CMR will likely require a correction for blood volume.

AN OVERVIEW OF MYOCARDIAL BOLD CARDIOVASCULAR MAGNETIC RESONANCE IMAGING IN THE PRECLINICAL SETTING To date, a number of in vivo animal and isolated heart studies have demonstrated that vasodilator interventions that alter the blood oxygenation levels result in myocardial signal changes.9–13,21–22,24–30 One of the first studies to demonstrate the feasibility of assessing myocardial oxygenation changes via BOLD effect was performed in a rat model. This study showed significant signal loss in both the left ventricular (LV) chamber and the myocardium during apnea.9 A subsequent study in an isolated rabbit heart model confirmed a substantial correlation between the gradient echo image intensity of the myocardium and deoxyhemoglobin concentration levels.13 In large animal models, gradient echo myocardial signal enhanced significantly after the infusion of dipyridamole,11 presumably because of an increase in myocardial blood flow in the absence of a corresponding increase in oxygen demand, resulting in a decrease in myocardial venous blood oxygen saturation. Other researchers have independently validated these studies in large animal models with T2-prepared methods.22,27–28 Canine studies also suggest that it is possible to derive myocardial oxygen extraction fraction and quantitative perfusion reserve values with BOLD imaging.24–26 Perhaps the greatest expectation of CMR BOLD imaging is in detecting regional differences in myocardial oxygenation caused by focal coronary artery disease. Figure 42-2 shows a schematic of the coronary vessels that supply the different myocardial territories of the mid-left ventricle in humans.18 Luminal narrowing of the coronary vessels (see Fig. 42-2) can lead to regional perfusion deficits in the associated myocardium territories during pharmacologic stress. Induction of stress can be achieved with the administration of a coronary vasodilator, such as dipyridamole or adenosine, which increases myocardial blood flow without much change in myocardial oxygen consumption. While Cardiovascular Magnetic Resonance 571

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demonstrated on human subjects with multi-gradient echo methods using two different pharmacologic stress agents: dipyridamole and dobutamine.14,16 Both agents induce an increase in myocardial blood flow but with differing effects on myocardial venous blood oxygenation. Dipyridamole is a potent coronary vasodilator that typically induces a threeto fourfold increase in myocardial perfusion but induces minimal effects on myocardial oxygen consumption.55 As a result, myocardial venous blood oxygen saturation increases as oxygen supply (blood flow) exceeds demand (oxygen consumption). In contrast, dobutamine is a potent beta-agonist with a primary pharmacologic effect to increase cardiac work.55 This results in an increase in myocardial oxygen consumption, which triggers an increase in myocardial perfusion. Thus, oxygen supply and demand remain largely balanced, and there is little change in myocardial venous blood oxygen saturation.56 Vasodilators have also allowed for direct demonstration of the in vivo correlation between myocardial MR signal response (via changes in R2* (or 1/T2*)) and venous blood oxygen saturation. In a well-controlled canine model, a wide range of global myocardial venous blood oxygen saturation levels were created. Hyperemic conditions were induced by the intravenous administration of dipyridamole and dobutamine. To induce hypoxemia, the oxygen content of the inspired gas was reduced by ventilating dogs with a mixture of 10% oxygen and 90% nitrogen, which reduced the oxygen saturation in both arteries and veins. To correlate myocardial R2* with global venous blood oxygenation, venous blood oxygen saturation levels were measured directly by coronary sinus sampling. Myocardial perfusion was quantified by the administration of radiolabeled microspheres. After animal preparation and scout scans to locate a short-axis view of the heart, measurements of myocardial R2* were obtained at baseline, during and after infusion of dipyridamole and dobutamine, and when the dogs were ventilated with hypoxic air. Paired arterial and coronary sinus blood samples were withdrawn at the six different stages of the study. Blood oxygen saturation levels were measured by using a blood gas analyzer interfaced with an oximeter. Coronary sinus blood oxygen saturation levels ranged from 9% to 80% with experimental interventions with dipyridamole, dobutamine, or hypoxic air. Administration of dipyridamole and dobutamine and ventilation of hypoxic air all increased myocardial perfusion significantly, but significant decrease in myocardial R2* occurred only after dipyridamole infusion, which indicates that myocardial R2* is not a direct function of myocardial perfusion. The relationship between the changes of myocardial R2* from baseline and the %O2 in the coronary sinus showed a linear regression (r ¼ 0.84), indicating a strong correlation between myocardial R2* and %O2 in the coronary sinus. Both dipyridamole and hypoxic air increase myocardial blood volume fraction in excess of 50%. However, their effects on R2* manifest differently. With administration of dipyridamole, %O2 in coronary sinus increases, which leads to a decrease in R2*. However, the blood volume fraction in the myocardium also increases, which increases the hematocrit content of a voxel, which tends to increase myocardial R2*; this is the opposite of the effect of increased oxygen saturation. Since a decrease in myocardial R2* was observed in these studies, increased

FUNCTIONAL CARDIOVASCULAR DISEASE

Anterior Anteroseptal Lateral

Inferoseptal Inferolateral

LAD RCA RCA, LCX when left dominant LCX, RCA, when right dominant

Inferior Figure 42-2 A schematic representation of the myocardial segments of the midventricular section with the assigned coronary territories. LAD, left anterior descending coronary artery; LCX, left circumflex coronary artery; RCA, right coronary artery. Source: Adapted from Friedrich MG, Niendorf T, Schulz-Menger J, Gross CM, Dietz R: Blood oxygen level-dependent magnetic resonance imaging in patients with stress-induced angina. Circulation 2003; 108(18):2219–2213.

vasodilators are effective at increasing blood flow to the myocardium perfused by healthy coronary vessels, myocardial blood flow to territories distal to the coronary artery stenosis is limited. The reduction in blood flow to the affected region decreases the oxygen saturation of hemoglobin and aids in delineating regional perfusion differences. Although it could be anticipated that myocardial oxygenation in the affected region may also be increased in comparison to the resting state, the increase may be smaller in normal myocardium. Over the past decade, animal and patient studies have been performed with the intent of validating the hypothesis that regional flow deficits due to coronary artery disease in the presence of adenosine or dipyridamole can provide an alternative to first pass perfusion technique based on the BOLD effect. Some of the earliest studies of the utility of BOLD imaging to understand the perfusion deficits associated with coronary occlusion were performed with gradient echo sequences at B0 ¼ 2 to 4.7 T. These studies showed that the occlusion of the left anterior descending artery (LAD) resulted in signal reduction in corresponding regions of the myocardium perfused by the LAD.11,12 These initial studies were followed by T2* mapping of the myocardium, which demonstrated that the myocardial regions with perfusion deficits also show preferential reduction in T2* under vasodilation. T2-prepared gradient echo methods in canine models have also shown that it may be possible to extend the myocardial BOLD technique to three-dimensional (3D) imaging and acquire multiple slices with BOLD contrast.21 However, poor image quality and signal-to-noise ratio (SNR), long scan times, and inadequate sensitivity are notable limitations of the gradient echo technique at 1.5 T. T2-prepared methods have been proposed as an approach to overcome some of the limitations of myocardial BOLD. However, the early efforts of this method employing spiral readout remains challenging, owing to off-resonancebased blurring.22–23 A subsequent T2-preparation method with a steady-state free precession (SSFP) readout has shown significant improvements in SNR, image quality, and scan times over the gradient echo techniques29–30 and good sensitivity for evaluating myocardial BOLD signal 572 Cardiovascular Magnetic Resonance

changes. This technique has been employed in canine models with varying degrees of stenoses of the left circumflex arteries (LCX) in the presence of systemic adenosine infusion to demonstrate the effect of perfusion abnormalities with BOLD effect.30 The degree of stenosis was adjusted on the basis of Doppler flow measurements obtained from a flow probe placed distal to the occluder. Results showed a visually discernible BOLD signal drop in the inferior and posterior walls supplied by the LCX as well as a close correspondence between flow reductions in the same region as identified by the first pass perfusion images and microsphere flow maps.

CLINICAL MYOCARDIAL BOLD IMAGING Successful demonstrations of the BOLD effect in animal models have provided the impetus for myocardial BOLD imaging in patients with coronary artery disease, although the method is mainly limited to gradient echo–based BOLD imaging. In one of the first patient studies, Friedrich and colleagues18 compared the T2*-weighted BOLD CMR method with single-photon emission computed tomography (SPECT) with the aim of visualizing regional oxygen deficits under pharmacologic stress. This study was performed in 25 patients (23 patients with >50% stenosis, 10 of whom had severe stenosis, that is, >75%). Results demonstrated that statistically significant changes in suspected regions of the myocardium were apparent in patients with severe stenosis. Receiver operating characteristic (ROC) analysis of both BOLD CMR and thallium SPECT methods showed that the area under the ROC curve is 0.66 and 0.73, respectively. These results indicate that both methods have limitations in identifying regional perfusion deficits and that the limitations are attributable to suboptimal resolution of SPECT images and the well-known image artifacts associated with gradient-recalled T2*-weighted methods in the heart, as discussed earlier. In addition to demonstrating the feasibility of the BOLD method for studying coronary artery disease in patients (Fig. 42-3), these studies demonstrated that improved CMR methods are needed. Another clinical study by Wacker and colleagues19 examined the possibility of evaluating oxygen-sensitive changes in patients with coronary artery disease without pharmacologic stress. Their study included a control group of 16 volunteers (age 31 10 years old) with no history of cardiovascular disease and 16 patients (63 9 years old) with single-vessel coronary artery disease (degree of stenosis > 70%) identified by X-ray coronary angiogram. Within the control group, T2* mapping of the mid-LV myocardium was homogenous (35 3 msec) and increased uniformly under the influence of dipyridamole (40 4 msec). However, in patients with coronary artery disease, the suspected regions of myocardium were at a lower T2* than the healthy regions (31% 9%). Furthermore, this noted deviation increased under the influence of dipyridamole (43% 21%). A representative set of short-axis T2 maps of two patients from this study is shown in Figure 42-4. This study also followed patients who underwent coronary bypass grafting or

Pat. #3 Critical stenosis RCA

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Pat. #2 Stenosis LAD, R. diag., and R. marg.

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Figure 42-3 Gradient echo planar BOLD-CMR images obtained before (top) and during (bottom) adenosine infusion and corresponding thallium SPECT images from three representative patients. Left, Patient without a 0% to 25% stenosis. The mean signal intensity (SI) change of the six segments during adenosine was 2.85 0.33%. Middle, Patient with a 50% stenosis of the left anterior descending coronary artery, a 90% stenosis in a large diagonal branch, and a 99% stenosis in the second marginal branch. Although there is no obvious visible change of SI, the quantitative evaluation showed a SI decrease of 0.2% to 7.6% in the segments related to the stenotic arteries. Right, Patient with a critical stenosis of the right coronary artery (RCA). The SI decrease in the inferior segment during adenosine was 7.2%. In this case, a signal drop can be visually evident (see arrow). Source: Adapted from Friedrich MG, Niendorf T, Schulz-Menger J, Gross CM, Dietz R: Blood oxygen level-dependent magnetic resonance imaging in patients with stress-induced angina. Circulation 2003;108(18):2219–2223.

stenting of the diseased vessel to reestablish flow to the affected regions of the myocardium. Their findings showed that following the intervention, the T2* homogeneity of the myocardium was improved, likely owing to the improved perfusion facilitated by the more patent coronary vessels. The finding that even in the absence of provocative stress, it is possible to visualize regional BOLD differences in patients with coronary artery disease as an important finding. This observation has been linked to capillary recruitment in the distal beds of stenotic coronary vessels, which results in an increase in blood volume and thus an elevation of deoxyhemoglobin concentration per unit of imaging voxel in the affected

50

15 Figure 42-4 T2* maps of two different patients with high-grade stenosis of the left anterior descending coronary artery (by X-ray) and septal wall motion abnormalities (by stress echocardiography). Areas with reduced T2* values in the anteroseptal and septal regions are clearly detectable. Bright to dark colors reflect T2* values from 15 to 50 msec. Source: Adapted from Wacker CM, Hartlep AW, Pfleger S, Schad LR, Ertl G, Bauer WR: Susceptibilitysensitive magnetic resonance imaging detects human myocardium supplied by a stenotic coronary artery without a contrast agent. J Am Coll Cardiol 2003;41(5):834–840.

territories.19,57 However, this hypothesis has yet to be validated through controlled studies. Moreover, it is important to note that while the elimination of the provocative stress is a welcome change for assessing regional myocardial perfusion deficits, this observation must be carefully evaluated. Regional signal differences in the myocardium, especially with gradient-recalled (T2*weighted) imaging, in addition to indicating perfusion deficits, may also stem from spatially varying static field inhomogeneities set up by the heart-lung interface or thalassemia,58 a genetic disorder that leads to the accumulation of iron in the myocardium. Myocardial perfusion reserve changes in hypertensive hypertrophic patients have also been studied with BOLD imaging using the T2* maps.20 This study included 10 patients (43 4 years old) with hypertension (blood pressure > 140/90 mmHg), hypertrophic hearts (assessed by two-dimensional echocardiography), and no coronary artery stenosis and 9 healthy subjects with no heart disease. The results showed that the change in T2* between basal and dipyridamole-infused states (DT2*) was approximately threefold lower in hypertrophic patients than in healthy controls. This observation is consistent with previous reports32 that myocardial perfusion reserves are reduced in people with hypertrophic hearts (refer to Fig. 42-1). From the perspective of myocardial oxygenation, in hypertrophic patients, vasodilatory mechanisms are compromised depending on the severity of the disease and are not fully effective at altering the oxygenation of the myocardium as in a healthy heart. This inability to alter blood oxygenation between the basal and dipyridamole-infused states in hypertrophic patients is likely the reason for the substantially different T2* between healthy subjects and hypertrophic patients. Cardiovascular Magnetic Resonance 573

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Pat. #1 No significant stenosis

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EMERGING TECHNIQUES FOR OXYGEN-SENSITIVE MYOCARDIAL CARDIOVASCULAR MAGNETIC RESONANCE IMAGING More recently, systematic studies have demonstrated that SSFP imaging can be used in the assessment of oxygen-sensitive imaging. Over the past decade, SSFP imaging has become invaluable in CMR. Superior temporal resolution and SNR are the distinguishing features of SSFP. Using this method, controlled in vitro studies with blood samples oxygenated to various levels revealed that when the off-resonance effects are minimized through the appropriate choice of flip angle and phase-cycling scheme, oxygen-sensitive contrast can be realized in whole blood in a TR-dependent manner.59 The results showed that relatively long TRs (compared to those that are used in conventional SSFP imaging) are necessary to establish oxygen-sensitive contrast in SSFP images in blood. These studies were extended to intravascular 3D peripheral angiographic methods aimed at discriminating arteries and veins based on %O2 differences between the vessels.60 The advantages of the SSFP method over other oxygen-sensitive CMR methods are the short scans, significant increase in SNR, reduced heat deposition, and improved oxygen sensitivity.

A

B

Studies demonstrating the feasibility of detecting oxygensensitive signal changes in microcirculations, specifically in kidneys and dorsal muscles of rabbits, have also been achieved by altering the systemic %O2. The results showed that BOLD SSFP contrast is directly linked to field strength, blood volume, and baseline microcirculatory oxygen saturation levels.61 The mechanisms that determine oxygen contrast were connected to a fast exchange of spins between the plasma and red blood cells, as well as the perivascular gradients due to the susceptibility shifts between the intravascular and extravascular pools of spins. To understand the relationship between oxygen sensitivity and CMR scan parameters for tissue, ischemic leg cuff studies have been performed. The results show that flip angle and TR are important parameters that determine oxygen contrast in SSFP images of skeletal muscle.62 In particular, results indicate that relatively long TRs are necessary for ensuring realizing oxygen contrast. These studies were followed up with SSFP-based myocardial BOLD CMR in canine models. Using a similar animal model as was described previously,30 controlled studies, which allow for the occlusion of the LCX in the presence of adenosine infusion were performed. Two-dimensional (2D) cine SSFP imaging with a relatively long TR (6.3 msec) was used to demonstrate that occlusion of the LCX leads to regional myocardial oxygen deficit in the coronary territory supplied by LCX. The results were compared to the first pass perfusion technique and validated with microsphere-based perfusion analysis (Fig. 42-5). Two-dimensional BOLD SSFP

C –5%

D

E

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F

Figure 42-5 End-systolic short axis images obtained from an instrumented canine showing regional myocardial BOLD contrast obtained with two-dimensional steady-state free precession (SSFP) imaging at baseline (no adenosine and no stenosis) (A), prestenosis (with adenosine and no stenosis) (B), mild stenosis (C), and severe stenosis (D). First pass perfusion image (E) at severe stenosis and percent microsphere flow difference between stenosis and nonstenosis states (F) are shown for reference. All stenosis studies were performed with adenosine. Note the discriminating signal loss in the LCX territories (subtended by arrows) in image D during stenosis of the LCX and the close correspondence between the first pass perfusion and microsphere-based flow difference map. Source: Adapted from Dharmakumar R, Mangalathu Arumana J, Larson AC, Chung Y, Wright D, Li D. Cardiac phase-resolved blood oxygen-sensitive steady-state free precession MRI for evaluating the functional significance of coronary artery stenosis. Invest Radiol 2007;42(3):180–188. 574 Cardiovascular Magnetic Resonance

MS

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Mild stenosis

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ES

Recently, 2D SSFP BOLD imaging has also been used in a feasibility study in patients (N ¼ 9) suspected of coronary artery disease.64 Patients were identified on the basis of positive thallium SPECT and coronary angiograms. Endsystolic frames obtained before and during administration of adenosine were analyzed in a segmental fashion by using the six-segment model recommended by the American Heart Association. On the basis of thallium SPECT, segments were

A Figure 42-6 Cardiac phase-resolved BOLD CMR with two-dimensional steady-state free precession (SSFP) imaging in canines. On the left panel (A), typical cardiac phase-resolved blood oxygen level dependent (BOLD) images (early systole (ES), midsystole (MS), late systole (LS), and late diastole (LD)) showing regional myocardial oxygen deficits in the left circumflex artery (LCX) territory during mild and severe stenosis of the LCX (with adenosine). Note that the extent of signal loss in the LCX territory is related to the extent of LCX stenosis. Baseline and prestenosis images are also shown at the same cardiac phases for reference. (Continued) Cardiovascular Magnetic Resonance 575

42 MAGNETIC RESONANCE ASSESSMENT OF MYOCARDIAL OXYGENATION

method accurately predicted the regional myocardial flow deficit region identified by the first pass technique employing an exogenous contrast media.63 SSFP method is expected to provide additional benefits over T2-prepared methods, because it has the capacity to allow for cardiac phaseresolved BOLD imaging, permitting increased confidence for evaluating myocardial BOLD signal changes in the presence of a coronary artery stenosis (Fig. 42-6).

% change in microsphere-based ? ? severe stenosis relative to prestenosis

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Figure 42-6—cont’d The plot on the right panel (B) shows the percent change in SSFP-based BOLD contrast at midsystole (open circles), end systole (open triangles), and late diastole (open squares) and the associated microsphere-based flow changes (closed squares) observed relative to prestenosis over all studies. Note the close correspondence between the magnetic resonance (MR) and microsphere data throughout the myocardium (sectors 1 through 8) at all the cardiac phases analyzed. The MR and microsphere measures of relative signal changes are plotted as mean standard error. The dotted black curves (MR) and solid gray curves (microsphere) are provided for visual guidance. Source: Adapted from Dharmakumar R, Mangalathu Arumana J, Larson AC, Chung Y, Wright D, Li D. Cardiac phase resolved blood oxygensensitive steady-state free precession MRI for evaluating the functional significance of coronary artery stenosis. Invest Radiol 2007;42(3):180–188.

classified as healthy, mildly affected, or severely affected. Segmental SSFP signal intensities at rest and stress were measured, and the signal intensity ratio (stress/rest) was calculated. Figure 42-7A shows representative examples of

SSFP BOLD

FPP

mid-LV short axis SSFP-based myocardial BOLD CMR images obtained under rest and stress from a patient with 70% narrowing of the LAD. For reference, the corresponding mid-LV short axis first pass perfusion and thallium SPECT images are also shown. Note the discriminating signal loss in the stress images (relative to rest) in the anterior zones of the myocardium and its close correspondence to first pass and thallium SPECT images. Statistical results (Fig. 427B) show that there are significant differences in stress/rest values computed from healthy, mildly affected, and severely affected segments (p < 0.05). Further studies are warranted to establish the sensitivity and specificity of the technique. Although the advantage of using high-field CMR for enhancing the sensitivity of myocardial BOLD imaging is well established in animals,35–36 it has not been validated in humans. To examine whether 3.0-T BOLD imaging can provide increased sensitivity for detecting myocardial oxygenation changes, theoretical and experimental canine studies were performed at 1.5 T and 3.0 T65 (Fig. 42-8). Theoretical results showed that at 3.0 T, a 3.0-fold increase in oxygen sensitivity could be expected in comparison to 1.5 T. Experimental canine studies showed that a 2.5-fold 0.2-fold increase in BOLD sensitivity is possible at 3.0 T relative to 1.5 T. On the basis of the relationship between BOLD signal changes and microsphere perfusion, it was found that the minimum regional perfusion difference that can be detected with SSFP-based myocardial BOLD imaging at 1.5 T and 3.0 T were 2.9 and 1.6, respectively. These findings suggest that SSFP-based myocardial BOLD imaging at 3.0 T may have the necessary sensitivity to detect the clinically required minimum flow difference of 2.0 that is achievable with first pass perfusion methods. However, additional improvements in shimming techniques are expected to be necessary prior to successful clinical translation.

Thallium 1.20

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FUNCTIONAL CARDIOVASCULAR DISEASE

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0.95 Severely affected Mildly affected

A

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Myocardial segments

Figure 42-7 Patient studies with two-dimensional steady-state free precession (SSFP) blood oxygen level dependent (BOLD) at 1.5 T. Midventricular short axis SSFP BOLD, first pass perfusion (FPP), and thallium SPECT images obtained from a patient with 70% stenosis of the left anterior descending (LAD) coronary artery at rest and stress (A). The windowing and leveling of images obtained at rest and stress are the same. Myocardial signal in the rest BOLD, FPP, and SPECT images are relatively homogenous. However, under stress, the territory supplied by the LAD (larger arc subtended by arrows) does not increase in the BOLD images as expected. This pattern of regional signal differences is also evident in the FPP and SPECT images. Statistical results from the myocardial BOLD signal analysis showed that significant differences in stress/rest values exist between healthy and affected regions (B). In comparison to healthy segments, the stress/rest values of affected regions are lower, consistent with previous findings in animals that SSFP signals obtained under pharmacologic stress are significantly reduced in myocardial territories supplied by stenotic arteries. Source: Adapted from Dharmakumar R, Green JD, Flewitt J, Voehringer M, Filipchuk NG, Li D, Friedrich MG. Blood oxygen-sensitive SSFP imaging for probing the myocardial perfusion reserves of patients with coronary artery disease: A feasibility study. SCMR 2008 (Los Angeles, USA). 576 Cardiovascular Magnetic Resonance

A

B

C

D

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Figure 42-8 Short axis two-dimensional steady-state free precession (SSFP) magnetic resonance images obtained at 1.5 T (top row, A–C) and at 3 T (bottom row, A0 –C0 ) in canines. Images A and A0 are SSFP images obtained without stenosis, images B and B0 are SSFP images at systole under severe stenosis of the LCX, and images C and C0 are the corresponding first pass images acquired under stenosis of similar extent as in B and B0 . Images D and D0 represent the spatial map (scale provided by the gray-scale bar) of percent difference in microsphere-based regional flow between prestenosis and severe stenosis in the presence of adenosine infusion at 1.5 T and 3.0 T, respectively. The arrows subtend the suspected regions (left circumflex artery territory) where the perfusion deficits are expected to develop as a result of left circumflex artery stenosis in dogs. Note the discriminating signal loss in these regions in images B and B0 and the close correspondence between the first pass perfusion (C and C0 ) and microsphere-based flow difference maps (D and D0 ). Source: Adapted from Dharmakumar R, Mangalathu Arumana J, Tang R, Harris K, Zhang Z, Li D. Assessment of regional myocardial oxygenation changes in the presence of coronary artery stenosis with balanced SSFP imaging at 3.0 T: theory and experimental evaluation in canines. J Magn Reson Imaging 2008;27(5):1037–1045.

FUTURE OF MYOCARDIAL BOLD CARDIOVASCULAR MAGNETIC RESONANCE IMAGING While recent advances in cardiac BOLD imaging in animal studies are promising, the newer techniques need to be validated and extended through prospective clinical studies in patients. Next, all myocardial BOLD imaging methods currently require multiple breath holds for full myocardial coverage. Since typical myocardial oxygenation changes are assessed in the presence of pharmacologic vasodilation, there are pragmatic limitations with 2D approaches that reduce the achievable myocardial coverage within the typical 4- to 6-minute adenosine infusion protocol. Moreover,

multiple breath holds require multiple recovery periods for patients during adenosine infusion, further reducing the time available for imaging. It is expected that free-breathing methods, such as self-gating,66 combined with parallel imaging strategies67–70 can significantly enhance the myocardial coverage and reduce patient discomfort associated with breath holding during the adenosine protocol. Finally, the preliminary findings by Wacker and colleagues19 that chronic coronary artery disease may be assessed without provocative stress needs to be systematically investigated with newer BOLD MRI methods that provide substantial improvement in image quality. Successful adoption of recent advances through clinical studies and additional technical improvements can propel myocardial BOLD MRI to becoming a powerful noninvasive diagnostic method in the early detection and post-treatment monitoring of ischemic heart disease.

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6. Gropler RJ, Siegel BA, Sampathkumaran KS. Dependence of recovery of contractile function on maintenance of oxidative metabolism after myocardial infarction. J Am Coll Cardiol. 1992;19:989–997. 7. Gropler RJ, Geltman EM, Sampathkumaran KS. Functional recovery after revascularization for chronic coronary artery disease is dependent on maintenance of oxidative metabolism. J Am Coll Cardiol. 1992;20:569–577. 8. Gropler RJ, Geltman EM, Sampathkumaran KS. Comparison of C-11 acetate with F-18 fluorodeoxyglucose for delineating viable myocardium by positron emission tomography. J Am Coll Cardiol. 1993;22:1587–1597. 9. Wendland MF, Saeed M, Lauerma K, de Crespigny A, Moseley ME, Higgins CB. Endogenous susceptibility contrast in myocardium during apnea measured using gradient recalled echo planar imaging. Magn Reson Med. 1993;29:273–276. 10. Atalay MK, Forder JR, Chacko VP, Kawamoto S, Zerhouni EA. Oxygenation in the Rabbit Myocardium: Assessment with SusceptibilityDependent MR imaging. Radiology. 1993;189:759. 11. Balaban RS, Taylor JF, Turner R. Effect of cardiac flow on gradient recalled echo images of the canine heart. NMR Biomed. 1994;7:89–95. 12. Stillman AE, Wilke N, Jerosch-Herold M. BOLD contrast of the heart during occlusion and reperfusion. In: Works in Progress Supplement, SMR 1st Meeting (Dallas); 1994:S24. 13. Atalay M, Reeder SB, Zerhouni E, Forder JR. Blood oxygenation dependence of T1 and T2 in the isolated, perfused rabbit heart at 4.7T. Magn Reson Med. 1995;34:623–627. 14. Li D, Dhawale P, Rubin PJ, Haacke EM, Gropler RJ. Myocardial signal response to dipyridamole and dobutamine: demonstration of the BOLD effect using a double-echo gradient-echo sequence. Magn Reson Med. 1996;36:16–20. 15. Niemi P, Poncelet BP, Kwong K, et al. Myocardial intensity changes associated with flow stimulation in blood oxygenation sensitive magnetic resonance imaging. Magn Reson Med. 1996;36:78–82. 16. Li D, Oellerich WF, Beck G, Gropler RJ. Assessment of myocardial response to pharmacologic interventions using an improved MR imaging technique to estimate T2* values. AJR Am J Roentgenol. 1999;172: 141–145. 17. Wacker CM, Bock M, Hartlep AM, et al. Changes in myocardial oxygenation and perfusion under pharmacological stress with dipyridamole: assessment using T2* and T1 measurements. Magn Reson Med. 1999;41:686–695. 18. Friedrich MG, Niendorf T, Schulz-Menger J, Gross CM, Dietz R. Blood oxygen level-dependent magnetic resonance imaging in patients with stress-induced angina. Circulation. 2003;108(18):2219–2223. 19. Wacker CM, Hartlep AW, Pfleger S, Schad LR, Ertl G, Bauer WR. Susceptibility-sensitive magnetic resonance imaging detects human myocardium supplied by a stenotic coronary artery without a contrast agent. J Am Coll Cardiol. 2003;41(5):834–840. 20. Beache GM, Herzka DA, Boxerman JL, et al. Attenuated myocardial vasodilator response in patients with hypertensive hypertrophy revealed by oxygenation-dependent magnetic resonance imaging. Circulation. 2001;104:1214–1217. 21. Wright KB, Klocke FJ, Deshpande VS, et al. Assessment of regional differences in myocardial blood flow using T2-weighted 3D BOLD imaging. Magn Reson Med. 2001;46:573–578. 22. Foltz WD, Huang H, Fort S, Wright GA. Vasodilator response assessment in porcine myocardium with magnetic resonance relaxometry. Circulation. 2002;106:2714–2719. 23. Foltz WD, Al-Kwifi O, Sussman MS, Stainsby JA, Wright GA. Optimized spiral imaging for measurement of myocardial T2 relaxation. Magn Reson Med. 2003;49:1089–1097. 24. Zheng J, Wang J, Rowold FE, Gropler RJ, Woodard PK. Relationship of apparent myocardial t2 and oxygenation: towards quantification of myocardial oxygen extraction fraction. J Magn Reson Imaging. 2004;20(2):233–241. 25. Zheng J, Wang J, Nolte M, Li D, Gropler RJ, Woodard PK. Dynamic estimation of the myocardial oxygen extraction ratio during dipyridamole stress by MRI: a preliminary study in canines. Magn Reson Med. 2004;51(4):718–726. 26. Zhang H, Gropler RJ, Li D, Zheng J. Assessment of myocardial oxygen extraction fraction and perfusion reserve with BOLD imaging in a canine model with coronary artery stenosis. J Magn Reson Imaging. 2007;26(1):72–79. 27. McCommis KS, Goldstein TA, Zhang H, Misselwitz B, Gropler RD, Zheng J. Quantification of myocardial blood volume during dipyridamole and dobutamine stress: a perfusion CMR study. J Cardiovasc Magn Reson. 2007;9(5):785–792.

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28. McCommis KS, Zhang H, Herrero O, Gropler RD, Zheng J. Feasibility study of myocardial perfusion and oxygenation by noncontrast MRI: comparison with PET study in a canine model. J Magn Reson Imaging. 2008;26(1):11–19. 29. Fieno DS, Shea SM, Li Y, Finn JP, Li D. Myocardial perfusion imaging based on the blood oxygen level-dependent effect using T2-prepared steady-state free precession magnetic resonance imaging. Circulation. 2004;110:1284–1290. 30. Shea SM, Fieno DS, Schirf BE, et al. T2-prepared steady-state free precession blood oxygen level-dependent MR imaging of myocardial perfusion in a dog stenosis model. Radiology. 2005;236(2):503–509. 31. Wilke N, Simm C, Zhang J, et al. Contrast-enhanced first pass myocardial perfusion imaging: correlation between myocardial blood flow in dogs at rest and during hyperemia. Magn Reson Med. 1993;29:485–497. 32. Klocke FJ. Measurements of coronary flow reserve: defining pathophysiology versus making decisions about patient care. Circulation. 1987;76(6):1183–1189. 33. Foltz WD, Merchant N, Downar E, Stainsby JA, Wright GA. Coronary venous oximetry using MRI. Magn Reson Med. 1999;42:837–848. 34. Wendland MF, Saeed M, Mausi T, Derugin N, Higgins CB. First pass of an MR susceptibility contrast agent through normal and ischemic heart: gradient-recalled echo-planar imaging. J Magn Reson Imaging. 1993;3:755–760. 35. Edelman RR, Li W. Contrast-enhanced echo-planar MR imaging of myocardial perfusion. preliminary study in humans. Radiology. 1994;190:771–777. 36. Wilke N, Jerosch-Herold M, Wang Y, et al. Myocardial perfusion reserve: assessment with multisection, quantitative, first-pass MR imaging. Radiology. 1997;204:373–384. 37. Pauling L, Coryell CD. The magnetic properties and structure of hemoglobin, oxyhemoglobin and carbonmonoxyhemoglobin. Proc Natl Acad Sci U S A. 1936;22:210–216. 38. Thulborn K, Waterton J, Matthews P, Radda G. Oxygenation dependence of the transverse relaxation time of water protons in whole blood at high field. Biochim Biophys Acta. 1982;714:265–270. 39. Li D, Wang Y, Waight DJ. Blood oxygen saturation assessment in vivo using T2* estimation. Magn Reson Med. 1998;39:685–690. 40. Wright GA, Hu BS, Macovski A. Estimating oxygen saturation of blood in vivo with MR imaging at 1.5T. J Magn Reson Imaging. 1991;1:275–283. 41. Brittain JH, Olcott EW, Szuba A, et al. Three-dimensional flowindependent peripheral angiography. Magn Reson Med. 1997;38 (3):343–354. Erratum in: Magn Reson Med. 1998;40(6):948–951. 42. Li KC, Dalman RL, Ch’en IY, et al. Chronic mesenteric ischemia: use of in vivo MR imaging measurements of blood oxygen saturation in the superior mesenteric vein for diagnosis. Radiology. 1997;204:71–77. 43. Nield LE, Qi XL, Valsangiacomo ER, et al. In vivo MRI measurement of blood oxygen saturation in children with congenital heart disease. Pediatr Radiol. 2005;35(2):179–185. 44. Kaul S, Jayaweera AR. Coronary and myocardial blood volumes: noninvasive tools to assess the coronary microcirculation? Circulation. 1997;96:719–724. 45. Bauer WR, Nadler W, Bock M, et al. Theory of the BOLD effect in the capillary region: an analytical approach for the determination of T2 in the capillary network of myocardium. Magn Reson Med. 1999;41:51–62. 46. Chien D, Levin DL, Anderson CM. MR gradient echo imaging of intravascular blood oxygenation: T2* determination in the presence of flow. Magn Reson Med. 1994;32:540–545. 47. Kim SG, Ugurbil K. Comparison of blood oxygenation and cerebral blood flow effects in fMRI: estimation of relative oxygen consumption change. Magn Reson Med. 1997;38:59–65. 48. Kennan RP, Scanley BE, Gore JC. Physiologic basis for BOLD MR signal changes due to hypoxia/hyperoxia: separation of blood volume and magnetic susceptibility effects. Magn Reson Med. 1997;37:953–956. 49. Boxerman JL, Hamberg LM, Rosen BR, Weisskoff RM. MR contrast due to intravascular magnetic susceptibility perturbations. Magn Reson Med. 1995;34(4):555–566. 50. Bauer WR, Nadler W, Bock M, et al. The relationship between the BOLD-induced T2 and T2*: a theoretical approach for the vasculature of myocardium. Magn Reson Med. 1999;42:1004–1010. 51. Belliveau JW, Kennedy DN, Mckinstry RC, et al. Functional mapping of human visual cortex by magnetic resonance imaging. Science. 1991;254:716–719. 52. Kwong KK, Belliveau JW, Chesler DA, et al. Dynamic magnetic resonance imaging of human brain activity during primary sensory stimulation. Proc Natl Acad Sci U S A. 1992;89:5675–5679.

62. Mangalathu Arumana J, Li D, Dharmakumar R. Deriving bloodoxygen-level-dependent contrast in MRI: an evaluation of T2*weighted, T2-prepared and phase-cycled SSFP methods at 1.5T and 3.0T. Magn Reson Med. 2008;59:561–570. 63. Dharmakumar R, Mangalathu Arumana J, Larson AC, Chung Y, Wright GA, Li D. Cardiac phase-resolved blood oxygen-sensitive steady-state free precession MRI for evaluating the functional significance of coronary artery stenosis. Invest Radiol. 2007;42(3):180–188. 64. Dharmakumar R, Green JD, Flewitt J, et al. Imaging for Probing the Myocardial Perfusion Reserves of Patients with Coronary Artery Disease: A Feasibility Study. Los Angeles, USA: SCMR; 2008. 65. Dharmakumar R, Mangalathu Arumana J, Tang R, Harris K, Zhang Z, Li D. Assessment of regional myocardial oxygenation changes in the presence of coronary artery stenosis with balanced SSFP imaging at 3.0T: theory and experimental evaluation in canines. J Magn Reson Imaging. 2008;27(5):1037–1045. 66. Larson AC, White RD, Laub G, McVeigh ER, Li D, Simonetti OP. Selfgated cardiac cine MRI. Magn Reson Med. 2004;51(1):93–102. 67. Griswold MA, Jakob PM, Heidemann RM, et al. Generalized autocalibrating partially parallel acquisitions (GRAPPA). Magn Reson Med. 2002;47:1202–1210. 68. Pruessmann KP, Weiger M, Scheidegger MB, Boesiger P. SENSE: sensitivity encoding for fast MRI. Magn Reson Med. 1999;42:952–962. 69. Sodickson DK, Manning WJ. Simultaneous acquisition of spatial harmonics (SMASH): fast imaging with radiofrequency coil arrays. Magn Reson Med. 1997;38:591–603. 70. Sodickson DK, McKenzie CA. A generalized approach to parallel magnetic resonance imaging. Med Phys. 2001;28:1629–1643.

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53. Ogawa S, Menon RS, Tank DW, et al. Functional brain mapping by blood oxygenation level-dependent contrast magnetic resonance imaging: a comparison of signal characteristics with a biophysical model. Biophys J. 1993;64:803–812. 54. Lai S, Hopkins AL, Haacke EM, et al. Identification of vascular structures as a major source of signal contrast in high resolution 2D and 3D functional activation imaging of the motor cortex as 1.5T: preliminary results. Magn Reson Med. 1993;30:387–392. 55. McGuinness ME, Talbert RL. Pharmacologic stress testing: experience with dipyridamole, adenosine, and dobutamine. Am J Hosp Pharm. 1994;51:328–346. 56. Massie BM, Schwartz GG, Garcia J, Wisneski JA, Weiner MW, Owens T. Myocardial metabolism during increased work states in the porcine left ventricle in vivo. Circ Res. 1994;74:64–73. 57. Klocke FJ, Li D. Testing coronary flow reserve without a provocative stress: a “BOLD” approach. J Am Coll Cardiol. 2003;41(5):841–842. 58. Westwood M, Anderson LJ, Firmin DN, et al. A single breath-hold multiecho T2* cardiovascular magnetic resonance technique for diagnosis of myocardial iron overload. J Magn Reson Imaging. 2003;18 (1):33–39. 59. Dharmakumar R, Hong J, Brittain J, Plewes DB, Wright GA. Oxygensensitive contrast in blood for steady-state free precession imaging. Magn Reson Med. 2005;53:574–583. 60. Brittain JH, Reeder SB, Shimakawa A, et al. Non-contrast-enhanced angiography at 3T using SSFP and “Dixon” fat-water separation. In: 2004 ISMRM Conference Proceedings (Kyoto, Japan); p. 11. 61. Dharmakumar R, Qi X, Hong J, Wright GA. Detecting microcirculatory changes in blood oxygen state with steady-state free precession imaging. Magn Reson Med. 2006;55:1372–1380.

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CHAPTER 43

Interventional Cardiovascular Magnetic Resonance Amish N. Raval and Robert J. Lederman

An ideal visual guidance system for cardiovascular, catheterbased interventional procedures would offer real-time, highresolution, three-dimensional (3D) imaging of important anatomic tissues and chambers, irrespective of respiratory, cardiac, or patient motion. Such tools would quickly enable novel minimally invasive alternatives to open surgical procedures. X-ray fluoroscopy (XRF) guides most contemporary catheter-based procedures. However, XRF has important limitations (Table 43-1). Iodinated radiocontrast, which provides the ability to outline chamber and vascular lumina, is injected only periodically. Tissue detail is minimal. Additionally, XRF provides only two-dimensional (2D) “projection” imaging, with limited depth perception. Iodinated radiocontrast is nephrotoxic in susceptible individuals. Ionizing radiation increases lifetime cancer risk, particularly in children.1–9 Both the patient and the operator as well as in-room personnel are exposed. Finally, operators and personnel risk disabling orthopedic injuries from the use of heavy and bulky protective lead apparel.10,11 Ultrasound is sometimes used in combination with certain XRF procedures. For example, transesophageal echocardiography (TEE) provides adjunctive imaging during XRF guided atrial septal defect closure, by assisting with device sizing, positioning, and post-deployment assessment. Ultrasound offers limited acoustic windows and has “shadowing” artifacts caused by devices. Advances in real-time 3D ultrasound may offer interventional imaging guidance in the future,12 although restricted acoustic windows and shadow artifacts will likely remain problematic. Cardiovascular magnetic resonance (CMR) imaging more closely approaches the idealized visual guidance system described earlier. CMR offers superior tissue imaging and the opportunity to acquire images directly in any arbitrary orientation (rather than reconstructions), without ionizing radiation or nephrotoxic contrast. Advances in rapid imaging techniques and modified CMR visible interventional devices can now permit real-time CMR-guided applications. In the future, the ability to visualize tissue space may enable catheter-based procedures that currently require open surgical exposure.

INTERVENTIONAL CARDIOVASCULAR MAGNETIC RESONANCE LABORATORY Early real-time CMR (RT-CMR) systems used low-magneticfield-strength magnets (0.2 to 0.5 Tesla), open- or closed580 Cardiovascular Magnetic Resonance

bore configurations, and sometimes incorporated XRF units.13,14 Major limitations included a relatively low signal-to-noise ratio (SNR); an inhomogeneous magnetic field, interfering with rapid imaging; and limited patient access when using long, closed bores. Short bore, 1.5 Tesla systems with high performance gradients provide excellent field homogeneity and a higher SNR. These scanners also include a large array of radiofrequency (RF) receivers (32 or more), intended for parallel imaging techniques (discussed later), to improve imaging speed. These additional receiver channels can also be used to attach “active” catheter devices to improve their visibility (discussed later). In addition, commercial CMR scanners now include interactive scan user interfaces that increasingly resemble echocardiography. These interfaces drive sophisticated image reconstruction hardware and software that permit images to be acquired with ease and to be displayed to operators with minimal delay. For investigational laboratories, newer CMR systems support better, highly versatile software environments that facilitate rapid development of investigational pulse sequence, image reconstruction, and user interface improvements. Combined XRF and RT-CMR laboratories, or XMR laboratories, are probably best for investigational CMR interventions. In these laboratories, conventional XRF guidance is immediately available for unanticipated CMR system failures or situations requiring emergency bailout. Additionally, early-stage CMR-guided procedures can be used as an adjunct to established XRF procedures. XMR laboratories are now commercially available. These systems are separated by RF-shielded doors and can be used independently or together for combined procedures (Fig. 43-1).15 An automated transport table moves patients between the two modalities. Scan plane and pulse sequences are modified either inside the scanner room or in the external control room. Wave guides and penetration panels are strategically positioned in the room so that new electronic communication and monitoring hardware can be installed without disrupting the RF shield system. Images can be displayed inside the scan room with liquid crystal display projectors or shielded liquid crystal display monitors (see Fig. 43-1B), and XRF and CMR datasets can be fused to guide interventions in both arenas.16

Communication and Monitoring Rapid CMR requires rapid gradient switching that causes substantial acoustic noise. This noise prevents verbal communication between operators, the patient, and staff. RF-filtered

Guidance System

Advantages

Disadvantages

RT-CMR

Excellent soft tissue imaging Simultaneous display of multiple arbitrary imaging perspectives Real-time 3D device tracking Image-based physiology assessment No ionizing radiation No lead aprons No nephrotoxic contrast Uninterrupted imaging to detect complications more efficiently

Clinical devices currently unavailable and require hardware attachments Acoustic noise Claustrophobia Rapid emergency response more difficult ECG monitoring for cardiac procedures more difficult Peripheral nerve stimulation from rapidly switching gradients Inferior spatial and temporal resolution in real-time mode

XRF

Excellent temporal resolution Widely available in most centers Clinical devices available Simple ECG monitoring Floating table permits remote imaging (i.e., groin site) Greater physician and nurse access to patient

Poor soft tissue imaging 2D “projection” imaging offers limited depth perception Cancer risk from ionizing radiation to patient and room personnel Nephrotoxic contrast required Orthopedic injuries to room personnel from long-term use of lead apron apparel Angiography interrupted during device deployment, resulting in detection of complications after contrast injection

Ultrasound

Good soft tissue imaging Excellent temporal resolution Widely available Imaging-based physiology assessment (i.e., Doppler)

Limited acoustic windows Device-related “shadow” artifacts No tip-tracking capability Transesophageal and intracardiac echo procedures are invasive Limited orientation-independent imaging

2D, two-dimensional; 3D, three-dimensional; ECG, electrocardiogram; RT-CMR, real-time cardiovascular magnetic resonance imaging; XRF, x-ray fluoroscopy.

