Cardiovascular Magnetic Resonance

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, Winston Salem, 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, Wake Forest 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 Publique Hopitaux 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 Bosch Krankenhaus, 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 car diac imaging, beginning with angiography and arteriogra phy, 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, infiltra tion, and fibrosis. Both valvular performance and ventri cular function can be measured. Magnetic resonance angiography allows assessment of vascular function and tis sue 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 non invasively by magnetic resonance spectroscopy, these tech niques 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. Man ning and Dudley J. Pennell and their talented contributors to Cardiovascular Magnetic Resonance have taken a signifi cant 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 car diac 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 com position 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, how ever, as this technology is not quickly learned, and high quality 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 moti vated practitioners in CMR for the benefit of clinical sci ence, patient care, and the advancement of the field. The reader should be forewarned that CMR is constantly devel oping, and no text can include all of the most recent devel opments; 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 indivi duals to whom we owe thanks. These include the outstand ing contributions of our primary authors and their collaborators; Rebecca Schmidt Gaertner, our editor at Else vier Health Sciences; Linda R. Garber, our production edi tor; 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, train ees, 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 appre ciation 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 fundamen tal 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 caus ing 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 imag ing. To determine the location of different magnetic spins (or 1 H), the magnetic field, B, is made to vary linearly with posi tion 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 direc tion 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 pre cess, or rotate, around the Z axis in a random, incoherent fash ion, 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 X Y plane will not detect any signal because the spins are in dif ferent 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 rotat ing 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 imag ing,4,5 late gadolinium enhancement, perfusion,6 and func tion.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) excita tion 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 fre quency 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 indi cates 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 dif ference 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 dia gram 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 mag netic 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 longitu dinal (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 equi librium value (M0). MzðtÞ ¼ M0 ð1  e

t=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 magne tization 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 sig nal 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 equilib rium 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 sys tem and return to equilibrium with the main magnetic field, B0, through T1 and T2 relaxation. T1 and T2 relaxa tion forms the foundation of MRI.

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

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 deoxygen ated 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 off resonance effects. Later we will describe how to minimize or use these effects to make in vivo measurements.

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 devel opment 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 dephas ing 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: SðtÞ ¼ S0 e

t=T2

Spin Echo Imaging 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 back to 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 “indi vidual” pace. The time it takes them to reach each other

(4)

T2 relaxation is also called spin spin, or transverse, relaxa tion. 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 fre quency, 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 through out the sample, then the frequency of the spins in different regions will vary (Equation 1). These spins are called off resonance 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 dimen sions (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 fre quency, amplitude, and phase. These can be used to deter mine 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 limi tation of MRI that results in a relatively slow image acquisi tion 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 infor mation 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 tomo graphic 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 fre quency 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 contin ually refocusing the magnetization with 180 pulses is called fast spin echo (or turbo spin echo), and is used in cardiac imag ing. 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 relax ation, 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 deter mining the optimal image sequence for gathering informa tion 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 con nective 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 Lar mor 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 fre quency 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 preces sing 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 gradi ent 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, depend ing 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 posi tion. 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 impor tant. The gradient “area” is the product of the gradient ampli tude, 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 trans verse 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 mag netic 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 fre quency 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 posi tion. 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 dur ing 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 gradi ent. The readout gradient is then ramped up in the oppo site direction and held constant for a time, and then the gradient is ramped back down to zero. This is the time dur ing 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 trans verse 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 posi tion. 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 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 com bined to create k space (Fig. 1 13). The k space of Fig ure 1 13 represents the raw data collected at each phase encoding step. The phase and frequency encoding direc tions 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 ampli tude 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 mea surement 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 ampli tude and imparts no phase. The signal measured is the bulk signal. In the next phase encoding step, a small gradi ent 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 loca tions. It is important to realize that for each phase encod ing, 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 exci tation pulse used in spin echo imaging, one must wait approximately five times the T1 value of the tissue to per mit the spins to completely regrow to their original Z mag netization (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 sig nal 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 struc ture 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 Fig ure 1 14B for a spin echo sequence. The effect of TR is illustrated for normal myocardium with a T1 of 0.8 sec onds 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 sam pled, this k space is converted into an image using a two dimensional fast Fourier transform. Ideally, the raw k space data contain enough information so that after the two dimensional fast Fourier transform an image of infi nite spatial resolution results. In reality, there are limits on the amount of data that can be collected. The resolu tion of the MRI image is directly related to the number of phase encoded echoes collected (Dy ¼ FOV/Ny); there fore, 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 excita tion. 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 usu ally restricted to relatively motion free phases of the car diac cycle, such as diastole. The inherently high signal to noise ratio (SNR) of these FSE approaches supports very high spatial resolution images of the heart. In addi tion, 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 cen tral k space (i.e., the longer the effective TE), the more T2 weighting will occur. The FSE technique is very com monly 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 rein verted 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 cos 1 ðe

TR=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 gradi ents are applied after each acquisition equal to and oppo site the phase encoding gradient. Furthermore, the read out 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* dephas ing, 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 col lected 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  e

Þ  MzSS e

t=T1

(10)

SS

The fast GRE pulse sequence is also used for contrast enhanced 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). How ever, 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 steady state 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.

t=T1

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 distinc tions. First, after each TR, the spins are rephased on each gradient axis (X, Y, and Z) to zero phase. No crushers, kill ers, 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 gradi ent. 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 exci tation 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 trans verse 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 electrocar diogram gated segmented EPI sequence for myocardial perfu sion 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 gadolin ium contrast agent through the heart. In this sequence, a non selective 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 conven tional 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 k space. Spiral image quality is affected by off resonance arti facts 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

–0.5

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 nor mal 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

1 BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

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

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

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.

18 Cardiovascular Magnetic Resonance

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 spe cific 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 clin ical 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 viabil ity/infarction) An additional 25% of the clinical volume is referred for via bility 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 underly ing 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 per formance 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 gado linium contrast (see Chapter 6). For those with moderate or severely depressed renal function (estimated glomerular fil tration 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 com bined to answer common diagnostic questions in cardio logy, 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 short and long axis views of the heart and to confirm the optimal position of the anterior and posterior elements of the tho racic coil. Because of patient to patient anatomic variation, Cardiovascular Magnetic Resonance 19

2 CLINICAL CARDIOVASCULAR MAGNETIC RESONANCE IMAGING TECHNIQUES

CHAPTER 2

700 CMR procedures (quarterly)

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

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 there fore 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 inter ventricular 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 steady state 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 pri mary advantages of SSFP imaging are rapid acquisition speed and resultant images that show very high signal to noise ratio (SNR) and contrast to noise ratio (CNR) in the blood pool and surrounding myocardium. Accord ingly, 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 pro cess 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 ven tricles are relatively quiescent. Acquiring the image data in diastole requires the scan ner 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 mini mized this issue.4 Assuming a heart rate of 60 bpm (R R duration of 1000 msec), the first “event” the scanner hard ware 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 cre ate 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 “bright blood” 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

RF

Slice

Read

Phase

Signal

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 phase encoding 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 two dimensional 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 imag ing 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 exam ination of morphology in that it allows one to clearly distin guish 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 prepara tory 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 slice selective RF pulse is applied that causes only the magneti zation 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) previ ously 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 sig nal, 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 recov ery) as well as the physiology of moving blood. Such images are helpful in detecting abnormal large vessel anat omy, 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

R

R

ECG

Delay time

Delay time All k-space lines (e.g., 100 lines) Scout 1

Trigger

All k-space lines (e.g., 100 lines) Scout 2

Trigger

Left

Head

Posterior

Right

Feet

Anterior

Sagittal

Axial

Coronal

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 beat ing heart to visualize its contractile function. Figure 2 14

shows a representative example of a midventricular short axis slice during eight different time points within the car diac 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

Posterior

Right

Feet

Anterior

Sagittal

Axial

Coronal

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 res olution during a single cardiac cycle in real time. For exam ple, 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, consider ably 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 car diac 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°

180°

180°

RF Slice

1st

2nd

3rd echo

Read Phase

Fast spin echo sequence

A 90°

180°

90°

180°

RF Slice Read Phase Signal TE/2

TE/2 TE TR

B

TE Spin echo sequence

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

R

Slice 1

R

Slice 2

R

ECG Delay time Trigger

Image slice 1

Delay time

Image slice 2

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 sec ond movie frame, and so on. In practice, cine imaging is performed during breath holding and takes 10 to 12 sec onds 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 (viabil ity) 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 adeno sine stress/rest myocardial perfusion CMR has gained clini cal attention. Perfusion CMR accurately diagnoses coronary artery disease with high sensitivity and specificity. Rest per fusion 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 adeno sine stress, perfusion defects appear as dark regions (i.e., hypoenhancement) surrounded by bright contrast enhanced normally perfused myocardium. In the corresponding rest perfusion images, perfusion to the hypoen hanced 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.

1

2

3

4

Heartbeat 6

5

ECG Trigger Acquired in segments:

1 2

30

1 2

30

1 2

30

1 2

30

1 2

30

1 2

30

30 segments acquired each heartbeat

Each segment contains 10 lines

10 lines

Ordered after cardiac phase: 1

2

normal, indicating normal resting perfusion. The reduced coro nary flow reserve is caused by hemodynamically significant cor onary 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 dimin ished 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 per formed 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 longi tudinal magnetization into the transverse plane. The trans verse magnetization then is dephased (spoiled) by a spoiler gradient. The contrast agent (e.g., gadolinium diethyl tria mine 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 hypoper fused myocardium. The optimal time for image acquisition is shown at location “b” in Figure 2 18. In practice, how ever, 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 lon gitudinal 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

1 Heartbeat 1

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 impor tant for avoiding image artifacts that can result from mixing of longitudinal and transverse magnetization. For stress myocardial perfusion imaging, a dose of aden osine 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 coro nary 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, aden osine is infused for at least 2 minutes. This minimum in fusion duration is chosen based on physiologic studies in humans, showing that, on average, maximum coronary

0

Scout imaging

Cine imaging

1

8

15

17

1 min

≈ 15 min

1 min

Wait or do additional cine imaging or valvular evaluation

Rest perfusion

2 min

Stress perfusion

Contrast injection 1

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 differ ence in myocardial contrast agent concentration and thus image intensity. Regional myocardium supplied by the dis eased 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

33

LGE imaging 40

45

Time [minutes]