Figure 43-1 A, National Heart, Lung and Blood Institute combined X-ray/cardiovascular magnetic resonance suite. B, Interventional cardiovascular magnetic resonance suite.

A

headsets (Magnacoustics, Atlantic Beach, NY) with directional fiberoptic microphones (Phone-Or, Yahuda, Israel) are used at the U.S. National Heart, Lung and Blood Institute (NHLBI). This versatile system offers multiple communication modes, such as operator-control room-only mode (the patient cannot hear), operator-patient mode, and universal communication mode. Gradient systems are now commercially available that are quieter by 20 dB or more.17 Some laboratories use CMR to create intermittently updated road map image sets and then use nearly silent pulse sequences for tracking catheter position.18 These permit open verbal communication among staff. Gradients, RF interference, and magnetohydrodynamic effects severely distort the electrocardiogram. Electronic

B

filters enable heart rate monitoring; however, electrocardiographic (ECG) waveform morphology is not interpretable for ST segment monitoring of myocardial ischemia or injury. Multichannel invasive pressure waveform monitoring, temperature monitoring, and oxygen saturation monitoring are implemented using commercial systems, without modification from XRF laboratories. Interventional catheterization suites require continuous operator monitoring of multiple information streams, including hemodynamics, imaging control, and interventional images. This is also true during interventional CMR procedures, especially with active devices. At the NHLBI, hemodynamic data, the scanner interface, and 3D-rendered images are shown separately on a single large projector screen. Cardiovascular Magnetic Resonance 581

43 INTERVENTIONAL CARDIOVASCULAR MAGNETIC RESONANCE

Table 43-1 Advantages and Disadvantages of Interventional Image Guidance Systems

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Interventional Cardiovascular Magnetic Resonance Scanner Interface The interventional CMR user interface has several custom features that are making their way into commercial packages. One of the most useful features is the colorized display of individual receiver channels that are attached to “active” catheter receiver coil devices. This makes catheter devices readily apparent and has proven extremely useful in vivo. Other custom features include interactive “projection-mode” imaging wherein the slice select feature is disabled and display of catheter devices as they would appear under conventional XRF. Operationally, it is useful to display multiple slices during interventional procedures, both separately and combined in a 3D-rendered image (Fig. 43-2).19 Scanning parameters, including temporal resolution, spatial resolution, and field of view, are altered interactively to suit the particular stage of the procedure. Other useful interactive features include digital subtraction during selective

arteriography and ECG gating to “freeze” motion. Other laboratories are experimenting with voice control, automatic adjustment of slice location to correspond to catheter position, and automatic adaptation of imaging parameters to the speed of catheter manipulation.

Safety Considerations All patient care staff must undergo formal training in CMR safety. For example, the hospital emergency medical response team is unlikely to be familiar with CMR operations, and team members may unwittingly jeopardize themselves, their colleagues, or the patient by rushing into the room with a steel oxygen tank or other ferromagnetic objects. Rapid patient evacuation from the magnet room must be practiced repeatedly during mock drills. Ferromagnetic objects that cannot be removed from the room must be firmly secured to the wall at all times. Oxygen, inhalational anesthetic, and vacuum for suction can be fed through RF-protective wall ports, eliminating the need for in-room tanks. The RF energy from CMR transmission coils will concentrate on long conductive devices, causing local heating,

Figure 43-2 Commercial interventional magnetic resonance imaging user interface design to accommodate interleaved multi-slice realtime cardiovascular magnetic resonance image acquisition (three panels on left), a volume rendering of the slices indicating their threedimensional relationship (center panel), “postage stamps” to store and recall important graphic slice prescriptions (bottom row), and interactive scanner parameter control (right panels). (Courtesy of Christine H. Lorenz, PhD, Siemens Medical Solutions, with permission.) 582 Cardiovascular Magnetic Resonance

Interventional Device Factors Length of conductive elements Geometric shape Orientation in the magnet Distance from the radiofrequency transmitter Physical proximity to tissues Insulation Patient Factors Body mass and surface area Convective cooling of intravascular devices by local blood flow General body temperature Implanted conductive devices Tissue thermosensitivity Scanner Factors Field strength Pulse sequence Flip angle Scanning duty cycle Position relative to bore isocenter (closer is better)

potentially leading to burns.20 Table 43-2 lists factors associated with heating, especially long conductive devices and cables. This complex problem is described later. Interventionists must also take care that connections to intravascular coils as well as surface coils do not inadvertently form loops, which can lead to patient burns.

Real-Time CMR Imaging Rapid CMR is required for invasive procedures, catheter visualization, and imaging of anatomic structures. Efficient image data sampling methods,21–31 parallel imaging,32–34 and coherent steady-state techniques35–40 have enabled real-time CMR without significant degradation of image quality in the field of interest. Frame rates of 10 to 15 images/sec or greater are now possible using multichannel (32 or more) RF systems that use parallel imaging techniques to achieve acceleration factors of 3 or more. In addition, interactive, real-time color flow imaging may supplement anatomic detail with critical physiologic features, such as leaks or gradients, during therapeutic procedures or interventions.41,42 Slice orientation can automatically follow the tip of catheter devices as they move, known as adaptive imaging, so that the catheter features are kept in view during manipulation. These methods also can be applied automatically to alter scanning parameters such as field of view and temporal resolution.43

Catheter Devices Interventional devices must be clearly and distinctly visible to conduct therapeutic procedures. Conventional XRF devices are generally unsuitable because most incorporate steel braids to increase X-ray attenuation for visibility and to enhance catheter performance characteristics, such as steerability, pushability, and trackability. The steel causes severe “blooming” or susceptibility artifacts that lead to artifacts and distort the CMR image (Fig. 43-3). Removing

Figure 43-3 A, Stainless steel braided Kumpe catheter. B, Cardiovascular magnetic resonance (CMR) of the catheter in a water phantom. Note the severe “blooming” signal void artifact in the image, rendering it useless for CMR-guided interventional procedures.

these ferrous components usually renders the catheter devices virtually invisible under CMR and usually renders them mechanically unsuitable (floppy) as catheters. Several approaches have been fielded to overcome these problems. Interventional CMR catheter devices are generally classified as passive or active. Passive devices have elements that cause discrete susceptibility imaging artifacts that “darken” images or T1-shortening elements that “brighten” images. Alternatively, active devices have built-in microcoils and electrical circuitry that allow the device to act as an RF receiver or transmitter.

Passive Devices Devices That Create Dark Signals Stents44 and guidewires45–48 made of copper, dysprosium, cobalt-chromium alloy, nitinol, titanium, and platinum are associated with less severe susceptibility artifacts than those made of iron alloys. They can be coated to improve biocompatibility. Carbon dioxide creates a dark CMR signal by excluding proton spins; carbon dioxide gas has been injected into humans for selective CMR angiography and has been used to fill balloon catheters for diagnostic CMR-guided catheter tracking in patients.49 A passive catheter tracking system based on an optically detuned resonance marker installed on the catheter tip has also been described in a canine model.50 Unfortunately, volume averaging makes it difficult to distinguish passive devices from neighboring anatomic features.

Devices That Create Bright Signals Devices that appear bright are less vulnerable to volume averaging effects and are more easily visualized against the anatomic background. Dilute gadolinium chelates, such as gadopentate dimeglumine (Gd-DTPA), have been used to fill51 and coat Cardiovascular Magnetic Resonance 583

43 INTERVENTIONAL CARDIOVASCULAR MAGNETIC RESONANCE

Table 43-2 Factors Influencing Heating-Related Injury by Cardiovascular Magnetic Resonance Interventional Devices

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catheters52,53 and balloons,54 offering bright device imaging. We have found this passive approach to be inferior in vivo compared with “active” catheter device techniques (Fig. 43-4). Investigational, off-resonance, chemical-selective visualization using alcohol, F-19,55 or C-13,56 and other hyperpolarizing agents has been proposed for catheter tracking. Catheters are filled with compounds (F-19, C-13) that resonate at a frequency other than proton frequency. These novel contrast technologies provide a very high SNR that can be used to enhance catheter visibility with short-acting intravascular contrast agents (Fig. 43-5).

Active Devices Highly sensitive, ultra-small receiver coils can be incorporated into devices to locate them (catheter tracking17,57), visualize them,58–60 or both.61 (Figs. 43-6 and 43-7) Typically,

these devices incorporate conductive wires along the length of the device. Signals from these devices are displayed on top of previously acquired anatomic road maps (device localization only) or combined into the array of receivers used to update real-time CMR images. The signal receiver can be color coded to distinguish the device position from greyscale background images of the anatomy. Devices displayed in this way are readily visible inside thick-slab projections that resemble projection X-ray images showing devices in profile. The tip of these devices can be tracked in 3D, which is ideal for intracardiac applications. Alternatively, direct current applied to conductive elements along the device induces magnetic field inhomogeneities, disrupts local signal, and creates dark impressions.62 Unfortunately, the long conductive transmission wires used to connect catheter coils to the CMR transmitter or receiver system are susceptible to RF heating, which may damage neighboring tissue. Multiple approaches to reduce

Figure 43-4 Passive and active device imaging during aortic coarctation repair. Left panel, Nitinol guide wire is positioned retrograde through the aortic valve into the left ventricular (LV) chamber with real-time, multi-slice and rendered steady-state free precession imaging. A partially filled balloon containing dilute gadolinium is positioned across the coarctation (yellow arrow). The wire and the balloon catheter shaft are poorly visualized against the anatomic background. However, the balloon itself is readily apparent. Right panel, An “active” guidewire coil is attached to a separate cardiovascular magnetic resonance receiver channel and the resulting signal is red. Both the gadolinium-filled balloon and the anatomic features are more conspicuous compared with solely “passive” catheter devices.

Figure 43-5 Passive device visualization using multispectral (chemical-selective) imaging of carbon 13-contrast-filled catheters combined with standard proton cardiovascular magnetic resonance in swine (A and B) and for selective renal arteriography (C). (Courtesy of S. Mansson et al., Malmo University Hospital, Sweden. Reproduced with permission of Springer-Verlag, GmBH.) 584 Cardiovascular Magnetic Resonance

A

43 INTERVENTIONAL CARDIOVASCULAR MAGNETIC RESONANCE

Figure 43-6 In vivo, real-time active tracking of an electrophysiologically mapping catheter in a three-dimensional surface-rendered cardiovascular magnetic resonance image of the left ventricle (St. Jude). A shows catheter within the ventricular from “endocardial” perspective; B shows epicardial perspective. Blue dots are individual microcoils along the catheter. (Courtesy of Charles Dumoulin, PhD, General Electric Medical Systems.)

B

Figure 43-7 Active transseptal puncture needle photograph (A) and cardiovascular magnetic resonance image in a water phantom (B) and in vivo across the atrial septum (C).

heating can be combined to make catheter devices safe, such as the use of circuitry to decouple and detune the transmission lines, intermittent chokes and transformers in the transmission lines, and insulation. Another hybrid approach is to incorporate closed-loop receiver coils into stents or catheter devices without connecting via transmission lines to the CMR system. These “inductively coupled” devices resonate at predetermined geometric shapes and thereby amplify the local RF signal.63,64 Unfortunately, inductively coupled catheter devices cannot easily be displayed in color during RTCMR, as can other actively visualized catheter devices.

Device Solutions for Cardiovascular Applications It is useful to use both active and passive devices during complex interventional procedures, albeit in animal models. Realtime, “active” visualization of the device tip in 3D is desirable during certain procedures, such as transseptal atrial puncture or myocardial wall injection. Similarly, active approaches

prove useful while surveying devices for common failure modes, such as buckling or kinking. Passive approaches unencumbered by electrical hardware attachments are sufficient for simple procedures, such as placing introducer sheaths that typically do not move once positioned. More importantly, combining passive devices (e.g., balloons filled with dilute Gd-DTPA) with active devices (e.g., active guidewire coils) generates wonderful and useful images (see Fig. 43-4).

APPLICATIONS Cardiac Applications Targeted Local Delivery of Cellular Agents to the Myocardium Investigational cell-based and other biologic therapies have attracted great interest in the treatment of myocardial and vascular disease. Intravenous, surgical, and catheter-based approaches to cell delivery have been used in animal as well as human studies in subjects with myocardial Cardiovascular Magnetic Resonance 585

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infarction and heart failure. Some investigators believe that precise anatomic targeting, to infarct border zones, for example, may improve the biologic response to cell therapy. If so, CMR-guided cell delivery may offer advantages over other imaging approaches, such as electromechanical mapping and intracardiac and 3D surface ultrasound. Minimally invasive procedures involving the myocardium are particularly amenable to CMR guidance because of the high contrast between myocardium and blood and because of the readily obtained contrast between normal and pathologic myocardial tissue. RT-CMR-guided endomyocardial cell delivery has achieved millimeter-scale precision using modified CMR needle catheters in animal models (Fig. 43-8).65–71 Major advantages with this technique include high anatomic detail of the myocardium combined with sufficiently rapid imaging to resolve nonperiodic cardiac and respiratory motion.

Electrophysiology: Atrial and Ventricular Mapping, Ablation, and Transseptal Puncture Therapeutic endomyocardial catheter ablation is widely performed using endomyocardial mapping systems and XRF to abolish atrial and ventricular tachyarrhythmia. In these procedures, a mapping catheter is advanced into the cardiac chambers, guided by endocardial electrogram patterns to localize the arrhythmia. Key targets are subjected to radiofrequency or cryoablation to create nonconductive zones to abolish the arrhythmia. Because available imaging modalities afford poor visualization of tissue and anatomic structures, these procedures can be challenging and time consuming. Road maps created using previous electromagnetic maps, CMR, or computed tomography can be used to “fuse” with updated catheter images72,73; however, these road maps are subject to intrinsic registration errors, nonperiodic cardiac and respiratory motion,74 alterations in volume as loading conditions change, and catheter-induced geometric distortion. One particularly attractive electrophysiologic procedure is catheter treatment of atrial fibrillation by creating lines of ablation to isolate all four pulmonary veins. Even in

experienced hands, these procedures, guided by XRF and electromagnetic mapping, usually require hours of radiation exposure. “Image-guided” treatment of atrial fibrillation, conducted under direct surgical exposure, can take minutes. It is tantalizing to speculate that comparable image-guided treatment of atrial fibrillation might be afforded by RT-CMR guidance without surgical exposure. The complex architecture of the pulmonary veins, atria, atrial appendices, and ventricles can clearly be visualized with CMR, permitting precise anatomic targeting. RT-CMR guidance systems using actively tracked catheters and filtered local electrograms75 are currently under development for use in these procedures.76 Interventional CMR in the electrophysiologic environment would also facilitate early detection of complications.77

Atrial Transseptal Procedures Atrial transseptal puncture is usually conducted as the first step in numerous cardiac procedures, such as pulmonary vein ablation. A needle is advanced from a vein through the right atrium (RA) and into left atrium (LA) across the interatrial septum. Currently, this procedure is conducted using subtle XRF visual cues and tactile feedback from sharp catheter devices, with or without adjunctive TEE or intracardiac echocardiography. Poor tissue visualization and limited acquisition windows, combined with unusual atrial anatomy, can lead to life-threatening perforation and pericardial tamponade in as many as 1% to 6% of cases, even in experienced laboratories. Using custom active needles, RT-CMR-guided atrial transseptal puncture has been performed successfully in swine models with modified transseptal puncture needles and with laser catheters (see Fig. 43-7).78–81 Related therapeutic procedures, such as closure of atrial septal defects and patent foramen ovale, have been reported using passively visualized, nitinol devices delivered with catheters in swine.82,83

Invasive Coronary Artery Imaging and Intervention Considered by some the “holy grail” of RT-CMR-guided interventions, percutaneous coronary selective angiography,84 Figure 43-8 A, Active endomyocardial injection guide catheter with deployed and undeployed Stiletto (Boston Scientific, Natick, MA) needles. B, Real-time cardiovascular magnetic resonance-guided endomyocardial delivery of ironlabeled cells (yellow arrow) into the border zone of a swine infarct model.

A

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B

Valve Replacement and Repair In aortic or mitral valve replacement, surgical incisions, along with cardiac arrest and cardiopulmonary bypass, are required to resect the native valve structures and implant a prosthetic or homograft valve. Alternative percutaneous approaches to valve replacement and repair are under intense investigation. Early investigational clinical percutaneous aortic valve replacement has been performed with XRF and TEE guidance in adult calcific aortic stenosis.87 Similarly, percutaneous pulmonary valve implantation has been performed in children with acquired and congenital pulmonary valve disease who previously would have required high-risk corrective surgery.88 XRF combined with TEE cannot fully resolve complex valve structure and mechanical device interactions, once again caused by poor tissue characterization, limited imaging windows, and problems with device-related acoustic shadowing. CMR can fully delineate all-important structures in multiple imaging planes and can provide physiologic assessment before and after treatment. Successful RT-CMR-guided transcatheter prosthetic aortic valve implantation has been demonstrated successfully in swine using passive susceptibility markers89 and active elements.90

Extracardiac Vascular Applications Invasive Arterial Imaging Endovascular “imaging” coils for arterial plaque characterization have also been developed and tested in both phantoms and animal models.91–93 The Intercept Internal CMR guidewire coil (Surgi-Vision, Gaithersburg, MD) is approved for marketing in the United States. This loopless coil design wire functions as a local receiver antenna, sensitive to excited spins within a radius of only a few millimeters. The feasibility of intravascular MRI has been tested in diseased human arteries, with mixed results.15,94,95 Dick and colleagues found in vivo that, although the procedure was feasible and safe, intravascular CMR with this simple device offered a poor SNR, poor spatial resolution, and imaging quality that was not improved over that obtained with surface coil imaging and was far inferior to intravascular ultrasound.15 Imaging coil motion

and unbalanced proximity to adjacent tissues resulted in poor and unpredictable contrast. Larger deployable loop, bird cage, or even opposed solenoid96 intravascular designs are under development and are expected to improve the performance of such devices.

Aortic Aneurysm and Aortic Dissection Repair Percutaneous endograft repair of thoracoabdominal aneurysms and aortic dissection is performed in patients with suitable anatomy who are considered to be at high risk for surgery. Aorta size, proximal and distal landing zones for stents and grafts, and vicinity to crucial arterial branches are vital measurements required for these procedures. These procedures are typically performed using XRF with adjunctive intravascular ultrasound. Bulky stent or graft devices may distort the native anatomy, preventing operator confidence in preacquired XRF road maps. Ultrasound scatter within stents or endograft offer limited external visualization. RT-CMR-guided endograft repair of abdominal aortic aneurysm97 and aortic dissection98 has been performed successfully in swine models using active and passive nitinol stents.99 Post-procedure assessment using phase contrast flow within and adjacent to the endograft showed the versatility of CMR-guided endograft therapy (Figs. 43-9 and 43-10).

Aortic Coarctation Stent Repair Advances in transcatheter therapy for many congenital cardiovascular conditions have reduced the need for invasive open surgery. Children are particularly sensitive to the harmful effects of X-ray radiation.2–6,100–102 Children with complex congenital cardiovascular disease often undergo multiple XRF procedures and therefore have greater cumulative radiation exposure. Radiation-sparing procedural guidance with CMR is attractive and forms the basis for a number of RT-CMR-guided research applications. Stent repair of aortic coarctation under RT-CMR guidance has successfully been performed in a swine coarctation model using commercially available clinical-grade devices (see Fig. 43-4).46,103

Transjugular Intrahepatic Portosystemic Shunt The transjugular intrahepatic portosystemic shunt (TIPS) procedure is performed in patients with liver cirrhosis and refractory portal hypertension. A needle/catheter is introduced via the internal jugular vein, traverses the liver parenchyma, and enters the portal vein. Stents are implanted to bridge a connection between the inferior vena cava and the portal vein, thereby relieving portal pressure. Major complications, such as liver capsule perforation leading to life-threatening hemorrhage, can occur. CMR is ideally suited to guide these procedures because all tissue structures involved are easily visualized. The TIPS procedure has been performed successfully using combination X-ray and MRI and wholly RT-CMR in healthy animals104 and humans.105 A hybrid open-CMR unit combined with XRF was successfully used to perform TIPS in patients with cirrhosis.105 The investigators reported reduced fluoroscopy time and fewer needle passes compared with conventional XRF-guided techniques. Cardiovascular Magnetic Resonance 587

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angioplasty,85 and stent insertion86 has been reported in large animals. Although this work is impressive, clinical translation of these coronary artery therapeutic procedures is hindered by seemingly insurmountable obstacles. Currently, XRF provides spatial resolution of 100 mm at a usual working temporal resolution of 66 msec to manipulate guidewires that are 350 mm wide and stent devices that are 600 mm across. It seems unlikely, barring an unforeseen technical breakthrough, that RT-CMR can provide comparable spatial and temporal resolution to that required for safe guidewire and catheter manipulation through delicate diseased human coronary arteries. Similarly, very low-profile, distinctly conspicuous, CMRcompatible catheter devices would be required for clinical implementation, and such devices are not currently available.

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0.5 0.4 0.3 0.2 0.2 0 cm/s –0.1 –0.2 –0.3 –0.3 –0.4

Figure 43-9 Real-time cardiovascular magnetic resonance-guided abdominal aortic aneurysm repair. Before (top left) and after (top center) repair of infrarenal abdominal aortic aneurysms in a swine model using an active endograft (top right) guided by real-time, multi-slice, and three-dimensional rendered imaging. Arrows reflect the distal and proximal stent limits. Figure 43-10 Real-time cardiovascular magnetic resonance-guided aortic dissection repair. Before (A) and after (C) repair of thoracic aortic dissection in a swine model using a passively visualized stent. B shows the stent delivery system being positioned. Fully deployed stent resolves the dissection (C). (Courtesy of Harold H. Quick, PhD, University Essen-Duisberg.)

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through the “invisible” occluded segment may cause perforation and hemorrhage. These procedures are often long and require excessive iodinated contrast. Successful RT-CMR-guided chronic occlusion recanalization and subsequent balloon angioplasty was performed successfully in a swine model of peripheral artery chronic occlusion, using modified “active” wires and support catheters (Fig. 43-11).118,119

Peripheral Artery Disease

Inferior Vena Cava Filter

Several groups have shown the feasibility of CMR-guided balloon angioplasty in healthy animals,107,108 animal models of arterial stenosis,109–112 and humans with obstructive peripheral artery disease.113 Dilute Gd-DTPA (bright signal), undiluted Gd-DTPA (dark signal), and carbon dioxide gas have been used to inflate balloons and provide balloon-tissue contrast, ensuring full inflation. Active receiver coils, either embedded on the balloon catheter or inserted through the wire port of the balloon filled with dilute Gd-DTPA, provide added contrast to adjacent tissue by enhancing further the bright signal within the inflated balloon. Radiopaque markers are typically added to balloon catheters to indicate the “shoulder” points of the balloon to assist with assessment of lesion length before and during deployment. These markers act as small susceptibility markers to assist with positioning and balloon deployment under CMR. Both balloon-expandable and self-expanding stents are implanted to prevent arterial recoil and alleviate flow-limiting dissection. Both stent designs have been successfully deployed under CMR guidance in animals114–116 and humans.117 Local susceptibility and shielding effects result in imaging voids within and adjacent to stents. Inductively coupled stents (described earlier) may ameliorate this problem.63 Chronic total arterial occlusion recanalization is particularly challenging under XRF. Only the patent inflow, occluded artery, and patent outflow distal to the artery beyond the obstruction can be visualized with conventional X-ray angiography. Traversal of the guidewires and catheter

Inferior vena cava filters are implanted scaffolds, usually self-expanding nitinol, designed to entrap migratory venous thromboemboli. They are placed in patients with lower extremity venous thrombosis in whom systemic anticoagulation is contraindicated or unsuccessful in preventing pulmonary embolism. Accurate positioning of these devices requires adequate visualization of important inflow branches, such as the renal and mesenteric veins, but is straightforward under XRF or ultrasound guidance. Successful RT-CMR-guided deployment of inferior vena cava filters in animals has been shown with passive imaging techniques.120–122 Concomitant CMR venography and thrombus imaging might have clinical value, especially in follow-up assessment.

Figure 43-11 Real-time cardiovascular magnetic resonance-guided carotid chronic total occlusion recanalization in a swine model. A, Active total occlusion wire and catheter. B, Sagittal realtime cardiovascular magnetic resonance showing total occlusion devices (red, wire; green, catheter) traversing the chronic occlusion. Simultaneous axial slice shows the wire tip confined within the arterial wall (yellow arrow). C, The contralateral carotid artery, which is patent, is shown (white arrow). The green arrow shows the occluded artery segment distal to the guidewire tip, and the yellow arrow shows the guidewire tip within the artery lumen.

A

B

CONCLUSION The combination of RT-CMR and CMR visible devices may offer a complete imaging solution for therapeutic cardiovascular interventions. The many advantages over existing guidance modalities include superior tissue imaging, no need for ionizing radiation or iodinated contrast, imagingbased physiology assessment, and 3D perspective. Important challenges remain to the clinical translation of RTCMR, especially the requirement for conspicuous, commercial-grade catheter devices. Minimally invasive and novel therapeutic interventions, once considered impossible with traditional imaging, may now be possible using this rapidly evolving technology.

C

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Mesocaval shunt is a variant of the TIPS procedure in which systemic veins are connected directly to the portal vein without traversing the liver. Guided entirely by RTCMR, Arepally and associates showed successful application of mesocaval shunt in healthy swine using modified “active” transseptal puncture needles and self-expanding stents.106 The investigators hope to conduct a range of novel procedures using RT-CMR access to the portal circulation.

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50. Weiss S, Kuehne T, Brinkert F, et al. In vivo safe catheter visualization and slice tracing using an optically detonable resonant marker. Magn Reson Med. 2004;52:860–868. 51. Omary RA, Unal O, Koscielski DS, et al. Real-time MR imagingguided passive catheter tracking with use of gadolinium-filled catheters. J Vasc Interv Radiol. 2000;11(8):1079–1085. 52. Unal O, Korosec FR, Frayne R, Strother CM, Mistretta CA. A rapid 2D time-resolved variable-rate k-space sampling MR technique for passive catheter tracking during endovascular procedures. Magn Reson Med. 1998;40(3):356–362. 53. Strother CM, Unal O, Frayne R, et al. Endovascular treatment of experimental canine aneurysms: feasibility with MR imaging guidance. Radiology. 2000;215(2):516–519. 54. Bakker CJ, Hoogeveen RM, Weber J, van Vaals JJ, Viergever MA, Mali WP. Visualization of dedicated catheters using fast scanning techniques with potential for MR-guided vascular interventions. Magn Reson Med. 1996;36(6):816–820. 55. Kozerke S, Hegde S, Schaeffter T, Lamerichs R, Razavi R, Hill DL. Catheter tracking and visualization using 19F nuclear magnetic resonance. Magn Reson Med. 2004;52(3):693–697. 56. Mansson S, Johansson E, Magnusson P, et al. 13C imaging: a new diagnostic platform. Eur Radiol. 2006;16(1):57–67. 57. Ackerman JL, Offutt MC, Buxton RB, Brady TJ. ISMRM. Rapid 3D tracking of small RF coils. In: Proc 5th Annual Meeting, ISMRM; 1986:1131. 58. McKinnon GC, Debatin JF, Leung DA, Wildermuth S, Holtz DJ, von Schulthess GK. Towards active guidewire visualization in interventional magnetic resonance imaging. MAGMA. 1996;4(1):13–18. 59. Ocali O, Atalar E. Intravascular magnetic resonance imaging using a loopless catheter antenna. Magn Reson Med. 1997;37(1):112–118. 60. Ladd ME, Zimmermann GG, Quick HH, et al. Active MR visualization of a vascular guidewire in vivo. J Magn Reson Imaging. 1998;8 (1):220–225. 61. Kocaturk O, Saikus CE, Guttman MA, Faranesh AZ, Ratnayaka K, Ozturk C, McVeigh ER, Lederman RJ. Whole shaft visibility and mechanical performance for active MR catheters using copper-nitinol braided polymer tubes. J Cardiovasc Magn Reson. 2009;11(1):29. 62. Glowinski A, Adam G, Bucker A, Neuerburg J, Vanvaals JJ, Gunther RW. Catheter visualization using locally induced actively controlled field inhomogeneities. Magn Reson Med. 1997; (38):253–258. 63. Quick HH, Kuehl H, Kaiser G, Bosk S, Debatin JF, Ladd ME. Inductively coupled stent antennas in MRI. Magn Reson Med. 2002;48 (5):781–790. 64. Quick HH, Zenge MO, Kuehl H, et al. Interventional magnetic resonance angiography with no strings attached: wireless active catheter visualization. Magn Reson Med. 2005;53(2):446–455. 65. Dick AJ, Guttman MA, Raman VK, et al. Magnetic resonance fluoroscopy allows targeted delivery of mesenchymal stem cells to infarct borders in Swine. Circulation. 2003;108(23):2899–2904. 66. Lederman RJ, Guttman MA, Peters DC, et al. Catheter-based endomyocardial injection with real-time magnetic resonance imaging. Circulation. 2002;105(11):1282–1284. 67. Kraitchman DL, Heldman AW, Atalar E, et al. In vivo magnetic resonance imaging of mesenchymal stem cells in myocardial infarction. Circulation. 2003;107(18):2290–2293. 68. Corti R, Badimon J, Mizsei G, et al. Real time magnetic resonance guided endomyocardial local delivery. Heart. 2005;91(3):348–353. 69. Krombach GA, Pfeffer JG, Kinzel S, Katoh M, Gunther RW, Buecker A. MR-guided percutaneous intramyocardial injection with an MR-compatible catheter: feasibility and changes in T1 values after injection of extracellular contrast medium in pigs. Radiology. 2005;235(2):487–494. 70. Saeed M, Lee R, Martin A, et al. Transendocardial delivery of extracellular myocardial markers by using combination X-ray/MR fluoroscopic guidance: feasibility study in dogs. Radiology. 2004;231(3):689–696. 71. Rickers C, Gallegos R, Seethamraju RT, et al. Applications of magnetic resonance imaging for cardiac stem cell therapy. J Interv Cardiol. 2004;17(1):37–46. 72. Sermesant M, Rhode K, Sanchez-Ortiz GI, et al. Simulation of cardiac pathologies using an electromechanical biventricular model and XMR interventional imaging. Med Image Anal. 2005;9(5):467–480. 73. Reddy VY, Malchano ZJ, Holmvang G, et al. Integration of cardiac magnetic resonance imaging with three-dimensional electroanatomic mapping to guide left ventricular catheter manipulation: feasibility in a porcine model of healed myocardial infarction. J Am Coll Cardiol. 2004;44(11):2202–2213.

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98. Eggebrecht H, Kuhl H, Kaiser GM, et al. Feasibility of real-time magnetic resonance-guided stent-graft placement in a swine model of descending aortic dissection. Eur Heart J. 2006;27(5):613–620. 99. Eggebrecht H, Heusch G, Erbel R, Ladd ME, Quick HH. Real-time vascular interventional magnetic resonance imaging: the future of aortic stent-graft placement? Basic Res Cardiol. 2007;102:1–8. 100. American Academy of Pediatrics. Committee on Environmental Health. Risk of ionizing radiation exposure to children: a subject review. Pediatrics. 1998;101(4 Pt 1):717–719. 101. McLaughlin JR, Kreiger N, Sloan MP, Benson LN, Hilditch S, Clarke EA. An historical cohort study of cardiac catheterization during childhood and the risk of cancer. Int J Epidemiol. 1993;22(4):584–591. 102. Spengler RF, Cook DH, Clarke EA, Olley PM, Newman AM. Cancer mortality following cardiac catheterization: a preliminary follow-up study on 4,891 irradiated children. Pediatrics. 1983;71(2):235–239. 103. Krueger JJ, Ewert P, Yilmaz S, et al. Magnetic resonance imagingguided balloon angioplasty of coarctation of the aorta: a pilot study. Circulation. 2006;113:1093–1100. 104. Kee ST, Rhee JS, Butts K, et al. Becker Young Investigator Award. MR-guided transjugular portosystemic shunt placement in a swine model. J Vasc Interv Radiol. 1999;10(5):529–535. 105. Kee ST, Ganguly A, Daniel BL, et al. MR-guided transjugular intrahepatic portosystemic shunt creation with use of a hybrid radiography/ MR system. J Vasc Interv Radiol. 2005;16(2):227–234. 106. Arepally A, Karmarkar PV, Weiss C, Atalar E. Percutaneous MR imaging-guided transvascular access of mesenteric venous system: study in swine model. Radiology. 2006;238(1):113–118. 107. Wildermuth S, Dumoulin CL, Pfammatter T, Maier SE, Hofmann E, Debatin JF. MR-guided percutaneous angioplasty: assessment of tracking safety, catheter handling and functionality. Cardiovasc Intervent Radiol. 1998;21(5):404–410. 108. Feng L, Dumoulin CL, Dashnaw S, et al. Feasibility of stent placement in carotid arteries with real-time MR imaging guidance in pigs. Radiology. 2005;234(2):558–562. 109. Yang X, Bolster Jr BD, Kraitchman DL, Atalar E. Intravascular MRmonitored balloon angioplasty: an in vivo feasibility study. J Vasc Interv Radiol. 1998;9(6):953–959. 110. Buecker A, Adam GB, Neuerburg JM, et al. Simultaneous real-time visualization of the catheter tip and vascular anatomy for MR-guided PTA of iliac arteries in an animal model. J Magn Reson Imaging. 2002;16(2):201–208.

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111. Godart F, Beregi JP, Nicol L, et al. MR-guided balloon angioplasty of stenosed aorta: in vivo evaluation using near-standard instruments and a passive tracking technique. J Magn Reson Imaging. 2000;12 (4):639–644. 112. Omary RA, Frayne R, Unal O, et al. MR-guided angioplasty of renal artery stenosis in a pig model: a feasibility study. J Vasc Interv Radiol. 2000;11(3):373–381. 113. Paetzel C, Zorger N, Bachthaler M, et al. Feasibility of MR-guided angioplasty of femoral artery stenoses using real-time imaging and intraarterial contrast-enhanced MR angiography. Rofo. 2004;176 (9):1232–1236. 114. Buecker A, Neuerburg JM, Adam GB, et al. Real-time MR fluoroscopy for MR-guided iliac artery stent placement. J Magn Reson Imaging. 2000;12(4):616–622. 115. Dion YM, Ben El Kadi H, Boudoux C, et al. Endovascular procedures under near-real-time magnetic resonance imaging guidance: an experimental feasibility study. J Vasc Surg. 2000;32(5):1006–1014. 116. Mahnken AH, Gunther RW, Tacke J. Radiofrequency ablation of renal tumors. Eur Radiol. 2004;14(8):1449–1455. 117. Manke C, Nitz WR, Djavidani B, et al. MR imaging-guided stent placement in iliac arterial stenoses: a feasibility study. Radiology. 2001;219(2):527–534. 118. Raval AN, Karmarkar PV, Guttman MA, et al. Real-time MRI guided endovascular recanalization of chronic total arterial occlusion in a swine model [In Press]. Circulation. 2006;. 119. Sampath S, Raval AN, Lederman RJ, McVeigh ER. High-resolution 3D arteriography of chronic total peripheral occlusions using a T1W turbo spin-echo sequence with inner-volume imaging. Magn Reson Med. 2007;57:40–49. 120. Frahm C, Gehl HB, Lorch H, et al. MR-guided placement of a temporary vena cava filter: technique and feasibility. J Magn Reson Imaging. 1998;8(1):105–109. 121. Bartels LW, Bos C, van Der Weide R, Smits HF, Bakker CJ, Viergever MA. Placement of an inferior vena cava filter in a pig guided by high-resolution MR fluoroscopy at 1.5 T. J Magn Reson Imaging. 2000;12(4):599–605. 122. Bucker A, Neuerburg JM, Adam GB, et al. Real-time MR guidance for inferior vena cava filter placement in an animal model. J Vasc Interv Radiol. 2001;12(6):753–756.

Pediatric Interventional Cardiovascular Magnetic Resonance Sanjeet R. Hegde and Reza S. Razavi

The last two decades have seen phenomenal advances made in the field of cardiovascular magnetic resonance (CMR), and these advances have fueled research into interventional applications for this remarkable imaging modality.1,2 Conventional X-ray fluoroscopically guided cardiac catheterization and interventions carry a substantial risk of exposure to ionizing radiation for both patients and staff. This is particularly relevant in younger patients, who are often required to undergo multiple procedures. The need for an imaging modality that offers multiplanar imaging, superior structural delineation of complex cardiac anatomy, and additional physiologic information, without the risk of ionizing radiation, has brought CMR guidance to the fore. In the last 4 years, clinical programs using CMR-guided cardiac catheterization have started and show promise.3 After the first MR images showing live human anatomy were produced,4–6 this technique evolved to enable a variety of clinical applications of MR.7,8 Over the years, improvements in signal detection, fast data handling, advanced understanding of spin systems, pulse sequences, and artifact suppression have resulted in much faster scan times and considerable improvements in image resolution.9–17 These ultrafast imaging techniques form the basis of real-time imaging, used for CMRguided cardiac catheterization. However, the first important step in making CMR cardiac catheterization a clinical reality is the design of a suitable interventional CMR system.18,19

INTERVENTIONAL CARDIOVASCULAR MAGNETIC RESONANCE SYSTEMS In the design of an interventional CMR suite, it is important to retain the full capabilities of a state-of-the-art diagnostic scanner without encumbering the interventionalist or creating a risk of high radiofrequency (RF) or switched magnetic field exposure. Open-magnet designs allow easier access to the patient, but typically are not available in field strengths higher than 1 Tesla (T). The cylindrical horizontal bore systems offer higher field strengths and gradient slew rates, allowing higherresolution imaging, shorter scan times, higher signal-to-noise ratio, reduced image distortion, and improved functionality with real-time imaging, all of which are of paramount importance when endovascular interventions are considered.20

A trade-off with the traditional cylindrical magnet design is access to the patient. More recently, magnets with shorter bores and flared margins have been introduced and offer better patient access, especially for cardiovascular interventions, without compromising the advanced CMR features of diagnostic scanners. Rapid improvements in the processing power of computers, along with the use of powerful and intuitive software, have allowed researchers to develop novel strategies for image data acquisition and reconstruction. It is now possible to achieve frame rates of as high as 20 images/sec with the aid of new parallel imaging techniques while maintaining suitable spatial resolution for interventional applications.21–24 Despite the inherent potential and promise of CMR-guided interventions and operations, there are still major obstacles associated with performing the complete procedure in the CMR scanner, particularly because of the lack of CMRcompatible catheters and devices. Therefore, the immediate future of interventional CMR lies in exploiting multi-modality imaging, such as X-ray and CMR (XMR) or XMR and ultrasound. Such hybrid units already in existence allow the use of separate modalities or a combination of them when needed. Cross-modality image integration, with spatial and temporal information about the anatomy, pathology, and therapy devices, can be provided to the users of these systems. A good example is the XMR system, which combines X-ray and CMR by having both modalities in the same room, with a tabletop design that allows patients to be moved from one modality to the other in less than 1 minute (Fig. 44-1).25–28

MERITS OF CARDIOVASCULAR MAGNETIC RESONANCE GUIDANCE Improved Visualization of Cardiac Anatomy A problem with X-ray-guided cardiac catheterization is the inherent poor contrast of soft tissues, such as the heart and great vessels. This makes it difficult for the cardiologist to manipulate or position guidewires, catheters, balloons, or interventional devices within the heart and surrounding Cardiovascular Magnetic Resonance 593

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CHAPTER 44

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Staff

Patient

A

B Figure 44-1 A, Schematic room plan of a typical X-ray and magnetic resonance (XMR) suite. B, XMR room with the X-ray and MR equipment joined by a movable tabletop. The c-arm of the X-ray unit is seen in the foreground, ceiling-mounted MR monitor and controls are seen in the distance, and the 5-gauss area is demarcated by a change in the floor coloring from the MR to the X-ray end of the room.

vessels. A skilled operator usually relies on recognizing anatomic structures from previous experience or on contrast angiographic images acquired earlier in the procedure. The lack of adequate visualization increases the risk of perforating the heart or great vessels, especially when performing complex interventional procedures. Certain interventional cardiac procedures involve selection of an appropriate cardiac device and its successful deployment within the heart, which requires accurate measurement of the size of defects and nearby anatomic structures. Such measurements are possible under XRF, but can be difficult. A successful interventional cardiac procedure therefore relies heavily on adequate visualization of the heart or vessel. This implies the need for superior imaging methods that provide excellent visualization without increasing the risk to the patient. This role fits CMR very well because it provides exceptional structural delineation of both the heart and its surrounding vasculature and therefore allows safe guidance of interventional procedures. 594 Cardiovascular Magnetic Resonance

Reduced Ionizing Radiation There is a pressing need for pediatric cardiac catheterization procedures to be made safer, especially in terms of ionizing radiation. According to the U.K. National Radiation Protection Board, the mean risk that a solid tumor will develop as a result of a single cardiac catheterization procedure is approximately 1 in 2500 in adults. This risk increases to 1 in 1000 in children if exposure occurs at 5 years of age.29–32 Also, the proportion of the body that is irradiated increases as the size of the patient decreases, and some procedures in patients with congenital heart disease often require much longer X-ray exposure. These risks are multiplied in children in particular because they often undergo multiple cardiac catheter procedures. In addition to the patients, there is also a significant risk from ionizing radiation to the staff in the catheter laboratory during these procedures, despite the use of protective shields.33,34

Cardiac catheterization is used not only to provide anatomic information and perform intervention but also to obtain functional information. Invasive pressures and blood gases are used to calculate systemic and pulmonary blood flow and resistance with the Fick principle. Cine angiography is also used to assess global ventricular function as well as regional wall motion abnormalities. The functional information obtained at cardiac catheterization is used alongside anatomic information to assess patient suitability for surgery or interventional cardiac catheterization or the need for long-term vasodilator therapy in patients with pulmonary vascular disease. The Fick principle to quantify flow is dependent on multiple measurements (hemoglobin, aortic/pulmonary artery oxygen saturation, partial pressure, oxygen consumption), which can be a considerable source of inaccuracy. In addition, in patients with large intracardiac shunts and high pulmonary blood flow, accuracy is further reduced.35–40 Therefore, there is a need for a method of flow quantification that allows accurate and reproducible measurement of pulmonary vascular resistance (PVR). Velocity encoded phase contrast CMR enables noninvasive quantification of blood flow in major vessels. Cardiac output and the pulmonary-to-systemic flow ratio (Qp:Qs) measured using this technique have been shown to be accurate.41–47 In addition, phase CMR has been validated in numerous phantom experiments, allowing for a novel method of quantification of PVR in patients with pulmonary hypertension by using invasive pressure measurements and MR flow data.48–50 Assessment of global and regional ventricular function can also be carried out much more accurately with cine steady-state free precession (SSFP) cardiovascular MR than with X-ray angiography. When using cardiovascular MR for assessing global ventricular function, there is no need to make assumptions about cardiac geometry, unlike with XRF or even echocardiography. This is particularly important when assessing right ventricular function (RV) and regional wall motion. Finally, combining invasive pressure measurement with CMR-derived blood flow and ventricular volumes also opens up interesting new ways of looking at pathophysiology. It allows for the study of pulmonary vascular compliance, derived ventricular pressure-volume loops, and assessment of load-independent ventricular function.51,52

MAGNETIC INSTRUMENTATION AND VISUALIZATION STRATEGIES Crucial to the success of interventional CMR is real-time tracking and visualization of catheters, guidewires, and devices in the CMR environment. Several groups around the world are putting considerable effort into developing CMR-suitable catheters and devices. Device localization under CMR is made possible by a variety of approaches that can be broadly classified as either electrically passive or electrically active.53

Passive Catheter Tracking and Visualization The passive tracking technique is commonly based on visualization of susceptibility artifacts or signal voids caused by the interventional device under CMR imaging. This is a well-studied technique and to date it is the most clinically feasible (see Fig. 44-3).54–58 Passive visualization often does not require any special hardware or software and therefore it can be performed on any commercial CMR system. The ideal passive tracking catheter or guidewire must be made of a material that provides adequate torque and allows tracking, but does not obscure the underlying anatomy. Ferromagnetic materials cause large susceptibility artifacts and therefore are not generally suitable for CMRguided procedures. This rules out most metals used for making cardiac devices. However, certain alloys, such as nitinol (nickel and titanium), have magnetic susceptibility close to that of tissue. Therefore, they are best suited for making guidewires and braided catheters that are MR compatible but not necessarily CMR safe. The polymeric materials used for making catheters typically have low magnetic susceptibility and therefore cannot be easily localized on CMR images.59 This implies that, if materials with higher susceptibility can be incorporated into the wall of the catheters or sheaths or the lumen filled with a suitable contrast agent, then improved visualization can be achieved. One approach to generating susceptibility artifacts is locally impregnating the catheter wall with gadolinium-like compounds, such as dysprosium oxide, in the form of rings or along the length of the catheter during the extrusion process (Fig. 44-2C).58 Another approach is to use gadolinium contrast agents in varying concentrations within catheter lumens60 or impregnated into catheter walls to create either a positive or negative signal on CMR imaging.61 Metallic devices and guidewires produce susceptibility artifacts that aid visualization by way of the artifacts, but different metals behave differently under CMR. Titanium alloys produce narrower artifacts compared with ferromagnetic or even certain other nonferromagnetic alloys, such as nickel-chromium, which can produce large RF and susceptibility artifacts. Guidewires with a fiberglass core and nonmetallic guidewires made of resin microparticle compound covered by polytetrafluoroethylene have been used for MRguided interventions in animals.62,63 In the case of balloon angiographic catheters, if the balloon is inflated with carbon dioxide, as is done conventionally with X-ray, then the inflated balloon creates a signal void in the CMR image, thus enabling visualization (see Fig. 44-2A and B). This method has been used successfully to guide catheters in patients under CMR (Fig. 44-3).3,64 Although this technique allows easy visualization of the tip, the length is impossible to visualize because the signal void from the catheter length is masked by volume averaging and dephasing effects of thicker slices.65 The success of passive visualization also relies on dedicated scan techniques. A dynamic gradient echo sequence, such as SSFP, has been shown to be ideal for passive catheter tracking, especially when signal voids or susceptibility artifacts are used for visualization.64,66 Cardiac catheterization under XRF guidance is usually performed at imaging Cardiovascular Magnetic Resonance 595

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Physiologic Information

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A

B

C

Figure 44-2 Passive tracking. A, Inflated balloon angiographic Bermann catheter filled with 0.8 mL CO2. B, CO2-filled balloon catheter manipulated in a phantom. C, Dysprosium catheter: a catheter impregnated with dysprosium oxide is manipulated in an in vitro set-up mimicking endovascular intervention. The catheter is clearly visualized along its full length, despite being orientated along B0. (A, Courtesy of Arrow International, Reading, PA.)

speeds of 25 to 30 frames/sec. The frame rates available for CMR-guided interventions are not comparable because of the post-processing of CMR images and their subsequent display, allowing a maximum of 10 to 14 frames routinely. Some of the proposed passive catheter tracking techniques require image subtraction or positive contrast to improve visualization of markers on the catheter, which means that, along with faster scan techniques, faster image processing algorithms are required.57,67–70

Active Catheter Tracking and Visualization The active catheter tracking and visualization method uses an electrical connection to the CMR scanner, and localization or tracking of the device requires the device itself, along with any additional hardware or software that comes with it. Typically, the device is equipped with a coil or an antenna that functions in either receive-only mode or transmit/receive mode. Active catheters that are used as receivers have a coil or an antenna that receives signal from tissue in its immediate vicinity.71 These devices do not transmit signal into the patient, but rely on the body coil to transmit into the patient. The signal received by these coils can then be used to pinpoint their position, for imaging of local tissue, or both. There are two important types of active catheters: those based on small coils positioned, for example, at the end of a catheter, and those based on a loopless antenna that can run along a catheter or can be made into a guidewire (Fig. 44-4).72–76 In addition, active designs in 596 Cardiovascular Magnetic Resonance

which signal voids along the catheter are created by electrically controlled magnetic field inhomogeneities have also been investigated.77 A small resonant coil at the tip of a catheter can be identified by a series of three one-dimensional projections along each axis.71 This can be done quickly (in three repetition times) and so could be repeated for very fast update of the catheter position, allowing real-time tracking of the catheter. The position of the catheter could then be projected over a previously acquired road map. Recently, similar techniques have been combined with fast/real-time sequences, imaging the heart or vessels using surface coils, and the combined (interleaved) sequence has allowed simultaneous localization of the catheter and imaging of the surrounding tissue. Further adaptation of these sequences has allowed automatic changing of the imaging plane to match the change in the position of the catheter. Another recent development of active catheter tracking by the group at the U.S. National Institutes of Health allows the visualization of two simultaneously acquired planes as well as visualization of the catheter or device positions in real time, thus reducing the major problem of the catheter moving through the plane when only one imaging plane is visualized.78–80 The great advantage of these active systems is that location of the catheter is unambiguous. Active visualization has great potential because it allows the whole length of the catheter or guidewire to be visualized and the imaging plane to be adapted to the moving catheter automatically. It may even allow high-resolution imaging of a small area of interest, such as a plaque in the vessel, when the coil or antenna is used in its imaging mode.81 However, the main disadvantage is concern with safety.82–87

44 PEDIATRIC INTERVENTIONAL CARDIOVASCULAR MAGNETIC RESONANCE

Figure 44-3 Manipulation of carbon dioxide-filled balloon catheter (arrows) from the inferior vena cava to the right pulmonary artery using solely magnetic resonance guidance. Real-time interactive images: repetition time 2.9 msec, echo time 1.45 msec, flip angle 45 , matrix 128 128, field of view 250 to 350, and temporal resolution 10 to 14 frames/sec. Arrows show the signal void of the catheter tip as it traverses the inferior vena cava, right atrium, tricuspid valve, and right ventricular outflow tract and enters the pulmonary artery.

These devices use intravascular coils as RF antennas, and the connection to the external circuits via a long wire in the strong magnetic field makes induction of an electrical current and heating possible. There have been recent developments to overcome this risk, such as electrical decoupling of loopless antennas and the use of optical coupling and long fiber optic connections.88 An innovative active catheter design that uses miniaturized transformers showed no

significant RF heating and holds promise for a safe transmission line for interventional applications (Fig. 44-5).89 Another promising approach to device localization is what some authors refer to as semi-active catheter tracking, implying passive localization of an electrically isolated resonant coil.53 These resonant coils locally enhance B1 and signal reception so that, for very low global flip angles, the signal from the fiducial is prominent.90–95 Cardiovascular Magnetic Resonance 597

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Figure 44-4 Active catheter designed for intramyocardial injection. (Courtesy of Dr. Parag Karmarkar, Johns Hopkins University, Baltimore, MD.)

The resonant coils can be interrogated by gradient echo sequences, such as SSFP with low flip angles. Catheters with multiple resonant coils can be tracked easily compared with passive catheters and have a relatively better safety profile compared with some of the active catheter designs (Fig. 44-6). Catheter visualization and localization using 19 F CMR in conjunction with proton imaging appears to be a promising alternative to existing methods that either are associated with safety concerns if active markers are used or have insufficient direction-dependent contrast if passive visualization is used (Fig. 44-7).96 Other multispectral CMR methods under investigation are catheter tracking and angiography using hyperpolarized gases.97,98

SAFETY ISSUES Bioeffects of Magnetic Fields The patient undergoing a CMR scan typically is exposed to three forms of electromagnetic radiation: static magnetic field, gradient magnetic field, and RF electromagnetic field. These can cause bioeffects at significantly high exposure levels. A health care worker in such a setting can also be exposed to electromagnetic fields, although exposure is more chronic and intermittent. However, numerous studies have shown no substantial risks to patients from the electromagnetic fields used in clinical CMR scanners.99–102 The risks to the health care worker, especially in a CMR setting, are fiercely debated, but the consensus is that more work needs to be carried out before occupational electromagnetic field exposure limits can be set.103,104 Furthermore, the bioeffects specifically related to the use of interventional CMR have not yet been fully investigated. Many reports in the literature regarding the bioeffects of static magnetic fields are conflicting. There is no strong evidence to suggest that there are any significant cardiac or neurologic effects from static magnetic fields of less than 2 T. In addition, several studies have shown that high static magnetic fields do not significantly alter skin and body temperature.105–110 598 Cardiovascular Magnetic Resonance

Gradient magnetic fields can induce electrical fields and current in conductive media, including biologic tissue, according to Faraday law of induction. The thermal effects of switched magnetic fields are considered negligible and are not believed to be clinically significant. Electrical stimulation of the retina is believed to cause magnetophosphenes, which are completely reversible, with no known residual side effects. Some volunteers have also reported experiencing a metallic taste and vertigo while undergoing imaging within 4 T magnets. These bioeffects caused by gradient fields are unusual in fields of less than 2 T.111 The exposure limits for RF radiation are set in terms of specific absorption rate in Wkg-1, which is the mass normalized rate at which RF power is coupled with biologic tissue. The main bioeffects associated with exposure to RF radiation relate to the generation of heat in tissues. Controversially, some researchers have reported that electromagnetic fields cause cancer and developmental abnormalities in animal models. However, the efficiency and absorption pattern of RF radiation is mainly determined by the physical dimensions of the tissue in relation to the incident wavelength, which implies that laboratory animal experiments cannot be simply scaled or extrapolated to humans.112–114

Heating and Electrical Safety of Interventional Equipment The heating of wires, devices, implants, and other instruments is an important safety issue that is holding back the rapid advance of interventional CMR. Heating as a result of RF radiation occurs by three mechanisms, according to Maxwell’s theory of electromagnetism.82 When a conductive device or instrument is moved through a magnetic field, small “rings” of current are induced that are called eddy currents and create internal magnetic fields opposing the change. The kinetic energy that goes into driving the eddy currents inside the metal will give off that energy as heat. Therefore, intravascular guidewires or device delivery systems with a metal core are unsafe in the CMR environment, with documented heating up to 74 C (165 F) of the tip.82,83,115,116 Electromagnetic induction heating has often been blamed for thermal injuries caused by monitoring cables used in CMR. RF electromagnetic fields and time-varying gradient magnetic fields can induce voltage in conductive media and cause current to flow. The circulating currents cause power loss by heating that is referred to as induction heating. A loop in a monitoring cable would increase the inductance of the circuit; therefore, larger currents would be induced, resulting in greater heating of the cable.87,117,118 If a circuit is in a resonant state, then there is maximum current induction such that significant electromagnetic induction heating occurs. Lengths of wire, for example, can behave as RF antennas that capture electromagnetic waves to extract power from them. The electromagnetic waves that enter the antennas have electrical charges and corresponding currents associated with them. When the antenna is approximately half a wavelength long, resonance

Coaxial Transformer

Transformer loop

A

Passive markers

Cross at tip coil

B Figure 44-5 A, Safe transmission line for active catheter tracking created with integrated miniaturized transformers. B, To evaluate the transformer concept for active tracking in vivo, a 6 Fr catheter was built for catheterization of the arterial and venous system of a swine. The catheter is seen being manipulated in the heart by the active tracking method. (Courtesy of Dr. Steffen Weiss, Philips Research, Hamburg, Germany.)

occurs and the electrical energy remains confined to the immediate vicinity of a given antinode. Hence, the highest electrical field of the antennas is believed to be at the tip. The electrical properties of the media surrounding the

antennas and the operating frequency also determine the wavelength.85,86 Newer designs of wires and cables aimed at reducing heating are currently being investigated, along with novel RF shielding technologies.119,120 Cardiovascular Magnetic Resonance 599

44 PEDIATRIC INTERVENTIONAL CARDIOVASCULAR MAGNETIC RESONANCE

6F = 2 mm

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A

B

C Figure 44-6 A, A 5 Fr balloon angiographic catheter with six prewound fiducial markers mounted onto the surface was manipulated in a 20-mm polyethylene tube taped to the chest of a volunteer. A real-time spoiled gradient echo sequence (fast field echo [FFE]: repetition time 2.3 msec, echo time 1.2 msec, flip angle 50 , slice thickness 20 mm) followed by an interactive FFE sequence with interleaving of scans with flip angles of 2 and 50 and a frame rate of 4 frames/sec was used. All six markers are visualized along the length of the catheter. B, Distal end of a 6 Fr catheter with an integrated self-resonant radiofrequency circuit. C, The active wireless catheter is shown being guided with real-time projection reconstruction steady-state free procession imaging into the celiac trunk. (A, Courtesy of Arrow International, Reading, PA. B and C, Courtesy of Dr. Harald H. Quick, University of Essen, Germany.)

Magnetic Force and Torque In addition to the bioeffects of CMR and heating and electrical safety of interventional devices, a significant risk to interventional procedures is magnetic force and torque exerted by the magnetic field on metallic devices.121,122 Conventional guidewires made of ferromagnetic materials, such as stainless steel, and catheters with metallic braiding, are inherently unsafe for use in the CMR environment. Interventional devices that are ferromagnetic will be subject 600 Cardiovascular Magnetic Resonance

to both deflection force (translational movement) and torque (rotational movement); therefore, they cannot be used for procedures within a CMR scanner. Hence, all CMR imaging facilities must have safeguards to ensure that ferromagnetic objects are not brought into the vicinity of the magnet. However, there are certain other metallic alloys, such as nitinol, that are CMR compatible. They produce minimal susceptibility artifacts and are not affected by the magnetic field in terms of deflection force and torque. This is an important consideration in developing suitable catheters and guidewires for use in interventional CMR procedures.

44 PEDIATRIC INTERVENTIONAL CARDIOVASCULAR MAGNETIC RESONANCE

Figure 44-7 Corresponding 1H (top left) and 19F (top right) images of a 7 Fr catheter containing PFOB. With a simple peak search algorithm in the image space, the catheter tip position was extracted and two orthogonal 19F projections were used to determine the position of the catheter tip (þ), as shown. Source: Kozerke S, Hedge S, Schaeffter T, Lamerichs R, Razavi R, Hill DL. Catheter tracking and visualization using 19F nuclear magnetic resonance. Magn Reson Med. 2004;52(3):693–697.

X-RAY AND CARDIOVASCULAR MAGNETIC RESONANCE GUIDANCE X-Ray and Cardiovascular Magnetic Resonance Facility Design The room design of a typical XMR facility is shown in Figure 44-1. There are many design features that make this room different from standard CMR facilities. The XMR suite is designed so that half of the room is outside the 5-gauss line of the magnet, permitting the use of traditional instruments and devices as well as echocardiography and RF ablation equipment when required. A movable

tabletop allows patients to be moved easily between modalities in less than 60 seconds. The paramount consideration in the design, construction, and operation of an XMR facility is safety, and a comprehensive safety protocol must be drawn up to minimize possible hazards (Table 44-1). Traditionally, CMR scans are planned and conducted from the control room, away from the magnet and the patient. However, during CMR-guided cardiac catheterization, there is a need for real-time changes to the scanning sequence parameters to follow catheter manipulation in the heart and great vessels. Also, the imager needs to have a clear view of the CMR images while performing the procedure. Therefore, it is useful to have a fully functional set of ceilingmounted, movable screens and scanner controls within the CMR scanner room that can be placed at either end of the bore of the scanner, in close proximity to the patient. The XMR suite includes appropriate CMR-compatible anesthetic equipment and monitoring equipment for invasive pressure monitoring via the catheter. A great deal of thought has been given to the safety of patients under Cardiovascular Magnetic Resonance 601

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Table 44-1 X-Ray and Cardiovascular Magnetic Resonance (CMR) Facility: Safety Features Compulsory safety training of all CMR interventional staff Specially designed clothes without pockets Safety officer restricting entry to the main room during XMR intervention Clear demarcation of ferromagnetic safe and unsafe areas within the room CMR-compatible anesthetic and monitoring equipment Noiseproof headphone systems for all staff within the room X-ray- and radiofrequency-shielded scrub room Positive pressure air handling and filtration system Tethering of all ferromagnetic equipment to the wall or floor Safety checks whenever a patient is transferred between X-ray and CMR to ensure that metallic instruments used for catheterization are not taken across to the CMR end of the room Written log of all safety infringements and regular review of safety procedures

anesthesia, especially during the transfer between the X-ray and CMR tables. All of the anesthetic and monitoring tubing and lines are designed with extra length and are secured to the movable tabletop to ensure smooth patient transfer. The electrocardiogram (ECG) and invasive pressure data are sent from the MR-compatible monitoring equipment via an optical network to a computer in the control room, where the cardiac technician is stationed. The appropriate measurement and recording of the data is made in the usual way. The technician has access to monitors that show the appropriate X-ray or CMR images of the procedure. The imagers in the room can view the CMR images and any monitoring data (e.g., ECG, invasive pressure data). Blood samples taken during the procedure are labeled in the room and passed to the technician in the control room via a wave guide. Reliable and accurate ECG synchronization is essential for CMR and in particular CMR-guided cardiac catheterization. When catheters are manipulated in the heart, there is the potential to cause arrhythmias (tachyarrhythmia or heart block). It is therefore important to perform accurate monitoring of the cardiac rhythm at all times during XMR catheterization. Obtaining a reliable ECG in the magnet, particularly during some CMR sequences, can be difficult. The magnetohydrodynamic effect and gradient noise can seriously disturb the ECG signal.123,124 This interference can reduce trigger signal reliability to less than 40%. Vector electrocardiogram (VCG) is a QRS detection algorithm that automatically adjusts to the actual electrical axis of the patient’s heart and the specific multidimensional QRS waveform. In our experience, this greatly improves the reliability of R-wave detection to nearly 100%. A reliable R-wave, with the P- and T-waves that are also always clearly seen with VCG, allows detection of nearly all arrhythmias. Unfortunately, there are no ECG systems that can reliably provide ST segment or T-wave morphologic information. In the future, using signal processing techniques, it may be possible to obtain ECG during CMR scanning that provides ST and T-wave information reliably. Another complication of performing cardiac catheterization under CMR guidance is the noise generated during scanning. There is a headphone and microphone system in the room that reduces the noise, but allows staff to communicate with each other in both the scanner and control rooms. Some CMR coils have X-ray-visible components and would need to be removed between CMR imaging and 602 Cardiovascular Magnetic Resonance

X-ray imaging of patients. It is therefore necessary to have specifically designed coils that are sufficiently radiotranslucent to be left in place during XRF without deterioration of image quality. We use these coils in our procedures so that patients do not have to be disturbed when moving from one imaging modality to the other.125 The XMR suite has positive-pressure air handling and filtration appropriate for a catheterization laboratory. There is a scrub room that is also RF and X-ray shielded and can be accessed both from the XMR suite and control room. This room acts as an RF lock, allowing access to the XMR suite during CMR scanning.

Performing X-Ray and Cardiovascular Magnetic Resonance Interventions In a typical XMR interventional procedure, after the induction of anesthesia, the patient is transferred from an MRcompatible trolley to the CMR end of the XMR facility and positioned on the CMR scanner tabletop (Fig. 44-8A). The monitoring and anesthetic equipment are attached. A three-lead ECG, separate from the VCG, is used for cardiac monitoring during MR scanning. The VCG electrodes are placed on the subcostal margin, outside the X-ray field of view, and the VCG is used for triggering CMR scans. An MR-compatible pulse oximeter and noninvasive blood pressure monitoring equipment are also attached. The exhaled anesthetic gases are monitored for end-tidal carbon dioxide as well as the concentration of the volatile anesthetic agents. Flexible phase array RF coils are used. These coils are relatively X-ray lucent and thus do not need to be removed between MR and X-ray imaging. The patient is then placed in the CMR scanner, and a multi-breath-hold three-dimensional (3D) SSFP scan of the heart and great vessels (echo time 2, repetition time 4, flip angle 50, 80 to 120 slices reconstructed to 1-mm cubic voxels) is obtained.126 Using an interactive SSFP sequence (8 to 10 frames/sec), with real-time manipulation of scan parameters, the likely imaging planes needed for subsequent catheter tracking, ventricular function, and flow quantification are stored. The patient is then transferred to the X-ray end of the room. Draping and vascular access are carried out as for routine cardiac catheterization; in addition, a second large drape is placed over the patient (see Fig. 44-8B). The patient is transferred back to the MR scanner after safety checks are performed, including an operating theater-style check of all metallic objects used under X-ray. The second drape is then lifted up and taped to the top of the magnet, which in effect provides sterile draping of the bore and sides of the magnet (see Fig. 44-8C). An end-hole or side-hole balloon angiographic catheter (4 to 7 Fr) is placed in the sheath, and with the balloon inflated with CO2 (see Fig. 44-3), the catheter tip is passively visualized using the interactive sequences described earlier. The previously stored imaging planes are used, along with interactive slice selection, to track the catheter. Because only the tip of the catheter is visualized, care is taken not to push the catheter too fast and thus beyond the CMR imaging plane. This also ensures that the catheter does not accidentally form loops and possible knots.

Early Experience in Humans

B

C Figure 44-8 X-ray and magnetic resonance intervention. A, Patient is placed on the magnetic resonance tabletop. B, Patient is slid across to the X-ray half of the room for sheath insertion. C, Passive catheter manipulation is performed under magnetic resonance guidance.

A duplicate CMR control console is positioned next to the bore of the magnet so that the interactive window can be easily visualized while the catheter is being manipulated. Therefore, this procedure requires two experienced operators, one to move the catheter and one to alter the CMR imaging planes to ensure that the catheter tip is tracked, using the real-time interactive sequence.

In our center, we have performed 53 diagnostic CMR-guided cardiac catheterizations. In most patients, the majority of the procedure was carried out under MR guidance, which allowed for a significant reduction of overall X-ray dose.3,50 We used CMR to assess pulmonary vascular resistance in the patients because it allowed for simultaneous measurement of pulmonary arterial flow and invasive pressures. We found moderate to good agreement between the Fick method and the CMR method of deriving PVR at baseline conditions. However, in the presence of nitric oxide, which is used to assess pulmonary vasoreactivity, there was less agreement between the two methods. There was not only worsening in agreement but also a large bias when PVR was measured in the presence of 100% oxygen and nitric oxide. We believe that this is the result of errors in the Fick method rather than the XMR method, which has important implications for patient management. This novel MR technique may prove to be a more accurate method to quantify PVR in humans; it also offers reduced exposure to ionizing radiation.50,127 Currently, RF ablation is used to treat some patients with symptomatic atrial or ventricular tachyarrhythmia. This is conventionally carried out under XRF or ultrasound guidance using electrical and electromagnetic mapping. XRF guidance is conventionally used to guide such procedures because it offers excellent temporal resolution and good visualization of catheters. However, as a projection imaging modality, more than one view is necessary to gain an appreciation of the 3D location and path of catheters. This implies moving the X-ray c-arm to obtain different projections. A few centers use a biplane X-ray system for the same purpose. The anatomic context of the acquired images can be difficult to interpret because soft tissues, such as the heart and blood vessels, are not visible during X-ray exposure. Therefore, we developed a real-time XMR guidance system for cardiovascular interventions that allows the use of both CMR and X-ray imaging for guidance, thereby overcoming some of the failings of exclusive XRF guidance.128 Cardiovascular Magnetic Resonance 603

44 PEDIATRIC INTERVENTIONAL CARDIOVASCULAR MAGNETIC RESONANCE

A

Once the catheter is positioned in the desired vessel or chamber, appropriate pressure data and saturation/blood gas samples are obtained, as for routine cardiac catheterization. In addition, ventricular function (short axis balanced SSFP) and flow (phase contrast) scans can be performed using the appropriate previously stored imaging planes. If catheter manipulation into a particular heart chamber or vessel using CMR guidance alone is difficult, the patient is transferred back to the X-ray end of the room, where catheterization can be continued under XRF (e.g., to use a guidewire or a braided catheter). The patient can be transferred back to the CMR scanner for further CMR measurements once the catheter is positioned satisfactorily. An interventional procedure or RF ablation of arrhythmias requires part of the procedure to be performed under X-ray fluoroscopy because the ablation catheters and delivery devices are not MR compatible. Therefore, MR imaging is performed at the beginning of the procedure for planning purposes, used in guiding the procedure, and performed again at the end of the procedure for evaluation.

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A combination of calibration and real-time tracking is used to achieve X-ray and CMR image registration. A Northern Digital Optotrak 3020 (NDI, Ontario, Canada) using infrared-emitting diodes optically tracks the X-ray c-arm and the X-ray table. The sliding table is tracked by the CMR system software while docked to the CMR scanner and becomes part of the X-ray table when docked to the X-ray system. Once calibrated, the system allows registration of 3D CMR image acquisitions to X-ray image acquisitions. During intervention, the guidance system can provide a real-time MR anatomy overlay onto X-ray images. One monitor is used to display the control interface and the

A

second monitor shows the image overlay. During fluoroscopy, the system acquires X-ray images and computes the registration matrix from the tracking data at 10 frames/sec and updates the overlay display at 3 frames/sec. Using this unique XMR technology, we have carried out RF ablation in pulmonary veins, atria, and ventricles to treat arrhythmias successfully in 30 patients (Fig. 44-9).128 This CMR to X-ray registration method also allows us to relate the position of measured electrophysiology data to cardiac motion data from 3D CMR images. The XMR technology is also being used to perform stent implantation in patients with coarctation of the aorta (Fig. 44-10).

B

C

Figure 44-9 A and B, Biplane X-ray views of the linear ablating catheter in the left atrial roof position. C, Posterior three-dimensional view of the left atrium derived from a gadolinium cardiovascular magnetic resonance angiography scan. The green dots show the mapped locations of the linear ablating catheter in three positions: (1) left atrial roof position; (2) left upper pulmonary vein to mitral valve annulus position; and (3) right upper pulmonary vein to mitral valve annulus position.

A

Prestent

B

Poststent

Figure 44-10 Cardiovascular magnetic resonance angiography image superimposed onto the X-ray cardiac catheter image during stent implantation. A, Undilated stent and guidewire across the coarctation site. B, The combined images show that the implanted open stent lies in a satisfactory position, distal to the origin of the left subclavian artery and across the coarctation narrowing. Stent implantation was performed in the X-ray half of the X-ray and magnetic resonance facility. Magnetic resonance imaging was used before stent insertion to acquire the three-dimensional cardiovascular magnetic resonance angiography images and after the procedure (guidewires removed) to confirm satisfactory position of the stent and relief of aortic obstruction. 604 Cardiovascular Magnetic Resonance

Recent Progress Several groups have shown the immense potential of interventional CMR in animal models. The interventions that were shown to be feasible with passive and active catheter techniques include balloon angioplasty of arterial stenoses,129–134 stenting of vessels,92,135–138 and atrial septal puncture/septostomy.139,140 Device closure of atrial septal defects is another application that has been explored.141–144 CMR-guided percutaneous pulmonary and aortic valve stent implantation have also been performed successfully (Fig. 44-11).135,145 More complex interventions, such as percutaneous coronary catheterization and intervention, have also been demonstrated in healthy animals using CMR.146–149 Recently, the group in Berlin performed balloon dilation of aortic coarctation in patients under CMR guidance.134 Although this required some modification of the standard technique, with the likely arrival of new CMR-compatible guidewires and catheters, it should soon be possible to perform a number of interventions under CMR guidance in patients with congenital heart disease.

Future Directions Novel catheters and guidewires have made possible targeted intramyocardial injection of progenitor stem cells in myocardial infarction in animal models.80,150,151 Using real-time CMR and direct apical access in porcine hearts, prosthetic aortic valves were implanted in the beating heart.152 This breakthrough application may allow CMR guidance of minimally invasive extra-anatomic bypass and beating-heart valve repair. MR guidance of intramyocardial gene therapy is another exciting field.153

A

Three-dimensional electromechanical models of the heart have been created that allow simulation of cardiovascular pathologies to test therapeutic strategies and plan interventions (Fig. 44-12).154,155 Newer catheter-tracking techniques using inductively coupled RF coils or hyperpolarized 13C and visualization strategies using novel k-space sampling also hold promise.156–158

Figure 44-12 Patient undergoing X-ray and cardiovascular magnetic resonance-guided biventricular pacing. Composite image showing one slice of a cardiovascular magnetic resonance cardiac anatomic scan with a superimposed surface model of the left ventricle. Cardiac electrical modeling was used to estimate myocardial conductivity for the left ventricle. The conductivity is represented by the color coding, with blue showing areas of low conductivity and yellow showing areas of normal conductivity. The white region shows the area of scarring segmented from late enhancement magnetic resonance imaging. There is good correspondence with predicted low conductivity and the region of the scar.

B

Figure 44-11 Percutaneous aortic valve stent implantation in a swine before (A) and after (B) valve stent implantation under cardiovascular magnetic resonance. (Courtesy of Dr. Titus Kuehne, German Heart Institute, Berlin, Germany.) Cardiovascular Magnetic Resonance 605

44 PEDIATRIC INTERVENTIONAL CARDIOVASCULAR MAGNETIC RESONANCE

INTERVENTIONAL CARDIAC APPLICATIONS

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CONCLUSION As outlined in this chapter, CMR guidance of cardiac catheterization is feasible and has been shown to be safe. The potential benefits of this new technique include reduction of X-ray dose, accurate assessment of pulmonary vascular resistance, and better visualization of complex anatomy for both diagnostic cardiac catheterization and interventional cardiac catheterization, such as RF ablation of arrhythmias, using XMR images. However, a number of issues must be resolved before exclusive interventional CMR becomes routine clinical practice. These include practical issues, such as reduction of noise in the CMR environment, improved access to the patient, and better CMR-compatible patient monitoring equipment. There is also a pressing need for catheter and device manufacturers to produce tools specifically designed for CMR-guided cardiac catheterization. This need, combined with the cost associated with installing expensive XMR suites, is holding back the rapid advance of interventional CMR. However, the potential benefits of 3D anatomic guidance for interventional cardiologists, radiologists, and surgeons, including the useful additional

physiologic information and the ability to assess tissue response to therapy with CMR, makes this remarkable imaging modality unique and one that offers great promise for safe guidance of complex cardiovascular interventions.79

ACKNOWLEDGMENTS Some of the work described in this chapter was performed at Guy’s Hospital, London, United Kingdom, by a team of academic and clinical staff. The authors thank Tobias Schaeffter, Vivek Muthurangu, Marc Miquel, Kawal Rhode, Redha Boubertakh, Andrew Taylor, Aaron Bell, Caroline Kehoe, Victoria Parish, Maxime Sermesant, Derek Hill, Stephen Keevil, David Hawkes, Jas Gill, Shakeel Qureshi, Eric Rosenthal, Gerald Greil, Philipp Beerbaum, and Edward Baker in the Departments of Imaging Sciences, Pediatric and Adult Cardiology. We would also like to acknowledge Michael Barnet and other members of the Anesthetic Department; and John Spence, Stephen Sinclair, and Rebecca Lund and other staff from the Radiology Department who have provided considerable support.

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33. Bashore TM, Bates ER, Berger PB, Clark DA, Cusma JT, Dehmer GJ, et al. American College of Cardiology/Society for Cardiac Angiography and Interventions Clinical Expert Consensus Document on Cardiac Catheterization Laboratory Standards: A report of the American College of Cardiology Task Force on Clinical Expert Consensus Documents endorsed by the American Heart Association and the Diagnostic and Interventional Catheterization Committee of the Council on Clinical Cardiology of the AHA. J Am Coll Cardiol. 2001;37(8):2170–2214. 34. McCullough PA, Wolyn R, Rocher LL, Levin RN, O’Neill WW. Acute renal failure after coronary intervention: incidence, risk factors, and relationship to mortality. Am J Med. 1997;103(5):368–375. 35. Cigarroa RG, Lange RA, Hillis LD. Oximetric quantitation of intracardiac left-to-right shunting: limitations of the Qp/Qs ratio. Am J Cardiol. 1989;64(3):246–247. 36. Dhingra VK, Fenwick JC, Walley KR, Chittock DR, Ronco JJ. Lack of agreement between thermodilution and fick cardiac output in critically ill patients. Chest. 2002;122(3):990–997. 37. Hillis LD, Firth BG, Winniford MD. Variability of right-sided cardiac oxygen saturations in adults with and without left-to-right intracardiac shunting. Am J Cardiol. 1986;58(1):129–132. 38. van den Berg Jr E, Pacifico A, Lange RA, Wheelan KR, Winniford MD, Hillis LD. Measurement of cardiac output without right heart catheterization: reliability, advantages, and limitations of a left-sided indicator dilution technique. Cathet Cardiovasc Diagn. 1986;12(3):205–208. 39. Hillis LD, Firth BG, Winniford MD. Analysis of factors affecting the variability of Fick versus indicator dilution measurements of cardiac output. Am J Cardiol. 1985;56(12):764–768. 40. Dehmer GJ, Firth BG, Hillis LD. Oxygen consumption in adult patients during cardiac catheterization. Clin Cardiol. 1982;5(8): 436–440. 41. Beerbaum P, Korperich H, Barth P, Esdorn H, Gieseke J, Meyer H. Noninvasive quantification of left-to-right shunt in pediatric patients: phase-contrast cine magnetic resonance imaging compared with invasive oximetry. Circulation. 2001;103(20):2476–2482. 42. Beerbaum P, Korperich H, Gieseke J, Barth P, Peuster M, Meyer H. Rapid left-to-right shunt quantification in children by phase-contrast magnetic resonance imaging combined with sensitivity encoding (SENSE). Circulation. 2003;108(11):1355–1361. 43. Beerbaum P, Korperich H, Gieseke J, Barth P, Peuster M, Meyer H. Blood flow quantification in adults by phase-contrast MRI combined with SENSE–a validation study. J Cardiovasc Magn Reson. 2005;7(2): 361–369. 44. Firmin DN, Nayler GL, Klipstein RH, Underwood SR, Rees RS, Longmore DB. In vivo validation of MR velocity imaging. J Comput Assist Tomogr. 1987;11(5):751–756. 45. Kilner PJ, Manzara CC, Mohiaddin RH, Pennell DJ, Sutton MG, Firmin DN, et al. Magnetic resonance jet velocity mapping in mitral and aortic valve stenosis. Circulation. 1993;87(4):1239–1248. 46. Hundley WG, Li HF, Hillis LD, Meshack BM, Lange RA, Willard JE, et al. Quantitation of cardiac output with velocity-encoded, phasedifference magnetic resonance imaging. Am J Cardiol. 1995;75(17): 1250–1255. 47. Hundley WG, Li HF, Lange RA, Pfeifer DP, Meshack BM, Willard JE, et al. Assessment of left-to-right intracardiac shunting by velocityencoded, phase-difference magnetic resonance imaging: a comparison with oximetric and indicator dilution techniques. Circulation. 1995;91(12):2955–2960. 48. Mousseaux E, Tasu JP, Jolivet O, Simonneau G, Bittoun J, Gaux JC. Pulmonary arterial resistance: noninvasive measurement with indexes of pulmonary flow estimated at velocity-encoded MR imaging–preliminary experience. Radiology. 1999;212(3):896–902. 49. Kondo C, Caputo GR, Masui T, Foster E, O’Sullivan M, Stulbarg MS, et al. Pulmonary hypertension: pulmonary flow quantification and flow profile analysis with velocity-encoded cine MR imaging. Radiology. 1992;183(3):751–758. 50. Muthurangu V, Taylor A, Andriantsimiavona R, Hegde S, Miquel ME, Tulloh R, et al. Novel method of quantifying pulmonary vascular resistance by use of simultaneous invasive pressure monitoring and phase-contrast magnetic resonance flow. Circulation. 2004;110 (7):826–834. 51. Muthurangu V, Atkinson D, Sermesant M, Miquel ME, Hegde S, Johnson R, et al. Measurement of total pulmonary arterial compliance using invasive pressure monitoring and MR flow quantification during MR-guided cardiac catheterization. Am J Physiol Heart Circ Physiol. 2005;289(3):H1301–H1306. 52. Kuehne T, Yilmaz S, Steendijk P, Moore P, Groenink M, Saaed M, et al. Magnetic resonance imaging analysis of right ventricular pressure-

FUNCTIONAL CARDIOVASCULAR DISEASE

77. 78. 79. 80. 81. 82. 83. 84. 85. 86. 87. 88. 89. 90.

91. 92.

93.

94. 95. 96. 97. 98. 99. 100. 101. 102.

intravascular MRI using a novel catheter-based, opposed-solenoid phased array coil. Magn Reson Med. 2004;51(4):668–675. Glowinski A, Adam G, Bucker A, Neuerburg J, van Vaals JJ, Gunther RW. Catheter visualization using locally induced, actively controlled field inhomogeneities. Magn Reson Med. 1997;38(2): 253–258. Guttman MA, Lederman RJ, Sorger JM, McVeigh ER. Real-time volume rendered MRI for interventional guidance. J Cardiovasc Magn Reson. 2002;4(4):431–442. Lederman RJ. Cardiovascular interventional magnetic resonance imaging. Circulation. 2005;112(19):3009–3017. Lederman RJ, Guttman MA, Peters DC, Thompson RB, Sorger JM, Dick AJ, et al. Catheter-based endomyocardial injection with real-time magnetic resonance imaging. Circulation. 2002;105(11):1282–1284. Atalar E, Bottomley PA, Ocali O, Correia LC, Kelemen MD, Lima JA, et al. High resolution intravascular MRI and MRS by using a catheter receiver coil. Magn Reson Med. 1996;36(4):596–605. Konings MK, Bartels LW, Smits HF, Bakker CJ. Heating around intravascular guidewires by resonating RF waves. J Magn Reson Imaging. 2000;12(1):79–85. Liu CY, Farahani K, Lu DS, Duckwiler G, Oppelt A. Safety of MRIguided endovascular guidewire applications. J Magn Reson Imaging. 2000;12(1):75–78. Yeung CJ, Atalar E. RF transmit power limit for the barewire loopless catheter antenna. J Magn Reson Imaging. 2000;12(1):86–91. Yeung CJ, Atalar E. A Green’s function approach to local rf heating in interventional MRI. Med Phys. 2001;28(5):826–832. Yeung CJ, Susil RC, Atalar E. RF heating due to conductive wires during MRI depends on the phase distribution of the transmit field. Magn Reson Med. 2002;48(6):1096–1098. Yeung CJ, Susil RC, Atalar E. RF safety of wires in interventional MRI: using a safety index. Magn Reson Med. 2002;47(1):187–193. Konings MK, Bartels LW, van Swol CF, Bakker CJ. Development of an MR-safe tracking catheter with a laser-driven tip coil. J Magn Reson Imaging. 2001;13(1):131–135. Weiss S, Vernickel P, Schaeffter T, Schulz V, Gleich B. Transmission line for improved RF safety of interventional devices. Magn Reson Med. 2005;54(1):182–189. Hegde S, Miquel ME, Boubertakh R, Gilderdale D, Muthurangu V, Keevil SF, et al. Interactive MR imaging and tracking of catheters with multiple tuned fiducial markers. J Vasc Interv Radiol. 2006;17(7): 1175–1179. Kuehne T, Fahrig R, Butts K. Pair of resonant fiducial markers for localization of endovascular catheters at all catheter orientations. J Magn Reson Imaging. 2003;17(5):620–624. Kuehne T, Weiss S, Brinkert F, Weil J, Yilmaz S, Schmitt B, et al. Catheter visualization with resonant markers at MR imaging-guided deployment of endovascular stents in swine. Radiology. 2004;233 (3):774–780. Quick HH, Zenge MO, Kuehl H, Kaiser G, Aker S, Massing S, et al. Interventional magnetic resonance angiography with no strings attached: wireless active catheter visualization. Magn Reson Med. 2005;53(2):446–455. Wong EY, Zhang Q, Duerk JL, Lewin JS, Wendt M. An optical system for wireless detuning of parallel resonant circuits. J Magn Reson Imaging. 2000;12(4):632–638. Weiss S, Kuehne T, Brinkert F, Krombach G, Katoh M, Schaeffter T, et al. In vivo safe catheter visualization and slice tracking using an optically detunable resonant marker. Magn Reson Med. 2004;52(4): 860–868. Kozerke S, Hegde S, Schaeffter T, Lamerichs R, Razavi R, Hill DL. Catheter tracking and visualization using 19F nuclear magnetic resonance. Magn Reson Med. 2004;52(3):693–697. Mansson S, Johansson E, Magnusson P, Chai CM, Hansson G, Petersson JS, et al. 13C imaging: a new diagnostic platform. Eur Radiol. 2006;16(1):57–67. Svensson J, Mansson S, Johansson E, Petersson JS, Olsson LE. Hyperpolarized 13C MR angiography using trueFISP. Magn Reson Med. 2003;50(2):256–262. McRobbie D, Foster MA. Thresholds for biological effects of timevarying magnetic fields. Clin Phys Physiol Meas. 1984;5(2):67–78. McRobbie D, Foster MA. Cardiac response to pulsed magnetic fields with regard to safety in NMR imaging. Phys Med Biol. 1985;30 (7):695–702. Budinger TF. MR safety: past, present, and future from a historical perspective. Magn Reson Imaging Clin N Am. 1998;6(4):701–714. Budinger TF. Emerging nuclear magnetic resonance technologies: health and safety. Ann N Y Acad Sci. 1992;649:1–18.