A Adenosine Contrast

2 min

1 min

Saline flush Breath holding Image acquisition

B

1

2 3

4

5

6

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 hyperenhance ment. 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 con centration 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 k space) is chosen such that the magnetization of viable (nor mal) myocardium is near its zero crossing, meaning that these regions appear dark (Fig. 2 23). Nonviable myocar dium (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 grow ing. Beside the assessment of acute and chronic myocardial infarction, viability imaging27–30 has been used for the pre diction of contractile improvement after revascularization,31 for measuring the response to beta blocker treatment,32 for the differentiation of ischemic versus nonischemic cardio myopathy,33 and for the diagnosis of various nonischemic cardiomyopathies.34–38 A large body of evidence shows that LGE can differenti ate between nonviable and viable tissue.27–30,38,39 Regions of irreversible injury exhibit high signal intensity (hyperen hancement) on T1 weighted images after administration of extracellular CMR contrast agent, such as gadolinium DTPA. The contrast agent significantly shortens local longi tudinal 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 gadolin ium 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 remain ing 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 dis played on the scanner in real time to allow observation of contrast agent arrival in the RV. If the scanner does not pro vide 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 gadolin ium 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 sever ity 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. Fig ure 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 back ground (black to dark gray). Analogous to cine imaging, each VENC image corre sponds 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 electrocardiogram triggered velocity encoded sequence. Gradients on the read axis are flow compensated. 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 through plane 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

RF

Read

Phase

Slice

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 lin ear 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 dis playing the difference on a grayscale, the images of Fig ure 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 quantita tive 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 vol ume, and pulmonic regurgitation can be readily measured and the pulmonic:systemic (Qp:Qs) ratio can also be deter mined. Mitral and tricuspid regurgitation can then be cal culated 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 ECG gated 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 T1 shortening 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 com monly 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 acquisi tion, 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 relaxa tion” 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

A

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

CONTRAST OPTIONAL IMAGING IMAGING

Selected brightblood imaging

≈ 2 min

≈ 6 min

≈ 5 min

≈ 5 min

Scout imaging

HASTE (2–3 planes)

Noncontrast 3D angio (GRE)

3D CE-MRA (arterial +/– venous phase)

16

23

0

1

6

12 Time (min)

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. Com pared 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 mod ern 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

Flow/ velocity

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 institu tions,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 histori cally 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

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

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.

36 Cardiovascular Magnetic Resonance

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 car diovascular applications of spiral imaging, radial imaging, and parallel imaging. Arguably, all three approaches earn the moniker of “advanced cardiovascular magnetic reso nance (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 three dimensional (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 under stand. 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, noth ing 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 acqui sition trajectories that have been studied extensively and used increasingly, particularly for rapid and real time CMR. Parallel magnetic resonance imaging (parallel MRI), mean while, 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 radiofre quency (RF) detector coils to supplement the spatial encod ing 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 tech niques 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 dif ferent 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 typi cal cardiovascular applications of spiral imaging, radial ima ging, and parallel imaging will be surveyed. A concluding section explores future directions.

SPIRAL IMAGING Principles Spiral trajectories1,2 have been used extensively for real time 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 accom plished using gradients that oscillate in a coordinated fash ion 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 Four ier 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

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

CHAPTER 3

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

B

A

C

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 spi ral 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 visi ble 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 acquisi tions. An example of spiral blurring as a result of off reso nance 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 off Acquisition

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 spi ral 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 out ward, 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 appli cation in the area of breath hold 2D coronary artery CMR, where the speed gain enabled (1
1
5) mm3 spatial res olution 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 cor onary 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 submil limeter spatial resolution with information for motion com pensation obtained directly from the coronary anatomy itself.16 This approach eliminates the need for cardiac Reconstruction

RF Regrid

Gx FFT

Gy

Gz

ADC

38 Cardiovascular Magnetic Resonance

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 two dimensional 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, three dimensional steady state free precession (SSFP) technique using spiral (left) and Cartesian (right) k space 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.)

A

B

C

D

Cardiovascular Magnetic Resonance 39

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

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 submillime ter 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 steady state 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 Carte sian sampling (see Fig. 3 3A). Variable density spiral trajec tories 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 func tion, guidance of interventional procedures, or flow mea surements. For example, the speed advantage of spiral imaging has been exploited for rapid left ventricular func tion assessment using free breathing cine CMR without car diac 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 underesti mation 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 resolu tion 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 tradi tional 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 dur ing 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 recon structions are more complex than for the Cartesian case, typically involving either regridding procedures or projec tion reconstruction algorithms (e.g., filtered back projec tion [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 pro jection. 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

Reconstruction

RF Regrid

Gx

Gz

ADC

40 Cardiovascular Magnetic Resonance

FFT

PR

Gy

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, projection reconstruction (PR) approaches similar to those used in X ray computed tomography (CT) may be used to generate an image. ADC, analog to digital 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.

A

C

positions, producing characteristic radial streaking arti facts (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 spa tial 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 con trast between blood and myocardium.30–32 Further accel erations have been achieved using combinations of radial undersampling with multi echo readouts.32 The under sampled 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 exten ding 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