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103. Hill DL, McLeish K, Keevil SF. Impact of electromagnetic field exposure limits in Europe: is the future of interventional MRI safe? Acad Radiol. 2005;12(9):1135–1142. 104. Keevil SF, Gedroyc W, Gowland P, Hill DL, Leach MO, Ludman CN, et al. Electromagnetic field exposure limitation and the future of MRI. Br J Radiol. 2005;78(935):973–975. 105. Tenforde TS. Biological interactions and potential health effects of extremely-low-frequency magnetic fields from power lines and other common sources. Annu Rev Public Health. 1992;13:173–196. 106. Schenck JF, Dumoulin CL, Redington RW, Kressel HY, Elliott RT, McDougall IL. Human exposure to 4.0-Tesla magnetic fields in a whole-body scanner. Med Phys. 1992;19(4):1089–1098. 107. Schenck JF. MR safety at high magnetic fields. Magn Reson Imaging Clin N Am. 1998;6(4):715–730. 108. Kelley DA, Schenck JF. Very-high-field magnetic resonance imaging: instrumentation and safety issues. Top Magn Reson Imaging. 1999;10(1):79–89. 109. Schenck JF. Safety of strong, static magnetic fields. J Magn Reson Imaging. 2000;12(1):2–19. 110. Schenck JF. Physical interactions of static magnetic fields with living tissues. Prog Biophys Mol Biol. 2005;87(2–3):185–204. 111. Budinger TF, Fischer H, Hentschel D, Reinfelder HE, Schmitt F. Physiological effects of fast oscillating magnetic field gradients. J Comput Assist Tomogr. 1991;15(6):909–914. 112. Guidelines on limits of exposure to static magnetic fields. International Commission on Non-Ionizing Radiation Protection. Health Phys. 1994;66(1):100–106. 113. Kanal E, Borgstede JP, Barkovich AJ, Bell C, Bradley WG, Felmlee JP, et al. American College of Radiology White Paper on MR Safety. Am J Roentgenol. 2002;178(6):1335–1347. 114. Medical magnetic resonance (MR) procedures: protection of patients. Health Phys. 2004;87(2):197–216. 115. Shellock FG, Shellock VJ. Cardiovascular catheters and accessories: ex vivo testing of ferromagnetism, heating, and artifacts associated with MRI. J Magn Reson Imaging. 1998;8(6):1338–1342. 116. Nitz WR, Oppelt A, Renz W, Manke C, Lenhart M, Link J. On the heating of linear conductive structures as guide wires and catheters in interventional MRI. J Magn Reson Imaging. 2001;13(1): 105–114. 117. Armenean C, Perrin E, Armenean M, Beuf O, Pilleul F, SaintJalmes H. RF-induced temperature elevation along metallic wires in clinical magnetic resonance imaging: influence of diameter and length. Magn Reson Med. 2004;52(5):1200–1206. 118. Dempsey MF, Condon B, Hadley DM. Investigation of the factors responsible for burns during MRI. J Magn Reson Imaging. 2001;13(4): 627–631. 119. Gray RW, Bibens WT, Shellock FG. Simple design changes to wires to substantially reduce MRI-induced heating at 1.5 T: implications for implanted leads. Magn Reson Imaging. 2005;23(8):887–891. 120. Helfer JL, Gray RW, MacDonald SG, Bibens TW. Can pacemakers, neurostimulators, leads, or guide wires be MRI safe? Technological concerns and possible resolutions. Minim Invasive Ther Allied Technol. 2006;15(2):114–120. 121. Luechinger R, Duru F, Scheidegger MB, Boesiger P, Candinas R. Force and torque effects of a 1.5-Tesla MRI scanner on cardiac pacemakers and ICDs. Pacing Clin Electrophysiol. 2001;24(2):199–205. 122. Buecker A, Spuentrup E, Grabitz R, Freudenthal F, Schaeffter T, van Vaals JJ, et al. Real-time-MR guidance for placement of a selfmade fully MR-compatible atrial septal occluder: in vitro test. Rofo. 2002;174(3):283–285. 123. Dimick RN, Hedlund LW, Herfkens RJ, Fram EK, Utz J. Optimizing electrocardiograph electrode placement for cardiac-gated magnetic resonance imaging. Invest Radiol. 1987;22(1):17–22. 124. Tenforde TS. Magnetically induced electric fields and currents in the circulatory system. Prog Biophys Mol Biol. 2005;87(2–3):279–288. 125. Rieke V, Ganguly A, Daniel BL, Scott G, Pauly JM, Fahrig R, et al. X-ray compatible radiofrequency coil for magnetic resonance imaging. Magn Reson Med. 2005;53(6):1409–1414. 126. Razavi RS, Hill DL, Muthurangu V, Miquel ME, Taylor AM, Kozerke S, et al. Three-dimensional magnetic resonance imaging of congenital cardiac anomalies. Cardiol Young. 2003;13(5):461–465. 127. Kuehne T, Yilmaz S, Schulze-Neick I, Wellnhofer E, Ewert P, Nagel E, et al. Magnetic resonance imaging guided catheterisation for assessment of pulmonary vascular resistance: in vivo validation and clinical application in patients with pulmonary hypertension. Heart. 2005;91(8):1064–1069.

144. Schalla S, Saeed M, Higgins CB, Weber O, Martin A, Moore P. Balloon sizing and transcatheter closure of acute atrial septal defects guided by magnetic resonance fluoroscopy: assessment and validation in a large animal model. J Magn Reson Imaging. 2005;21(3): 204–211. 145. Kuehne T, Yilmaz S, Meinus C, Moore P, Saeed M, Weber O, et al. Magnetic resonance imaging-guided transcatheter implantation of a prosthetic valve in aortic valve position: feasibility study in swine. J Am Coll Cardiol. 2004;44(11):2247–2249. 146. Spuentrup E, Ruebben A, Schaeffter T, Manning WJ, Gunther RW, Buecker A. Magnetic resonance–guided coronary artery stent placement in a swine model. Circulation. 2002;105(7):874–879. 147. Zhang S, Rafie S, Chen Y, Hillenbrand CM, Wacker FK, Duerk JL, et al. In vivo cardiovascular catheterization under real-time MRI guidance. J Magn Reson Imaging. 2006;24(4):914–917. 148. Serfaty JM, Yang X, Aksit P, Quick HH, Solaiyappan M, Atalar E. Toward MRI-guided coronary catheterization: visualization of guiding catheters, guidewires, and anatomy in real time. J Magn Reson Imaging. 2000;12(4):590–594. 149. Serfaty JM, Yang X, Foo TK, Kumar A, Derbyshire A, Atalar E. MRIguided coronary catheterization and PTCA: a feasibility study on a dog model. Magn Reson Med. 2003;49(2):258–263. 150. Dick AJ, Guttman MA, Raman VK, Peters DC, Pessanha BS, Hill JM, et al. Magnetic resonance fluoroscopy allows targeted delivery of mesenchymal stem cells to infarct borders in swine. Circulation. 2003;108(23):2899–2904. 151. Dick AJ, Lederman RJ. MRI-guided myocardial cell therapy. Int J Cardiovasc Intervent. 2005;7(4):165–170. 152. McVeigh ER, Guttman MA, Lederman RJ, Li M, Kocaturk O, Hunt T, et al. Real-time interactive MRI-guided cardiac surgery: aortic valve replacement using a direct apical approach. Magn Reson Med. 2006;56(5):958–964. 153. Yang X, Atalar E. MRI-guided gene therapy. FEBS Lett. 2006;580(12): 2958–2961. 154. Sermesant M, Coudiere Y, Moreau-Villeger V, Rhode KS, Hill DL, Razavi RS. A fast-marching approach to cardiac electrophysiology simulation for XMR interventional imaging. Med Image Comput Comput Assist Interv Int Conf Med Image Comput Comput Assist Interv. 2005;8(Pt 2):607–615. 155. Sermesant M, Rhode K, Sanchez-Ortiz GI, Camara O, Andriantsimiavona R, Hegde S, et al. Simulation of cardiac pathologies using an electromechanical biventricular model and XMR interventional imaging. Med Image Anal. 2005;9(5):467–480. 156. Celik H, Uluturk A, Tali T, Atalar EA. Catheter tracking method using reverse polarization for MR-guided interventions. Magn Reson Med. 2007;58(6):1224–1231. 157. Magnusson P, Johansson E, Mansson S, et al. Passive catheter tracking during interventional MRI using hyperpolarized 13C. Magn Reson Med. 2007;57(6):1140–1147. 158. Peng H, Draper JN, Frayne R. Rapid passive MR catheter visualization for endovascular therapy using nonsymmetric truncated k-space sampling strategies. Magn Reson Imaging. 2008;26(3):293–303.

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128. Rhode KS, Sermesant M, Brogan D, Hegde S, Hipwell J, Lambiase P, et al. A system for real-time XMR guided cardiovascular intervention. IEEE Trans Med Imaging. 2005;24(11):1428–1440. 129. Wildermuth S, Dumoulin CL, Pfammatter T, Maier SE, Hofmann E, Debatin JF. MR-guided percutaneous angioplasty: assessment of tracking safety, catheter handling and functionality. Cardiovasc Intervent Radiol. 1998;21(5):404–410. 130. Yang X, Atalar E. Intravascular MR imaging-guided balloon angioplasty with an MR imaging guide wire: feasibility study in rabbits. Radiology. 2000;217(2):501–506. 131. Yang X, Bolster BD, Kraitchman DL, Atalar E. Intravascular MR-monitored balloon angioplasty: an in vivo feasibility study. J Vasc Interv Radiol. 1998;9(6):953–959. 132. Godart F, Beregi JP, Nicol L, Occelli B, Vincentelli A, Daanen V, et al. MR-guided balloon angioplasty of stenosed aorta: in vivo evaluation using near-standard instruments and a passive tracking technique. J Magn Reson Imaging. 2000;12(4):639–644. 133. Omary RA, Frayne R, Unal O, Warner T, Korosec FR, Mistretta CA, et al. MR-guided angioplasty of renal artery stenosis in a pig model: a feasibility study. J Vasc Interv Radiol. 2000;11(3):373–381. 134. Krueger JJ, Ewert P, Yilmaz S, Gelernter D, Peters B, Pietzner K, et al. Magnetic resonance imaging-guided balloon angioplasty of coarctation of the aorta: a pilot study. Circulation. 2006;113(8):1093–1100. 135. Kuehne T, Saeed M, Higgins CB, Gleason K, Krombach GA, Weber OM, et al. Endovascular stents in pulmonary valve and artery in swine: feasibility study of MR imaging-guided deployment and postinterventional assessment. Radiology. 2003;226(2):475–481. 136. Raval AN, Telep JD, Guttman MA, Ozturk C, Jones M, Thompson RB, et al. Real-time magnetic resonance imaging-guided stenting of aortic coarctation with commercially available catheter devices in Swine. Circulation. 2005;112(5):699–706. 137. Buecker A, Neuerburg JM, Adam GB, Glowinski A, Schaeffter T, Rasche V, et al. Real-time MR fluoroscopy for MR-guided iliac artery stent placement. J Magn Reson Imaging. 2000;12(4):616–622. 138. Quick HH, Kuehl H, Kaiser G, Bosk S, Debatin JF, Ladd ME. Inductively coupled stent antennas in MRI. Magn Reson Med. 2002;48(5): 781–790. 139. Arepally A, Karmarkar PV, Weiss C, Rodriguez ER, Lederman RJ, Atalar E. Magnetic resonance image-guided trans-septal puncture in a swine heart. J Magn Reson Imaging. 2005;21(4):463–467. 140. Raval AN, Karmarkar PV, Guttman MA, Ozturk C, Desilva R, Aviles RJ, et al. Real-time MRI guided atrial septal puncture and balloon septostomy in swine. Catheter Cardiovasc Interv. 2006;67(4): 637–643. 141. Buecker A, Spuentrup E, Grabitz R, Freudenthal F, Muehler EG, Schaeffter T, et al. Magnetic resonance-guided placement of atrial septal closure device in animal model of patent foramen ovale. Circulation. 2002;106(4):511–515. 142. Rickers C, Jerosch-Herold M, Hu X, Murthy N, Wang X, Kong H, et al. Magnetic resonance image-guided transcatheter closure of atrial septal defects. Circulation. 2003;107(1):132–138. 143. Schalla S, Saeed M, Higgins CB, Martin A, Weber O, Moore P. Magnetic resonance–guided cardiac catheterization in a swine model of atrial septal defect. Circulation. 2003;108(15):1865–1870.

Analogous CMR Terminology Used by Various Vendors Sequence Type

Philips

Siemens

GE

Hitachi

Toshiba

Spin echo Multispin echo Fast spin echo Ultrafast spin echo

SE Multi SE TSE (Turbo SE) SSH TSE

SE Multi echo/MS Turbo SE SSTSE/HASTE

SE SE FSE (Fast SE) SS-FSE

SE SE Fast SE FSE-ADA

Inversion recovery Fast inversion recovery STIR FLAIR Gradient recalled echo Spoiled gradient echo Ultrafast gradient echo

IR IR TSE STIR FLAIR FFE T1 FFE T1-TFE T2-TFE BLISS, THRIVE THRIVE, eTHRIVE

IR/IRM Turbo IR/TIRM STIR FLAIR GRE FLASH TurboFLASH

IR FIR STIR FLAIR GE RSSG SARGE

Ultrafast 3D gradient echo

3D TFE

MPRAGE

Ultrafast gradient echo with magnetization preparation Volume-interpolated gradient echo Steady-state gradient echo Contrast-enhanced steady state gradient echo Balanced gradient echo Fast balanced gradient echo Spin echo—echo planar Gradient echo—echo planar Hybrid echo Multi echo Spin echo black-blood (cardiac)

IR TFE

T1/T2 Turbo FLASH

THRIVE, eTHRIVE

VIBE/LAVA/LAVA XV

IR FSE-IR STIR FLAIR GRE SPGR/MPSPGR FGRE Fast SPGR VIBRANT LAVA, LAVA XV 3D FGRE, 3D fast SPGR IR-prepped/ DE-SPGR FAME

SE Multi echo Fast SE (Super) FASE-DIET IR Fast IR STIR FLAIR FE RF spoiled/FE Fast FE

FFE T2 FFE

FISP PSIF

MPGR/GRE SSFP

TRSG

FE FE

BFFE BTFE SE-EPI FFE-EPI, TFE-EPI GRASE mFFE Black-blood prepulse

TrueFISP

FIESTA

BASG

True SSFP

EPI SE EPIFl TGSE

EPI SE GRE EPI

SE EPI SG-EPI

SE EPI FE-EPI

Spin echo black-blood null fat (cardiac) Single shot black blood

Black-blood prepulse with SPIR or SPAIR BB-SSh

TOF (time-of-flight) MR angiography Time-resolved time-of-flight with contrast Contrast-enhanced MR angiography Noncontrast angiography Contrast-enhanced MR angiography with moving table Real-time interactive scan Susceptibility-weighted imaging

Inflow MRA

Dark-blood prepared TSE, HASTE TRIM

QUICK 3D MPRAGE Fast FE

Merge/Cosmic Double IR FSE with blood suppression Double IR FSE with blood suppression

Single shot 2D TrueFISP TOF

TRACS

TWIST

Bolus Trak

Care Bolus

TRANCE MobiTrak, MobiFlex

NATIVE

Interactive Venous BOLD

CARE Susceptibility weighted

TRICKS

SmartStep iDRIVE SWAN

Cardiovascular Magnetic Resonance 611

I ANALOGOUS CMR TERMINOLOGY USED BY VARIOUS VENDORS

APPENDIX I

CMR SCREENING FORM—BETH ISRAEL DEACONESS MEDICAL CENTER—CMR CENTER

APPENDIX II

CMR Screening Form—Beth Israel Deaconess Medical Center—CMR Center 1/6/2010-V5

Cardiac MR Center East Campus/Gryzmish 4 Telephone – 617-667-8555 Fax – 617-975-5480

Date of Birth ___/____/19___

Date __/___/2010 Name_____________________________ Height_________m_ Weight_______Kg An MRI involves the use of a very strong magnet. For your safety, the presence of certain metallic objects must be determined before you enter the exam room. Please place a check in the appropriate column for each item below. Yes 1.

No

Have you had an MRI/CMR before?

No

17. Do you have a Port-a-cath or Hickman device? If yes, is it accessed?

If yes, did you receive a contrast injection? 2.

Yes

Pacemaker/Pacer wires/Implantable defibrillator

18. Are you on dialysis? If yes, how often:___________________________

3.

Metallic heart valve or any metallic stents

19. Please list all surgeries:

4.

Intracranial or brain aneurysm clip/brain surgery

_________________________________________________________

5.

Bio or neurostimulator, electronic device, or implant

_________________________________________________________

6.

Tattoo(s), Tattooed eyeliner

_________________________________________________________ _________________________________________________________

If yes, location(s):_____________________________ 7.

20. Please circle if you have any of the following medical conditions:

Body piercing

If yes, location(s):____________________________

Asthma/Hay fever

Heart Disease

8.

Metal injury to the eye requiring medical attention

Thyroid Disease

Pheochromocytoma Sickle Cell Disease

9.

Shrapnel/gunshot (metal in body)

FEMALES

10. Eye prosthesis/surgery on eye

Y E S

Multiple Myeloma

MALES

N O

11. Ear prosthesis/surgery on ear

21. Possibility of

24. Do you have a

12. Limb or joint replacement or pinning

pregnancy?

penile implant?

13. Tissue expander (e.g. breast implant)

22. IUD (Intra

14. Implanted pump (insulin, pain med, chemotherapy)

Uterine Device) or

Y E S

N O

If yes, make and model:

Diaphragm

___________________________

15. Are you wearing a patch that delivers medication?

23. Pessary (in

___________________________

16. Do you have a history of difficult IV starts?

pelvis)?

CMR Staff will speak to you about the need for removing the following items: Removable dental work

Eyeglasses

Wallet/keys

Watch/Jewelry

Credit and ATM cards

Hearing aids

Wigs/hairpieces or bobby pins

How do you describe your ethnic background? (Answering this question is optional.) American Indian or Alaskan Native

Hispanic

Asian or Pacific Islander

White (not of Hispanic origin)

Black (not of Hispanic origin)

Prefer not to answer

Patient Signature _________________________________________ Relationship (if not the patient)_________ Date __/__/___ Signature of Nurse or Technologist__________________________________________________________________Date __/__/___

612 Cardiovascular Magnetic Resonance

East Campus/Gryzmish 4 Telephone – 617-667-8555 Fax – 617-975-5480

Date of Birth ___/____/19___

Yes

No

26. Have you ever been told you have kidney problems? 27. Have you ever been told you have protein in your urine? 28. Have you ever had high blood pressure? 29. Do you have diabetes? 30. Have you ever had gout? 31. Have you ever had kidney surgery? 32. Please list below any allergies to medications, food, or latex:

NONE

Reaction

33. Please list below all prescription and over-the-counter medications you take:

Last Dose Date & Time /

/

:

/

/

:

/

/

:

/

/

:

/

/

:

/

/

:

/

/

:

/

/

:

/

/

:

/

/

:

/

/

:

34. Are you taking any of the following medications? sildenafil

tadalafil

vardenafil

Date of Last Dose: ___/__/10

These drugs can interfere with certain aspects of some Cardiac MRI Examinations. If you take any of these drugs for erectile dysfunction you should refrain from taking these medications for 48 hours (2 days) prior to your Cardiac MRI. If you take any of these drugs for pulmonary hypertension you should NOT stop taking your medication but MUST inform the Cardiac MR Center staff before your examination.

Patient Signature_______________________________________ Relationship (if not the patient)__________ Date ___/___/_____ Signature of Nurse or Technologist_____________________________________________________________ Date ___/___/_____

Discharge instructions to patient:

Resume your usual medications

Special instructions:_____________________________________________________________________________________________________ RN/MD/RT Signature______________________________________________________________________________________ Date ___/___/_____

CMR Screening Form in use (2010) at the Beth Israel Deaconess Medical Center (BIDMC), Cardiac MR Center, Boston, MA. (Provided by Kraig V. Kissinger, RT)

Cardiovascular Magnetic Resonance 613

II CMR SCREENING FORM—BETH ISRAEL DEACONESS MEDICAL CENTER—CMR CENTER

Cardiac MR Center

CMR WORKSHEET AND SEQUENCE PROTOCOL DATAFORM IN USE (2010) AT THE BETH ISRAEL DEACONESS MEDICAL CENTER (BIDMC)—CMR CENTER

APPENDIX III

CMR Worksheet and Sequence Protocol Dataform in use (2010) at the Beth Israel Deaconess Medical Center (BIDMC)—CMR Center Beth Israel Deaconess Medical Center - Cardiac MRI 330 Brookline Avenue, Boston, MA 02215 Patient Name:

(617) 667-8555

fax: (617) 975-5480

BIDMC MRN:

Date of Birth:

/

/19

Height:

m

M F

Weight:

kg

Indication/History:

Scan Date:

BSA:

/

/201

m2

Referring MD: Heart Rate:

Analysis of Left Ventricular Function

bpm

Blood pressure: LV Cavity Dimension (diastole):

mm

LV Cavity Dimension (systole):

mm

Anteroseptal wall thickness:

mm

Inferolateral wall thickness:

mm

LV End Diastolic Volume:

ml

LV End Systolic Volume:

ml

LV Stroke Volume: LV ED Mass:

Rhythm SR AF Other:_____

/

mm/m2

mmHg

Comments:

7 2

3

g

g/m2

ml

16 11

5

4

Comments: ml/m2

ml RVEF:

RV Stroke Volume:

15

RCA

Analysis of Right Ventricular Function ml

17

9

6

12

10

%

RV End Systolic Volume:

LCX

13

8 14

ml/m2

ml LVEF:

RV End Diastolic Volume:

1

LAD

RV fatty infiltrate

%

Analysis of Aorta/Pulmonary Artery Flow PA total flow PA stroke vol PR TR

(ml) (ml) (ml) (ml) Qp/Qs

Ao total flow (ml) Ao stroke vol (ml) % AR (ml) % MR (ml) Cardiac output

Heart rate bpm Effective LVEF: % % l/min Cardiac index

%

l/min/m2

Measurements Ascending Aorta

Diameter=

mm

Transverse Aorta

Diameter=

mm

Descending Aorta

Diameter=

mm

Abdominal Aorta

Diameter=

mm

Pulmonary Artery

Diameter=

mm

PLA LA Dimension=

mm

4Chamber LA Length=

mm

4Chamber RA Length=

mm

2 Pulmonary veins dimension (mm) Area (m )

mm/m2

Plaque

2

Plaque

Plaque mm/m

Plaque mm/m2 cm2______ cm2/m2

Aortic valve area= pericardium thickness=

mm

Coronary sinus=

mm

Coronary origins/lengths

Left lower

RCA

Norm.

Abn.

Not Seen

Left upper

LM

Norm.

Abn.

Not Seen

Right lower

RCA Length (mm) _______

Disease: _________________

Right upper

LM Length (mm) _______

Disease: _________________

LGE:

LAD Length (mm) _______

Disease: _________________

LCX Length (mm) _______

Disease: _________________

1-24%:

ⱖ=50%:

25-49%:

mid/epi:

7

Dobutamine viability Stress Findings

2

614 Cardiovascular Magnetic Resonance

Pericardial effusion

Pleural effusion

R/L

9

17 15 10

RCA

LCX

13

8 14

3

Additional Findings

1

LAD

4

12

6

16 11

5

A. LV cines (2Ch, HLA, SA stack, 4CH)

H1. 2D Late Gd-enhancement (short axis stack, 2Ch, 4Ch)

A1. LA cine (SAX extension)

H2. 3D Hi-res late Gd-enhancement (LA views)

A2. LV cine (4Chamber Stack)

H3. 3D Hi-res late Gd-enhancement (LV)

A3. Real time-LV cines (2Ch, 4Ch, SA mid LV) if Afib

H4. Early (5min) 3D Hi-res late Gd-enhancement (LV)

A4. C-SPAMM – 4Ch, 2Ch, mid-ventricular 3 SA

I1. Cor MRI (navigator, 3D targeted TFE w/ Isordil 5 mg)

B. LVOT cines, Aortic valve cine

I2. Low res Cor (nav, 3D targeted axial SSFP w/ Isordil 5 mg)

B2. LVOT stack

I3. Low res Cor (wide nav, 3D whole H, SSFP w/ Isordil 5 mg)

B3. RVOT/main PA sagittal series, pulm valve cine

J1. Thoracic Axial T1w TSE without fat saturation

C. Aortic Q-flow (axial at level of PA bifurcation)

J2. Thoracic Axial T1w TSE with fat saturation

D. Pulmonary artery Q-flow (oblique coronal)

J3. Thoracic axial T2w TSE

E1. Resting myocardial perfusion (SA stack)

J4. Thoracic axial T1w TSE without fat sat post Gd

E2. Stress myocardial perfusion (SA stack)

J5. Thoracic Sagittal T1w TSE without fat saturation (aorta)

F. Low dose dobutamine viability 5mcg/kg/min

K. T2* (hemachromatosis)

G1. Pulmonary vein 3D BH CE-MRA

L. Coronary vein imaging

G2. Aorta 3D BH CE-MRA Comments/Notes: Signature _______________________________________________________________

Allergies:

Medications: Gd-DTPA Stress

Isordil

2.5mg SL

Magnevist

Multihance ____ ml _____mmol/kg

Dobut 0.05mmol/kg/min

5mg SL Atropine ___ mg

Injection Site:

IV Fluid

Cr ____mg/dl on ___/___/___

eGFR____

Dobut max ____ mmol/kg/min

Persantine _____

Catheter Size:

Problems/Reactions: Nurse: _____________________ Technologist: ________________

Protocols: 1. LV/RV function only

A, B, C, D, J1

2. ARVC

A, B, C, D, J1, J2 (H3 if at least mod probability)

3a. Mitral Valve (Prolapse)

A, B, B2, C, D, J1

3b. Pulmonic Valve

A, B, B3, C, D, J1

3c. Tricuspid/Aortic Valve

A, B, C, D, J1

4. Pericardial Disease

A, A4, A3 (4Ch, mid-SA), B, C, D, J1, J3, H1, H3

5a. Cardiomyopathy-dilated

(if 1st CMR - I1) A, B, C, D, J1, H1, H3 (if preCRT - A4,L)

5b. Cardiomyopathy-HCM, amyloid

A, A4, B, C, D, H1, (5min H1 if Amyloid), (H3 if HCM), J1

5c. Cardiomyopathy-sarcoid, myocarditis, Hemo A, B, C, D, H1, J1, J4 (sarcoid-J3, H3; myocarditis J3, J4, H4; iron depo-K) 5d. Cardiomyopathy-VT

A, B, C, D, J1, J3, H1, H3

6. CAD Viability - delayed enhancement

A, B, C, D, (E if CAD), H1, H3, J1, if WMA -->Low dose Dobut x 5 min --> A

7. CAD wall motion stress (dobutamine)

A, B, C, D, J1; Dobutamine (stages - A3, A)

8. CAD perfusion stress (dipyridamole)

A, B, C, D, J1, E2, E1

7a. Pulm vein isolation (pre-ablation)

A, (Afib-A3), B, (Sinus-C, D), G1, H2, J1

7b. Pulm vein isolation (F/u)

A, (AFib-A3), (Sinus-C, D), G1, H2.

8. Coronaries

CAD - I1; Anomalous - I2; , CABG grafts - I3, A, B, C, D, J1

9. Congenital

A, (ASD-A1), A2, B, C, D, J1, special views

10. Cardiac mass/tumor

A, (A1), A2, B, C, D, E1, J1, (J2), J3, J4, (H1) special views

11. Aorta

A, B, C, D, J1, J5, G2

CMR Worksheet and Sequence Protocol dataform in use at the Beth Israel Deaconess Medical Center (BIDMC), Cardiac MR Center, Boston, MA. (Provided by Warren J. Manning, MD) 2D, two-dimensional; 3D, three-dimensional; AR, aortic regurgitation; ARVC, arrhythmogenic right ventricular cardiomyopathy; BH, breath-hold; BSA, body surface area; CAD, coronary artery disease; CE, contrast enhanced; Ch, chamber; ED, end-diastolic; eGFR, estimated glomerular filtration rate; F, female; Gd, gadolinium; HCM, hypertrophic cardiomyopathy; LA, left atrium; LAD, left anterior descending; LCX, left circumflex; LGE, late gadolinium enhancement, LM, left main; LV, left ventricular; LVEF, left ventricular ejection fraction; LVOT, left ventricular outflow tract; M, male; MR, mitral regurgitation; MRA, magnetic resonance angiography; MRN, medical record number; PA, pulmonary artery; PLA, parasternal long axis; PR, pulmonary regurgitation; RA, right atrium; RCA, right coronary artery; RV, right ventricular; RVOT, right ventricular outflow tract; T1w, T1 weighted; T2w, T2 weighted; TR, tricuspid regurgitation; TSE, fast spin echo; VT, ventricular tachycardia.

Cardiovascular Magnetic Resonance 615

III CMR WORKSHEET AND SEQUENCE PROTOCOL DATAFORM IN USE (2010) AT THE BETH ISRAEL DEACONESS MEDICAL CENTER (BIDMC)—CMR CENTER

12/28/09 v10

CMR Sequences

INDEX

Index Note: Page numbers followed by f indicate figure and those followed by t indicate table.

A

Abdominal aorta, MRA of, 468–470 Abdominal ischemia, 472 Ablation, interventional CMR for, 586 Ablavar. See Gadofosveset trisodium (MS-325, Vasovist, Ablavar). Absolute tissue perfusion, in stress myocardial perfusion imaging, 221–222 Accept-reject algorithm, for navigator echoes, 131, 131t, 132f ACE-I (angiotensin-converting enzyme inhibitor), for ventricular remodeling animal studies of, 260, 261f human studies of, 262 Acorn cardiac support device, for ventricular remodeling, 261 Active catheter tracking and visualization, 584–585, 585f, 596–598 advantage of, 596–597 for aortic coarctation repair, 584f for endomyocardial injection, 586f 19 F, 598, 601f for intramyocardial injection, 598f with multiple resonant coils, 597–598, 600f safe transmission line for, 599f semi-, 597–598 for transseptal puncture, 585f Acute myocardial infarction (AMI), 241–252 appropriateness of indications for CMR for, 250t cine CMR for, 241–242, 242f complications of, 248–249, 249f coronary artery CMR for, 248 late gadolinium enhancement for, 242, 243f validation of, 243–244, 244t microvascular obstruction after pathophysiology of, 253 prognostic significance of, 246 and regional recovery of function, 246–247 residual coronary occlusion vs., 245–246, 245f, 246f myocardial viability after, 269–274 contrast-enhanced studies for, 269–274 late gadolinium enhancement for, 269–271 adjustment of T1 in, 271, 273f ischemic bed at risk in, 271, 272f nulling of signal intensity of normal myocardium in, 269, 270f and recovery of function, 273–274 T1-weighted segmented inversion recovery pulse sequence for, 269, 269f time between contrast injection and imaging in, 269–271, 271f no-reflow phenomenon and, 271–272, 274f T2-weighted images for, 269 wall thickness and, 274 stress perfusion imaging after, 247, 247f T2-weighted CMR of, 247–248, 248f ventricular remodeling after, 253–266

Acute myocardial infarction (AMI) (Continued) CMR spectroscopy of energetics during, 256–257, 257f contrast-enhanced CMR and predictors of, 257–259, 258f, 259f early phase of, 253 late phase of, 253, 254f pathophysiology of, 253 regional left ventricular function during, 255–256 therapy for animal studies of, 259–262, 261f human studies of, 262–264, 263f ventricular anatomy during, 253–255, 254f Adenosine contraindications and termination criteria for, 236, 236t, 237t drug interactions with, 236 pharmacologic effects of, 231–232, 231t route and duration of administration of, 231, 238 safety aspects of, 232–234 stress-inducible perfusion abnormalities with, 232 stress-inducible wall motion abnormalities with, 231–232 Adenosine stress CMR, 198t, 208–209 abnormalities induced by, 231–232 duration of, 237 Adenosine stress myocardial perfusion imaging, 29–30, 214, 216f abnormalities induced by, 232 combined dobutamine wall motion CMR with diagnostic performance of, 238–239, 238f protocol for, 232, 233f, 234f in comprehensive CMR assessment of coronary artery disease, 159–160 safety of, 104 Adenosine triphosphate (ATP) transfer, in 31 P-CMRS, 557–558, 560 Advanced CMR technique(s), 37–56 for increased speed, 37, 52f, 53–54 parallel imaging as, 42–53 applications of, 49–53 for assessment of global and regional cardiac function, 47f, 49–51 for coronary artery, 50f, 51–53 for detection of myocardial infarction and assessment of myocardial viability, 50f, 51 for first-pass myocardial perfusion imaging, 50f, 51 for imaging of cardiac anatomy and structure, 49, 50f artifacts in, 48–49, 48f coil arrays for, 46 coil sensitivity calibration strategies in, 46, 47f data acquisition and image reconstruction in, 45–46, 45f dynamic, 47–48 multi-detector-row CT vs., 42–45, 44f

Advanced CMR technique(s) (Continued) principles of, 42–49 signal-to-noise ratio in, 46–47, 47f undersampling in, 46 radial imaging as, 40–42 applications of, 42, 43f principles of, 39f, 40–42, 40f undersampling in, 41, 41f spiral imaging as, 37–40 applications of, 38–40, 39f off-resonance effects in, 38, 39f principles of, 37–38, 38f Adventitia, of artery, 362–363 Aliasing artifacts in parallel imaging, 46 in radial imaging, 41, 41f Allograft rejection, 548–555 CMR of experimental animal models of, 548–549, 549f patient studies of, 549–550 coronary artery CMR in, 550 CMR spectroscopy of animal studies of, 550–551, 551f patient studies of, 551–553, 552f, 553f, 562–563 comparison of diagnostic modalities for, 553, 554t Alpha pulse, in gradient echo imaging, 14 AMI. See Acute myocardial infarction (AMI). AMI-227 (ferumoxtran), 83 Amyloidosis, cardiac, 523–524, 523f, 524f morphology and function in, 523 tissue characterization in, 523–524, 524f Anatomic imaging of coarctation of the aorta, 458–459, 458f of congenital heart disease, 112–113 double inversion recovery dark-blood imaging for, 113, 114f, 115f gadolinium-based CMR for, 116–118 of arrhythmogenic right ventricular cardiomyopathy, 118 of coronary artery, 117, 119f late gadolinium enhancement imaging as, 117, 118f myocardial and blood tagging in, 116, 119f myocardial perfusion imaging as, 116, 118f, 120 saturation band in, 116, 119f static, 116, 117f technical considerations with, 120 time-resolved, 116, 118f of tumor or mass, 117 multiplanar reconstruction for, 113, 114f steady-state free precession for, 112–113, 112f of coronary artery bypass graft, 330–332 conventional spin echo and gradient echo imaging as, 330, 331f, 332f imaging strategy for, 332–333, 337f three-dimensional contrast-enhanced breath hold MRA as, 332, 335f, 336f Cardiovascular Magnetic Resonance 617

INDEX

Anatomic imaging (Continued) three-dimensional respiratory gated MRA as, 331 two-dimensional breath hold MRA as, 330–331, 333f, 334f parallel imaging for, 49, 50f in pediatric CMR, 112–113 double inversion recovery dark-blood imaging for, 113, 114f, 115f gadolinium-based, 116–118 of arrhythmogenic right ventricular cardiomyopathy, 118 of coronary artery, 117, 119f late gadolinium enhancement imaging as, 117, 118f myocardial and blood tagging in, 116, 119f myocardial perfusion imaging as, 116, 118f, 120 saturation band in, 116, 119f static, 116, 117f technical considerations with, 120 time-resolved, 116, 118f of tumor or mass, 117 multiplanar reconstruction for, 113, 114f steady-state free precession for, 112–113, 112f Anesthesia, for pediatric CMR, 395–396 Aneurysm(s) aortic, 374f, 375f causes of, 469 classification of, 469 clinical features of, 469 interventional CMR for, 587, 588f MRA of, 470 mycotic, 469–470 pseudo-, 469 thoracic, 456–457, 456f, 457f coronary artery, 299–301, 300f in Kawasaki disease, 353, 354f pseudoof thoracic aorta, 456 ventricular, due to acute myocardial infarction, 249, 249f ventricular, due to acute myocardial infarction, 249 Angiogenesis, in atherosclerotic plaques of aorta and carotid artery, 345 of coronary artery, 358–359 Angiography CT magnetic resonance vs., 466–467 of pulmonary embolism, 480 of renal artery stenosis, 471 magnetic resonance (See Magnetic resonance angiography [MRA]) phase contrast, 94 radionuclide, for right ventricular assessment, 382 X-ray of mesenteric arteries, 472 of peripheral vascular disease, 473–474 of pulmonary embolism, 480 of renal artery stenosis, 471 Angiosarcoma, 537, 540f, 544t Angiotensin type 1 receptor (AT1R), in ventricular remodeling, 260 Angiotensin type 1 receptor blocker (ARB), for ventricular remodeling, 261f Angiotensin type 2 receptor (AT2R), in ventricular remodeling, 260 Angiotensin-converting enzyme inhibitor (ACE-I), for ventricular remodeling animal studies of, 260, 261f human studies of, 262

618 Cardiovascular Magnetic Resonance

Anomalous coronary artery disease, 299, 300f, 300t Anxiolytic, in CMR stress tests, 198t Aorta abdominal, MRA of, 468–470 ascending, morphology of, 409 atherosclerotic plaque imaging in, 341–350 future directions for, 347 molecular, 344–347 angiogenesis in, 345 endothelial dysfunction in, 344–345 extracellular matrix in, 345–346 plaque inflammation in, 346–347, 347f thrombus in, 345, 346f multicontrast CMR for, 342–344, 342f with automatic segmentation, 342, 344f with complications, 343–344 imaging sequences for, 342, 342t with pharmacologic therapy, 343 in subclinical disease, 342–343 validation of, 342 rationale for, 342–347 coarctation of (See Aortic coarctation) MRA of, 468–470 after repair of transposition of the great vessels, 485, 486f thoracic, 450–462 aneurysm of, 456–457, 456f, 457f aortitis of, 459, 459f coarctation of, 458–459, 458f dissection of, 452–454, 452f, 453f, 454f flow mapping of, 450–451, 451f gradient echo CMR imaging of, 450–451 interventional CMR imaging of, 459–460 intramural hematoma of, 454–455, 455f MRA of, 451–452, 452f, 468–470 penetrating ulcer of, 455–456, 455f spin echo CMR imaging of, 450, 451f trauma to, 457–459, 457f trauma to, 457–459, 457f Aortic aneurysm(s), 374f, 375f causes of, 469 classification of, 469 clinical features of, 469 interventional CMR for, 587, 588f magnetic resonance angiography of, 470 mycotic, 469–470 pseudo-, 469 thoracic, 456–457, 456f, 457f Aortic arch, interrupted, 428–430 vs. aortic arch atresia, 428 epidemiology of, 428 in infant and pediatric patients classification of, 428, 429f evaluation of, 429 postoperative assessment of, 430, 430f preoperative assessment of, 429, 429f with truncus arteriosus, 426, 427f, 428f surgical repair of, 428–429 Aortic arch atresia, 428 Aortic arch hypoplasia, truncus arteriosus with, 426, 427f Aortic arch imaging, with single ventricle, 123–124, 123f, 124f Aortic coarctation, 124–125, 126f, 400–403, 458–459 adult, 458 anatomic imaging for, 458–459, 458f cine CMR of, 113, 458–459 clinical manifestations of, 400 CMR-guided intervention for, 604, 604f, 605 contrast-enhanced MRA of, 400, 401f dark-blood imaging of, 115f defined, 124–125 epidemiology of, 458

Aortic coarctation (Continued) etiology of, 400, 458 fast spin echo image of, 400f flow mapping of, 400–402, 402f, 458–459 gross morphologic features of, 400 infantile, 458 postductal, 458 preductal, 458 repair of, 402, 459 imaging after, 402 passive and active catheter devices during, 584f, 587 truncus arteriosus with, 426, 427f velocity encoded cine CMR of, 417 Aortic compliance, 362 CMR of regional, 364, 364f and coronary artery disease, 368, 368f flow wave velocity and, 368, 368f fluvastatin and, 368–369 menotropin and, 369 Aortic dissection, 452–454 acute phase of, 452–453 characteristics of, 452–453 choice of imaging modality for, 453 chronic, 469–470 classification of, 452–453, 452f, 453f CMR vs. TEE for, 453–454 contrast-enhanced MRA of, 454, 454f etiology of, 470 imaging sequence for, 452–453 interventional CMR for, 587, 588f pathophysiology of, 470 therapeutic management of, 470 Aortic distensibility cilazapril and, 369 in Marfan syndrome, 369 Aortic elastic modulus, 369 Aortic flow, 152, 153f Aortic flow wave velocity, 366, 366f Aortic recesses, normal anatomy of, 489–490, 490f Aortic regurgitation, 403–404, 509 measurement of, 509, 510f severity of, 502t Aortic root, total coronary flow reserve from measurements in, 313–314 Aortic stenosis, 502t, 507 Aortic stiffness. See Aortic wall stiffness. Aortic ulcer, penetrating, 455–456, 455f Aortic valve bicuspid, 403–404, 403f cine CMR of, 113, 116f normal function of, 152–154, 154f papillary fibroelastoma of, 534, 536f, 544t in tetralogy of Fallot, 420 Aortic valve replacement, interventional CMR for, 587, 605, 605f Aortic valve stenosis, 502t, 507 Aortic wall stiffness and future cardiac events, 370, 372f in Marfan syndrome, 369 Aortitis, 459, 459f Aortocoronary bypass graft. See Coronary artery bypass graft (CABG). Aortography, for aortic dissection, 453 Aortopulmonary collateral vessels, in tetralogy of Fallot, 421f Apical rotation, 71f, 73–74, 73f, 73t ARB (angiotensin type 1 receptor blocker), for ventricular remodeling, 260, 261f Array spatial sensitivity encoding technique (ASSET), 45f, 46 Arrhythmogenic right ventricular cardiomyopathy (ARVC), 118, 520–521, 520f diagnostic criteria for, 520, 521, 521t

Atherosclerotic plaque imaging (Continued) challenge(s) in, 351 cardiac motion as, 351–352 respiratory motion as, 352–353, 353f clinical studies of, 359 contrast-enhanced, 354–356, 355f molecular, 356, 356f angiogenesis in, 358–359 inflammation in, 358 thrombosis in, 356–358, 357f, 358f noncontrast, 353–354, 354f, 355f outlook for, 359 future directions for, 347 rationale for, 342–347 Atherothrombosis of aorta and carotid artery, 345, 346f of coronary artery, 356–358, 357f, 358f Athlete’s heart, CMR spectroscopy of, 560–561 ATP (adenosine triphosphate) transfer, in 31 P-CMRS, 557–558, 560 Atrial fibrillation epidemiology of, 445 interventional CMR for, 586 pulmonary veins in pathophysiology of, 445 Atrial fibrillation ablation, pulmonary vein imaging before and after, 445–446, 445f, 446f Atrial isomerism, 409 Atrial mapping, interventional CMR for, 586 Atrial morphology, 408–409 Atrial septal defect (ASD), 116, 398–400 advantages of CMR for, 399 clinical manifestations of, 398 CMR-guided catheterization for, 399 with Ebstein anomaly, 414f location of, 398, 398f, 399f post-closure evaluation of, 400 shunt quantification in, 399, 399f types of, 398, 398f, 399f uses of CMR for, 398–399 Atrial septum, lipomatous hypertrophy of, 142, 146f, 533–534, 535f, 544t Atrial situs inversus, 408 Atrial situs solitus, 408, 409 Atrial switch procedure, 415–416 postoperative assessment of, in infant and pediatric patients, 423–425 contrast-enhanced CMR of, 425f ECG-gated SSFP imaging of, 423, 423f navigator-gated imaging of, 423, 424f Atrial tachyarrhythmia, CMR-guided RF ablation for, 603 Atrial transseptal procedures, interventional CMR for, 585f, 586 Atrial-esophageal fistula, 445–446 Atriopulmonary anastomosis, 432, 432f Atriopulmonary Fontan connection, steadystate free precession images of, 112f Atrioventricular (AV) connection abnormalities of, 409 concordant, 409 discordant, 409 double-inlet, 409 Atrioventricular plane descent (AVPD), normal values for, 385–388, 388t in females, 385–388, 387t in males, 385–388, 387t Atrioventricular (AV) septal defect, 398, 398f Atrioventricular (AV) valve(s), straddling, 409 Atrioventricular (AV) valve atresia, 409 Atropine, in dobutamine CMR stress test, 198, 199f, 202–203 Atropine stress CMR pharmacokinetics of, 196 safety of, 196–197, 198t

Auditory considerations, with CMR, 103 Automatic segmentation, for atherosclerotic plaques, of aorta and carotid arteries, 342, 344f Automatic triggering, for contrast-enhanced MRA, 465 AV. See Atrioventricular (AV). AVMs (arteriovenous malformations), pulmonary, 486 Axial plane scout image in, 140, 141f uses for, 140–142, 143f Axial scout image, 20, 21f