B

D

Cardiovascular Magnetic Resonance 41

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

reconstruction that effectively merges these representations. Somewhat paradoxically for a sequence that is often identi fied with rapid imaging, however, this same redundancy also represents an underlying inefficiency in data acquisi tion. In principle, the number of acquired data points needed to gather the outer portions of k space with a req uisite separation is generally larger than for Cartesian acquisitions. Nevertheless, radial trajectories have had a recent resurgence for rapid and real time imaging applica tions, 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 regu larly spaced points in the image plane based only on the undersampled spatial frequency data. The regular pat tern of Cartesian undersampling results in coherent over lap 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 applica tions, 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 com parable 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 blood myocardium contrast, which improves the accuracy of the cardiac functional assessment. The scan efficiency of radial imaging, together with real time sliding window recon struction, 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 tempo ral resolution compared with conventional segmented radial techniques. Self gated radial acquisition strategies were intro duced 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 gat ing 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 com pared 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 coro nary 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 encom passing 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 con trast medium application. This approach provides images well suited for vessel geometry assessment and reliable ste nosis detection in renal arteries (see Fig. 3 6D).47 A signifi cant 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 imag ing 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 enhance ment (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 arti fact 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 imag ing approach that employs parallelism, namely, multi detector 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

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

A

B

C

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 jux taposition of spatial encoding mechanisms shown at the bottom. Both modalities use arrays of detectors to acceler ate 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 rota tion, with additional simultaneous projection information provided by multiple detector rows with coverage of dis tinct 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 pres ence 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 infor mation 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 under sampled acquisitions with information derived from multi ple 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 strate gies 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 pat tern, 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 miss ing 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 over lap. (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 sensi tivity 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 suit able 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 acquisi tion of spatial harmonics (SMASH) technique.50 SMASH was the first parallel imaging approach demonstrated in vivo, although various proposals for simultaneous acqui sition had been made earlier,51–57 dating back to the time of the first uses of coil arrays for MRI. Various improve ments 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 reconstruc tion strategies have since been described. The most com monly used today is the GeneRalized Autocalibrating Partially Parallel Acquisition (GRAPPA) technique.58 Com mercial 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) cur rently provide the array spatial sensitivity encoding tech nique (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 techni ques.) Techniques such as GRAPPA or parallel imaging with augmented radius in k space (PARS)60 may be used to com bine undersampled non Cartesian datasets (or irregularly undersampled Cartesian datasets) into fully sampled Carte sian 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 strate gies 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 bur den 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 tar geted 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 con structed, 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 commer cially available receiver channels has grown steadily, moti vated 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 recon struction, 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 refer ence 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 suf fice for an entire series of accelerated acquisitions. However, motion of the patient or the coil array between the time of cal ibration 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 prob lem, self calibrating parallel imaging strategies may be used58,68–70 (see Fig. 3 9B). Self calibrating approaches incor porate a sensitivity reference directly into the accelerated acquisition (in the form of a small set of additional fully sam pled central k space lines). Therefore, they allow accurate par allel imaging even in the presence of vigorous motion. Among the commonly used techniques currently available from ven dors, 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 p g R

(1)

Here, R represents the acceleration factor and the square root dependence in Equation 1 reflects the reduced SNR aver aging 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 charac terizing 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 gener ally 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 func tion of time with otherwise identical acquisition parameters and to take advantage of resulting spatiotemporal correla tions. This is the concept behind techniques such as Un aliasing 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 imag ing 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

Parallel image reconstruction

A

g (R = 2)

10

Separate accelerated acquisition

5 0

g (R = 3)

10 5 0

g (R = 4)

10 5 0

Parallel image reconstruction Single variable-density

B acquisition

C

Figure 3 9 Coil sensitivity calibration strategies and noise propagation behavior for parallel imaging. A, External calibration. A low resolution 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 1 to R 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 exist ing imaging studies as well as to enable new ones previ ously precluded by constraints on scan time. Today, commercial parallel imaging implementations are available for all modern MR scanners, and parallel imaging acquisi tions 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 deposi tion, 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 simulta neously 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 misregistra tion compared with the conventional approach.82 Mean while, 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 capa bility 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 tempo ral resolution, and short scan times. Parallel imaging strate gies 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 sin gle 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 pre cisely the exact time point of maximal systolic contrac tion 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 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 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 24 cm). Image artifacts are not present in the accelerated scans if the target field of view is increased sufficiently (right, FOV 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 com mon 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 res olution 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 k t 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 clin ical 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 con trast 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 refer ence 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 improv ing coronary artery CMR by minimizing the effect of physi ologic motion. The use of parallel imaging strategies enables 3D free breathing navigator techniques to be accel erated (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 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 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 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, three dimensional 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 scan ning the entire heart in a single breath hold.91,92 Acceleration of real time imaging methods for the assess ment 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 mini mized 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 facili tates the monitoring of transient hemodynamics and car diac 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 undersam pling strategies. The combination of UNFOLD73 with SENSE has been used to achieve higher acceleration fac tors 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 multi slice 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

52 Cardiovascular Magnetic Resonance

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 several decade history of MRI is that imaging speed is likely to con tinue 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 recon struction is shown on the right. Clearly, the use of eight projections is insufficient for traditional FBP reconstruc tion, and the image is uninterpretable. However, HYPR incorporates some principles of the k t techniques dis cussed earlier, using information from a time series of undersampled radial acquisitions to constrain each time frame. As the fidelity of the HYPR image suggests, extraor dinary levels of acceleration may be possible, as long as conditions appropriate for HYPR exist, namely, that tempo ral 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 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.)

Cardiovascular Magnetic Resonance 53

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

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 com peting constraints of spatial resolution, anatomic coverage, and breath hold time. The efficiency advantage of parallel imaging strategies minimizes the effect of cardiac and respi ratory motion and facilitates shorter data acquisition win dows 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 acquisi tion window. Parallel imaging also allows reductions in the number of cardiac cycles used for in plane segmenta tion. 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 trans lated 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 imag ing.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.

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

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 num bers 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 acceler ation caused by fundamental electrodynamic constraints on the shapes of RF coil sensitivities.116,117 However, with suffi cient 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 com bine 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 vol umetric scans, which have been viewed increasingly as a challenge to CMR. As we have seen, however, CMR is capa ble 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 com plete 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 3 12D. Of course, comprehensive volumetric CMR boasts a number of advantages over its kindred tomographic modal ity. 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 diffi cult to imagine the comprehensive cardiovascular examina tion 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.

References 1. Ahn BC, Kim JH, Cho ZH. High-speed spiral-scan echo planar NMR imaging. IEEE Trans Med Imaging. 1986;MI-5:2–7. 2. Meyer CH, Hu BS, Nishimura DG, Macovski A. Fast spiral coronary artery imaging. Magn Reson Med. 1992;28:202–213. 3. O’Sullivan JD. A fast sinc function gridding algorithm for Fourier inversion in computer tomography. IEEE Trans Med Imaging. 1985;4:200–207. 4. Jackson JI, Meyer CH, Nishimura DG, Macovski A. Selection of a convolution function for Fourier inversion using gridding. IEEE Trans Med Imaging. 1991;10:473–478. 5. Kerr AB, Pauly JM, Hu BS, et al. Real-time interactive MRI on a conventional scanner. Magn Reson Med. 1997;38:355–367. 6. Noll DC, Pauly JM, Meyer CH, Nishimura DG, Macovski A. Deblurring for non-2D Fourier transform magnetic resonance imaging. Magn Reson Med. 1992;25:319–333. 7. Irarrazabal P, Meyer CH, Nishimura DG, Macovski A. Inhomogeneity correction using an estimated linear field map. Magn Reson Med. 1996;35:278–282. 8. Man LC, Pauly JM, Macovski A. Improved automatic off-resonance correction without a field map in spiral imaging. Magn Reson Med. 1997;37:906–913. 9. Meyer CH, Pauly JM, Macovski A, Nishimura DG. Simultaneous spatial and spectral selective excitation. Magn Reson Med. 1990;15: 287–304. 10. Block W, Pauly J, Nishimura D. RARE spiral T2-weighted imaging. Magn Reson Med. 1997;37:582–590. 11. Bornert P, Aldefeld B, Eggers H. Reversed spiral MR imaging. Magn Reson Med. 2000;44:479–484. 12. Irarrazabal P, Nishimura DG. Fast three dimensional magnetic resonance imaging. Magn Reson Med. 1995;33:656–662. 13. 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;46:789–794. 14. Foo TK, Ho VB, Hood MN. Vessel tracking: prospective adjustment of section-selective MR angiographic locations for improved coronary artery visualization over the cardiac cycle. Radiology. 2000;214: 283–289. 15. Saranathan M, Ho VB, Hood MN, Foo TK, Hardy CJ. Adaptive vessel tracking: automated computation of vessel trajectories for improved efficiency in 2D coronary MR angiography. J Magn Reson Imaging. 2001;14:368–373.

54 Cardiovascular Magnetic Resonance

16. Sussman MS, Stainsby JA, Robert N, Merchant N, Wright GA. Variable-density adaptive imaging for high-resolution coronary artery MRI. Magn Reson Med. 2002;48:753–764. 17. Santos JM, Cunningham CH, Lustig M, et al. Single breath-hold whole-heart MRA using variable-density spirals at 3T. Magn Reson Med. 2006;55:371–379. 18. Katoh M, Spuentrup E, Buecker A, Manning WJ, Gunther RW, Botnar RM. MR coronary vessel wall imaging: comparison between radial and spiral k-space sampling. J Magn Reson Imaging. 2006;23: 757–762. 19. Kuhl HP, Spuentrup E, Wall A, et al. Assessment of myocardial function with interactive non-breath-hold real-time MR imaging: comparison with echocardiography and breath-hold cine MR imaging. Radiology. 2004;231:198–207. 20. Liao JR, Pauly JM, Brosnan TJ, Pelc NJ. Reduction of motion artifacts in cine MRI using variable-density spiral trajectories. Magn Reson Med. 1997;37:569–575. 21. Spuentrup E, Schroeder J, Mahnken AH, et al. Quantitative assessment of left ventricular function with interactive real-time spiral and radial MR imaging. Radiology. 2003;227:870–876. 22. Nayak KS, Hargreaves BA, Hu BS, Nishimura DG, Pauly JM, Meyer CH. Spiral balanced steady-state free precession cardiac imaging. Magn Reson Med. 2005;53:1468–1473. 23. Nayak KS, Cunningham CH, Santos JM, Pauly JM. Real-time cardiac MRI at 3 tesla. Magn Reson Med. 2004;51:655–660. 24. Nezafat R, Kellman P, Derbyshire JA, McVeigh ER. Real-time blood flow imaging using autocalibrated spiral sensitivity encoding. Magn Reson Med. 2005;54:1557–1561. 25. Lauterbur P. Image formation by induced local interactions: examples employing nuclear magnetic resonance. Nature. 1973;242:190–191. 26. Lauterbur P, Lai C-M. Zeugmatography by reconstruction from projections. IEEE Trans Nucl Sci. 1980;NS-27:1227–1231. 27. Glover GH, Pauly JM. Projection reconstruction techniques for reduction of motion effects in MRI. Magn Reson Med. 1992;28: 275–289. 28. Peters DC, Korosec FR, Grist TM, et al. Undersampled projection reconstruction applied to MR angiography. Magn Reson Med. 2000;43:91–101. 29. Peters DC, Epstein FH, McVeigh ER. Myocardial wall tagging with undersampled projection reconstruction. Magn Reson Med. 2001;45:562–567.