B

B0 field, 3–5, 4f B1 field, 3–5, 4f B22956. See Gadocoletic acid (B22956). Background phase offset, 94–95 BACSPIN (breathing autocorrection with spiral interleaves), 136–137 Balanced steady-state free precession, 15, 15f BAR (brachial artery reactivity) testing, 370–371, 372f, 373f Becker muscular dystrophy, CMR spectroscopy for, 565 Benign cardiac tumor(s), 532–536 fibroma as, 534–536, 537f, 544t hemangioma as, 536, 538f, 544t leiomyomatosis with intracardiac extension as, 536, 539f, 544t lipoma as, 533–534, 535f, 544t myxoma as, 533, 534f, 544t other, 536, 539f papillary fibroelastoma as, 534, 536f, 544t rhabdomyoma as, 536, 538f, 544t Beta-blockers, for ventricular remodeling, 262 Bicuspid valve, 403 aortic, 403–404, 403f cine CMR of, 113, 116f pulmonary, 404, 404f Bing-Taussig anomaly, 412, 425, 426 Biological effects, safety of, 100 Biplanar approach, for cardiomyopathy, 516 Biventricular pacing, X-ray and CMR-guided, 605f Biventricular volume and function, in valvular heart disease, 501 Black blood imaging for cardiac allograft rejection, 550 for congenital heart disease, 396 for coronary artery CMR, 289t with atherosclerotic plaques, 353–354, 355f fast spin echo, 13, 13f in morphology scanning cardiac gating for, 26f goal of, 21–22, 23f physiology of, 22, 26f for right ventricular assessment, 383, 383f of thoracic aorta, 450, 451f, 468–469 uses for, 140–142, 143f Blalock-Taussig shunt, 433f BLAST. See Broad-use linear acquisition speed-up technique (BLAST). Blood flow velocity assessment, 91–99 history of, 91 phase flow imaging methods for, 91–93 Fourier flow imaging as, 91, 93 velocity phase encoding in, 93, 93f visualizing flow data in, 96, 96f phase contrast velocity mapping as, 91, 92–93 flow vector map in, 96, 97f flow velocity images in, 92–93, 93f and Fourier velocity imaging, 94

Cardiovascular Magnetic Resonance 619

INDEX

Arrhythmogenic right ventricular cardiomyopathy (ARVC) (Continued) function and morphology in, 520, 521 right ventricular assessment in, 391 tissue characterization in, 521 Arterial blood, T1 and T2 values for, 7t Arterial compliance, 363 Arterial imaging, invasive, 587 Arterial input function, in quantitative evaluation of myocardial perfusion, 65–66, 65f Arterial spin labeling, for stress myocardial perfusion imaging, 215 Arterial stenosis, interventional CMR for, 589, 589f Arterial stiffness. See Arterial wall stiffness. Arterial structure, 362–363 Arterial switch procedure, 120–121, 121f for transposition of the great arteries, 410–411, 416 Arterial wall biophysical mechanical properties of, 362–378 structure of, 362–363 Arterial wall compliance, 363 Arterial wall shear stress, 371–374, 374f, 375f Arterial wall stiffness clinical use of CMR for assessing, 368–370, 368f, 370f, 371f, 372f defined, 363 measurement of, 363 Arteriovenous malformations (AVMs), pulmonary, 486 Arteritis, Takayasu, 459, 459f Artifacts, 142–149 cardiac motion, 142–145, 146, 147f chemical shift, 142–145, 147f, 148–149 with high field CMR, 170, 171f metal, 142–145, 146–148, 147f in parallel imaging, 48–49, 48f respiratory motion, 142–145, 146, 147f ARVC. See Arrhythmogenic right ventricular cardiomyopathy (ARVC). ASD. See Atrial septal defect (ASD). ASSET (Array spatial sensitivity encoding technique), 45f, 46 AT1R (angiotensin type 1 receptor), in ventricular remodeling, 260 AT2R (angiotensin type 2 receptor), in ventricular remodeling, 260 Atherogenesis, 213, 214f Atherosclerosis defined, 341, 351 development of, 213, 214f epidemiology of, 341 natural history of, 341 pathobiology of, 341–342 risk factors for, 341 subclinical, 342–343 Atherosclerotic plaque, rupture of, 341–342 Atherosclerotic plaque imaging of aorta and carotid artery, 341–350 molecular, 344–347 angiogenesis in, 345 endothelial dysfunction in, 344–345 extracellular matrix in, 345–346 plaque inflammation in, 346–347, 347f thrombus in, 345, 346f multicontrast CMR for, 342–344, 342f, 343f with automatic segmentation, 342, 344f with complications, 343–344 imaging sequences for, 342, 342t with pharmacologic therapy, 343 in subclinical disease, 342–343 validation of, 342 of coronary arteries, 351–361

INDEX

Blood flow velocity assessment (Continued) improving accuracy of, 93–97, 94f, 95f rapid, 95–96 validation of, 95 principles of, 91–93, 92f rapid, 95–96, 96f time-of-flight methods for, 91, 92f visualizing flow data in, 96–97 flow pressure maps in, 96–97, 97f flow vector map in, 96, 97f Fourier velocity imaging in, 96, 96f phase contrast velocity images in, 96, 97f three-dimensional, 97 wall shear stress in, 96 Blood oxygen level dependent (BOLD) imaging in clinical setting, 572–573, 573f emerging techniques for, 574–576, 574f, 575f, 576f, 577f future of, 577 high field, 576, 577f for myocardial perfusion imaging, 62, 215–216 myocardial perfusion reserve in, 573 without pharmacologic stress, 572–573, 573f in preclinical setting, 571–572, 572f rationale for, 570 vs. SPECT, 572, 573f SSFP-based, 574–576, 574f, 575f, 576f Blood pool contrast agents, 80–81, 80f, 81t for three-dimensional magnetic resonance angiography, 464, 465f Blood tagging, for congenital heart disease, 116, 119f BOLD. See Blood oxygen level dependent (BOLD). Bolus chase technique, for peripheral vascular disease, 474, 475f Bolus tagging, for coronary artery velocity measurement, 314 Bolus timing, for contrast-enhanced magnetic resonance angiography, 465–466 Bolus tracking approach in pediatric CMR, 120 in stress myocardial perfusion imaging, 221 BOOST trial, 263 Brachial artery reactivity (BAR) testing, 370–371, 372f, 373f Breath holding alternative to, 129 in coronary artery CMR, 286, 286t, 289–290, 289f with atherosclerotic plaques, 352 with coronary artery bypass grafts, 289–290, 290f three-dimensional contrast-enhanced, 332, 335f, 336f two-dimensional, 330–331, 333f, 334f for coronary artery velocity measurement, 315–316 with native vessel stenosis, 301–302, 302t for coronary sinus flow assessment, 311 limitations of, 129 multiple, 129–130 mean diaphragm displacement in, 130f respiratory trace data for, 130f in pediatric CMR, 120 in phase contrast velocity mapping, 95 for ventricular function assessment, 149 Breathing autocorrection with spiral interleaves (BACSPIN), 136–137 Bright blood imaging of aorta, 468–469 for congenital heart disease, 396 for coronary artery CMR, 289t in morphology scanning, 21–22, 24f of myocardial function, 141f, 142, 143f for right ventricular assessment, 383, 384f

620 Cardiovascular Magnetic Resonance

Bright signals, catheter devices that create, 583–584, 584f Broad-use linear acquisition speed-up technique (BLAST), k-t applications of, 50f, 51 to assess cardiac function, 185 for myocardial perfusion imaging, 61 principles of, 47–48 Bulboventricular foramen, 430 Bypass graft. See Coronary artery bypass graft (CABG).

C

CABG. See Coronary artery bypass graft (CABG). CAD. See Coronary artery disease (CAD). Calcification, in atherosclerosis, 341 Calcium, epicardial, 304–305, 305f Captopril, for ventricular remodeling, 260 Captopril renography, 471 Carbon dioxide (CO2)-filled balloon, in interventional CMR, 595, 596f, 597f Carbon-13 (13C), hyperpolarized, 87 Carbon-13 (13C) CMR spectroscopy, 87 experimental studies with, 559 Cardiac allograft rejection, 548–555 CMR of experimental animal models of, 548–549, 549f patient studies of, 549–550 coronary artery CMR in, 550 CMR spectroscopy of animal studies of, 550–551, 551f patient studies of, 551–553, 552f, 553f, 562–563 comparison of diagnostic modalities for, 553, 554t Cardiac amyloidosis, 523–524, 523f, 524f morphology and function in, 523 tissue characterization in, 523–524, 524f Cardiac anatomy anatomic variants of, 142, 146f artifacts in imaging of, 142–149 cardiac motion, 142–145, 146, 147f chemical shift, 142–145, 147f, 148–149 metal, 142–145, 146–148, 147f respiratory motion, 142–145, 146, 147f breath hold scout image for, 140, 141f imaging planes and, 140, 141f, 143f main structures in, 140, 145f parallel imaging of, 49, 50f preparatory prepulses for, 140 Cardiac catheterization and intervention, CMR-guided for atrial septal defect, 399 for congenital heart disease, 396, 415–417 for patent ductus arteriosus, 400 for ventricular septal defect, 397 Cardiac disease, congenital. See Congenital heart disease (CHD). Cardiac dysfunction, population impact of, 181 Cardiac function aortic flow in, 152, 153f assessment of, 179–195 CMR for, 183–185 accuracy and reproducibility of, 185–186, 185f, 186f advantages of, 183 of diastolic function, 190, 192f ECG gating in, 188 end-diastolic and end-systolic volumes in, 183–184, 183f, 184f FISP vs. FLASH in, 185, 187

Cardiac function (Continued) future of, 193 left ventricular, 183–185, 183f, 184f practical guide to, 186–189, 187f reference ranges for, 189, 189t, 190t, 191f of regional function, 190–192 right ventricular, 185 scanning time for, 185 short axis slices in, 183–184, 184f, 188 Simpson’s rule in, 183–184, 184f of systolic function, 189–190 computed tomography for, 182–183 echocardiography for, 181–182, 182f importance of, 181 nuclear cardiology for, 182 slice-thickness for, 188 in coronary artery disease, 158 left ventricular, 141f, 142, 149, 150f left ventricular mass in, 142, 150 effect of imaging sequence and magnetic field strength on, 150–152, 151t, 152t myocardial, 141f, 142, 143f parallel imaging of, 49–51, 50f pulmonary artery flow in, 152 right ventricular, 150 stroke volume in, 150 effect of imaging sequence and magnetic field strength on, 150–152, 151t, 152t systolic and diastolic, 149–150 valvular, 152–155, 154f Cardiac gating for black-blood imaging, 26f for cine CMR, 24–25, 27f for coronary artery CMR, 286–287, 287f with slice tracking, 287–288 for morphology scanning, 22, 25f for scout scanning, 21, 23f for vascular angiography, 34–35 for velocity-encoded CMR imaging, 32 for viability imaging, 31, 32f Cardiac hemangioma, 536, 538f, 544t Cardiac masses, 532–547 contrast agents for, 532 due to intracardiac thrombus formation, 540–542, 542f, 544t pediatric, 117 technical considerations with, 532, 533t tumors as (See Cardiac tumor[s]) Cardiac morphology, in coronary artery disease, 158 Cardiac motion, 69–70 assessment of CMR methods of, 69–70 non-CMR methods of, 69 in coronary artery and vein CMR, 284–285, 285f with coronary artery atherosclerotic plaque imaging, 351–352 Cardiac motion artifacts, 142–145, 146, 147f Cardiac output, 150, 189–190 Cardiac pacemakers, safety of CMR with, 101, 102, 107–108, 108f Cardiac phase to order reconstruction (CAPTOR), 74 Cardiac rejection, 548–555 CMR of experimental animal models of, 548–549, 549f patient studies of, 549–550 coronary artery CMR in, 550 CMR spectroscopy of animal studies of, 550–551, 551f

Cardiomyopathy(ies) (CMPs) (Continued) tissue characterization in, 523–524, 524f myocardial siderosis as, 524 morphology and function in, 524 tissue characterization in, 524 sarcoidosis as, 523 morphology and function in, 523 tissue characterization in, 523 metabolism in, 516 morphology and function in, 515–516 myocarditis as, 517–518, 517f combined protocols for, 517f, 518 early enhancement in, 517f, 518 edema in, 517f, 518 follow-up for, 518 function and morphology in, 517f, 518 late gadolinium enhancement in, 517f, 518 tissue characterization in, 517f, 518 noncompaction, 522, 522f characteristics of, 522 function and morphology in, 522 tissue characterization in, 522 primary, 515 restrictive, 521–522 vs. constrictive pericarditis, 493 morphology and function in, 522 tissue characterization in, 522 stress-induced (Tako-Tsubo), 525f, 526 function and morphology in, 525f, 526 tissue characterization in, 526 tissue characterization in, 516 transthoracic echocardiography for, 515 Cardiovascular magnetic resonance (CMR) alignment with main magnetic field in, 3 balanced steady-state free precession in, 15, 15f basic principles of, 1–18 contraindications to, 105–108 with coronary stents, 106 general, 105–108 with pacemakers and implantable cardioverter defibrillators, 101, 102, 107–108, 108f with valvular prosthesis, 106 echo planar imaging, spiral, and radial in, 16–17, 16f frequency encoding: position in X in, 9–10, 9f, 10f gradient echo imaging in, 13–14, 14f gradients in, 7–8 image creation in, 7, 7f inversion recovery fast gradient recalled echo: late gadolinium enhancement in, 15 phase encoding: position in Y in, 10–11, 10f pulse sequences and contrast in, 12–13, 12f radiofrequency and magnet strength in, 3–5, 4f raw k-space data and fast Fourier transform in, 11–12, 11f screening form for, 612–613 selective excitation: position in Z in, 8–9, 8f, 9f signal detection in, 3–17 spin echo imaging in, 6–7, 6f, 7t fast (turbo), 6–7, 13 double inversion recovery (black-blood), 13, 13f T1 relaxation, 5 T2 relaxation and spin phase in, 5–6, 5f terminology used by various vendors for, 611 three-dimensional fast gradient echo: magnetic resonance angiography in, 15

Cardiovascular magnetic resonance (CMR) (Continued) worksheet and sequence protocol dataform for, 614–615 Cardiovascular magnetic resonance spectroscopy (CMRS), 556–568 atomic nucleic used in, 556, 557t for cardiac allograft rejection animal studies of, 550–551, 551f vs. other imaging modalities, 554t patient studies of, 551–553, 552f, 553f, 562–563 for cardiomyopathy, 516 dilated, 517 hypertrophic, 519 clinical studies of, 559–565 in athlete’s heart and hypertension, 560–561 in diabetes and obesity, 561 in healthy subjects, 560, 561f with heart failure and cardiac transplantation, 561–563, 562f in ischemic heart disease, 563–565 for myocardial viability assessment, 564–565, 564f for stress testing, 563–564, 563f methodologic considerations in, 559–560, 560f, 561f with specific gene defects with cardiac pathology, 565 in valvular heart disease, 563 of energetics during left ventricular remodeling, 256–257, 257f experimental foundations of, 557–559 other nuclei in, 559 31 P-CMRS in, 557–558, 558f high field, 175 for myocardial viability, 268, 278–280 perspectives on, 565–566 physical principles of, 556–557, 557f at 3-Tesla, 565, 565f Cardiovascular magnetic resonance (CMR) tagging assessment, of left ventricular function, 69–75 apical rotation in, 71f, 73–74, 73f, 73t cardiac motion and, 69–70 complementary spatial modulation of magnetization for, 70–71 echo planar imaging in, 70–71 grid-tagged images in, 70, 71f, 72f limitations of, 74 slice thickness in, 70 slice-following principle in, 70, 70f spiral imaging in, 70–71 steady-state free precession in, 70–71 with strain measurements, 74 3D, 71 in dobutamine stress CMR, 208f, 209 evaluation of motion data from, 71–72, 73f, 73t limitations of, 74 methods for, 70–72 during remodeling, 255–256 results of, 72–74 strain measurement in, 74 Carotid arteries atherosclerotic plaque imaging in, 341–350 future directions for, 347 molecular, 344–347 angiogenesis in, 345 endothelial dysfunction in, 344–345 extracellular matrix in, 345–346 plaque inflammation in, 346–347, 347f thrombus in, 345, 346f multicontrast CMR for, 342–344, 343f

Cardiovascular Magnetic Resonance 621

INDEX

Cardiac rejection (Continued) patient studies of, 551–553, 552f, 553f, 562–563 comparison of diagnostic modalities for, 553, 554t Cardiac rotation, 69–70 assessment of CMR methods of, 69–70 non-CMR methods of, 69 Cardiac structure, parallel imaging of, 49, 50f Cardiac tamponade, 492 Cardiac thrombus, 540–542, 542f, 544t Cardiac transplantation, 548–555 CMR of experimental animal models of, 548–549, 549f patient studies of, 549–550 coronary artery CMR in, 550 CMR spectroscopy of animal studies of, 550–551, 551f patient studies of, 551–553, 552f, 553f, 562–563 comparison of diagnostic modalities for, 553, 554t Cardiac triggering. See Cardiac gating. Cardiac tumor(s), 532–547 benign, 532–536 fibroma as, 534–536, 537f, 544t hemangioma as, 536, 538f, 544t leiomyomatosis with intracardiac extension as, 536, 539f, 544t lipoma as, 533–534, 535f, 544t myxoma as, 533, 534f, 544t other, 536, 539f papillary fibroelastoma as, 534, 536f, 544t rhabdomyoma as, 536, 538f, 544t contrast agents for, 532 malignant, 537–540 lymphoma as, 540, 541f, 544t metastatic, 540 sarcoma as, 537–540 angio-, 537, 540f, 544t leiomyo-, 537, 544t lipo-, 540, 541f, 544t pediatric, 117 prognosis of, 545 step-by-step procedure for assessment of, 545 technical considerations with, 532, 533t tissue characterization for, 543–544, 544t Cardiomyopathy(ies) (CMPs), 515–531 arrhythmogenic right ventricular, 118, 515, 520f diagnostic criteria for, 520, 521, 521t function and morphology in, 520, 521 right ventricular assessment in, 391 tissue characterization in, 521 classification of, 515 clinical presentation of, 515 CMR-derived information in, 515, 516t dilated, 516–517 CMR spectroscopy in, 561–563, 562f function and morphology in, 516–517 metabolic CMR in, 517 endomyocardial diseases as, 524–526, 525f morphology and function in, 526 tissue characterization in, 526 hypertrophic, 518–520, 519f CMR spectroscopy for, 565 follow-up for, 520 function and morphology in, 519 LVOT obstruction in, 518, 519–520 tissue characterization in, 519 infiltrative secondary, 515, 523 cardiac amyloidosis as, 523–524, 523f, 524f morphology and function in, 523

INDEX

Carotid arteries (Continued) with automatic segmentation, 342, 344f with complications, 343–344 imaging sequences for, 342, 342t with pharmacologic therapy, 343 in subclinical disease, 342–343 validation of, 342 rationale for, 342–347 extracranial, magnetic resonance angiography of, 467–468, 467f, 468f Carotid artery stenosis, 467f, 468, 468f Cartesian pattern, 37 Carvedilol, for ventricular remodeling, 262, 263f Catheter tracking and visualization, 583, 595–598 active, 584–585, 585f, 596–598 advantage of, 596–597 for aortic coarctation repair, 584f for endomyocardial injection, 586f 19 F, 598, 601f for intramyocardial injection, 598f with multiple resonant coils, 597–598, 600f safe transmission line for, 599f semi-, 597–598 for transseptal puncture, 585f application(s) of for aortic coarctation repair, 584f for atrial septal defect, 399 for congenital heart disease, 396, 415–417 for endomyocardial injection, 586f for intramyocardial injection, 598f for patent ductus arteriosus, 400 for transseptal puncture, 585f for ventricular septal defect, 397 vs. conventional XRF devices, 583, 583f passive, 583–584, 595–596 chemical-selective visualization of, 584, 584f CO2-filled balloon in, 595, 596f, 597f dysprosium oxide, 595, 596f ideal material for, 595 that create bright signals, 583–584, 584f that create dark signals, 583 strategy for, 585 Cavity dilation, after acute myocardial infarction, 253 CE-CTA (contrast-enhanced computed tomography angiography) CE-MRA vs., 466–467 of pulmonary embolism, 480 Cellular agents, endomyocardial delivery of, 585–587, 586f CE-MRA. See Contrast-enhanced magnetic resonance angiography (CE-MRA). Central volume principle, in quantitative evaluation of myocardial perfusion, 62, 63 Cerebral aneurysm clips, CMR with, 105–106 CFD (computational fluid dynamic) simulation, for arterial wall shear stress, 371–372 CFR. See Coronary flow reserve (CFR). CHD. See Congenital heart disease (CHD). Chemical shift, in CMR spectroscopy, 557, 560 Chemical shift artifacts, 142–145, 147f, 148–149 Chemical-selective visualization, of passive catheter device, 584, 584f Chest pain cine CMR for, 241 in CMR stress tests, 198t Chest x-ray of constrictive pericarditis, 493 of pericardial disease, 488 Children, CMR in. See Pediatric CMR. Chordae tendineae, 381

622 Cardiovascular Magnetic Resonance

CHRISTMAS trial, 262, 263f Chronic myocardial infarction (CMI), 242, 244t myocardial viability in, 274–280 contractile reserve and, 275–278 late gadolinium enhancement for, 276, 277f, 278f, 279f vs. dobutamine CMR, 277–278 vs. other imaging modalities, 276–277, 281f thickness of epicardial rim and recovery of function in, 276, 280f wall thickness and, 275, 275f Cilazapril and aortic distensibility, 369 for ventricular remodeling, 260 Cine CMR, 22–25 acquisition time in, 24 for acute myocardial infarction, 241–242, 242f of aortic coarctation, 113, 458–459 cardiac gating for, 24–25, 27f of congenital heart disease, 113, 115f, 408 for evaluation of function, 417 after repair, 416–417 due to transposition of the great arteries, 120–122, 121f valvular, 113, 116f of coronary artery disease for disease detection, 160, 161, 161f for viability studies, 165f, 166f goal of, 23–24, 27f of myocardial function, 141f, 142 for right ventricular assessment, 383, 384f with single ventricle, 124, 124f for stress tests, 198, 199t of thoracic aorta, 450–451, 451f of valvular heart disease, 501, 502f in infants and children, 113, 116f visualization and planimetry of jets by, 504–505, 505f, 510f of ventricular remodeling, 255, 256 CK (creatine kinase) flux, in 31P-CMRS, 557–558, 560 CK/PCr (creatine kinase/phosphocreatine) energy shuttle, in 31P-CMRS, 557–558, 558f Claustrophobia, during CMR, 103 Clinical technique(s), 19–36 cine CMR as, 22–25 acquisition time in, 24 cardiac gating for, 24–25, 27f goal of, 23–24, 27f combination of, 35, 35f morphology scanning as, 21–22 black-blood imaging in cardiac gating for, 26f goal of, 21–22, 23f physiology of, 22, 26f bright-blood imaging in, 21–22, 24f cardiac gating for, 22, 25f goal of, 21–22, 23f, 24f HASTE in, 21–22, 23f pulse sequence in, 21–22, 25f steady-state free precession imaging in, 21–22, 24f myocardial perfusion scanning as, 25–31 goal of, 25–27, 28f image acquisition in, 27–28, 29f magnetization recovery in, 28, 29f pulse sequence in, 28–29, 30f rest, 31 stress adenosine for, 29–30

Clinical technique(s) (Continued) specific steps in, 31 timeline for, 30–31, 30f scout scanning as, 19–21 cardiac gating for, 21, 23f goal of, 19–20 image acquisition in, 20, 21f k-space filling in, 20–21, 22f pulse sequence in, 20, 22f vascular angiography as, 34–35 cardiac gating for, 34–35 goal of, 34, 34f pulse sequence in, 35 timing of image acquisition in, 34, 35f velocity-encoded imaging as, 32–34 cardiac gating for, 32 goal of, 32, 33f pulse sequence in, 33–34, 33f viability imaging as, 31 cardiac gating for, 31, 32f goal of, 31, 31f inversion recovery in, 31, 32f CMI. See Chronic myocardial infarction (CMI). CMRS. See Cardiovascular magnetic resonance spectroscopy. CNR (contrast-to-noise ratio), in stress myocardial perfusion imaging, 218, 222 CO2 (carbon dioxide)-filled balloon, in interventional CMR, 595, 596f, 597f Coarctation of the aorta, 124–125, 126f, 400–403, 458–459 adult, 458 anatomic imaging for, 458–459, 458f cine CMR of, 113, 458–459 velocity encoded, 417 clinical manifestations of, 400 contrast-enhanced MRA of, 400, 401f dark-blood imaging of, 115f defined, 124–125 epidemiology of, 458 etiology of, 400, 458 fast spin echo image of, 400f flow mapping of, 400–402, 402f, 458–459 gross morphologic features of, 400 infantile, 458 postductal, 458 preductal, 458 repair of, 402, 459 CMR-guided, 604, 604f, 605 imaging after, 402 passive and active catheter devices during, 584f, 587 truncus arteriosus with, 426, 427f Cochlear implants, as contraindication to CMR, 105–106 Coil arrays, in parallel imaging, 46 Coil sensitivity calibration strategies, in parallel imaging, 46, 47f Collagen, type I, contrast agent specific to, 85 Collagen fibers, in arterial wall, 362–363 Collateral vessels, in tetralogy of Fallot, 421f Column orientations, multiple, for navigator echoes, 133–136, 135f Column positioning, for navigator echoes, 133, 135f, 135t Column selection, for navigator echoes, 131–132 Comb-excited Fourier velocity-encoded measurement, of aortic flow wave velocity, 366–367, 367f Combidex (ferumoxtran), 83 Communication, in interventional CMR laboratory, 580–581 Compartmental model, in quantitative evaluation of myocardial perfusion, 63–65, 64f

Congenital heart disease (CHD) (Continued) late gadolinium enhancement imaging as, 117, 118f myocardial and blood tagging in, 116, 119f myocardial perfusion imaging as, 116, 118f, 120 saturation band in, 116, 119f static, 116, 117f technical considerations with, 120 time-resolved, 116, 118f of tumor or mass, 117 multiplanar reconstruction for, 113, 114f steady-state free precession for, 112–113, 112f atrial morphology and determination of situs in, 408–409 atrial septal defect as, 398–400 advantages of CMR for, 399 clinical manifestations of, 398 CMR-guided catheterization for, 399 location of, 398, 398f post-closure evaluation of, 400 shunt quantification in, 399, 399f types of, 398, 399f uses of CMR for, 398–399 atrial switch for, 423–425 contrast-enhanced CMR of, 425f ECG-gated SSFP for, 423, 423f navigator-gated imaging of, 423, 424f cardiac catheterization and intervention for, 395, 415–417 cine CMR of, 113, 115f, 116f gradient echo, 408 CMR vs. transesophageal echocardiography for, 408, 415 coarctation of the aorta as, 124–125, 126f, 400–403 cine CMR of, 113 velocity encoded, 417 clinical manifestations of, 400 contrast-enhanced MRA of, 400, 401f dark-blood imaging of, 115f defined, 124–125 etiology of, 400 fast spin echo image of, 400f flow mapping of, 400–402, 402f gross morphologic features of, 400 post-repair imaging of, 402 repair of, 402 complex, 408–419 in infant and pediatric patients, 420–438 ventricular, 413–415 contrast-enhanced 3D MRA for, 408 after correction, 395 double-outlet right ventricle as, 410f, 412 in infant and pediatric patients, 425–426 postoperative assessment of, 426 preoperative assessment of, 426, 426f Ebstein anomaly of tricuspid valve as, 413, 414f, 415f ECG-gated spin echo CMR of, 408 epidemiology of, 395, 396t evaluation of function in, 417 functional fetal CMR for, 125–127, 126f future of, 125–127, 126f general protocol for, 112–118, 112f imaging techniques in, 396–404 in infant and pediatric patients, 420–438 double-outlet right ventricle as, 425–426 for postoperative assessment, 426 for preoperative assessment, 426, 426f interrupted aortic arch as, 428–430 classification of, 428, 429f for evaluation, 429 for postoperative assessment, 430, 430f

Congenital heart disease (CHD) (Continued) for preoperative assessment, 429, 429f postoperative atrial switch, 423–425 contrast-enhanced, 425f ECG-gated SSFP, 423, 423f navigator-gated, 423, 424f single ventricle as, 430–435 for evaluation, 432 Fontan procedure for, 432, 432f left, 430, 431f post-Fontan, 434–435, 434f right, 430, 431f during staged palliation, 433, 433f, 434f tetralogy of Fallot as, 420–422 for evaluation, 421 for postoperative assessment, 421–422, 422f for preoperative assessment, 421, 421f transposition of the great arteries as, 422–423 truncus arteriosus as, 426–428 classification of, 426, 427f contrast-enhanced, 428f for postoperative assessment, 427–428 for preoperative assessment, 427, 428f interrupted aortic arch as, in infant and pediatric patients, 428–430 classification of, 428, 429f evaluation of, 429 postoperative assessment of, 430, 430f preoperative assessment of, 429, 429f late gadolinium enhancement for, 408 limitations and challenges of, 111 patent ductus arteriosus as, 400 postoperative evaluation of, 415–417 right ventricular assessment in, 391 for shunt evaluation, 395 simple, 395–407 single ventricle as, 122–124, 413–415 aortic arch imaging for, 123–124, 123f, 124f cine CMR for, 124, 124f with Fontan baffle, 123, 123f, 125f in infant and pediatric patients, 430–435 evaluation of, 432 Fontan procedure for, 432, 432f left, 430, 431f post-Fontan, 434–435, 434f right, 430, 431f during staged palliation, 433, 433f, 434f late gadolinium enhancement for, 124f perfusion imaging for, 124f pulmonary artery imaging for, 123, 123f, 125f time-resolved gadolinium imaging for, 118f velocity mapping for, 123–124 venous pathway imaging for, 123f, 124 with ventricular outflow obstruction, 125f surgical procedures for, 411t technical consideration(s) in, 118–120 with gadolinium-based techniques, 120 inability to hold breath as, 120 spatial and temporal resolution as, 118–120 tetralogy of Fallot as, 413, 420–422 in infant and pediatric patients, 420–422 evaluation of, 421 postoperative assessment of, 421–422, 422f preoperative assessment of, 421, 421f surgical repair of, 416 transposition of the great arteries as, 120–122, 409–411, 410f, 411f cine SSFP of, 120–122, 121f dark-blood CMR of, 120–122, 121f

Cardiovascular Magnetic Resonance 623

INDEX

Compartmentalization, and contrast-enhanced tissue relaxation, 84–85 Complementary spatial modulation of magnetization (CSPAMM), 70–71 echo planar imaging in, 70–71 grid-tagged images in, 70, 71f, 72f limitations of, 74 slice thickness in, 70 slice-following principle in, 70, 70f spiral imaging in, 70–71 steady-state free precession in, 70–71 with strain measurements, 74 3D, 71 Comprehensive CMR assessment, of coronary artery disease, 159–166 analysis of studies with, 164–166 for detection of disease, 164, 167f for viability studies, 164–166, 167f contrast agent delivery for, 160 defined, 159 detection of disease in analysis of studies with, 164, 167f protocols for, 160–162, 161f, 162f, 163f future directions for, 167–168 historical background of, 158 protocols for, 160–164, 160t, 161f, 162f sensitivity and specificity of, 161, 163f suggested, 161, 165f selection of methods for, 159–160 viability studies in analysis of studies with, 164–166, 167f protocols for, 163–164, 165f, 166f, 167f Computational fluid dynamic (CFD) simulation, for arterial wall shear stress, 371–372 Computed tomography (CT) of aortic dissection, 453 of aortic intramural hematoma, 454–455 to assess cardiac function, 182–183 of constrictive pericarditis, 494, 494f multidetector to assess cardiac function, 182–183 vs. coronary artery CMR, 304–305, 305f, 305t vs. parallel MRI, 42–45, 44f of pericardial disease, 488 multislice, for right ventricular assessment, 382 of pericardial disease, 488 single photon emission (See Single photon emission computed tomography [SPECT]) Computed tomography angiography (CTA) MRA vs., 466–467 of pulmonary embolism, 480 of renal artery stenosis, 471 Computer architecture, for navigator echoes, 137 Concordant atrioventricular connection, 409 Congenital anomalies of coronary arteries, 299, 300f, 300t of pulmonary arteries, 485–486, 485f, 486f Congenital heart disease (CHD), 111–128 abnormalities of atrioventricular connection as, 409 abnormalities of ventriculoarterial connections as, 409–413 advantages of CMR for, 111 anatomic imaging of, 112–113 double inversion recovery dark-blood imaging for, 113, 114f, 115f gadolinium-based CMR for, 116–118 of arrhythmogenic right ventricular cardiomyopathy, 118 of coronary artery, 117, 119f

INDEX

Congenital heart disease (CHD) (Continued) in infant and pediatric patients, 422–423 three-dimensional contrast-enhanced MRA of, 117f, 120–122, 122f truncus arteriosus as, 412–413, 412f in infant and pediatric patients, 426–428 classification of, 426, 427f contrast-enhanced CMR of, 428f postoperative assessment of, 427–428 preoperative assessment of, 427, 428f valvular, 403–404 with atresia, 402f, 403 bicuspid valve as, 403 aortic, 403–404, 403f pulmonary, 404, 404f gradient recalled echo CMR of, 403f, 404f pressure and volume overload in, 404 quantification of regurgitation volume in, 403f, 405f after repair, 403, 405f transverse spin echo CMR of, 402f velocity mapping for, 113–115 ventricular morphology and isomerism in, 409 ventricular septal defect as, 396–397 anatomic delineation of, 397, 397f clinical manifestations of, 396–397 CMR-guided catheterization and intervention for, 397 location of, 396, 396f shunt quantification in, 397, 398f spontaneous closure of, 396–397 surgical management of, 396–397, 397f Congenital pericardial defects, 490 Conotruncal anomalies, 420 Conoventricular septal defect, 420 Constrictive pericarditis, 493–495 chest x-ray of, 493 clinical presentation of, 493 CMR of, 494, 494f CT of, 494, 494f effusive-, 493 etiology of, 493 pericardial thickening in, 493 vs. restrictive cardiomyopathy, 493 transthoracic echocardiography of, 493 ventricular filling pattern in, 495 Contractile function, cine CMR of, 22–25 acquisition time in, 24 cardiac gating for, 24–25, 27f goal of, 23–24, 27f Contractile reserve in chronic myocardial infarction, 275–278 in dobutamine stress CMR studies, 204f, 206f of viable myocardium, 267–268 Contraindications, 105–108 with coronary stents, 106 general, 105–108 with pacemakers and implantable cardioverter defibrillators, 101, 102, 107–108, 108f with valvular prosthesis, 106 Contrast, 12–13, 12f endogenous, for assessment of myocardial perfusion, 61–62 Contrast agents, 76–90 biophysics of, 76–81, 77f, 78f blood pool, 80–81, 80f, 81t for cardiac and paracardiac masses, 532 for comprehensive CMR assessment of coronary artery disease, 160 contrast-enhanced tissue relaxation with, 84–85

624 Cardiovascular Magnetic Resonance

Contrast agents (Continued) in development, 85–87, 86f effects on signal intensity of, 77, 78f extracellular, 78–79, 79f, 79t gadolinium, 78–79, 79f, 79t history of, 76, 77–78 hyperpolarized, 87 iron oxide–based cross-linked, 86–87 in development, 86–87 relaxivity with, 83 structure of, 83 uses of, 81 for myocardial oxygenation assessment, 570 newer, 85 positive and negative, 76 relaxation rate with, 76–77 relaxivity of, 81–83 correlation time in, 82 effect and definition of, 76, 77f effect of correlation time and field strength on, 82, 83, 83f electronic relaxation in, 82 inner- and outer-sphere, 81–82 iron oxide–based, 83 longitudinal and transverse, 82, 83f magnetic field dependence on, 82, 82f magnetic moment in, 82 molecule size and, 82 for selected media, 83, 83t safety of, 87–88, 103–104 with spoiled gradient recalled echo, 77, 78f for stress myocardial perfusion imaging endogenous, 215–216 exogenous, 216–218 extravascular, 216–217, 217f hyperpolarized, 217–218 intravascular, 217 T1 and T2, 76–81, 77f, 78f Contrast echocardiography, to assess cardiac function, 181–182 Contrast enhancement, water exchange and its effects on myocardial, 61 Contrast-enhanced CMR of acute myocardial infarction, 269–274 of coarctation of the aorta, 400, 401f of coronary artery, 292–294, 293f with atherosclerosis, 354–356, 355f with native vessel stenosis, 304, 304f of postoperative atrial switch, 425f and predictors of left ventricular remodeling, 257–259, 258f, 259f Contrast-enhanced computed tomography angiography (CE-CTA) CE-MRA vs., 466–467 of pulmonary embolism, 480 Contrast-enhanced magnetic resonance angiography (CE-MRA), 34–35 advantages of, 463 of aorta, 468–470 of aortic aneurysms, 470 of aortic dissection, 452–453, 454, 454f, 469–470 bolus timing for, 465–466 cardiac gating for, 34–35 of congenital heart disease, 116, 117f, 396, 408 due to transposition of the great arteries, 117f, 120–122, 122f contrast agents for, 80–81, 80f, 81t of coronary artery bypass graft, 332, 335f, 336f vs. CT angiography, 466–467 of extracranial carotid arteries, 467–468, 468f goal of, 34, 34f of mesenteric arteries, 473

Contrast-enhanced magnetic resonance angiography (CE-MRA) (Continued) parallel imaging in, 466 of peripheral vascular disease, 474, 475f for assessment of bypass graft patency, 474 bolus chase technique in, 474, 475f of hand and wrist, 475 sensitivities and specificities with, 474 time-resolved technique in, 474 for pulmonary artery hypertension, 482–483, 482f, 483f of pulmonary embolism, 481, 481f, 482f pulse sequence in, 35 of renal artery stenosis, 472 technique of, 463 of thoracic aortic aneurysm, 456–457, 457f three-dimensional, 464–465 gadolinium chelates for, 464, 465f pulse sequences for, 464 time-resolved, 466 vessel brightness (T1) vs. gadolinium dose in, 464–465 timing of image acquisition in, 34, 35f Contrast-enhanced tissue relaxation, 84–85 Conus arteriosus, 381 Coronal plane scout image in, 140, 141f uses for, 140–142, 143f Coronal scout image, 20, 21f, 140, 141f Coronary arterial pressure, in myocardial oxygenation assessment, 569, 570f Coronary artery(ies) atherosclerotic plaque imaging of, 351–361 challenge(s) in, 351 cardiac motion as, 351–352 respiratory motion as, 352–353, 353f clinical studies of, 359 contrast-enhanced, 354–356, 355f molecular, 356, 356f angiogenesis in, 358–359 inflammation in, 358 thrombosis in, 356–358, 357f, 358f noncontrast, 353–354, 354f, 355f outlook for, 359 congenital anomalies of, 299, 300f, 300t physiology and pathophysiology of, 229 in tetralogy of Fallot, 420 Coronary artery aneurysms, 299–301, 300f Coronary artery blood flow. See Coronary artery flow. Coronary artery bypass graft (CABG), 329–340 coronary artery CMR for, 305–306, 329–330, 330t anatomic imaging techniques of, 330–332 conventional spin echo and gradient echo imaging as, 330, 331f, 332f imaging strategy for, 332–333, 337f three-dimensional contrast-enhanced breath hold MRA as, 332, 335f, 336f three-dimensional respiratory gated MRA as, 331 two-dimensional breath hold MRA as, 330–331, 333f, 334f breath holding for, 289–290, 290f diagnostic accuracy of, 306, 307f, 307t fast spin echo for, 305–306, 306f indications for, 338 limitations of, 306, 307f, 337–338 for quantification of flow and flow reserve, 333–337, 337f, 338f sensitivity and specificity of, 305–306, 306t demographics of, 329 occlusion of, 329 other imaging modalities for, 329

Coronary artery CMR (Continued) suppression of signal from surrounding tissue as, 288, 288f three-dimensional acquisition sequence for, 290–292 of native vessel stenosis with navigator gating, 302–303, 303t targeted, 301–302, 301f whole heart, 302f, 303, 303t targeted acquisition sequence for, 291, 291f of native vessel stenosis, 301–302, 301f thin-slab, 291–292 whole heart acquisition sequence for, 291, 291f, 292f of native vessel stenosis, 302f, 303, 303t Coronary artery disease (CAD), 158–169 aortic compliance and, 368, 368f CMR in, 158–159, 159t for coronary artery imaging, 159, 159t of ischemia, 158–159, 159t of morphology and function, 158, 159t of viability, 159, 159t comprehensive CMR assessment of, 159–166 analysis of studies with, 164–166 for detection of disease, 164, 167f for viability studies, 164–166, 167f contrast agent delivery for, 160 defined, 159 for detection of disease analysis of studies with, 164, 167f protocols for, 160–162, 161f, 162f, 163f future directions for, 167–168 historical background of, 158 protocols for, 160–164, 160t, 161f, 162f sensitivity and specificity of, 161, 163f suggested, 161, 165f selection of methods for, 159–160 for viability studies analysis of studies with, 164–166, 167f protocols for, 163–164, 165f, 166f, 167f development of, 213, 214f dobutamine stress CMR of, 198, 201f contemporary studies of, 202–203, 202t, 203f, 203t early studies of, 201–202, 202f 3-T, 203 epidemiology of, 284 Coronary artery flow, 310–328 with coronary artery bypass graft, 333–337, 337f, 338f direct assessment of, 320 future developments in, 324–327 patient studies of, 324, 325f, 326f validation and feasibility studies of, 320–324, 322f, 323f effect of nitrates on, 285, 286f indirect assessment of, 310–324, 312f, 313f in myocardial oxygenation assessment, 569, 570f Coronary artery flow velocity mapping. See Coronary artery velocity mapping. Coronary artery flow velocity reserve, 322–323, 323f Coronary artery MRA, for coronary artery disease, 159 Coronary artery stenosis coronary CMR of, 301–304 contrast-enhanced, 304, 304f three-dimensional with navigator gating, 302–303, 303t targeted, 301–302, 301f

Coronary artery stenosis (Continued) whole heart, 302f, 303, 303t two-dimensional, 301–302, 302f, 302t and plaque rupture, 213, 214f after stent implantation, 324, 326f Coronary artery velocity mapping, 314 bolus tagging for, 314 with coronary artery bypass graft, 333–337, 337f, 338f echo planar time-of-flight technique for, 314–315, 315f future developments in, 324–327 gradient echo phase, 315–317 breath holding techniques for, 315–316 in-plane, 316–317, 317f navigator techniques for, 316–317, 317f, 318f through-plane, 317, 317f, 318f interleaved spiral phase, 317–320, 319f, 320f, 321f patient studies of, 324, 325f, 326f validation and feasibility studies of, 320–324, 322f, 323f Coronary artery wall, high field CMR of, 174, 174f Coronary autoregulation, 229 Coronary blood flow. See Coronary artery flow. Coronary flow. See Coronary artery flow. Coronary flow reserve (CFR), 213–214, 224, 230 with coronary artery bypass graft, 333–337, 337f, 338f defined, 310 direct assessment of, 320 future developments in, 324–327 patient studies of, 324, 325f, 326f validation and feasibility studies of, 320–324, 322f, 323f indirect assessment of from coronary sinus flow, 310–313, 313f from measurements in aortic root, 313–314 Coronary flow velocity measurement. See Coronary artery velocity mapping. Coronary sinus defect, 398f Coronary sinus flow, 310–313, 312f, 313f Coronary stenosis. See Coronary artery stenosis. Coronary stents restenosis after implantation of, 324, 326f safety of CMR with, 106 Coronary vein CMR, 295, 295f Coronary venous outflow, velocity mapping of, 310–313, 312f, 313f Correction factors, for navigator echoes, 132–133, 133f, 134f Creatine kinase (CK) flux, in 31P-CMRS, 557–558, 560 Creatine kinase/phosphocreatine (CK/PCr) energy shuttle, in 31P-CMRS, 557–558, 558f Creatine phosphate to adenosine triphosphate (CP-to-ATP) ratio, during left ventricular remodeling, 256, 257, 257f Crusher gradient, 14, 14f CSPAMM. See Complementary spatial modulation of magnetization (CSPAMM). CT. See Computed tomography (CT). CTA (computed tomography angiography) MRA vs., 466–467 of pulmonary embolism, 480 of renal artery stenosis, 471 Cyst(s) hydatid, 536, 539f pericardial, 490–491, 491f, 543, 543f, 544t