53. Kelton JR, Magin RL, Wright SM. An algorithm for rapid image acquisition using multiple receiver coils. In: Proceedings of the 8th Annual Meeting of the Society for Magnetic Resonance in Medicine. Amsterdam, Netherlands: 1989:1172. 54. Kwiat D, Einav S, Navon G. A decoupled coil detector array for fast image acquisition in magnetic resonance imaging. Med Phys. 1991;18:251–265. 55. Ra JB, Rim CY. Fast imaging using subencoding data sets from multiple detectors. Magn Reson Med. 1993;30:142–145. 56. Carlson JW, Minemura T. Imaging time reduction through multiple receiver coil data acquisition and image reconstruction. Magn Reson Med. 1993;29:681–687. 57. Kwiat D, Einav S. Preliminary experimental evaluation of an inverse source imaging procedure using a decoupled coil detector array in magnetic resonance imaging. Med Eng Phys. 1995;17:257–263. 58. Griswold MA, Jakob PM, Heidemann RM, et al. Generalized autocalibrating partially parallel acquisitions (GRAPPA). Magn Reson Med. 2002;47:1202–1210. 59. Pruessmann KP, Weiger M, Scheidegger MB, Boesiger P. SENSE: Sensitivity encoding for fast MRI. Magn Reson Med. 1999;42:952–962. 60. Yeh EN, McKenzie CA, Ohliger MA, Sodickson DK. Parallel magnetic resonance imaging with adaptive radius in k-space (PARS): constrained image reconstruction using k-space locality in radiofrequency coil encoded data. Magn Reson Med. 2005;53:1383–1392. 61. Pruessmann KP, Weiger M, Bornert P, Boesiger P. Advances in sensitivity encoding with arbitrary k-space trajectories. Magn Reson Med. 2001;46:638–651. 62. Kyriakos WE, Panych LP, Kacher DF, et al. Sensitivity profiles from an array of coils for encoding and reconstruction in parallel (SPACE RIP). Magn Reson Med. 2000;44:301–308. 63. Sodickson DK, McKenzie CA. A generalized approach to parallel magnetic resonance imaging. Med Phys. 2001;28:1629–1643. 64. Bydder M, Larkman DJ, Hajnal JV. Generalized SMASH imaging. Magn Reson Med. 2002;47:160–170. 65. Cohen S, Grant AK, Yeh EN, Joshi S, Sodickson DK. A parallel computing solution for rapid reconstruction of highly-accelerated volumetric parallel MRI Data. In: Proceedings of the 12th Scientific Meeting of the International Society for Magnetic Resonance in Medicine. Kyoto: 2004:2631. 66. Ohliger MA, Sodickson DK. An introduction to coil array design for parallel MRI. NMR Biomed. 2006;19:300–315. 67. Sodickson DK. Tailored SMASH image reconstructions for robust In vivo parallel MR imaging. Magn Reson Med. 2000;44:243–251. 68. Jakob PM, Griswold MA, Edelman RR, Sodickson DK. AUTOSMASH: a self-calibrating technique for SMASH imaging. MAGMA. 1998;7: 42–54. 69. Heidemann RM, Griswold MA, Haase A, Jakob PM. VD-AUTOSMASH imaging. Magn Reson Med. 2001;45:1066–1074. 70. McKenzie CA, Yeh EN, Ohliger MA, Price MD, Sodickson DK. Selfcalibrating parallel imaging with automatic coil sensitivity extraction. Magn Reson Med. 2002;47:529–538. 71. Hunold P, Maderwald S, Ladd ME, Jellus V, Barkhausen J. Parallel acquisition techniques in cardiac cine magnetic resonance imaging using TrueFISP sequences: comparison of image quality and artifacts. J Magn Reson Imaging. 2004;20:506–511. 72. Sodickson DK, Griswold MA, Jakob PM, Edelman RR, Manning WJ. Signal-to-noise ratio and signal-to-noise efficiency in SMASH imaging. Magn Reson Med. 1999;41:1009–1022. 73. Madore B, Glover GH, Pelc NJ. Unaliasing by fourier-encoding the overlaps using the temporal dimension (UNFOLD), applied to cardiac imaging and fMRI. Magn Reson Med. 1999;42:813–828. 74. Tsao J, Boesiger P, Pruessmann KP. k-t BLAST and k-t SENSE: dynamic MRI with high frame rate exploiting spatiotemporal correlations. Magn Reson Med. 2003;50:1031–1042. 75. Hansen MS, Kozerke S, Pruessmann KP, Boesiger P, Pedersen EM, Tsao J. On the influence of training data quality in k-t BLAST reconstruction. Magn Reson Med. 2004;52:1175–1183. 76. Tsao J. K-t SENSE for efficient dynamic parallel imaging by jointspace time consideration. In: Second International Workshop on Parallel MRI. Zurich, Switzerland: 2004:51. 77. Madore B. UNFOLD-SENSE: a parallel MRI method with selfcalibration and artifact suppression. Magn Reson Med. 2004;52: 310–320. 78. Kellman P, Epstein FH, McVeigh ER. Adaptive sensitivity encoding incorporating temporal filtering (TSENSE). Magn Reson Med. 2001;45:846–852.