Cardiovascular Magnetic Resonance 625

INDEX

Coronary artery CMR, 284–298 acquisition sequence for, 289–292, 289f, 289t breath hold two-dimensional segmented k-space gradient echo, 289–290, 289f with coronary artery bypass grafts, 289–290, 290f free breathing spin echo, 289 three-dimensional, 290–292 targeted, 291, 291f thin-slab, 291–292 whole heart, 291, 291f, 292f for acute myocardial infarction, 248 advanced methods for, 292–296 contrast-enhanced, 292–294, 293f with intracoronary stents, 294–295, 294f spiral and radial, 292, 292f, 293f 3-Tesla, 294, 294f of aneurysms and Kawasaki disease, 299–301, 300f for cardiac allograft rejection, 550 clinical results of, 299–309 of congenital heart disease, 299, 300f, 300t navigator-gated, ECG-gated, 117, 119f perfusion imaging in, 116 of coronary artery bypass grafts, 305–306, 329–330, 330t anatomic imaging techniques of, 330–332 conventional spin echo and gradient echo imaging as, 330, 331f, 332f imaging strategy for, 332–333, 337f three-dimensional contrast-enhanced breath hold MRA as, 332, 335f, 336f three-dimensional respiratory gated MRA as, 331 two-dimensional breath hold MRA as, 330–331, 333f, 334f breath holding for, 289–290, 290f diagnostic accuracy of, 306, 307f, 307t fast spin echo for, 305–306, 306f indications for, 338 limitations of, 306, 307f, 337–338 for quantification of flow and flow reserve, 333–337, 337f, 338f sensitivity and specificity of, 305–306, 306t for coronary artery disease, 159 future technical developments in, 295–296 high field, 172–173, 173f history of, 284 imaging planes for, 142 invasive and interventional, 586–587 vs. multidetector CT, 304–305, 305f, 305t of native vessel stenosis, 301–304 contrast-enhanced, 304, 304f three-dimensional with navigator gating, 302–303, 303t targeted, 301–302, 301f whole heart, 302f, 303, 303t two-dimensional, 301–302, 302f, 302t parallel imaging for, 50f, 51–53 rationale for, 284 technical challenge(s) of, 284–288 cardiac motion as, 284–285, 285f effect of nitrates on coronary artery blood flow as, 285, 286f respiratory motion as, 285–286, 286t breath hold methods for, 286, 286t free breathing methods for, 286, 286t navigators for gating and slice tracking for, 287–288 navigators for triggering alone for, 286–287, 287f spatial resolution as, 288, 288f

INDEX

D

DANTE schemes, for cardiac allograft rejection, 548–549 Dark signals, catheter devices that create, 583 Dark-blood imaging, of congenital heart disease, 113, 114f, 115f due to transposition of the great arteries, 120–122, 121f DCM. See Dilated cardiomyopathy (DCM). DCMR. See Dobutamine stress CMR (DCMR). DCMRC (Duke Cardiovascular Magnetic Resonance Center) clinical volume at, 19, 20f overall makeup of, 19, 20f exam menu for, 19, 20f DeBakey classification, of aortic dissection, 452–453, 452f Deconvolution analysis, in quantitative evaluation of myocardial perfusion, 63 Defibrillators, implantable cardioverter, safety of CMR with, 102, 107–108 Delayed enhancement. See Late gadolinium enhancement. Delayed gadolinium enhancement. See Late gadolinium enhancement. DENSE (displacement encoding with stimulated echoes), for left ventricular systolic function, 149 Deoxyhemoglobin, in BOLD technique, 62 Dephasing, 5f, 6 Dephasing gradient, 10 Diabetes, CMR spectroscopy in, 560–561 Diastolic function, 149–150 assessment of, 190, 191f Diastolic strain, 74 Diffusion tensor magnetic resonance imaging (DTMRI), after acute myocardial infarction, 255 Dilated cardiomyopathy (DCM), 516–517 CMR spectroscopy in, 561–563, 562f function and morphology in, 516–517 metabolic CMR in, 517 Diminishing variance algorithm, for navigator echoes, 131, 131t, 132f Diphenhydramine, in CMR stress tests, 198t 2,3-Diphosphoglycerate (2,3-DPG), in 31PCMRS, 559, 561f Dipyridamole for coronary sinus flow assessment, 311, 312f for myocardial oxygenation assessment, 570–571 safety considerations with, 104 during stress CMR, 208–209 for stress myocardial perfusion studies, 214, 216f Discordant atrioventricular connection, 409 Displacement encoding with stimulated echoes (DENSE), for left ventricular systolic function, 149 Distensibility, of blood vessel, 363 D-loop transposition of the great arteries, 409–410, 410f, 422–423 Dobutamine contraindications and termination criteria for, 236, 236t, 237t drug interactions with, 236 for myocardial oxygenation assessment, 570–571 pharmacologic effects of, 231t, 232 route and duration of administration of, 238 safety considerations with, 104, 234–236 stress-inducible perfusion abnormalities with, 232 stress-inducible wall motion abnormalities with, 232 626 Cardiovascular Magnetic Resonance

Dobutamine stress CMR (DCMR) abnormalities induced by, 232 apical and short axis views in, 198, 200f atropine in, 198, 199f, 202–203 cine GRE or SSFP bright-blood images in, 198, 199t combined adenosine perfusion and diagnostic performance of, 238–239, 238f protocol for, 232, 233f, 234f of coronary artery disease, 158–159 in comprehensive CMR assessment for disease detection, 159–160, 160t for viability studies, 164, 165f delineation of orthogonal left ventricular myocardial segments in, 198, 200f vs. dobutamine stress echocardiography, 202–203, 202t accuracy of, 197 safety of, 196–197, 197t technique of, 197 uses of, 197 for viability studies, 204 dopamine infusion protocol for, 198, 199f duration of, 237 facilities for, 198, 199f inducible ischemia during, 198, 201f contemporary studies of, 202–203, 202t, 203f, 203t early studies of, 201–202, 202f late gadolinium enhancement for, 205, 207f for myocardial infarction acute, 241 chronic, 275–278 in myocardial perfusion studies, 214, 216f pharmacokinetics of, 196 for prognosis, 205, 207f vs. radionuclide studies, 204 safety of, 196–197, 197t, 198t, 234–236 sensitivity and specificity of, 203, 203f, 203t vs. stress myocardial perfusion imaging, 229–240 technique for, 197–200 3-T, 203 tissue tagging during, 206–208, 208f for viability studies, 204–205, 207f in viability studies, 204–205 contractile reserve in, 204f, 206f end-systolic wall thickening in, 204, 204f intramyocardial segment shortening in, 205 vs. late gadolinium enhancement, 277–278 low-dose, 204–205, 237–238 short-axis basal views in, 205f tissue tagging in, 204–205, 207f Dobutamine stress echocardiography (DSE) accuracy of, 197 vs. dobutamine stress CMR, 202–203, 202t safety of, 196–197, 197t technique of, 197 uses of, 197 for viability studies, 204 Doppler ultrasonography for monitoring during CMR, 105 of peripheral vascular disease, 473 of renal artery stenosis, 471–472 DORV. See Double-outlet right ventricle (DORV). Dotarem (gadoterate), 79f, 79t compartmentalization and relaxivity with, 84 safety of, 87–88 Double inversion recovery dark-blood imaging, for congenital heart disease, 113, 114f, 115f Double inversion recovery fast spin echo imaging, 13, 13f

Double oblique planes, 19–20 Double-chambered right ventricle, 508 jet flow in, 504–505, 505f Double-outlet right ventricle (DORV), 410f, 412 defined, 425 epidemiology of, 425 features of, 425 in infant and pediatric patients, 425–426 postoperative assessment of, 426 preoperative assessment of, 426, 426f surgical management of, 425 late complications after, 425–426 ventricular septal defect in, 412, 425 2,3-DPG (2,3-diphosphoglycerate), in 31 P-CMRS, 559, 561f Dressler syndrome, 491–492 DSE. See Dobutamine stress echocardiography (DSE). DTMRI (diffusion tensor magnetic resonance imaging), after acute myocardial infarction, 255 D-transposition of the great arteries, 409–410, 410f, 422–423 Dual-bolus approach, in stress myocardial perfusion imaging, 222 Dual-T1-sensitivity method, in stress myocardial perfusion imaging, 222 Ductus arteriosus, 400 patent, 400 CMR evaluation of, 400 dark-blood imaging of, 115f interventional closure of, 400 pathogenesis of, 400 pulmonary arteries in, 485, 485f Duke Cardiovascular Magnetic Resonance Center (DCMRC) clinical volume at, 19, 20f overall makeup of, 19, 20f exam menu for, 19, 20f D-ventricular loop, 409 Dysprosium oxide catheter, for interventional CMR, 595, 596f

E

Early gadolinium hypoenhancement, and myocardial viability, 268 Earplugs, for CMR, 103 Ebstein anomaly, of tricuspid valve, 413, 414f, 415f ECF (extracellular fluid) contrast agents, 78–79, 79f, 79t ECG. See Electrocardiographic (ECG). Echinococcus, 536, 539f Echo(es), 6–7, 6f gradient recalled (See Gradient recalled echo [GRE]) navigator (See Navigator echoes) spin (See Spin echo imaging) Echo planar imaging (EPI), 16, 16f for chemical shift artifact, 147f, 148–149 for coronary artery velocity measurement, 314–315, 315f in CSPAMM, 70–71 for myocardial perfusion imaging, 58–59 for stress myocardial perfusion imaging, 218 Echo time (TE), 11f, 12 effect on signal of, 12, 12f for myocardial perfusion imaging, 58 Echocardiography to assess cardiac function, 181–182, 182f for cardiac allograft rejection, 554t dobutamine stress accuracy of, 197 vs. dobutamine stress CMR, 202–203, 202t safety of, 196–197, 197t

End-diastolic volume (EDV) (Continued) measurement of, 183–184, 183f, 184f intraobserver, interobserver, and interstudy variability in, 187f normal values for, 385–388, 388t in females, 385–388, 387t, 390f in males, 385–388, 387t, 389f for right ventricular assessment, 383–385, 386f Endocardial cushion defects, 398 Endocardial trabeculae, in left ventricular mass, 188 Endocarditis, Loeffler’s eosinophilic, 524–526, 525f morphology and function in, 526 tissue characterization in, 526 Endogenous contrast, for assessment of myocardial perfusion, 61–62 Endomyocardial biopsy, for cardiac allograft rejection, 554t Endomyocardial catheter ablation, 586 Endomyocardial delivery, of cellular agents, 585–587, 586f Endomyocardial diseases, 524–526, 525f morphology and function in, 526 tissue characterization in, 526 Endomyocardial fibrosis, 524–526, 525f morphology and function in, 526 tissue characterization in, 526 Endomyocardial mapping, 586 Endorem (ferumoxide), 83 Endothelial dysfunction, in atherosclerosis, 341 of aorta and carotid artery, 344–345 Endothelial function, assessment of, 370–371, 372f, 373f End-systolic volume (ESV) after acute myocardial infarction, 253–254, 255 measurement of, 183–184, 183f, 184f intraobserver, interobserver, and interstudy variability in, 187f normal values for, 385–388, 388t in females, 385–388, 387t, 390f in males, 385–388, 386t, 389f for right ventricular assessment, 383–385, 386f Eovist (gadoxetic acid), 85 EP-1242, for thrombus formation in aorta and carotid arteries, 345 EP-1873, for thrombus formation in coronary artery, 356–357, 357f EP-2104R, 85, 86f for thrombus formation in aorta and carotid arteries, 345, 346f in coronary artery, 356–357, 358f EPI. See Echo planar imaging (EPI). Epicardial calcium, 304–305, 305f Epicardial rim, thickness of viable, 276, 280f Ernst angle, in gradient echo imaging, 13–14, 59 E-selectin, in atherosclerosis of aorta and carotid artery, 345 of coronary artery, 358 ESV. See End-systolic volume (ESV). Exchange time, in contrast-enhanced tissue relaxation, 84 Excitation field, 3–5, 4f Exercise, for ventricular remodeling, 264 Exercise stress CMR, left ventricular wall motion during, 209–210, 210f Extracellular fluid (ECF) contrast agents, 78–79, 79f, 79t Extracellular matrix, in atherosclerotic plaques of aorta and carotid artery, 345–346

Extravascular contrast media, for stress myocardial perfusion imaging, 216–217, 217f

F 19

F (fluorine-19) CMR, catheter visualization and localization using, 598, 601f Fast acquisition relaxation mapping (FARM), T1, for myocardial perfusion imaging, 59–60, 60f Fast exchange, in contrast-enhanced tissue relaxation, 84 Fast Fourier transform, 11–12 Fast gradient recalled echo, 14, 14f inversion recovery, 15 three-dimensional, 15 Fast imaging with steady-state precession (FISP), to assess cardiac function, 185, 187 Fast low angle shot (FLASH) imaging to assess cardiac function, 185 for coronary artery velocity mapping, 315–316, 317–320, 319f, 320f turbo, 14, 14f for valvular heart disease, 505 Fast spin echo (FSE) imaging, 6–7, 13 of aorta, 468 of aortic coarctation, 400, 400f of coronary artery atherosclerotic plaques, 353 of coronary artery bypass graft, 300t, 305–306, 306f double inversion recovery (black-blood), 13, 13f pulse sequence diagram for, 13, 13f Fat, T1 and T2 values for, 7t Fat saturation, in coronary artery CMR, 288, 288f Fat-excitation acquisition, for coronary artery velocity mapping, 317–320, 321f Fatty streak, in atherosclerosis, 341 FDG. See Fluorodeoxyglucose (FDG). Ferromagnetic materials, 83 Ferromagnetism, safety of, 101, 101f, 102f Ferucarbotran (Resovist), relaxivity with, 83 Ferumoxide (Endorem, Feridex), 83 Ferumoxtran (AMI-227, Sinerem, Combidex), 83 Ferumoxytol (Feraheme), 85 for 3D MRA, 465f Fetal CMR, functional real-time, 125–127, 126f 18 F-fluorodeoxyglucose positron emission tomography, to assess cardiac function, 182 FFR (fractional flow reserve), 230 FHS (Framingham Heart Study), aortic atherosclerosis in, 342–343 Fibrin-targeted contrast agent, 85, 86f Fibroelastoma, papillary, 534, 536f, 544t Fibroma(s), 534–536, 537f, 544t pediatric, 117, 118f pericardial, 495 Fibromuscular dysplasia, renal artery stenosis due to, 470 Fibrous cap, in atherosclerosis, 341 ruptured, 341–342 CMR imaging of, 343–344 Fick principle, 595 Fick’s law, 569 FID (free induction decay), 4f, 5 in CMR spectroscopy, 556–557 Field of view (FOV), 10 Field strength high (See High field CMR)

Cardiovascular Magnetic Resonance 627

INDEX

Echocardiography (Continued) technique of, 197 uses of, 197 for viability studies, 204 for pulmonary artery hypertension, 482t for right ventricular assessment, 382 transesophageal for aortic dissection, 453–454 of aortic intramural hematoma, 454–455 of pericardial disease, 488 of thoracic aortic aneurysm, 457 transthoracic of atrial septal defect, 398–399 of cardiomyopathy, 515 of congenital heart disease, 408, 415 of constrictive pericarditis, 493 of pericardial disease, 488 of pericardial effusions, 492 Eddy currents, 94–95 with interventional CMR, 598 Edema in acute myocardial infarction, 269 and myocardial viability, 268 in myocarditis, 517f, 518 EDV. See End-diastolic volume (EDV). Effusive-restrictive pericarditis, 493 Ejection fraction (EF) intraobserver, interobserver, and interstudy variability for measurement of, 187f normal values for, 385–388, 388t in females, 385–388, 387t, 390f in males, 385–388, 386t, 389f reference ranges for, 190t for right ventricular assessment, 383–385 Elastic limit, 363 Elastic modulus aortic, 369 of vascular wall, 363 Elasticity, of blood vessel, 363 Elastin fibers, in arterial wall, 362–363 Electrical safety, with interventional CMR, 598–599 Electrocardiographic (ECG) electrodes, safety of, 103, 104–105, 105f Electrocardiographic (ECG) gating for atherosclerotic plaques of coronary artery, 351–352 for congenital heart disease, 408 for coronary sinus flow assessment, 311 in pediatric CMR, 117, 119f in PET, 182 for postoperative atrial switch, 423, 423f prospective vs. retrospective, 188 for right ventricular assessment, 382–383 in SPECT, 182 for thoracic aorta, 450, 451f, 468 Electrocardiographic (ECG) setup, for patient monitoring, 104–105, 105f Electrocardiographic (ECG) synchronization, for interventional CMR, 602 Electronic implants, as contraindication to CMR, 105–106 Electrophysiology, interventional CMR for, 586 Embolism, pulmonary, 480–481 catheter-based X-ray pulmonary angiography of, 480 CE-CTA for, 480 CE-MRA of, 481, 481f, 482f CMR lung perfusion imaging of, 481, 481f current workup for, 480 incidence of, 480 pulmonary artery hypertension with, 482f ventilation/perfusion scanning for, 480 Enalaprilat, for ventricular remodeling, 262 End-diastolic volume (EDV) after acute myocardial infarction, 255

INDEX

Field strength (Continued) for stress myocardial perfusion imaging, 219 FISP (fast imaging with steady-state precession), to assess cardiac function, 185, 187 FLASH imaging. See Fast low angle shot (FLASH) imaging. Flip angle, in gradient echo imaging, 13–14, 59 Flow data, visualizing, 96–97 Flow enhancement methods, 91, 92f Flow mapping of coarctation of the aorta, 400–402, 402f, 458–459 of congenital heart disease, 396 of thoracic aorta, 450–451, 451f Flow pressure maps, 96–97, 97f Flow vector map, 96, 97f Flow velocity, phase of signal and, 92 Flow velocity images, in phase contrast velocity mapping, 92–93, 93f Flow wave, foot of, 366, 366f Flow wave velocity aortic compliance and, 368, 368f CMR of, 364–367, 365f, 366f, 367f defined, 364–366 Flow-related enhancement, 463 Flow-related signal loss, 94, 95f Flow/velocity imaging, 32–34 cardiac gating for, 32 goal of, 32, 33f pulse sequence in, 33–34, 33f Fluorine-19 (19F) CMR, catheter visualization and localization using, 598, 601f Fluorodeoxyglucose positron emission tomography (FDG PET), to assess cardiac function, 182 Fluorodeoxyglucose (FDG) uptake, in chronic myocardial infarction vs. late gadolinium enhancement, 276–277, 281f and myocardial wall thickness, 275, 275f and viable epicardial rim thickness, 276, 280f Fluoroscopy, X-ray, vs. interventional CMR, 580, 581t Fluvastatin, and aortic compliance, 368–369 Fontan baffle, with single ventricle, 123, 123f, 125f Fontan procedure, 416 aortopulmonary, 432, 432f classic, 432, 432f extracardiac, 432, 432f fenestrated, 432, 432f for single ventricle, 125f, 432, 432f postoperative assessment of, 434–435, 434f variations on, 432, 432f Fossa ovalis defect, 398, 398f, 399f Fourier flow imaging, 91, 93 velocity phase encoding in, 93, 93f visualizing flow data in, 96, 96f Fourier velocity imaging, 96, 96f phase contrast velocity mapping and, 94 for valvular heart disease, 507 Fourier velocity-encoded measurement, of aortic flow wave velocity, 366–367, 367f FOV (field of view), 10 Fractional flow reserve (FFR), 230 Frame rate, for ventricular function, 149 Framingham Heart Study (FHS), aortic atherosclerosis in, 342–343 Free breathing methods, 129–131 in coronary artery CMR, 286, 286t, 289 with native vessel stenosis, 302–303, 303t for coronary sinus flow assessment, 311–312 mean diaphragm displacement in, 130f respiratory trace data for, 130f

628 Cardiovascular Magnetic Resonance

Free induction decay (FID), 4f, 5 in CMR spectroscopy, 556–557 Frequency encoding, 7, 7f, 9–10, 9f pulse sequence diagram with, 10, 10f Frequency encoding direction, 9–10 Frequency encoding gradients, 7–8, 9–10, 9f Friedreich ataxia, CMR spectroscopy for, 562 FSE imaging. See Fast spin echo (FSE) imaging. Functional imaging, high field CMR for, 170, 171f Functional real-time fetal CMR, 125–127, 126f Fundamental law of magnetostimulation, 102

G

Gadobenate dimeglumine (Gd-BOPTA, MultiHance), 81, 85 binding and relaxivity features of, 81t, 83t chemical structure of, 80f for coronary artery CMR, 293, 293f safety of, 87–88 Gadobutrol (Gd-DO3A-butrol, Gadovist), 79f, 79t safety of, 87–88 Gadocoletic acid (B22956), 81 binding and relaxivity features of, 81, 81t chemical structure of, 80f factors affecting relaxivity of, 84 Gadodiamide (Gd-DTPA-BMA, Omniscan), 79f, 79t relaxivity with, 83t safety of, 87–88 Gadofluorine(s), 85, 86f Gadofluorine M, for atherosclerotic plaques of aorta and carotid arteries, 345–346 Gadofosveset trisodium (MS-325, Vasovist, Ablavar), 80, 85 binding of, 81, 81t chemical structure of, 80f for coronary artery CMR, 293–294 relaxivity of, 81, 81t, 83t factors affecting, 84 magnetic field dependence on, 82, 82f safety of, 87 for stress myocardial perfusion imaging, 217 for three-dimensional MRA, 464 Gadolinium (Gd)-based CMR, for pediatric imaging, 116–118 of arrhythmogenic right ventricular cardiomyopathy, 118 of coronary artery, 117, 119f late gadolinium enhancement, 117, 118f myocardial and blood tagging in, 116, 119f myocardial perfusion imaging as, 116, 118f, 120 saturation band in, 116, 119f static, 116, 117f technical considerations with, 120 time-resolved, 116, 118f of tumor or mass, 117 Gadolinium (Gd) chelates for stress myocardial perfusion imaging, 216–217, 217f for three-dimensional MRA, 464 Gadolinium (Gd) contrast agents, 78–79, 79f, 79t safety of, 87–88, 103–104 Gadolinium diethylenetriamine pentaacetic acid (Gd-DTPA, gadopentetate dimeglumine, Magnevist), 79f, 79t for atherosclerotic plaques of aorta and carotid arteries, 346

Gadolinium diethylenetriamine pentaacetic acid (Gd-DTPA, gadopentetate dimeglumine, Magnevist) (Continued) for cardiac allograft rejection, 548, 550 relaxivity with, 83t compartmentalization and, 84 magnetic field dependence on, 82, 82f safety of, 87–88, 103 for stress myocardial perfusion imaging, 217f for three-dimensional MRA, 464 for ventricular remodeling, 257–258 Gadolinium enhancement, late. See Late gadolinium enhancement (LGE). Gadolinium (Gd)-mesoporphyrin, after acute myocardial infarction, 259 Gadolinium-enhanced T1-weighted CMR, for cardiomyopathy, 516 Gadolinium-enhanced three-dimensional cardiovascular MRA, 464–465 gadolinium chelates for, 464, 465f pulse sequences for, 464 vessel brightness (T1) vs. gadolinium dose in, 464–465 Gadomer-17 binding of, 81, 81t chemical structure of, 80f relaxivity of, 81, 81t, 83t molecule size and, 82 Gadopentetate dimeglumine. See Gadolinium diethylenetriamine pentaacetic acid (GdDTPA, gadopentetate dimeglumine, Magnevist). Gadoterate (Gd-DOTA, Dotarem), 79f, 79t compartmentalization and relaxivity with, 84 safety of, 87–88 Gadoteridol (Gd-HPDO3A, ProHance), 79f, 79t relaxivity with, 83t safety of, 87 Gadoversetamide (Gd-DTPA-BMEA, Optimark), 79f, 79t safety of, 87–88 Gadovist (gadobutrol), 79f, 79t safety of, 87–88 Gadoxetic acid (Gd-EOB-DTPA, Primovist, Eovist), 85 Gating, cardiac. See Cardiac gating. Gd. See Gadolinium (Gd). Gd-BOPTA. See Gadobenate dimeglumine (Gd-BOPTA, MultiHance). Gd-DO3A-butrol (gadobutrol), 79f, 79t safety of, 87–88 Gd-DOTA (gadoterate), 79f, 79t compartmentalization and relaxivity with, 84 safety of, 87–88 Gd-DTPA. See Gadolinium diethylenetriamine pentaacetic acid (Gd-DTPA, gadopentetate dimeglumine, Magnevist). Gd-DTPA-BMA (gadodiamide), 79f, 79t relaxivity with, 83t safety of, 87–88 Gd-DTPA-BMEA (gadoversetamide), 79f, 79t safety of, 87–88 Gd-EOB-DTPA (gadoxetic acid), 85 Gd-HPDO3A (gadoteridol), 79f, 79t relaxivity with, 83t safety of, 87 General anesthesia, for pediatric CMR, 120 General Electric (GE), CMR terminology used by, 611 Generalized autocalibrating partially parallel acquisition (GRAPPA), 45–46, 45f

H

Half-Fourier single-shot fast spin echo (HASTE) for coronary artery bypass graft, 330–331, 334f in morphology scanning goal of, 21–22, 23f physiology of, 22, 26f Hand, peripheral vascular disease of, 475 Harmonic phase (HARP) technique for dobutamine stress CMR, 207–208 for left ventricular systolic function, 149 HCM. See Hypertrophic cardiomyopathy (HCM). 1 H-CMRS. See Proton CMR spectroscopy (1H-CMRS). HDL (high-density lipoprotein), in atherosclerotic plaques of aorta and carotid arteries, 346–347, 347f HDL (high-density lipoprotein)-based contrast agent, 85 Headphones, for CMR, 103 Hearing impairment, due to CMR, 103 Heart disease congenital (See Congenital heart disease [CHD]) valvular (See Valvular heart disease)

Heart failure CMR spectroscopy in, 561–563, 562f right ventricular assessment in, 388–391 Heart rates, in pediatric CMR, 120 Heart transplantation. See Cardiac transplantation. Heart valve(s) disease of (See Valvular heart disease) mechanical (prosthetic), 512–513 safety of CMR with, 101, 101f, 106 Heating-related injury from CMR imaging, 102 from interventional CMR, 582–583, 583t, 598–599 Hemangioma(s) cardiac, 536, 538f, 544t pericardial, 495 Hematoma, aortic intramural, 454–455, 455f Hemi-Fontan procedure, for single ventricle, 124, 124f Hemopericardium, 491–492 Hemorrhagic infarcts, 259 Hemorrhagic pericardial effusions, 495 Hibernating myocardium, 267 High field CMR, 170–178 of coronary artery, 172–173, 173f, 294, 294f of coronary artery wall, 174, 174f for dobutamine stress CMR, 203 for functional imaging, 170, 171f for late gadolinium enhancement imaging, 174 limitations of, 170, 171f for magnetic resonance angiography, 466, 466f for magnetic resonance spectroscopy, 175, 565, 565f for myocardial stress perfusion imaging, 174 for myocardial tagging, 170–171, 172f for oxygen-sensitive myocardial MRI, 576, 577f for parallel imaging, 175 rationale for, 170 High-density lipoprotein (HDL), in atherosclerotic plaques of aorta and carotid arteries, 346–347, 347f High-density lipoprotein (HDL)-based contrast agent, 85 High-energy phosphates, and myocardial viability, 268 Highly constrained back projection (HYPR), 52f, 53 Hitachi, CMR terminology used by, 611 HMG-CoA (3-Hydroxy-methylglutaryl coenzyme A) reductase inhibition, in ventricular remodeling, 261 Hooke’s law, 363 Horizontal long axis (HLA) image, 140, 141f, 186–187 for left ventricular function and size, 142 Hydatid cysts, 536, 539f Hydrogen spins, 3, 4f 3-Hydroxy-methylglutaryl coenzyme A (HMGCoA) reductase inhibition, in ventricular remodeling, 261 Hyperpolarized contrast media, for stress myocardial perfusion imaging, 217–218 Hypertension CMR spectroscopy in, 560–561 pulmonary artery (See Pulmonary artery [PA] hypertension) Hypertrophic cardiomyopathy (HCM), 518–520, 519f CMR spectroscopy for, 565 follow-up for, 520 function and morphology in, 519

Hypertrophic cardiomyopathy (HCM) (Continued) LVOT obstruction in, 518, 519–520 tissue characterization in, 519 Hypoplastic left heart syndrome aortic arch, pulmonary artery, and venous pathway imaging with, 123–124, 123f myocardial and blood tagging for, 119f staged palliation of, 433, 433f, 434f HYPR (highly constrained back projection), 52f, 53

I

ICAM-1 (intercellular adhesion molecule 1), in atherosclerosis of aorta and carotid artery, 345 of coronary artery, 358 IMA (internal mammary artery) graft, 329 Image-selected in vivo spectroscopy (ISIS), for cardiac allograft rejection, 551–552, 552f Imaging planes, and cardiac anatomy, 140, 141f, 143f IMH (intramural hematoma), aortic, 454–455, 455f Implantable cardioverter defibrillators (ICDs), safety of CMR with, 102, 107–108 Impulse response, in quantitative evaluation of myocardial perfusion, 62–63 Inducible nitric oxide synthase (iNOS), in ventricular remodeling, 260 Induction heating, with interventional CMR, 598 Infants, CMR in. See Pediatric CMR. Infarct(s), hemorrhagic vs. nonhemorrhagic, 259 Infarct expansion, after acute myocardial infarction, 253, 274 Infarct resorption, after acute myocardial infarction, 258, 258f Inferior vena cava filter, interventional CMR for, 589 Inflammation, with atherosclerotic plaques of aorta and carotid arteries, 346–347, 347f of coronary artery, 358 Infundibulum, of right ventricle, 381 Inner-sphere relaxivity, 81–82 Inorganic phosphate (Pi), in 31P-CMRS, 557, 558, 558f iNOS (inducible nitric oxide synthase), in ventricular remodeling, 260 Inotropic stress, for comprehensive CMR assessment of coronary artery disease, 159–160 In-plane velocity mapping of congenital heart disease, 113 of coronary artery, 316–317, 317f Interatrial septum, lipomatous hypertrophy of, 142, 146f, 533–534, 535f, 544t Intercellular adhesion molecule 1 (ICAM-1), in atherosclerosis of aorta and carotid artery, 345 of coronary artery, 358 Intercept Internal CMR guidewire coil, 587 Interleaved spiral imaging, for coronary artery velocity mapping, 317–320, 319f, 320f, 321f Intermediate exchange, in contrast-enhanced tissue relaxation, 84 Internal mammary artery (IMA) graft, 329 Interrupted aortic arch, 428–430 vs. aortic arch atresia, 428 epidemiology of, 428 in infant and pediatric patients classification of, 428, 429f evaluation of, 429

Cardiovascular Magnetic Resonance 629

INDEX

Generalized encoding matrix (GEM), 45f, 46 Glagov effect, in atherosclerosis, 341, 353–354 Glenn shunt, 416 for single ventricle, 123, 124, 433, 433f Gradient(s), 7–8 Gradient coils, 3, 7–8 Gradient magnetic fields, bioeffects of, 598 Gradient recalled echo (GRE), 10, 13–14, 14f of aorta, 468–469 thoracic, 450–451 for cardiac and paracardiac masses, 532 for congenital heart disease, 396, 408 of coronary artery, 289–290 with coronary artery bypass graft, 305–306, 306t conventional, 330, 332f three-dimensional respiratory gated, 331 for coronary artery velocity mapping, 315–317 breath hold technique for, 315–316 navigator techniques for, 316–317, 317f, 318f with native vessel stenosis, 301–302, 301f, 303t fast, 14, 14f inversion recovery, 15 for stress myocardial perfusion imaging, 218 three-dimensional, 15 for myocardial perfusion imaging single-shot, 58, 59f with steady-state free precession, 59 for right ventricular assessment, 383, 384f spoiled, with contrast agents, 77, 78f for stress tests, 198, 199t for ventricular volumes, 150–151, 151t Gradient strengths, 7–8 GRAPPA (generalized autocalibrating partially parallel acquisition), 45–46, 45f GRE. See Gradient recalled echo (GRE) Great arteries relationship of, 409 transposition of (See Transposition of the great arteries [TGA]) Grid-tagged images, in CSPAMM, 70, 71f, 72f Guidewires, in interventional CMR, 600 Gyromagnetic ratio, 3

INDEX

Interrupted aortic arch (Continued) postoperative assessment of, 430, 430f preoperative assessment of, 429, 429f with truncus arteriosus, 426, 427f, 428f surgical repair of, 428–429 Interventional CMR, 580–592 application(s) of cardiac, 585–587, 605 for atrial and ventricular mapping, 586 for atrial septal defect, 399 for atrial transseptal procedures, 585f, 586 for congenital heart disease, 396, 415–417 future directions in, 605, 605f for invasive coronary artery imaging and intervention, 586–587 for patent ductus arteriosus, 400 recent progress in, 605, 605f for RF ablation, 586, 603, 604, 604f for targeted local delivery of cellular agents to myocardium, 585–586, 586f for valve replacement and repair, 587, 605f for ventricular septal defect, 397 extracardiac, 587–589 for aortic aneurysm and aortic dissection repair, 587, 588f for aortic coarctation stent repair, 584f, 587 for inferior vena cava filter, 589 for invasive arterial imaging, 587 for peripheral artery disease, 589, 589f for transjugular intrahepatic portosystemic shunt, 587–589 catheter devices for, 583, 595–598 active, 584–585, 585f, 596–598 advantage of, 596–597 for aortic coarctation repair, 584f for endomyocardial injection, 586f 19 F, 598, 601f for intramyocardial injection, 598f with multiple resonant coils, 597–598, 600f safe transmission line for, 599f semi-, 597–598 for transseptal puncture, 585f vs. conventional XRF devices, 583, 583f passive, 583–584, 595–596 chemical-selective visualization of, 584, 584f CO2-filled balloon in, 595, 596f, 597f dysprosium oxide, 595, 596f ideal material for, 595 that create bright signals, 583–584, 584f that create dark signals, 583 strategy for, 585 communication and monitoring with, 580–581 ECG synchronization for, 602 laboratory for, 580–585, 581f magnetic instrumentation and visualization strategies for, 595–598 merit(s) of, 593–595 improved visualization of cardiac anatomy as, 593–594 physiologic information as, 595 reduced ionizing radiation as, 594 pediatric, 593–610 reduced ionizing radiation in, 594 real-time, 581t, 583

630 Cardiovascular Magnetic Resonance

Interventional CMR (Continued) safety consideration(s) with, 582–583, 598–600 bioeffects of magnetic fields as, 598 heating and electrical safety as, 465, 583t, 598–599 magnetic force and torque as, 600 scanner interface for, 582, 582f system design for, 593, 594f in thoracic aorta, 459–460 vs. ultrasound, 580, 581t in XMR system, 593, 601–604 for biventricular pacing, 605f early experience in humans with, 603–604, 604f facility design for, 594f, 601–602 image registration in, 604 performance of, 602–603, 603f safety features for, 594f, 601 vs. X-ray fluoroscopy, 580, 581t Interventricular septum, rhabdomyoma of, 538f Intestinal ischemia, 472 Intima, of artery, 362–363 Intracardiac thrombus, 540–542, 542f, 544t Intracoronary stents coronary artery CMR with, 294–295, 294f safety of CMR with, 106 Intramural hematoma (IMH), aortic, 454–455, 455f Intramyocardial injection, active catheter tracking and visualization for, 598f Intramyocardial segment shortening, in dobutamine stress CMR, 205 Intravascular contrast media, for stress myocardial perfusion imaging, 217 Inversion recovery curve, in viability imaging, 31, 32f Inversion recovery fast gradient recalled echo, 15 Inversion recovery pulse sequence, 14f, 15 for acute myocardial infarction, 269, 269f Inversion recovery technique for acute myocardial infarction, 242 for cardiac and paracardiac masses, 532 Inversion time (TI), 14f, 15 Iron oxide–based contrast agents cross-linked, 86–87 in development, 86–87 relaxivity with, 83 structure of, 83 uses of, 81 Ischemia myocardial (See Myocardial ischemia) subendocardial vulnerability to, 244 Ischemic bed at risk, in acute myocardial infarction, 271, 272f Ischemic cascade, 230 Ischemic heart disease. See also Myocardial ischemia. CMR spectroscopy in, 563–565 for myocardial viability assessment, 564–565, 564f for stress testing, 563–564, 563f right ventricular assessment in, 391 ISIS (image-selected in vivo spectroscopy), for cardiac allograft rejection, 551–552, 552f Isolated ventricular inversion, 409 Isomerism, 409

J

Jatene procedure, for transposition of the great arteries, 410–411, 416 Jet flow, 504, 505f Jet velocity mapping, for stenotic valvular heart disease, 501, 506

K

Kawasaki disease, 299–301, 300f mural thrombosis in, 353, 354f Killer gradient, 14, 14f k-space, 11–12, 11f k-space data, raw, 11–12, 11f in cine CMR, 24–25, 27f in scout scanning, 20–21, 22f k-space gradient echo imaging, of coronary artery, 289–290 k-space ordering, for navigator echoes, 131, 131t, 132f k-space signal, 11f k-t broad-use linear acquisition speed-up technique (k-t BLAST) applications of, 50f, 51 to assess cardiac function, 185 for myocardial perfusion imaging, 61 principles of, 47–48 k-t sensitivity encoding (k-t SENSE) applications of, 51 to assess cardiac function, 185 for stress myocardial perfusion imaging, 218

L

LA (left atrium), morphology of, 408 Larmor equation, 3 Late gadolinium enhancement (LGE) imaging, 31 cardiac gating for, 31, 32f of cardiomyopathy dilated, 516–517 hypertrophic, 519 of congenital heart disease, 117, 118f, 408 of coronary artery atherosclerotic plaques, 354–355, 355f in coronary artery disease for detection of disease analysis of, 164, 167f protocols for, 160–164, 160t, 161f, 162f, 163f for viability assessment analysis of, 164–166, 167f protocols for, 163–164, 165f, 166f, 167f for dobutamine stress CMR, 205, 207f goal of, 31, 31f high field CMR for, 174 in infarcted tissue, 268 inversion recovery in, 31, 32f of myocardial infarction acute, 242, 243f, 269–271 adjustment of T1 in, 271, 273f ischemic bed at risk in, 271, 272f and left ventricular remodeling, 258, 258f, 259f nulling of signal intensity of normal myocardium in, 269, 270f and recovery of function, 273–274 T1-weighted segmented inversion recovery pulse sequence for, 269, 269f time between contrast injection and imaging in, 269–271, 271f validation of, 243–244, 244t chronic, 276, 277f, 278f, 279f of myocarditis, 517f, 518 in parallel imaging, 50f, 51 for pulmonary artery hypertension, 484–485, 485f of pulmonary veins, 447, 447f with single ventricle, 124f LDL (low-density lipoprotein), in atherosclerosis, 341 Lecompte maneuver, 120–121, 121f Left atrium (LA), morphology of, 408

Left ventricular (LV) stroke volume, validation of, 186f Left ventricular (LV) systolic function, reference ranges for, 189t, 191f Left ventricular (LV) thrombosis, due to acute myocardial infarction, 248, 249f Left ventricular (LV) volume after acute myocardial infarction, 253–255 measurement of, 185f reference ranges for, 189t, 191f Left ventricular (LV) wall motion assessment cine CMR of, 22–25 acquisition time in, 24 in acute myocardial infarction, 241, 242f cardiac gating for, 24–25, 27f goal of, 23–24, 27f wall motion stress CMR for (See Wall motion stress CMR) Left ventricular (LV) wall thickness, and myocardial viability, 267 Left-sided isomerism, 409 Leiomyomatosis, with intracardiac extension, 536, 539f, 544t Leiomyosarcoma, 537, 544t Levo-transposition of the great arteries, 410–411, 411f, 423 LGE. See Late gadolinium enhancement (LGE). Lipoma(s), 533–534 atrial, 535f endocardial, 533–534 epidemiology of, 533–534 pediatric, 117 pericardial, 495 subepicardial, 533–534 tissue characterization of, 544t Lipomatous hypertrophy, of atrial septum, 142, 146f, 533–534, 535f, 544t Liposarcoma, 540, 541f, 544t L-loop transposition of the great arteries, 410–411, 411f, 423 L-NAME (N-methyl-L-arginine methyl ester), for ventricular remodeling, 261 Loeffler’s eosinophilic endocarditis, 524–526, 525f morphology and function in, 526 tissue characterization in, 526 Longitudinal direction, 3 Low-density lipoprotein (LDL), in atherosclerosis, 341 L-transposition of the great arteries, 410–411, 411f, 423 Lung, T1 and T2 values for, 7t Lung perfusion imaging for pulmonary artery hypertension, 483–484, 483f for pulmonary embolism, 481, 481f Lung transplantation, right ventricular assessment in, 391–392 LV. See Left ventricle (LV). LVOT. See Left ventricular outflow tract (LVOT). Lymphoma, 540, 541f, 544t

M

Macrophage(s) in atherosclerotic plaques of aorta and carotid arteries, 346 in cardiac allograft rejection, 548 MPIO-labeled, 549, 549f USPIO-labeled, 548–549, 549f Macrophage Scavenger Receptor-A (MSR-A), for atherosclerotic plaques of aorta and carotid arteries, 346 Magnet strength, 3–5, 4f

Magnetic field(s) alignment with main, 3 bioeffects of, 598 in CMR spectroscopy, 556–557 safety of, 100, 598 radiofrequency, 102–103 rapidly switched, 101–102 Magnetic force, in interventional CMR, 600 Magnetic resonance angiography (MRA), 34–35, 463–479 of aorta, 468–470 thoracic, 451–452, 452f of aortic aneurysms, 470 thoracic, 456–457, 457f of aortic dissection, 452–453, 454, 454f, 470 basic principles and techniques of, 463–466 for cardiac allograft rejection, 554t cardiac gating for, 34–35 contrast-enhanced (See Contrast-enhanced magnetic resonance angiography [CE-MRA]) of coronary artery bypass graft three-dimensional contrast-enhanced breath hold, 332, 335f, 336f three-dimensional respiratory gated, 331 two-dimensional breath hold, 330–331, 333f, 334f of coronary artery disease, 159, 160t, 161, 162f of extracranial carotid arteries, 467–468, 467f, 468f fast gradient echo in, 15 gadolinium-enhanced 3D, 464–465 gadolinium chelates for, 464, 465f pulse sequences for, 464 vessel brightness (T1) vs. gadolinium dose in, 464–465 goal of, 34, 34f of mesenteric arteries, 472–473 non-contrast approaches to, 463 of peripheral vessels, 473–475, 475f phase contrast, 94, 464 pulse sequence in, 35 of renal arteries, 470–472 of renal artery stenosis, 472 at 3-Tesla, 466 time-of-flight, 463 of extracranial carotid arteries, 467–468, 467f timing of image acquisition in, 34, 35f Magnetic resonance imaging (MRI) alignment with main magnetic field in, 3 balanced steady-state free precession in, 15, 15f basic principles of, 1–18 echo planar imaging, spiral, and radial in, 16–17, 16f frequency encoding: position in X in, 9–10, 9f, 10f gradient echo imaging in, 13–14, 14f gradients in, 7–8 image creation in, 7, 7f inversion recovery fast gradient recalled echo: late gadolinium enhancement in, 15 phase encoding: position in Y in, 10–11, 10f pulse sequences and contrast in, 12–13, 12f radiofrequency and magnet strength in, 3–5, 4f raw k-space data and fast Fourier transform in, 11–12, 11f selective excitation: position in Z in, 8–9, 8f, 9f

Cardiovascular Magnetic Resonance 631

INDEX

Left ventricle (LV) morphology of, 409 single, 430, 431f Left ventricular (LV) anatomy, during remodeling, 253–255, 254f Left ventricular (LV) diastolic function CMR tagging assessment of, 69–75 apical rotation in, 71f, 73–74, 73f, 73t cardiac motion and, 69–70 complementary spatial modulation of magnetization for, 70–71 echo planar imaging in, 70–71 grid-tagged images in, 70, 71f, 72f limitations of, 74 slice thickness in, 70 slice-following principle in, 70, 70f spiral imaging in, 70–71 steady-state free precession in, 70–71 with strain measurements, 74 3D, 71 in dobutamine stress CMR, 208f, 209 evaluation of motion data from, 71–72, 73f, 73t limitations of, 74 methods for, 70–72 during remodeling, 255–256 results of, 72–74 strain measurement in, 74 normal values for, 190, 192f Left ventricular ejection fraction (LVEF), after acute myocardial infarction, 255 Left ventricular (LV) function, 149 in acute myocardial infarction, 241–242, 242f CMR assessment of, 183–185, 183f short axis slices in, 183–184, 184f, 188 Simpson’s rule method for, 181–182, 184f global, 149 imaging planes for, 141f, 142 regional systolic, 149, 150f during remodeling, 255–256 Left ventricular (LV) hypertrophy, CMR spectroscopy for, 563 Left ventricular (LV) mass, 142, 150 after acute myocardial infarction, 254–255 effect of imaging sequence and magnetic field strength on, 150–152, 151t papillary muscles and endocardial trabeculae in, 188 reference ranges for, 189t, 191f validation of, 186f Left ventricular outflow tract (LVOT), volume flow through, in aortic stenosis, 507 Left ventricular outflow tract (LVOT) obstruction, in hypertrophic cardiomyopathy, 518, 519–520 Left ventricular outflow tract (LVOT) view, 140, 141f Left ventricular (LV) remodeling, after acute myocardial infarction, 253–266 CMR spectroscopy of energetics during, 256–257, 257f contrast-enhanced CMR and predictors of, 257–259, 258f, 259f early phase of, 253 late phase of, 253, 254f pathophysiology of, 253 regional left ventricular function during, 255–256 therapy for animal studies of, 259–262, 261f human studies of, 262–264, 263f ventricular anatomy during, 253–255, 254f

INDEX

Magnetic resonance imaging (MRI) (Continued) signal detection in, 3–17 spin echo imaging in, 6–7, 6f, 7t fast (turbo), 13 double inversion recovery (black-blood), 13, 13f T1 relaxation, 5 T2 relaxation and spin phase in, 5–6, 5f three-dimensional fast gradient echo: magnetic resonance angiography in, 15 Magnetic resonance (MR) signal, detection of, 3–17 Magnetic resonance spectroscopy (MRS), cardiovascular. See Cardiovascular magnetic resonance spectroscopy (CMRS). Magnetization preparation, for stress myocardial perfusion imaging, 218–219, 219f Magnetization recovery, in myocardial perfusion scanning, 28, 29f Magnetization transfer contrast for assessment of myocardial perfusion, 61–62 in coronary artery CMR, 288, 288f Magnetohydrodynamic effect, 104–105, 105f Magnetostimulation, fundamental law of, 102 Magnevist. See Gadolinium diethylenetriamine pentaacetic acid (Gd-DTPA, gadopentetate dimeglumine, Magnevist). Malignant cardiac tumor(s), 537–540 lymphoma as, 540, 541f, 544t metastatic, 540 sarcoma as, 537–540 angio-, 537, 540f, 544t leiomyo-, 537, 544t lipo-, 540, 541f, 544t Marfan syndrome, aortic distensibility and stiffness in, 369 Matrix metalloproteinases (MMPs), in atherosclerotic plaques of aorta and carotid arteries, 346 Maximal intensity projection (MIP), for pulmonary veins, 441, 442f Maxwell gradients, 94–95 MBF (myocardial blood flow), 57 measurement of, 311, 313f MBF (myocardial blood flow) reserves, 311–312, 313f MDCT. See Multidetector computed tomography (MDCT). Mean transit time (MTT), in stress myocardial perfusion imaging, 221 Mechanical heart valves, 512–513 safety of CMR with, 101, 101f, 106 Mechanical restraint devices, for ventricular remodeling, 261 Media, of artery, 362–363 Menotropin, and aortic compliance, 369 MERIT-HF study, 262 MESA (Multiethnic Study of Atherosclerosis), aortic atherosclerosis in, 342–343 Mesenteric arteries, MRA of, 472–473 Mesenteric artery stenosis, 472–473 intestinal ischemia due to, 472 MRA of, 472–473 contrast-enhanced, 473 noncontrast CMR of, 473 X-ray angiography of, 472 Mesenteric ischemia, 472 Mesocaval shunt, interventional CMR for, 589 Mesoscopic inhomogeneities, in contrastenhanced tissue relaxation, 84–85 Metal artifacts, 142–145, 146–148, 147f

632 Cardiovascular Magnetic Resonance

Metallic shard injuries, as contraindication to CMR, 105–106 Metastasis, pericardial, 542–543 Metastatic cardiac tumors, 540 Metoprolol for ventricular remodeling, 260, 262 MI. See Myocardial infarction. Micrometer-sized paramagnetic iron oxide (MPIO), for cardiac allograft rejection, 549, 549f Microvascular obstruction (MO), after acute myocardial infarction contrast-enhanced CMR of, 257–258 late gadolinium enhancement of, 258–259, 259f no-reflow phenomenon and, 271–272, 274f pathophysiology of, 253 prognostic significance of, 246 and regional recovery of function, 246–247 residual coronary occlusion vs., 245–246, 245f, 246f Microvessels, in coronary flow reserve, 230 MIP (maximal intensity projection), for pulmonary veins, 441, 442f Missile effect, with ferromagnetic objects, 101 Mitral regurgitant fraction, 512 Mitral regurgitant volume, 512 Mitral regurgitation, 510–512 asymmetric, 510, 512f central, 510, 511f quantification of, 512 severity of, 502t Mitral stenosis, 502t, 508 Mitral valve replacement, interventional CMR for, 587 Mitral valve stenosis, 502t, 508 M-mode echocardiography, to assess cardiac function, 181–182, 182f MMPs (matrix metalloproteinases), in atherosclerotic plaques of aorta and carotid arteries, 346 MO. See Microvascular obstruction (MO). Moderator band, 381 Modified Simpson’s rule, 149 Molecular imaging, of atherosclerotic plaques in aorta and carotid arteries, 344–347 angiogenesis in, 345 endothelial dysfunction in, 344–345 extracellular matrix in, 345–346 plaque inflammation in, 346–347, 347f thrombus in, 345, 346f in coronary arteries, 356, 356f angiogenesis in, 358–359 inflammation in, 358 thrombosis in, 356–358, 357f, 358f Monitoring during CMR, 104–105, 105f in interventional CMR laboratory, 580–581 Monophosphate esters (MPE), in 31P-CMRS, 557 Morphology scanning, 21–22 black-blood imaging in cardiac gating for, 26f goal of, 21–22, 23f physiology of, 22, 26f bright-blood imaging in, 21–22, 24f cardiac gating for, 22, 25f goal of, 21–22, 23f, 24f HASTE in, 21–22, 23f pulse sequence in, 21–22, 25f steady-state free precession imaging in, 21–22, 24f Motion artifacts cardiac, 142–145, 146, 147f respiratory, 142–145, 146, 147f navigator echoes for (See Navigator echoes) in pediatric CMR, 120 Motion models, for navigator echoes, 137

MPEs (monophosphate esters), in 31P-CMRS, 557 MPIO (micrometer-sized paramagnetic iron oxide), for cardiac allograft rejection, 549, 549f MPO (myeloperoxidase), contrast agent sensitive to, 85–86 MPR (multiplanar reconstruction), for congenital heart disease, 113, 114f MR (magnetic resonance) signal, detection of, 3–17 MRA. See Magnetic resonance angiography (MRA). MRI. See Magnetic resonance imaging (MRI). MR-Imaging for Myocardial Perfusion Assessment in Coronary Artery Disease Trial (MR-IMPACT), 224–226 MRS (magnetic resonance spectroscopy), cardiovascular. See Cardiovascular magnetic resonance spectroscopy (CMRS). MS-325. See Gadofosveset trisodium (MS-325, Vasovist, Ablavar). MSCT (multislice computed tomography), for right ventricular assessment, 382 MSR-A (Macrophage Scavenger Receptor-A), for atherosclerotic plaques of aorta and carotid arteries, 346 MTT (mean transit time), in stress myocardial perfusion imaging, 221 Multicontrast CMR, of atherosclerotic plaques of aorta and carotid artery, 342–344, 342f, 343f with automatic segmentation, 342, 344f with complications, 343–344 imaging sequences for, 342, 342t with pharmacologic therapy, 343 in subclinical disease, 342–343 validation of, 342 Multidetector computed tomography (MDCT) to assess cardiac function, 182–183 vs. coronary artery CMR, 304–305, 305f, 305t vs. parallel MRI, 42–45, 44f of pericardial disease, 488 Multi-echo imaging, 16, 16f Multiethnic Study of Atherosclerosis (MESA), aortic atherosclerosis in, 342–343 MultiHance. See Gadobenate dimeglumine (GdBOPTA, MultiHance). Multiplanar reconstruction (MPR), for congenital heart disease, 113, 114f Multiple breath hold methods, 129–130 mean diaphragm displacement in, 130f respiratory trace data for, 130f Multiple column orientations, for navigator echoes, 133–136, 135f Multiple excitations, in pediatric CMR, 120 Multislice computed tomography (MSCT), for right ventricular assessment, 382 Muscular dystrophy, Becker, CMR spectroscopy for, 565 Mustard procedure, 415, 416–417 MVO2. See Myocardial oxygen consumption (MVO2). Mycotic aortic aneurysms, 469–470 Myeloperoxidase (MPO), contrast agent sensitive to, 85–86 Myocardial blood flow (MBF), 57 measurement of, 311, 313f Myocardial blood flow (MBF) reserves, 311–312, 313f Myocardial BOLD MRI in clinical setting, 572–573, 573f emerging techniques for, 574–576, 574f, 575f, 576f, 577f

Myocardial oxygenation assessment (Continued) high field, 576, 577f myocardial perfusion reserve in, 573 without pharmacologic stress, 572–573, 573f in preclinical setting, 571–572, 572f rationale for, 570 vs. SPECT, 572, 573f SSFP-based, 574–576, 574f, 575f, 576f myocardial perfusion reserve in, 569, 570f vasodilators in, 570–571 Myocardial perfusion imaging, 25–31, 57–68 acceleration techniques for, 60–61 applications of, 57 of congenital heart disease, 116, 118f, 120 contrast agent injection in, 66 contrast residue detection in, 66 echo planar imaging for, 58–59 endogenous contrast for, 61–62 first pass imaging with exogenous tracers in, 58–60 goal of, 25–27, 28f gradient echo imaging for single-shot, 58, 59f with steady-state free precession, 59 image acquisition in, 27–28, 29f magnetization recovery in, 28, 29f measurement of arterial input in, 66 vs. other imaging modalities, 57–58, 60f parallel imaging for, 50f, 51, 60–61, 66 physiologic basis for, 57–58 practical aspects of, 66 pulse sequence in, 28–29, 30f quantitative evaluation of, 62–65, 63f, 64f arterial input function in, 65–66, 65f rest, 31 with single ventricle, 124f spatial resolution in, 66 stress (See Stress myocardial perfusion imaging) T1 fast acquisition relaxation mapping for, 59–60, 60f T1-weighted techniques for, 58 T2*-weighted techniques for, 58–59 temporal resolution of measurements in, 66 up-slope method in, 66 water exchange and its effect on contrast enhancement in, 61 Myocardial perfusion reserve, for myocardial oxygenation assessment, 569, 570f, 573 Myocardial siderosis, 524 morphology and function in, 524 tissue characterization in, 524 Myocardial stunning, 267 Myocardial tagging for congenital heart disease, 116, 119f high field CMR for, 170–171, 172f for left ventricular diastolic function, 69–75 apical rotation in, 71f, 73–74, 73f, 73t cardiac motion and, 69–70 complementary spatial modulation of magnetization for, 70–71 echo planar imaging in, 70–71 grid-tagged images in, 70, 71f, 72f limitations of, 74 slice thickness in, 70 slice-following principle in, 70, 70f spiral imaging in, 70–71 steady-state free precession in, 70–71 with strain measurements, 74 3D, 71 evaluation of motion data from, 71–72, 73f, 73t limitations of, 74 methods for, 70–72

Myocardial tagging (Continued) during remodeling, 255–256 results of, 72–74 strain measurement in, 74 for left ventricular systolic function, 149 during remodeling, 255–256 for right ventricular assessment, 388 Myocardial viability, 267–283 in acute myocardial infarction, 269–274 contrast-enhanced studies for, 269–274 late gadolinium enhancement for, 269–271 adjustment of T1 in, 271, 273f ischemic bed at risk in, 271, 272f nulling of signal intensity of normal myocardium in, 269, 270f and recovery of function, 273–274 T1-weighted segmented inversion recovery pulse sequence for, 269, 269f time between contrast injection and imaging in, 269–271, 271f no-reflow phenomenon and, 271–272, 274f T2-weighted images for, 269 wall thickness and, 274 assessment of, 31 cardiac gating for, 31, 32f CMR spectroscopy for, 564–565, 564f goal of, 31, 31f inversion recovery in, 31, 32f parallel imaging for, 50f, 51 in chronic myocardial infarction, 274–280 contractile reserve and, 275–278 late gadolinium enhancement for, 276, 277f, 278f, 279f vs. dobutamine CMR, 277–278 vs. other imaging modalities, 276–277, 281f thickness of epicardial rim and recovery of function in, 276, 280f wall thickness and, 275, 275f CMR spectroscopy for, 278–280 contractile reserve and, 267–268 defined, 267 feature(s) of, 267–269 contractile reserve as, 267–268 early hypoenhancement with gadolinium as, 268 late gadolinium enhancement in infarcted tissue as, 268 left ventricular wall thickness as, 267 no-reflow phenomenon as, 268 scar formation as, 267 tissue edema as, 268 high-energy phosphates and, 268–269 sodium and potassium CMR for, 268–269 Myocardial wall thickness, and myocardial viability, in chronic myocardial infarction, 275, 275f Myocarditis, 517–518, 517f combined protocols for, 517f, 518 early enhancement in, 517f, 518 edema in, 517f, 518 follow-up for, 518 function and morphology in, 517f, 518 late gadolinium enhancement in, 517f, 518 tissue characterization in, 517f, 518 Myocardium hibernating, 267 T1 and T2 values for, 7t targeted local delivery of cellular agents to, 585–587, 586f vulnerable, 351 Myxoma, 533, 534f, 544t

Cardiovascular Magnetic Resonance 633

INDEX

Myocardial BOLD MRI (Continued) future of, 577 high field, 576, 577f myocardial perfusion reserve in, 573 without pharmacologic stress, 572–573, 573f in preclinical setting, 571–572, 572f rationale for, 570 vs. SPECT, 572, 573f SSFP-based, 574–576, 574f, 575f, 576f for stress myocardial perfusion imaging, 215–216 Myocardial contrast enhancement, water exchange and its effects on, 61 Myocardial edema and myocardial viability, 268 in acute myocardial infarction, 269 in myocarditis, 517f, 518 Myocardial function, 141f, 142, 143f Myocardial infarction acute (See Acute myocardial infarction [AMI]) chronic (See Chronic myocardial infarction [CMI]) parallel imaging of, 50f, 51 viability imaging for, 31 cardiac gating for, 31, 32f goal of, 31, 31f inversion recovery in, 31, 32f Myocardial ischemia, 57–58 absolute vs. relative, 230–231 in coronary artery disease, 158–159 during dobutamine stress CMR, 198, 201f contemporary studies of, 202–203, 202t, 203f, 203t early studies of, 201–202, 202f ischemic cascade in, 230 noninvasive imaging of, 230–231 spatiotemporal disparity in, 230 stress testing for, 231–239 contraindications and termination criteria for, 236, 236t, 237t cost of, 238 coverage with, 237 diagnostic performance of, 238–239, 238f drug interactions in, 236 duration of examination with, 237 functional assessment of viable myocardium with, 237–238 image display and analysis for, 237 imaging protocols for, 232, 233f, 234f monitoring during, 236–237 patient evacuation and emergency equipment for, 237 pharmacologic effects of, 231–232, 231t pitfalls and advanced issues with, 237–239 practicability of, 236–237 route and duration of administration in, 238 safety aspects of, 232–236 Myocardial oxygen consumption (MVO2) estimation of, 569 maximal, 229 measurement of, 569 resting, 229 Myocardial oxygenation assessment, 569–579 contrast for, 570 coronary flow and coronary arterial pressure in, 569, 570f historical background of, 569 myocardial BOLD MRI for in clinical setting, 572–573, 573f emerging techniques for, 574–576, 574f, 575f, 576f, 577f future of, 577

INDEX

N 23

Na CMR spectroscopy. See Sodium-23 (23Na) CMR spectroscopy. Navigator(s), in coronary artery CMR with atherosclerotic plaques, 352–353, 353f for gating and slice tracking, 287–288 of native vessel stenosis, 302–303, 303t other forms of, 136–137 for triggering alone, 286–287, 287f for velocity mapping, 316–317, 317f, 318f Navigator acceptance window, 129–130 Navigator echoes, 129–139 accept-reject algorithm for, 131, 131t, 132f column positioning for, 133, 135f, 135t column selection for, 131–132 computer architecture for, 137 correction factors for, 132–133, 133f, 134f diminishing variance algorithm for, 131, 131t, 132f free breathing methods for, 129–131 mean diaphragm displacement in, 130f respiratory trace data for, 130f history of, 129 implementation of, 131–133 k-space ordering for, 131, 131t, 132f more recent approaches to, 136–137 motion models for, 137 multiple breath hold methods for, 129–130 mean diaphragm displacement in, 130f respiratory trace data for, 130f multiple column orientations for, 133–136, 135f navigator timing for, 133–136, 136f other forms of navigators for, 136–137 precision of measurement with, 136 uses for, 129–131, 130f Navigator timing, 133–136, 136f Navigator-based respiratory gating, in pediatric CMR, 117, 119f, 120 Navigator-gated imaging, of postoperative atrial switch, 423, 424f Negative contrast agents, 76 Neovascularization, in atherosclerosis, 341 of aorta and carotid artery, 345 of coronary artery, 358–359 Nephrogenic systemic fibrosis (NSF), gadolinium contrast agents and, 87 Neurostimulation systems, as contraindication to CMR, 105–106 Nicorandil, for ventricular remodeling, 260–261 90 pulse, 3–5 Nitrates effect on coronary artery blood flow of, 285, 286f for ventricular remodeling, 260 Nitric oxide synthase (NOS), inducible, in ventricular remodeling, 260 Nitric oxide synthase (NOS) inhibitor, for ventricular remodeling, 260, 261 N-methyl-L-arginine methyl ester (L-NAME), for ventricular remodeling, 261 Noise pixels, 95 Noise reduction, during CMR, 103 Non-Cartesian paths, 37 Noncompaction cardiomyopathy, 522, 522f characteristics of, 522 function and morphology in, 522 tissue characterization in, 522 Nonselective radiofrequency pulse, 8 No-reflow phenomenon after acute myocardial infarction, 257–258 and myocardial viability, 271–272, 274f defined, 268 and myocardial viability, 268

634 Cardiovascular Magnetic Resonance

No-reflow phenomenon (Continued) after acute myocardial infarction, 271–272, 274f NOS (nitric oxide synthase), inducible, in ventricular remodeling, 260 NOS (nitric oxide synthase) inhibitor, for ventricular remodeling, 260, 261 NSF (nephrogenic systemic fibrosis), gadolinium contrast agents and, 87 Nuclear cardiology, to assess cardiac function, 182 Nuclear spins, 3, 4f Nulling, of signal intensity of normal myocardium, in acute myocardial infarction, 269, 270f

O

Obesity, CMR spectroscopy with, 560–561 Oblique sagittal planes, uses for, 140–142, 143f Off-resonance effects, in spiral imaging, 38, 39f Off-resonance spins, 6 Omniscan (gadodiamide), 79f, 79t relaxivity with, 83t safety of, 87–88 One-dimensional chemical shift imaging (1DCSI), for cardiac allograft rejection, 551–552 Optical pumping, with contrast agents, 87 Optimark (gadoversetamide), 79f, 79t safety of, 87–88 Ostium primum defect, 398, 398f Ostium secundum defect, 398, 398f, 399f Outer-sphere relaxivity, 81–82 Oxygen consumption, myocardial maximal, 229 resting, 229 Oxygen-sensitive myocardial imaging, 569–579 contrast for, 570 coronary flow and coronary arterial pressure in, 569, 570f historical background of, 569 myocardial BOLD MRI for in clinical setting, 572–573, 573f emerging techniques for, 574–576, 574f, 575f, 576f, 577f future of, 577 high field, 576, 577f myocardial perfusion reserve in, 573 without pharmacologic stress, 572–573, 573f in preclinical setting, 571–572, 572f rationale for, 570 vs. SPECT, 572, 573f SSFP-based, 574–576, 574f, 575f, 576f myocardial perfusion reserve in, 569, 570f vasodilators in, 570–571 Oxyhemoglobin, in BOLD technique, 62

P 31

P CMR spectroscopy. See Phosphorus-31 (31P) CMR spectroscopy. P947, for atherosclerotic plaques of aorta and carotid arteries, 346 PA. See Pulmonary artery(ies) (PA). Pacemakers, safety of CMR with, 101, 102, 107–108, 108f Papillary fibroelastoma, 534, 536f, 544t Papillary muscles, in left ventricular mass, 188 Parallel imaging, 42–53 applications of, 49–53 for assessment of global and regional cardiac function, 47f, 49–51, 185 for coronary artery, 50f, 51–53

Parallel imaging (Continued) for detection of myocardial infarction and assessment of myocardial viability, 50f, 51 for imaging of cardiac anatomy and structure, 49, 50f for myocardial perfusion imaging, 50f, 51, 60–61, 66 for ventricular function, 149 artifacts in, 48–49, 48f coil arrays for, 46 coil sensitivity calibration strategies in, 46, 47f for contrast-enhanced MRA, 466 data acquisition and image reconstruction in, 45–46, 45f dynamic, 47–48 high field, 175 multi-detector-row CT vs., 42–45, 44f principles of, 42–49 signal-to-noise ratio in, 46–47, 47f undersampling in, 46 Parallel imaging with augmented radius in k-space (PARS), 45f, 46 Parasagittal planes, uses for, 140–142, 143f Partial saturation, in CMR spectroscopy, 556–557 Passive catheter tracking and visualization, 583–584, 595–596 chemical-selective, 584, 584f CO2-filled balloon in, 595, 596f, 597f dysprosium oxide, 595, 596f ideal material for, 595 that create bright signals, 583–584, 584f that create dark signals, 583 Patent ductus arteriosus (PDA), 400 CMR evaluation of, 400 dark-blood imaging of, 115f interventional closure of, 400 pathogenesis of, 400 pulmonary arteries in, 485, 485f Patient monitoring during CMR, 104–105, 105f in interventional CMR laboratory, 580–581 PC. See Phase contrast (PC). PCI (percutaneous coronary intervention) CMR-guided, 605 for ventricular remodeling, 262–263 PCr (phosphocreatine), in 31P-CMRS, 557–558, 558f PCr/ATP (phosphocreatine/adenosine triphosphate) ratio in 31P-CMRS, 557, 558, 559 in cardiac allograft rejection, 551–553, 553f PDA. See Patent ductus arteriosus (PDA). PDE (phosphodiesters), in 31P-CMRS, 559, 561f Peak filling rate (PFR), normal values for, 385–388, 388t in females, 385–388, 387t, 390f in males, 385–388, 386t, 387t, 389f Peak flow velocity, 189–190 Pediatric CMR, 111–128, 395–396 adult vs., 111, 395–396 advantages of, 111 anatomic imaging in, 112–113 double inversion recovery dark-blood imaging for, 113, 114f, 115f gadolinium-based, 116–118 of arrhythmogenic right ventricular cardiomyopathy, 118 of coronary artery, 117, 119f late gadolinium enhancement imaging as, 117, 118f myocardial and blood tagging in, 116, 119f

Pediatric CMR (Continued) spatial and temporal resolution as, 118–120 velocity mapping in, 113–115 Percutaneous coronary intervention (PCI) CMR-guided, 605 for ventricular remodeling, 262–263 Perfluorocarbon-based contrast agents, 85, 86f Perfusion defect, after acute myocardial infarction prognostic significance of, 246 and regional recovery of function, 246–247 with residual coronary occlusion vs. microvascular obstruction, 245–246, 245f, 246f Perfusion imaging. See Myocardial perfusion imaging. Perfusion status, after acute myocardial infarction prognostic significance of, 246 and regional recovery of function, 246–247 with residual coronary occlusion vs. microvascular obstruction, 245–246, 245f, 246f Perfusion-related parameters, in stress myocardial perfusion imaging, 220, 221f Pericardial agenesis, 490 Pericardial cavity, 489 Pericardial cysts, 490–491, 491f, 543, 543f, 544t Pericardial defects, 490 Pericardial effusions, 491–493, 493f hemorrhagic, 495 malignant, 495 Pericardial fluid, 489 Pericardial lesions, 542–543 contrast agents for, 532 cysts as, 490–491, 491f, 543, 543f, 544t metastases as, 542–543 technical considerations with, 532, 533t tumors as, 542 Pericardial metastasis, 542–543 Pericardial recess, superior, 142, 146f Pericardial sinuses, 489 Pericardial thickness, 490 Pericardial tubes, 489 Pericardial tumors, 495, 542 primary, 495 secondary malignant, 495 Pericarditis, 491, 492f constrictive, 493–495 chest x-ray of, 493 clinical presentation of, 493 CMR of, 494, 494f CT of, 494, 494f effusive-, 493 etiology of, 493 pericardial thickening in, 493 vs. restrictive cardiomyopathy, 493 transthoracic echocardiography of, 493 ventricular filling pattern in, 495 Pericardium, 488–498 in cardiac homeostasis, 489 fibrous, 489 imaging modalities for, 488–489 normal anatomy of, 489–490, 489f aortic recesses in, 489–490, 490f serous, 489 Peripheral artery disease, interventional CMR for, 589, 589f Peripheral bypass graft patency, 474 Peripheral vascular disease, 473–475 diagnosis of Doppler ultrasonography for, 473 noninvasive techniques for, 473 3D contrast-enhanced MRA for, 474, 475f

Peripheral vascular disease (Continued) for assessment of bypass graft patency, 474 bolus chase technique in, 474, 475f of hand and wrist, 475 sensitivities and specificities with, 474 time-resolved technique in, 474 2D time-of-flight MRA for, 474 X-ray angiography for, 473–474 etiology of, 473 interventional CMR for, 589, 589f risk factors for, 473 Peripheral vessels, MRA of, 473–475, 475f PET. See Positron emission tomography (PET). PFR (peak filling rate), normal values for, 385–388, 388t in females, 385–388, 387t, 390f in males, 385–388, 386t, 387t, 389f Phase of signal, and flow velocity, 92 of spin, 5–6, 5f Phase contrast (PC) MRA, 94, 464 of aorta, 468 of renal artery stenosis, 472 Phase contrast (PC) velocity mapping, 91, 92–93 flow vector map in, 96, 97f flow velocity images in, 92–93, 93f and Fourier velocity imaging, 94 improving accuracy of, 93–97, 94f, 95f rapid, 95–96 of thoracic aorta, 450–451, 451f validation of, 95 for valvular heart disease, 501, 502f, 506 Phase diagram, 5–6, 5f Phase encoding, 7, 7f, 10–11, 10f pulse sequence diagram with, 11–12, 11f Phase encoding gradients, 7–8 Phase flow imaging methods, 91–93 Fourier flow imaging as, 91, 93 velocity phase encoding in, 91, 93 visualizing flow data in, 96, 96f phase contrast velocity mapping as, 91, 92–93 flow vector map in, 96, 97f flow velocity images in, 92–93, 93f and Fourier velocity imaging, 94 improving accuracy of, 93–97, 94f, 95f rapid, 95–96 validation of, 95 principles of, 91–93, 92f rapid, 95–96, 96f Phase ordering, for navigator echoes, 131, 131t, 132f Phase reconstruction, 92 Phase sensitive reconstruction of inversion recovery (PSIR), in parallel imaging, 50f, 51 Phase unwrapping, 92, 93–94, 94f Phase velocity encoding, principles of, 91–93, 92f Phase velocity mapping, for coronary artery velocity, 315–317 breath hold technique for, 315–316 navigator techniques for, 316–317, 317f, 318f Phased-array coils, in stress myocardial perfusion imaging, 220 Philips, CMR terminology used by, 611–611 Phosphates, high-energy, and myocardial viability, 268 Phosphocreatine (PCr), in 31P-CMRS, 557–558, 558f Phosphocreatine/adenosine triphosphate (PCr/ATP) ratio in 31P-CMRS, 557, 558, 559 in cardiac allograft rejection, 551–553, 553f

Cardiovascular Magnetic Resonance 635

INDEX

Pediatric CMR (Continued) myocardial perfusion imaging as, 116, 118f, 120 saturation band in, 116, 119f static, 116, 117f technical considerations with, 120 time-resolved, 116, 118f of tumor or mass, 117 multiplanar reconstruction for, 113, 114f steady-state free precession for, 112–113, 112f anesthesia for, 395–396 cine, 113, 115f, 116f for coarctation of the aorta, 115f, 124–125, 126f of congenital heart disease, 420–438 double-outlet right ventricle as, 425–426 for postoperative assessment, 426 for preoperative assessment, 426, 426f interrupted aortic arch as, 428–430 classification of, 428, 429f for evaluation, 429 for postoperative assessment, 430, 430f for preoperative assessment, 429, 429f postoperative atrial switch, 423–425 contrast-enhanced, 425f ECG-gated SSFP, 423, 423f navigator-gated, 423, 424f single ventricle as, 122–124, 430–435 aortic arch imaging for, 123–124, 123f, 124f cine CMR for, 124, 124f for evaluation, 432 with Fontan baffle, 123, 123f, 125f Fontan procedure for, 432, 432f late gadolinium enhancement for, 124f left, 430, 431f perfusion imaging for, 124f post-Fontan, 434–435, 434f pulmonary artery imaging for, 123, 123f, 125f right, 430, 431f during staged palliation, 433, 433f, 434f time-resolved gadolinium imaging for, 118f velocity mapping for, 123–124 venous pathway imaging for, 123f, 124 with ventricular outflow obstruction, 125f tetralogy of Fallot as, 420–422 for evaluation, 421 for postoperative assessment, 421–422, 422f for preoperative assessment, 421, 421f transposition of the great arteries as, 120–122, 422–423 cine SSFP as, 120–122, 121f dark-blood CMR as, 120–122, 121f three-dimensional contrast-enhanced MRA as, 117f, 120–122, 122f truncus arteriosus as, 426–428 classification of, 426, 427f contrast-enhanced, 428f for postoperative assessment, 427–428 for preoperative assessment, 427, 428f duration of, 395 functional fetal, 125–127, 126f future of, 125–127, 126f general protocol for, 112–118, 112f interventional, 593–610 reduced ionizing radiation in, 594 limitations and challenges of, 111 technical consideration(s) in, 118–120 with gadolinium-based techniques, 120 inability to hold breath as, 120

INDEX

Phosphodiesters (PDE), in 31P-CMRS, 559, 561f Phosphorus-31 (31P) CMR spectroscopy, 556, 557t for cardiac allograft rejection, 551–553, 552f for cardiomyopathy, 516 dilated, 517 clinical studies of, 559–565 in athlete’s heart and hypertension, 560–561 in diabetes and obesity, 561 in healthy subjects, 560, 561f with heart failure and cardiac transplantation, 561–563, 562f in ischemic heart disease, 563–565 for myocardial viability assessment, 564 for stress testing, 563–564, 563f methodologic considerations in, 559–560, 560f, 561f with specific gene defects with cardiac pathology, 565 in valvular heart disease, 563 experimental foundations of, 557–558, 558f during left ventricular remodeling, 257 and myocardial viability, 268, 278–279 physical principles of, 556–557, 557f Pi (inorganic phosphate), in 31P-CMRS, 557, 558, 558f Plaque inflammation, in atherosclerotic plaques of aorta and carotid artery, 346–347, 347f Plaque rupture, coronary stenosis and, 213, 214f Poly-lysine-Gd compounds, for stress myocardial perfusion imaging, 217 Positive contrast agents, 76 Positive remodeling, in atherosclerosis, 341 Positron emission tomography (PET) to assess cardiac function, 182 for cardiac allograft rejection, 554t of myocardial infarction acute, 243–244 chronic vs. late gadolinium enhancement, 276–277, 281f and myocardial wall thickness, 275, 275f and viable epicardial rim thickness, 276, 280f Postoperative atrial switch, 423–425 contrast-enhanced CMR of, 425f ECG-gated SSFP imaging of, 423, 423f navigator-gated imaging of, 423, 424f Post-pericardiotomy syndrome, 491–492 Postprocessing adaptive motion correction technique, 133 Potassium CMR, for myocardial viability, 268–269 PR (projection-reconstruction) imaging, 40, 40f undersampled, 41–42, 41f Precession, steady-state free. See Steady-state free precession (SSFP). Precessional frequency, 3 Primovist (gadoxetic acid), 85 ProHance (gadoteridol), 79f, 79t relaxivity with, 83t safety of, 87 Projectile effect, with ferromagnetic objects, 101 Projection-reconstruction (PR) imaging, 40, 40f undersampled, 41–42, 41f Prosthetic heart valves, 512–513 safety of CMR with, 101, 101f, 106 Proton CMR spectroscopy (1H-CMRS), 556, 557t for cardiomyopathy, 516 dilated, 517

636 Cardiovascular Magnetic Resonance

Proton CMR spectroscopy (1H-CMRS) (Continued) experimental studies with, 559 during left ventricular remodeling, 257 for myocardial viability assessment, 279–280, 564 P-selectin, in atherosclerosis of coronary artery, 358 Pseudoaneurysm of thoracic aorta, 456, 469 ventricular, due to acute myocardial infarction, 249, 249f PSIR (phase sensitive reconstruction of inversion recovery), in parallel imaging, 50f, 51 Psychological effects, of CMR, 103 Pulmonary arteriovenous malformations, 486 Pulmonary artery(ies) (PA), 480–487 in arterial switch procedure, 120–121, 121f CE-MRA of, 481, 481f congenital anomalies of, 485–486, 485f, 486f morphology of, 409 pulse wave velocity of, 370 in tetralogy of Fallot, 420 in truncus arteriosus, 426, 427f, 428f Pulmonary artery (PA) compliance, 362 Pulmonary artery (PA) distensibility, 369, 370f, 371f Pulmonary artery (PA) flow, 152 Pulmonary artery (PA) hypertension, 481–485 CE-MRA for, 482–483, 482f, 483f clinical features of, 481–482 CMR perfusion imaging for, 483–484, 483f CMR vs. other diagnostic techniques for, 481–482, 482t defined, 481–482 early diagnosis of, 481–482 echocardiography for, 482t etiology of, 482 late gadolinium enhancement for, 484–485, 485f with pulmonary embolism, 482f right heart catheterization for, 482t right ventricular assessment in, 391–392 in tetralogy of Fallot, 483f thromboembolic vs. nonthromboembolic, 482 velocity-encoded CMR for, 484–485, 484f Pulmonary artery (PA) imaging, with single ventricle, 123, 123f, 125f Pulmonary artery (PA) stenosis, 410f after repair of transposition of the great vessels, 485, 486f in tetralogy of Fallot, 413 Pulmonary atresia repair of, 416 tetralogy of Fallot with, 420–421, 421f Pulmonary embolism, 480–481 catheter-based X-ray pulmonary angiography of, 480 CE-CTA for, 480 CMR lung perfusion imaging of, 481, 481f contrast-enhanced MRA of, 481, 481f, 482f current workup for, 480 incidence of, 480 pulmonary artery hypertension with, 482f ventilation/perfusion scanning for, 480 Pulmonary regurgitation, 405f, 509–510 etiology of, 509–510 free, 509–510, 511f measurement of, 509–510, 511f severity of, 502t Pulmonary valve, bicuspid, 404, 404f Pulmonary valve implantation, interventional CMR for, 587

Pulmonary valve stenosis, 502t, 508 Pulmonary vascular resistance (PVR), CMR measurement of, 595, 603 Pulmonary vein(s) anatomy of normal, 442–443, 442f variant, 443, 443f, 444f congenital anomalies of, 443–445, 444f embryology of, 441–442 and pathophysiology of atrial fibrillation, 445 quantification of size of, 442f, 446–447, 446f Pulmonary vein atresia, congenital, 444 Pulmonary vein imaging, 439–449 before and after atrial fibrillation ablation, 445–446, 445f, 446f of congenital anomalies, 443–445, 444f image display for, 441, 442f imaging method for, 441 late gadolinium enhancement, 447, 447f of normal anatomy, 442–443, 442f for quantification of size, 442f, 446–447, 446f of variant anatomy, 443, 443f, 444f Pulmonary vein stenosis after atrial fibrillation ablation, 446, 446f congenital, 444 Pulmonary venous return, partial anomalous, 444f, 445 Pulmonary-to-systemic flow ratio (Qp/Qs) with atrial septal defect, 399, 399f with single ventricle, 123–124 with ventricular septal defect, 113, 397, 398f Pulse sequence(s), 12–13, 12f for CMR spectroscopy, 559, 560f for three-dimensional MRA, 464 for vascular angiography, 35 Pulse sequence diagram for fast spin echo imaging, 13, 13f with frequency encoding, 10, 10f for gradient recalled echo imaging, 14, 14f for morphology scanning, 21–22, 25f for myocardial perfusion scanning, 28–29, 30f with phase encoding, 11–12, 11f for scout scanning, 20, 22f with slice selection, 8, 9f for velocity-encoded CMR imaging, 33–34, 33f Pulse wave velocity (PWV) age-related increase in, 368 assessment of, 365f, 366 of pulmonary artery, 370 PVR (pulmonary vascular resistance), CMR measurement of, 595, 603

Q

Qp/Qs (pulmonary-to-systemic flow ratio) with atrial septal defect, 399, 399f with single ventricle, 123–124 with ventricular septal defect, 113, 397, 398f Quantitative coronary angiography (QCA), vs. stress myocardial perfusion imaging, 222, 223t, 224 Quantitative evaluation, of myocardial perfusion, 62–65, 63f, 64f

R

RA (right atrium), morphology of, 408 RA (right atrial) ridge, 142, 146f RACE (real-time acquisition and velocity evaluation), 96 Radial imaging, 16–17, 16f, 40–42 applications of, 42, 43f of coronary artery, 292

Renal arteries, magnetic resonance angiography of, 470–472 Renal artery stenosis, 470–472 diagnosis of, 471 captopril test for, 471 Doppler ultrasonography for, 471 MRA for, 472 contrast-enhanced, 472 phase contrast, 472 radionuclide renography for, 471 X-ray angiography for, 471 epidemiology of, 471 etiology of, 470 renovascular hypertension due to, 470 treatment of, 471 Renography captopril, 471 radionuclide, 471 Renovascular hypertension, 470, 471 Reperfusion, after acute myocardial infarction prognostic significance of, 246 and regional recovery of function, 246–247 with residual coronary occlusion vs. microvascular obstruction, 245–246, 245f, 246f Repetition, 11 Repetition time (TR), 11, 11f, 12 effect on signal of, 12–13, 12f for myocardial perfusion imaging, 58 in pediatric CMR, 120 Rephasing, 6–7, 6f Rephasing gradient, 8–9, 9f Resistance vessels, in coronary flow reserve, 230 Resovist (ferucarbotran), relaxivity with, 83 Respiratory drift, 130 Respiratory efficiency, with navigator echoes, 129–130 Respiratory feedback monitor, 130 Respiratory gating for coronary artery bypass graft, 331 retrospective, 130–131, 131t, 132f for coronary artery velocity mapping, 317, 318f Respiratory motion artifacts, 142–145, 146, 147f in coronary artery CMR, 285–286, 286t with atherosclerotic plaques, 352–353, 353f navigator echoes for (See Navigator echoes) in pediatric CMR, 120 Rest function, in comprehensive CMR assessment of coronary artery disease, 160t Rest myocardial perfusion imaging, 31 Rest perfusion imaging, of coronary artery disease for disease detection, 160t, 161 for viability studies, 166f Restrictive cardiomyopathy (RCM), 521–522 vs. constrictive pericarditis, 493 morphology and function in, 522 tissue characterization in, 522 Retrospective respiratory gating (RRG), 130–131, 131t, 132f for coronary artery velocity mapping, 317, 318f Revascularization, predicting myocardial response to, 276, 277f, 278f, 279f RF. See Radiofrequency (RF). Rhabdomyoma, 536, 538f, 544t Right atrial (RA) ridge, 142, 146f Right atrium (RA), morphology of, 408 Right heart catheterization, for pulmonary artery hypertension, 482t Right ventricle (RV) in congenital heart disease, 417 double-chambered, 508 jet flow in, 504–505, 505f

Right ventricle (RV) (Continued) double-outlet, 410f, 412 defined, 425 epidemiology of, 425 features of, 425 in infant and pediatric patients, 425–426 postoperative assessment of, 426 preoperative assessment of, 426, 426f surgical management of, 425 late complications after, 425–426 ventricular septal defect in, 412, 425 morphology of, 409 single, 430, 431f Right ventricular (RV) anatomy assessment of (See Right ventricular [RV] assessment) normal, 381–385 during remodeling, 253–254, 255 Right ventricular (RV) assessment advantages of CMR for, 382 in arrhythmogenic RV cardiomyopathy, 391 in congenital heart disease, 391 in heart failure, 388–391 imaging strategies for, 382–385 ECG gating in, 382–383 gradient echo (white-blood) cine imaging as, 383, 384f myocardial tagging as, 388 other techniques in, 385 Simpson’s rule in, 383–385, 386f spin echo (black-blood) sequences in, 383, 383f transaxial plane in, 383, 384f importance of, 381 in ischemic heart disease, 391 normal values in, 385–388 for RV diastolic function and atrioventricular plane descent, 385–388 for females, 385–388, 387t, 390f for males, 385–388, 387t, 389f for RV volumes, systolic function, and mass, 385–388 for females, 385–388, 387t, 390f for males, 385–388, 386t, 389f summary data of, 385–388, 388t in pulmonary hypertension and lung transplantation, 391–392 techniques for, 382 Right ventricular (RV) dimensions, assessment of. See Right ventricular (RV) assessment. Right ventricular ejection fraction (RVEF) after acute myocardial infarction, 255 normal values for, 385–388, 388t in females, 385–388, 387t, 390f in males, 385–388, 386t, 389f Right ventricular (RV) failure, causes of, 381 Right ventricular (RV) function, 150, 185 assessment of (See Right ventricular [RV] assessment) prognostic value of, 381 Right ventricular (RV) infarction, 391 Right ventricular (RV) mass normal values for, 385–388, 388t in females, 385–388, 387t, 390f in males, 385–388, 386t, 389f reference ranges for, 190t validation of, 186f Right ventricular outflow tract (RVOT) aneurysm, in tetralogy of Fallot, 421–422, 422f Right ventricular outflow tract (RVOT) obstruction, in tetralogy of Fallot, 420–421 Right ventricular outflow tract (RVOT) stenosis, 508