Cardiovascular Magnetic Resonance 55

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

30. Shankaranarayanan A, Simonetti OP, Laub G, Lewin JS, Duerk JL. Segmented k-space and real-time cardiac cine MR imaging with radial trajectories. Radiology. 2001;221:827–836. 31. Schaeffter T, Weiss S, Eggers H, Rasche V. Projection reconstruction balanced fast field echo for interactive real-time cardiac imaging. Magn Reson Med. 2001;46:1238–1241. 32. Larson AC, Simonetti OP. Real-time cardiac cine imaging with SPIDER: steady-state projection imaging with dynamic echo-train readout. Magn Reson Med. 2001;46:1059–1066. 33. Barger AV, Block WF, Toropov Y, Grist TM, Mistretta CA. Timeresolved contrast-enhanced imaging with isotropic resolution and broad coverage using an undersampled 3D projection trajectory. Magn Reson Med. 2002;48:297–305. 34. Larson AC, White RD, Laub G, McVeigh ER, Li D, Simonetti OP. Selfgated cardiac cine MRI. Magn Reson Med. 2004;51:93–102. 35. Larson AC, Kellman P, Arai A, et al. Preliminary investigation of respiratory self-gating for free-breathing segmented cine MRI. Magn Reson Med. 2005;53:159–168. 36. Krombach GA, Pfeffer JG, Kinzel S, Katoh M, Gunther RW, Buecker A. MR-guided percutaneous intramyocardial injection with an MRcompatible catheter: feasibility and changes in T1 values after injection of extracellular contrast medium in pigs. Radiology. 2005;235:487–494. 37. Weiss S, Kuehne T, Brinkert F, et al. In vivo safe catheter visualization and slice tracking using an optically detunable resonant marker. Magn Reson Med. 2004;52:860–868. 38. 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;105:874–879. 39. Spuentrup E, Katoh M, Buecker A, et al. Free-breathing 3D steadystate free precession coronary MR angiography with radial k-space sampling: comparison with cartesian k-space sampling and cartesian gradient-echo coronary MR angiography–pilot study. Radiology. 2004;231:581–586. 40. McLeish K, Kozerke S, Crum WR, Hill DL. Free-breathing radial acquisitions of the heart. Magn Reson Med. 2004;52:1127–1135. 41. Stehning C, Bornert P, Nehrke K, Dossel O. Free breathing 3D balanced FFE coronary magnetic resonance angiography with prolonged cardiac acquisition windows and intra-RR motion correction. Magn Reson Med. 2005;53:719–723. 42. 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. 43. Park J, Larson AC, Zhang Q, Simonetti O, Li D. 4D radial coronary artery imaging within a single breath-hold: cine angiography with phase-sensitive fat suppression (CAPS). Magn Reson Med. 2005;54: 833–840. 44. Du J, Carroll TJ, Brodsky E, et al. Contrast-enhanced peripheral magnetic resonance angiography using time-resolved vastly undersampled isotropic projection reconstruction. J Magn Reson Imaging. 2004;20:894–900. 45. Du J, Korosec FR, Thornton FJ, Grist TM, Mistretta CA. High-resolution multistation peripheral MR angiography using undersampled projection reconstruction imaging. Magn Reson Med. 2004;52:204–208. 46. Fain SB, Browning FJ, Polzin JA, et al. Floating table isotropic projection (FLIPR) acquisition: a time-resolved 3D method for extended field-of-view MRI during continuous table motion. Magn Reson Med. 2004;52:1093–1102. 47. Katoh M, Spuentrup E, Stuber M, Hoogeveen R, Gunther RW, Buecker A. Free-breathing renal magnetic resonance angiography with steady-state free-precession and slab-selective spin inversion combined with radial k-space sampling and water-selective excitation. Magn Reson Med. 2005;53:1228–1233. 48. Katoh M, Stuber M, Buecker A, Gunther RW, Spuentrup E. Spinlabeling coronary MR angiography with steady-state free precession and radial k-space sampling: initial results in healthy volunteers. Radiology. 2005;236:1047–1052. 49. Peters DC, Botnar RM, Kissinger KV, Yeon SB, Appelbaum EA, Manning WJ. Inversion recovery radial MRI with interleaved projection sets. Magn Reson Med. 2006;55:1150–1156. 50. Sodickson DK, Manning WJ. Simultaneous acquisition of spatial harmonics (SMASH): fast imaging with radiofrequency coil arrays. Magn Reson Med. 1997;38:591–603. 51. Carlson JW. An algorithm for NMR imaging reconstruction based on multiple RF receiver coils. J Magn Reson. 1987;74:376–380. 52. Hutchinson M, Raff U. Fast MRI data acquisition using multiple detectors. Magn Reson Med. 1988;6:87–91.