Cardiovascular Magnetic Resonance 637

INDEX

Radial imaging (Continued) principles of, 39f, 40–42, 40f undersampling in, 41–42, 41f Radiofrequency (RF) ablation, CMR guidance for, 603, 604, 604f Radiofrequency (RF) excitation, 3–5, 4f Radiofrequency (RF) magnetic fields, safety of, 102–103 Radiofrequency (RF) pulse, nonselective, 8 Radiofrequency (RF) radiation, bioeffects of, 598 Radionuclide angiography, for right ventricular assessment, 382 Radionuclide renography, 471 Radionuclide ventriculography, to assess cardiac function, 182 Rapid phase flow imaging methods, 95–96, 96f Rapid volumetric imaging, 52f, 54 RCM. See Restrictive cardiomyopathy (RCM). Readout gradient, 9–10, 9f Real-time acquisition and velocity evaluation (RACE), 96 Real-time CMR imaging, 581t, 583 Real-time prospective slice following, 132–133, 134f Receiver coil, 3 Reflected waves, 367–368 Refocusing, 6–7, 6f Refocusing echo, 7, 7f Regional function, assessment of, 190–192 Regional left ventricular function, during remodeling, 255–256 Regurgitant valvular heart disease, 509–512 aortic, 502t, 509, 510f classification of severity of, 502t flow measurements for, 506–507 general principles for, 509 mitral, 502t, 510–512, 511f, 512f pulmonary, 502t, 509–510, 511f quantification of regurgitation volume in, 403f, 405f for multiple valves, 509 for single valve, 509 surgical intervention for, 509 tricuspid, 502t, 512, 513f Relaxation, 5 Relaxation rate, with contrast agents, 76–77 Relaxivity, of contrast agents, 81–83 correlation time in, 82 effect and definition of, 76, 77f effect of correlation time and field strength on, 82, 83, 83f electronic relaxation in, 82 inner- and outer-sphere, 81–82 iron oxide–based, 83 longitudinal and transverse, 82, 83f magnetic field dependence on, 82, 82f magnetic moment in, 82 molecule size and, 82 for selected media, 83, 83t Remodeling, after acute myocardial infarction, 253–266 CMR spectroscopy of energetics during, 256–257, 257f contrast-enhanced CMR and predictors of, 257–259, 258f, 259f early phase of, 253 late phase of, 253, 254f pathophysiology of, 253 regional left ventricular function during, 255–256 therapy for animal studies of, 259–262, 261f human studies of, 262–264, 263f ventricular anatomy during, 253–255, 254f

INDEX

Right ventricular (RV) stroke volume after acute myocardial infarction, 255 measurement of, 189–190 validation of, 186f Right ventricular (RV) systolic function, reference ranges for, 190t Right ventricular (RV) volume measurement of, 150–151, 152t normal values for, 190t, 385–388 in females, 385–388, 387t, 390f in males, 385–388, 386t, 389f Right ventricular (RV) wall motion assessment, cine CMR for, 22–25 acquisition time in, 24 cardiac gating for, 24–25, 27f goal of, 23–24, 27f Right-sided isomerism, 409 R-R interval, in pediatric CMR, 120 RRG (retrospective respiratory gating), 130–131, 131t, 132f for coronary artery velocity mapping, 317, 318f RV. See Right ventricle (RV). RVEF. See Right ventricular ejection fraction (RVEF). RVOT (right ventricular outflow tract) aneurysm, in tetralogy of Fallot, 421–422, 422f RVOT (right ventricular outflow tract) obstruction, in tetralogy of Fallot, 420–421 RVOT (right ventricular outflow tract) stenosis, 508

S

Safety consideration(s), 100–103 auditory, 103 with biologic effects, 100 in interventional CMR, 598 with contrast agents, 87–88, 103–104 with ferromagnetism, 101, 101f, 102f general, 100 with interventional CMR, 582–584, 583t bioeffects of magnetic fields as, 598 heating and electrical safety as, 465, 583t, 598–599 magnetic force and torque as, 600 with pacemakers and implantable cardioverter defibrillators, 101, 102, 107–108, 108f with psychological effects, 103 with radiofrequency time varying field, 102–103 with rapidly switched magnetic fields, 101–102 during stress conditions, 104 with superconducting system, 103 Sagittal plane scout image in, 140, 141f uses for, 140–142, 143f Sagittal scout image, 20, 21f Saphenous vein graft (SVG), 306, 307f, 307t, 329 Sarcoidosis, myocardial involvement in, 523 morphology and function in, 523 tissue characterization in, 523 Sarcoma, 537–540 angio-, 537, 540f, 544t leiomyo-, 537, 544t lipo-, 540, 541f, 544t Saturation, in time-of-flight methods, 91 Saturation band, for congenital heart disease, 116, 119f Saturation factors, in CMR spectroscopy, 556–557

638 Cardiovascular Magnetic Resonance

Saturation pulse, in myocardial perfusion scanning, 28, 29f Scan efficiency, with navigator echoes, 129–130 Scar formation after acute myocardial infarction, 253 and myocardial viability, 267 Scimitar syndrome, 444f, 445 Scout images, for cardiac anatomy, 140, 141f Scout scanning, 19–21 cardiac gating for, 21, 23f goal of, 19–20 image acquisition in, 20, 21f k-space filling in, 20–21, 22f pulse sequence in, 20, 22f Screening form, for CMR, 612–613 Segmentation, automatic, for atherosclerotic plaques of aorta and carotid arteries, 342, 344f Selectins, in atherosclerosis of aorta and carotid artery, 345 of coronary artery, 358 Selective excitation, 8–9, 8f, 9f Semi-active catheter tracking, 597–598 Senning procedure, 415, 416–417 Sensitivity encoding (SENSE) applications of, 50f, 51 high field, 175 k-t applications of, 51 to assess cardiac function, 185 for stress myocardial perfusion imaging, 218 for myocardial perfusion imaging, 60 stress, 218 principle of, 45f, 46 for stress tests, 199t time-adaptive applications of, 51 for stress myocardial perfusion imaging, 218 Sensitivity profiles from an array of coils for encoding and reconstruction in parallel space (SPACE RIP), 46 Septal leaflets, 409 Sequence protocol dataform, for CMR, 614–616 Sequence type, terminology used by various vendors for, 611–611 Shim gradients, in CMR spectroscopy, 556–557 Short axis slices, to assess cardiac function, 183–184, 184f, 188 Short T1 inversion recovery techniques, for cardiomyopathy, 516 SHU555C, relaxivity with, 83 Shunt quantification, for ventricular septal defect, 397, 398f Siderosis, myocardial, 524 morphology and function in, 524 tissue characterization in, 524 Siemens, CMR terminology used by, 611 Signal intensity, effect of contrast agents on, 77, 78f Signal intensity–contrast media concentration relationship, 221 Signal intensity–time curves, in stress myocardial perfusion imaging, 220, 221f Signal loss, flow-related, 94, 95f Signal misregistration, in time-of-flight methods, 91 Signal-to-noise ratio (SNR) in CMR spectroscopy, 556–557 with high field CMR, 170 in parallel imaging, 46–47, 47f in pediatric CMR, 118–120 in stress myocardial perfusion imaging, 218

Simpson’s rule method, 149, 183–184, 184f for right ventricular assessment, 383–385, 386f Simultaneous acquisition of spatial harmonics (SMASH), 45–46, 45f high field, 175 for myocardial perfusion imaging, 60 sinc function, 8, 8f Sinerem (ferumoxtran), 83 Single photon emission computed tomography (SPECT) of acute myocardial infarction, 243–244 to assess cardiac function, 182 for cardiac allograft rejection, 554t vs. dobutamine stress CMR, 201 vs. myocardial BOLD MRI, 572, 573f vs. stress myocardial perfusion imaging, 222, 224–226 Single ventricle, 122–124, 413–415 aortic arch imaging for, 123–124, 123f, 124f cine CMR for, 124, 124f defined, 430 epidemiology of, 430 with Fontan baffle, 123, 123f, 125f in infant and pediatric patients, 430–435 evaluation of, 432 Fontan procedure for, 432, 432f left, 430, 431f post-Fontan, 434–435, 434f right, 430, 431f during staged palliation, 433, 433f, 434f late gadolinium enhancement for, 124f perfusion imaging for, 124f pulmonary artery imaging for, 123, 123f, 125f time-resolved gadolinium imaging for, 118f velocity mapping for, 123–124 venous pathway imaging for, 123f, 124 with ventricular outflow obstruction, 125f Single-ventricle heart disease, 420 Sinus venosus defect, 398, 398f, 399f Skeletal muscle, T1 and T2 values for, 7t Slice following, real-time prospective, 132–133, 134f Slice following principle, in CSPAMM, 70, 70f Slice selection, 7, 7f, 8–9, 8f pulse sequence diagram with, 8, 9f Slice selection gradients, 7–8, 8f in pulse sequence diagram, 8, 9f Slice thickness, 8, 8f to assess cardiac function, 188 in CSPAMM, 70 Slice tracking, 132–133, 134f for coronary artery CMR, 287–288 SLOOP (spectral localization with optimum point spread function), in 31P-CMRS, 560 Slow exchange, in contrast-enhanced tissue relaxation, 84 Small particle iron oxide (SPIO) contrast agents relaxivity with, 83 structure of, 83 uses of, 81 SMASH (simultaneous acquisition of spatial harmonics), 45–46, 45f high field, 175 for myocardial perfusion imaging, 60 SNR. See Signal-to-noise ratio (SNR). Sodium CMR, for myocardial viability, 268–269 Sodium-23 (23Na) CMR spectroscopy, 556, 557t experimental studies with, 559 during left ventricular remodeling, 257 for myocardial viability assessment, 564, 564f SPACE RIP (sensitivity profiles from an array of coils for encoding and reconstruction in parallel space), 46

Steady-state free precession (SSFP) (Continued) in coronary artery stenosis, 301f, 302–303, 302f in CSPAMM, 70–71 in morphology scanning, 21–22, 24f of myocardial function, 141f, 142 for myocardial oxygenation assessment, 574–576, 575f, 576f, 577f for myocardial perfusion imaging, 59 stress, 218 of postoperative atrial switch, 423, 423f for right ventricular assessment, 384f in scout scanning, 20, 22f for stress tests, 198, 199t of valvular heart disease, 502f, 504–505, 505f for ventricular volumes, 150–151, 151t Stem cell transplantation, for ventricular remodeling, 261–262, 263 Stenotic valvular heart disease, 507–508 aortic, 502t, 507 classification of severity of, 502t jet velocity mapping for, 501, 506 mitral and tricuspid, 502t, 508 pulmonary, 502t, 508 of right ventricular outflow tract, 508 subaortic, 508 Stents, safety of CMR with, 106 Sternal wires, safety of CMR with, 101, 102f, 106 Strain measurements, 74 Strain-encoded CMR imaging, for myocardial tagging, 171 Stress agents, 231–239 for comprehensive CMR assessment of coronary artery disease, 159–160 contraindications and termination criteria for, 236, 236t, 237t cost of, 238 coverage with, 237 diagnostic performance of, 238–239, 238f drug interactions with, 236 duration of examination with, 237 functional assessment of viable myocardium with, 237–238 image display and analysis for, 237 imaging protocols for, 232, 233f, 234f monitoring during, 236–237 patient evacuation and emergency equipment for, 237 pharmacologic effects of, 231–232, 231t pitfalls and advanced issues with, 237–239 practicability of, 236–237 route and duration of administration in, 238 safety aspects of, 232–236 Stress CMR, 196–212 adenosine, 198t, 208–209 atropine pharmacokinetics of, 196 safety of, 196–197, 198t dipyridamole, 208–209 dobutamine for acute myocardial infarction, 241 apical and short axis views in, 198, 200f atropine in, 198, 199f, 202–203 cine GRE or SSFP bright-blood images in, 198, 199t delineation of orthogonal left ventricular myocardial segments in, 198, 200f vs. dobutamine stress echocardiography, 202–203, 202t accuracy of, 197 safety of, 196–197, 197t technique of, 197 uses of, 197 for viability studies, 204

Stress CMR (Continued) dopamine infusion protocol for, 198, 199f facilities for, 198, 199f inducible ischemia during, 198, 201f contemporary studies of, 202–203, 202t, 203f, 203t early studies of, 201–202, 202f late gadolinium enhancement for, 205, 207f pharmacokinetics of, 196 for prognosis, 205, 207f vs. radionuclide studies, 204 safety of, 196–197, 197t, 198t sensitivity and specificity of, 203, 203f, 203t technique for, 197–200 3-T, 203 tissue tagging during, 206–208, 208f for viability studies, 204–205, 207f in viability studies, 204–205 contractile reserve in, 204f, 206f end-systolic wall thickening in, 204, 204f intramyocardial segment shortening in, 205 low-dose, 204–205, 237–238 short-axis basal views in, 200f tissue tagging in, 204–205, 207f during exercise, 209–210, 210f on-site medications for, 198t Stress conditions, patient safety during, 104 Stress echocardiography, dobutamine accuracy of, 197 vs. dobutamine stress CMR, 202–203, 202t safety of, 196–197, 197t technique of, 197 uses of, 197 for viability studies, 204 Stress myocardial perfusion imaging, 213–228 after acute myocardial infarction, 247, 247f adenosine for, 29–30 analysis of data from, 219–222, 220f quantitative approach for, 220 absolute tissue perfusion in, 221–222 clinical performance of, 224, 224f perfusion-related parameters in, 220, 221f visual assessment for, 219–220 clinical performance of, 222–226, 223t clinical performance of, 222–226 multicenter studies of, 223t, 224–226, 225f single-center studies of quantitative semiautomatic analysis of, 224, 224f visual interpretation of, 222–223, 223t CMR data readout from, 218 CMR spectroscopy for, 563–564, 563f combined dobutamine wall motion CMR with diagnostic performance of, 238–239, 238f protocol for, 232, 233f, 234f contrast media for endogenous, 215–216 exogenous, 216–218 extravascular, 216–217, 217f hyperpolarized, 217–218 intravascular, 217 of coronary artery disease, 160, 160t, 161 field strength for, 219 high field CMR for, 174 magnetization preparation for, 218–219, 219f options for inducing stress in, 214, 215f, 216f perspectives on, 226 protocol for, 213–214

Cardiovascular Magnetic Resonance 639

INDEX

Spatial encoding, 7–8 Spatial modulation of magnetization (SPAMM), 69–70 for cardiac allograft rejection, 548–549 complementary, 70–71 echo planar imaging in, 70–71 grid-tagged images in, 70, 71f, 72f limitations of, 74 slice thickness in, 70 slice-following principle in, 70, 70f spiral imaging in, 70–71 steady-state free precession in, 70–71 with strain measurements, 74 3D, 71 Spatial resolution in coronary artery CMR, 288, 288f in pediatric CMR, 118–120 Spatiotemporal disparity, in myocardial ischemia, 230 SPECT. See Single photon emission computed tomography (SPECT). Spectral localization with optimum point spread function (SLOOP), in 31P-CMRS, 560 Spectroscopy, CMR. See Cardiovascular magnetic resonance spectroscopy (CMRS). Speed, increased, 37, 52f, 53–54 Spin(s), 3, 4f off-resonance, 6 Spin echo imaging, 6–7, 6f, 7t of acute myocardial infarction, 269–274 of cardiac anatomy, 140, 143f for congenital heart disease, 396, 408 of coronary artery, 12–13 of coronary artery bypass graft, 330, 331f fast (turbo), 6–7, 13 double inversion recovery (black-blood), 13, 13f pulse sequence diagram for, 13, 13f for right ventricular assessment, 383, 383f of thoracic aorta, 450, 451f Spin echo sequence, 11f, 12 Spin exchange, with contrast agents, 87 Spin labeling, for assessment of myocardial perfusion, 62 Spin phase, 5–6, 5f Spin-lattice relaxation, 5 Spin-spin relaxation, 6 SPIO (small particle iron oxide) contrast agents relaxivity with, 83 structure of, 83 uses of, 81 Spiral imaging, 16–17, 16f, 37–40 applications of, 38–40, 39f of coronary artery, 292, 292f, 293f for coronary artery velocity mapping, 317–320, 319f, 320f, 321f in CSPAMM, 70–71 off-resonance effects in, 38, 39f principles of, 37–38, 38f Spoiler gradient, 14, 14f SSFP. See Steady-state free precession (SSFP). Stainless steel implants, safety of, 101 Stanford classification, of aortic dissection, 452–453, 453f Starr-Edwards prosthetic valve, safety of CMR with, 101f, 106 Static magnetic fields, bioeffects of, 598 Steady-state free precession (SSFP) in acute myocardial infarction, 241 balanced, 15, 15f for cardiomyopathy, 515–516 for congenital heart disease, 112–113, 112f, 396 due to transposition of the great arteries, 120–122, 121f

INDEX

Stress myocardial perfusion imaging (Continued) rationale for, 213, 214f specific steps in, 31 stress-only vs. stress-rest, 213–214, 215t subendocardial, 224, 224f timeline for, 30–31, 30f vs. wall motion CMR, 229–240 Stress-induced cardiomyopathy, 525f, 526 function and morphology in, 525f, 526 tissue characterization in, 526 Stroke volume(s) (SV), 150 calculation of, 150 defined, 150 effect of imaging sequence and magnetic field strength on, 150–152, 151t, 152t normal values for, 385–388, 388t in females, 385–388, 387t, 390f in males, 385–388, 386t, 389f for right ventricular assessment, 383–385 validation of, 186f Subaortic stenosis, 508 Subendocardial perfusion data, 224, 224f Subendocardial vulnerability, to ischemia, 244 Subinfundibular stenosis, 505f, 508 Superconducting system, safety issues with, 103 Superior pericardial recess, 142, 146f Superparamagnetism, 83 Suppression of signal, from surrounding tissues, in coronary artery CMR, 288, 288f Susceptibility contrast, 84–85 SV. See Stroke volume(s) (SV). SVG (saphenous vein graft), 306, 307f, 307t, 329 Swan-Ganz catheters, as contraindication to CMR, 106 Systolic function, 149–150 assessment of, 189–190 Systolic strain, 74

T

3-T (3-Tesla) systems. See High field CMR. T1 contrast agents, 76–81, 77f, 78f T1 fast acquisition relaxation mapping (T1-FARM), for myocardial perfusion imaging, 59–60, 60f T1 relaxation, 5 effects of contrast agent on, 61 T1 values, 7, 7t T1-weighted segmented inversion recovery pulse sequence, for acute myocardial infarction, 269, 269f T1-weighted techniques, for myocardial perfusion imaging, 58 T2 contrast agents, 76–81, 77f, 78f T2 preparation prepulses, in coronary artery CMR, 288, 288f T2 relaxation, 5–6 T2 values, 7, 7t T2-weighted CMR of acute myocardial infarction, 247–248, 248f and left ventricular remodeling, 259–260 signal intensity changes on, 269 for cardiomyopathy, 516 T2*, 6–7 T2*-weighted techniques, for myocardial perfusion imaging, 58–59 Tagged time-of-flight methods, 91, 92f Takayasu arteritis, 459, 459f Tako-Tsubo cardiomyopathy, 525f, 526 function and morphology in, 525f, 526 tissue characterization in, 526 Tantalum stents, safety of CMR with, 106

640 Cardiovascular Magnetic Resonance

Targeted 3D coronary artery CMR acquisition sequence for, 291, 291f of atherosclerotic plaques, 353–354, 354f, 355f of native vessel stenosis, 301–302, 301f Target-specific imaging, of atherosclerotic plaques in aorta and carotid artery, 344–347 angiogenesis in, 345 endothelial dysfunction in, 344–345 extracellular matrix in, 345–346 plaque inflammation in, 346–347, 347f thrombus in, 345, 346f in coronary arteries, 356, 356f angiogenesis in, 358–359 inflammation in, 358 thrombosis in, 356–358, 357f, 358f Taussig-Bing anomaly, 412, 425, 426 TE (time to echo), 11f, 12 effect on signal of, 12, 12f for myocardial perfusion imaging, 58 TEE. See Transesophageal echocardiography (TEE). Telemetric ECG, for monitoring during CMR, 105 Temporal resolution in pediatric CMR, 118–120 for ventricular function, 149 Temporary pacing wires, as contraindication to CMR, 106 Terminology, used by various vendors, 611–611 3-Tesla (3-T) systems. See High field CMR. Tetralogy of Fallot (TOF), 408–409 characteristics of, 409 cine CMR of, 115f clinical presentation of, 420 dark-blood imaging of, 115f epidemiology of, 409 etiology of, 409 in infant and pediatric patients, 420–422 evaluation of, 421 postoperative assessment of, 421–422, 422f preoperative assessment of, 421, 421f “pink,” 420 pulmonary artery hypertension in, 483f with pulmonary atresia, 420–421, 421f surgical repair of, 416, 420–421 pulmonary regurgitation after, 509–510, 511f TFE (turbo field echo), 14, 14f TGA. See Transposition of the great arteries (TGA). Thallium-201 uptake, and myocardial viability, 276–277 The Open Artery Trial (TOAT), 263 Thermal effects of CMR imaging, 102 of interventional CMR, 582–583, 583t, 598–599 Thin-slab 3D coronary artery CMR, 291–292 Thoracic aorta, 450–462 aneurysm of, 456–457, 456f, 457f aortitis of, 459, 459f coarctation of, 458–459, 458f dissection of, 452–454, 452f, 453f, 454f flow mapping of, 450–451, 451f gradient echo CMR imaging of, 450–451 interventional CMR imaging of, 459–460 intramural hematoma of, 454–455, 455f MRA of, 451–452, 452f penetrating ulcer of, 455–456, 455f spin echo CMR imaging of, 450, 451f trauma to, 457–459, 457f 3-Tesla (3-T) systems. See High field CMR.

Three-dimensional (3D) contrast-enhanced MRA of coronary artery bypass graft, 332, 335f, 336f of transposition of the great arteries, 117f, 120–122, 122f Three-dimensional (3D) coronary artery CMR, 290–292 acquisition sequence for, 290–292 of native vessel stenosis with navigator gating, 302–303, 303t targeted, 301–302, 301f whole heart, 302f, 303, 303t targeted acquisition sequence for, 291, 291f of atherosclerotic plaques, 353–354, 354f, 355f of native vessel stenosis, 301–302, 301f thin-slab, 291–292 whole heart acquisition sequence for, 291, 291f, 292f of native vessel stenosis, 302f, 303, 303t Three-dimensional dobutamine stress CMR, 208 Thromboembolic pulmonary artery hypertension, 482 Thrombosis, left ventricular, due to acute myocardial infarction, 248, 249f Thrombus in atherosclerotic plaques of aorta and carotid artery, 345, 346f of coronary artery, 356–358, 357f, 358f intracardiac, 540–542, 542f, 544t Through-plane velocity mapping of congenital heart disease, 113–115 of coronary artery, 317, 317f, 318f of valvular heart disease, 506–507 TI (time to inversion), 14f, 15 Time to echo (TE), 11f, 12 effect on signal of, 12, 12f for myocardial perfusion imaging, 58 Time to inversion (TI), 14f, 15 Time to repetition (TR), 11, 11f, 12 effect on signal of, 12–13, 12f for myocardial perfusion imaging, 58 in pediatric CMR, 120 Time-adaptive sensitivity encoding (TSENSE) applications of, 51 for stress myocardial perfusion imaging, 218 Time-of-flight (TOF) methods, 91, 92f for coronary artery velocity measurement, 314–315, 315f for MRA, 463 of extracranial carotid arteries, 467–468, 467f for peripheral vascular disease, 474 Time-resolved gadolinium imaging of congenital heart disease, 116, 118f with single ventricle, 118f Time-resolved imaging of contrast kinetics (TRICKS) technique, for extracranial carotid arteries, 467 Time-resolved technique, for peripheral vascular disease, 474 TIPS (transjugular intrahepatic portosystemic shunt), interventional CMR for, 587 Tissue edema, and myocardial viability, 268 in acute myocardial infarction, 269 Tissue relaxation, contrast-enhanced, 84–85 Tissue tagging, during dobutamine stress CMR studies, 206–208, 208f of viability, 204–205, 207f TMLR (transmyocardial laser revascularization), for ventricular remodeling, 261 TOAT (The Open Artery Trial), 263

TSE imaging. See Turbo spin echo (TSE) imaging. TSENSE (time-adaptive sensitivity encoding) applications of, 51 for stress myocardial perfusion imaging, 218 TTE. See Transthoracic echocardiography (TTE). Tumor(s) cardiac (See Cardiac tumor(s)) pericardial, 495, 542 primary, 495 secondary malignant, 495 Turbo field echo (TFE), 14, 14f Turbo FLASH, 14, 14f Turbo spin echo (TSE) imaging, 6–7, 13 of cardiac and paracardiac masses, 532 double inversion recovery (black-blood), 13, 13f of transposition of the great arteries, 410f Two-chamber view, 140, 141f

U

Uhl anomaly cine CMR of, 115f myocardial and blood tagging for, 119f Ulcer, penetrating aortic, 455–456, 455f Ultra-fast flow imaging techniques, 95–96 Ultrasmall particle iron oxide (USPIO) contrast agents for atherosclerotic plaques, of aorta and carotid arteries, 346 for cardiac allograft rejection, 548–549, 549f relaxivity with, 83 structure of, 83 uses of, 81 Ultrasonography Doppler for monitoring during CMR, 105 of peripheral vascular disease, 473 of renal artery stenosis, 471–472 vs. interventional CMR, 580, 581t Unaliasing by Fourier encoding the overlaps using the temporal dimension (UNFOLD) applications of, 51 to assess cardiac function, 185 for myocardial perfusion imaging, 61 principles of, 47–48 Undersampling in parallel imaging, 46 in radial imaging, 41–42, 41f Untwisting time, 71–72, 73t Untwisting velocity, 71–72, 73t Upslope parameter, in stress myocardial perfusion imaging, 220, 221f USPIO contrast agents. See Ultrasmall particle iron oxide (USPIO) contrast agents.

V

Valve replacement and repair, invasive and interventional CMR for, 587 Valvular anatomy cine CMR of, 113, 116f velocity mapping for, 113 Valvular atresia, 402f, 403 Valvular function, normal, 152–155, 154f Valvular heart disease, 403–404, 499–514 advantages of CMR for, 501 with atresia, 402f, 403 bicuspid valve as, 403 aortic, 403–404, 403f pulmonary, 404, 404f biventricular volume and function in, 501 cine imaging of, 501, 502f

Valvular heart disease (Continued) in infants and children, 113, 116f visualization and planimetry of jets by, 504–505, 505f, 510f classification of severity of, 502t CMR spectroscopy for, 563 FLASH imaging for, 505 Fourier CMR velocity traces for, 507 gradient recalled echo CMR of, 403f, 404f invasive and interventional CMR for, 587 mechanical heart valves for, 512–513 phase contrast velocity mapping for, 501, 502f, 506 pressure and volume overload in, 404 regurgitant, 509–512 aortic, 502t, 509, 510f classification of severity of, 502t flow measurements for, 506–507 general principles for, 509 mitral, 502t, 510–512, 511f, 512f pulmonary, 502t, 509–510, 511f quantification of regurgitation volume in, 403f, 405f for multiple valves, 509 for single valve, 509 surgical intervention for, 509 tricuspid, 502t, 512, 513f after repair, 403, 405f slice thickness and visualization of thin structures in, 503, 504f SSFP imaging of, 502f, 504–505, 505f stenotic, 507–508 aortic, 502t, 507 classification of severity of, 502t jet velocity mapping for, 501, 506 mitral and tricuspid, 502t, 508 pulmonary, 502t, 508 of right ventricular outflow tract, 508 subaortic, 508 transverse spin echo CMR of, 402f weaknesses of CMR for, 501 Valvular insufficiency, 403 Valvular prostheses, safety of CMR with, 101, 101f, 106 Valvular regurgitation, 403 Valvular stenosis, 403 Vascular angiography, 34–35 cardiac gating for, 34–35 goal of, 34, 34f pulse sequence in, 35 timing of image acquisition in, 34, 35f Vascular cell adhesion molecule (VCAM), in atherosclerosis of aorta and carotid artery, 344–345 of coronary artery, 358 Vascular Doppler, for monitoring during CMR, 105 Vascular smooth muscle cells (VSMCs), in atherosclerosis, 341–342 Vascular wall stiffness, defined, 363 Vasodilator(s), for myocardial oxygenation assessment, 570–571 Vasodilator stress for comprehensive CMR assessment of coronary artery disease, 159–160 for myocardial perfusion studies, 214, 216f Vasovist. See Gadofosveset trisodium (MS-325, Vasovist, Ablavar). Vastly undersampled isotropic projection reconstruction (VIPR), 41–42 VCAM (vascular cell adhesion molecule), in atherosclerosis of aorta and carotid artery, 344–345 of coronary artery, 358 Vectorcardiography, 105, 146

Cardiovascular Magnetic Resonance 641

INDEX

TOF (tetralogy of Fallot). See Tetralogy of Fallot (TOF). TOF methods. See Time-of-flight (TOF) methods. Torque, in interventional CMR, 600 Toshiba, CMR terminology used by, 611–611 TR. See Time to repetition (TR). Trabeculae carneae, 381 Tracer kinetic model, in quantitative evaluation of myocardial perfusion, 63–65, 64f Transesophageal echocardiography (TEE) of aortic dissection, 453–454 of aortic intramural hematoma, 454–455 dobutamine, of chronic myocardial infarction, 276 of pericardial disease, 488 of thoracic aortic aneurysm, 457 Transjugular intrahepatic portosystemic shunt (TIPS), interventional CMR for, 587 Transmyocardial laser revascularization (TMLR), for ventricular remodeling, 261 Transposition of the great arteries (TGA), 120–122, 409–411 arterial switch procedure for, 410–411, 416 atrial switch procedure for, 415–416 postoperative assessment of, 423–425 contrast-enhanced CMR for, 425f ECG-gated SSFP imaging for, 423, 423f navigator-gated imaging for, 423, 424f cine SSFP of, 120–122, 121f congenitally (physiologically) corrected, 410–411, 411f, 423 D- (complete), 409–410, 410f, 422–423 dark-blood CMR of, 120–122, 121f defined, 422 in infant and pediatric patients, 422–423 L- (levo-, L-loop), 410–411, 411f, 423 pulmonary arteries after repair of, 485, 486f steady-state free precession images of, 411f three-dimensional contrast-enhanced MRA of, 117f, 120–122, 122f types of, 422 Transseptal needle puncture, interventional CMR for, 585f, 586 Transthoracic echocardiography (TTE) of atrial septal defect, 398–399 of cardiomyopathy, 515 of congenital heart disease, 408, 415 of constrictive pericarditis, 493 of pericardial disease, 488 of pericardial effusions, 492 Transverse plane, 3 Transverse relaxation, 6 TRICKS (time-resolved imaging of contrast kinetics) technique, for extracranial carotid arteries, 467 Tricuspid regurgitation, 182f, 417, 512 etiology of, 512 measurement of, 512, 513f severity of, 502t Tricuspid stenosis, 502t, 508 Tricuspid valve, Ebstein anomaly of, 413, 414f, 415f Tricuspid valve atresia, 402f Tricuspid valve stenosis, 502t, 508 Triggering, cardiac. See Cardiac gating. Truncus arteriosus, 412–413, 412f associated anomalies with, 426–427 cine CMR of, 116f defined, 426 epidemiology of, 426 in infant and pediatric patients, 426–428 classification of, 426, 427f contrast-enhanced CMR of, 428f postoperative assessment of, 427–428 preoperative assessment of, 427, 428f surgical repair of, 427

INDEX

Velocity encoded CMR of congenital heart disease for evaluation of function, 417 after repair, 416–417 for pulmonary artery hypertension, 484–485, 484f for valvular heart disease, 403 Velocity encoding value (VENC), for aortic flow, 152 Velocity mapping of congenital heart disease, 113–115 due to single ventricle, 123–124 due to transposition of the great arteries, 121–122 of coronary artery velocity, 314 bolus tagging for, 314 with coronary artery bypass graft, 333–337, 337f, 338f echo planar time-of-flight technique for, 314–315, 315f gradient echo phase, 315–317 breath holding techniques for, 315–316 navigator techniques for, 316–317, 317f, 318f in-plane, 316–317, 317f interleaved spiral phase, 317–320, 319f, 320f, 321f through-plane, 317, 317f, 318f of diastolic function, 190, 192f phase contrast, 91, 92–93 flow vector map in, 96, 97f flow velocity images in, 92–93, 93f and Fourier velocity imaging, 94 improving accuracy of, 93–97, 94f, 95f rapid, 95–96 of thoracic aorta, 450–451, 451f validation of, 95 for valvular heart disease, 501, 502f, 506 with single ventricle, 123–124 of systolic function, 189–190 Velocity phase encoding, in Fourier flow imaging, 91, 93 Velocity phase sensitivity, 92 Velocity sensitivity, in phase contrast velocity mapping, 93–94 Velocity-encoded CMR imaging (VENC), 32–34 cardiac gating for, 32 goal of, 32, 33f pulse sequence in, 33–34, 33f of thoracic aorta, 450–451, 451f VENC. See Velocity-encoded CMR imaging. VENC (velocity encoding value), for aortic flow, 152 Venous pathway imaging, with single ventricle, 123f, 124 Ventilation/perfusion scanning, for pulmonary embolism, 480 Ventricle left (See Left ventricle (LV)) right (See Right ventricle (RV)) single, 122–124, 413–415 aortic arch imaging for, 123–124, 123f, 124f cine CMR for, 124, 124f defined, 430 epidemiology of, 430 with Fontan baffle, 123, 123f, 125f in infant and pediatric patients, 430–435 evaluation of, 432 Fontan procedure for, 432, 432f left, 430, 431f post-Fontan, 434–435, 434f right, 430, 431f during staged palliation, 433, 433f, 434f late gadolinium enhancement for, 124f perfusion imaging for, 124f

642 Cardiovascular Magnetic Resonance

Ventricle (Continued) pulmonary artery imaging for, 123, 123f, 125f time-resolved gadolinium imaging for, 118f velocity mapping for, 123–124 venous pathway imaging for, 123f, 124 with ventricular outflow obstruction, 125f Ventricular abnormalities, complex, 413–415 Ventricular anatomy, during remodeling, 253–255, 254f Ventricular aneurysm, due to acute myocardial infarction, 249 Ventricular fibroma, 118f Ventricular filling patterns, in constrictive pericarditis, 495 Ventricular function, 149–150 left, 149, 150f right, 150 Ventricular inversion, isolated, 409 Ventricular loop, 409 Ventricular mapping, interventional CMR for, 586 Ventricular mass, for right ventricular assessment, 383–385, 386f Ventricular morphology, 409 Ventricular noncompaction (VNC), 522, 522f characteristics of, 522 function and morphology in, 522 tissue characterization in, 522 Ventricular outflow obstruction with single ventricle, 125f with transposition of the great arteries, 121f, 122, 122f Ventricular pseudoaneurysm, due to acute myocardial infarction, 249, 249f Ventricular remodeling, after acute myocardial infarction, 253–266 CMR spectroscopy of energetics during, 256–257, 257f contrast-enhanced CMR and predictors of, 257–259, 258f, 259f early phase of, 253 late phase of, 253, 254f pathophysiology of, 253 regional left ventricular function during, 255–256 therapy for animal studies of, 259–262, 261f human studies of, 262–264, 263f ventricular anatomy during, 253–255, 254f Ventricular septal defect (VSD), 396–397 due to acute myocardial infarction, 249, 249f anatomic delineation of, 397, 397f cine CMR of, 113 clinical manifestations of, 396–397 CMR-guided catheterization and intervention for, 397 with double-outlet right ventricle, 412, 425 location of, 396, 396f saturation band for, 116 shunt quantification in, 397, 398f spontaneous closure of, 396–397 surgical management of, 396–397, 397f in tetralogy of Fallot, 420 in truncus arteriosus, 426 velocity mapping of, 113 Ventricular tachyarrhythmia, CMR-guided RF ablation for, 603 Ventricular volumes effect of imaging sequence and magnetic field strength on, 150–152, 151t, 152t for right ventricular assessment, 383–385, 386f

Ventriculoarterial connection disorder(s), 409–413 double-outlet right ventricle as, 410f, 412 tetralogy of Fallot as, 413 transposition of the great arteries as, 409–411, 410f, 411f truncus arteriosus as, 412–413, 412f Vertical long axis (VLA) image, 140, 141f, 186–187 Viability imaging, 31, 267–283 in acute myocardial infarction, 269–274 contrast-enhanced studies for, 269–274 late gadolinium enhancement for, 269–271 adjustment of T1 in, 271, 273f ischemic bed at risk in, 271, 272f nulling of signal intensity of normal myocardium in, 269, 270f and recovery of function, 273–274 T1-weighted segmented inversion recovery pulse sequence for, 269, 269f time between contrast injection and imaging in, 269–271, 271f no-reflow phenomenon and, 271–272, 274f T2-weighted images for, 269 wall thickness and, 274 cardiac gating for, 31, 32f in chronic myocardial infarction, 274–280 contractile reserve and, 275–278 late gadolinium enhancement for, 276, 277f, 278f, 279f vs. dobutamine CMR, 277–278 vs. other imaging modalities, 276–277, 281f thickness of epicardial rim and recovery of function in, 276, 280f wall thickness and, 275, 275f CMR spectroscopy for, 564–565, 564f for coronary artery disease CMR for, 159 comprehensive CMR assessment for analysis of studies with, 164–166, 167f protocols for, 163–164, 165f, 166f, 167f dobutamine stress CMR for, 204–205 contractile reserve in, 204f, 206f end-systolic wall thickening in, 204, 204f intramyocardial segment shortening in, 205 low-dose, 204–205, 237–238 short-axis basal views in, 205f tissue tagging in, 204–205, 207f feature(s) of, 267–269 contractile reserve as, 267–268 early hypoenhancement with gadolinium as, 268 late gadolinium enhancement in infarcted tissues as, 268 left ventricular wall thickness as, 267 no-reflow phenomenon as, 268 scar formation as, 267 tissue edema as, 268 goal of, 31, 31f high-energy phosphates and, 268–269 inversion recovery in, 31, 32f magnetic resonance spectroscopy for, 278–280 parallel imaging for, 50f, 51 sodium and potassium CMR for, 268–269 Viable myocardium, 267–283 in acute myocardial infarction, 269–274 contrast-enhanced studies for, 269–274 late gadolinium enhancement for, 269–271

W

Wall motion stress CMR, 196–212 adenosine, 198t, 208–209 atropine pharmacokinetics of, 196 safety of, 196–197, 198t dipyridamole, 208–209 dobutamine apical and short axis views in, 198, 200f atropine in, 198, 199f, 202–203 cine GRE or SSFP bright-blood images in, 198, 199t combined adenosine perfusion and diagnostic performance of, 238–239, 238f protocol for, 232, 233f, 234f delineation of orthogonal left ventricular myocardial segments in, 198, 200f vs. dobutamine stress echocardiography, 202–203, 202t accuracy of, 197 safety of, 196–197, 197t technique of, 197 uses of, 197 for viability studies, 204 dopamine infusion protocol for, 198, 199f facilities for, 198, 199f inducible ischemia during, 198, 201f contemporary studies of, 202–203, 202t, 203f, 203t early studies of, 201–202, 202f late gadolinium enhancement for, 205, 207f pharmacokinetics of, 196 for prognosis, 205, 207f vs. radionuclide studies, 204 safety of, 196–197, 197t, 198t sensitivity and specificity of, 203, 203f, 203t technique for, 197–200 3-T, 203 tissue tagging during, 206–208, 208f for viability studies, 204–205, 207f in viability studies, 204–205 contractile reserve in, 204f, 206f end-systolic wall thickening in, 204, 204f intramyocardial segment shortening in, 205 low-dose, 204–205, 237–238 short-axis basal views in, 200f tissue tagging in, 204–205, 207f during exercise, 209–210, 210f on-site medications for, 198t vs. stress myocardial perfusion CMR, 229–240 Wall shear stress (WSS) arterial, 371–374, 374f, 375f in blood flow velocity assessment, 96 Wall thickness, and myocardial viability, 267 after acute myocardial infarction, 274

Wall thickness, and myocardial viability (Continued) in chronic myocardial infarction, 275, 275f Washin/washout methods, 91, 92f Water exchange in contrast-enhanced tissue relaxation, 84 effects on myocardial contrast enhancement of, 61 in stress myocardial perfusion imaging, 222 Water-excitation acquisition, for coronary artery velocity mapping, 317–320, 321f Whole heart 3D coronary artery CMR acquisition sequence for, 291, 291f, 292f of native vessel stenosis, 302f, 303, 303t Windowing, for navigator echoes, 131, 131t, 132f Worksheet, for CMR, 614–616 Wrist, peripheral vascular disease of, 475 WSS (wall shear stress) arterial, 371–374, 374f, 375f in blood flow velocity assessment, 96

X

X position, 7, 7f selection of, 9–10, 9f, 10f X-ray and CMR (XMR) guidance system, 593, 601–604 for biventricular pacing, 605f early experience in humans with, 603–604, 604f facility design for, 594f, 601–602 image registration in, 604 laboratory for, 580 performance of intervention using, 602–603, 603f safety features for, 594f, 601 X-ray angiography of mesenteric arteries, 472 of peripheral vascular disease, 473–474 of pulmonary embolism, 480 of renal artery stenosis, 471 X-ray coronary angiography, vs. dobutamine stress echocardiography and dobutamine stress CMR, 202t, 203 X-ray fluoroscopy (XRF), vs. interventional CMR, 580, 581t

Y

Y position, 7, 7f selection of, 10–11, 10f

Z

Z position, 7, 7f selection of, 8–9, 8f, 9f Z-axis direction, 3

Cardiovascular Magnetic Resonance 643

INDEX

Viable myocardium (Continued) adjustment of T1 in, 271, 273f ischemic bed at risk in, 271, 272f nulling of signal intensity of normal myocardium in, 269, 270f and recovery of function, 273–274 T1-weighted segmented inversion recovery pulse sequence for, 269, 269f time between contrast injection and imaging in, 269–271, 271f no-reflow phenomenon and, 271–272, 274f T2-weighted images for, 269 wall thickness and, 274 assessment of, 31 cardiac gating for, 31, 32f CMR spectroscopy for, 564–565, 564f goal of, 31, 31f inversion recovery in, 31, 32f parallel imaging for, 50f, 51 in chronic myocardial infarction, 274–280 contractile reserve and, 275–278 late gadolinium enhancement for, 276, 277f, 278f, 279f vs. dobutamine CMR, 277–278 vs. other imaging modalities, 276–277, 281f thickness of epicardial rim and recovery of function in, 276, 280f wall thickness and, 275, 275f CMR spectroscopy for, 278–280 contractile reserve of, 267–268 defined, 267 feature(s) of, 267–269 contractile reserve as, 267–268 early hypoenhancement with gadolinium as, 268 late gadolinium enhancement in infarcted tissues as, 268 left ventricular wall thickness as, 267 no-reflow phenomenon as, 268 scar formation as, 267 tissue edema as, 268 high-energy phosphates and, 268–269 sodium and potassium CMR for, 268–269 VIPR (vastly undersampled isotropic projection reconstruction), 41–42 Visceral-atrial rule, 409 Viscoelasticity, 363 VLA (vertical long axis) image, 140, 141f, 186–187 VNC. See Ventricular noncompaction (VNC). VSD. See Ventricular septal defect (VSD). VSMCs (vascular smooth muscle cells), in atherosclerosis, 341–342 Vulnerable myocardium, 351 Vulnerable patient, 351 Vulnerable plaques defined, 351 markers of, 351, 352t noninvasive diagnosis of, 351, 352t