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

79. Tsao J, Kozerke S, Boesiger P, Pruessmann KP. Optimizing spatiotemporal sampling for k-t BLAST and k-t SENSE: application to highresolution real-time cardiac steady-state free precession. Magn Reson Med. 2005;53:1372–1382. 80. Kozerke S, Tsao J, Razavi R, Boesiger P. Accelerating cardiac cine 3D imaging using k-t BLAST. Magn Reson Med. 2004;52:19–26. 81. Madore B. Using UNFOLD to remove artifacts in parallel imaging and in partial-Fourier imaging. Magn Reson Med. 2002;48:493–501. 82. Priatna A, Zhu D. Double/triple IR dual contrast FSE of the heart with ASSET. In: Proceedings of the 11th Scientific Meeting of the International Society of Magnetic Resonance in Medicine. Toronto, Ontario, Canada: 2003:1562. 83. Gutberlet M, Noeske R, Schwinge K, Freyhardt P, Felix R, Niendorf T. Comprehensive cardiovascular MR imaging at 3.0 Tesla: feasibility and implications for clinical applications. Invest Radiol. 2006;41. 84. Busse RF. Reduced RF power without blurring: correcting for modulation of refocusing flip angle in FSE sequences. Magn Reson Med. 2004;51:1031–1037. 85. Henning J, Scheffler K. Hyperechoes. Magn Reson Med. 2001;46: 6–12. 86. Hennig J, Weigel M, Thiel T. Optimizing SAR-reduction for high-field TSE with asymmetric hyperechoes combined with partial Fourier parallel imaging. In: Proceedings of the 12th Meeting of the International Society of Magnetic Resonance in Medicine. Kyoto, Japan: 2004:539. 87. Zhu Y, Hardy CJ, Sodickson DK, et al. Highly parallel volumetric imaging with a 32-element RF coil array. Magn Reson Med. 2004;52:869–877. 88. Sodickson DK, Hardy CJ, Zhu Y, et al. Rapid volumetric MRI using parallel imaging with order-of-magnitude accelerations and a 32element RF coil array: feasibility and implications. Acad Radiol. 2005;12:626–635. 89. Noeske R, King KF, Friedrich MG, Niendorf T. Assessment of cardiac function using 2D CINE FIESTA imaging in combination with parallel imaging. In: Proceedings of the 10th Scientific Meeting of the International Society of Magnetic Resonance in Medicine. Honolulu, Hawaii, USA: 2002:1659. 90. Noeske R, Messroghli D, Friedrich MG, Niendorf T. 2D CINE FIESTA imaging for the assessment of cardiac function in combination with parallel imaging: a comparison study. In: Proceedings of the 11th Scientific Meeting of the International Society of Magnetic Resonance in Medicine. Toronto, Ontario, Canada: 2003:1563. 91. Wintersperger BJ, Nikolaou K, Dietrich O, et al. Single breath-hold real-time cine MR imaging: improved temporal resolution using generalized autocalibrating partially parallel acquisition (GRAPPA) algorithm. Eur Radiol. 2003;13:1931–1936. 92. Rettmann DW, Saranathan M, Wu KC, Azevedo CF, Bluemke DA, Foo TK. High temporal resolution breath-held 3D FIESTA CINE: validation of ventricular function in patients with chronic myocardial infarction. In: Proceedings of the 12th Scientific Meeting of the International Society of Magnetic Resonance in Medicine. Kyoto, Japan: 2004:3823. 93. Weiger M, Pruessmann KP, Boesiger P. Cardiac real-time imaging using SENSE: SENSitivity Encoding scheme. Magn Reson Med. 2000;43:177–184. 94. Pruessmann KP, Weiger M, Boesiger P. Sensitivity encoded cardiac MRI. J Cardiovasc Magn Reson. 2001;3:1–9. 95. Zhang Q, Park J, Li D, Simonetti OP. Improving True-FISP Parallel CINE imaging using a new data acquisition scheme for coil sensitivity calibration. In: Proceedings of the 11th Scientific Meeting of the International Society of Magnetic Resonance in Medicine. Toronto, Ontario, Canada: 2003:2329. 96. McVeigh ER. Regional myocardial function. Cardiol Clin. 1998; 16:189–206. 97. Herzka DA, Derbyshire A, Kellman P, McVeigh ER. Myocardial tagging in a single heart-beat with EPI-SSFP and TSENSE. In: Proceedings of the 11th Scientific Meeting of the International Society of Magnetic Resonance in Medicine. Toronto, Ontario, Canada: 2003:374.

56 Cardiovascular Magnetic Resonance

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 mag netic resonance (CMR). The use of a gadolinium based con trast 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 myocar dial 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 compen sate under resting conditions for up to 80% diameter coro nary 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 sig nificance of a lesion is most apparent during maximal vaso dilation.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 angi ography 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 perfu sion 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 perfu sion imaging detects subclinical disease and silent ischemia in subjects without a history or symptoms of atheroscle rotic 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 increas ing 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 indi cators of regional myocardial ischemia. Blood flow across the myocardial wall is not uniform, but instead favors the suben docardium 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 approxi mately 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 sus ceptible 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

4 MYOCARDIAL PERFUSION IMAGING THEORY

CHAPTER 4

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

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 quanti tatively 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 abnormal ities starts with an understanding of the contrast mechan isms 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 inhomogene ities and also shortens the T2* relaxation rate of water. The contrast reagent also shortens the T1 magnetization recov ery after a radiofrequency pulse has disturbed the equilib rium state. The CMR methods for perfusion imaging can be subdi vided 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 vascu lar 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 gradi ent echo imaging with short echo times and a magnetiza tion preparation for optimal T1 weighting have therefore steadily gained preference. This chapter provides a brief overview of three major sequence techniques for T1 weighted 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 acqui sition only needs to be delayed by 10 to 20 msec after a sat uration 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 relaxa tion rate constant. (The relaxivity of the contrast reagent is not appreciably different between blood and tissue.) Even tually, 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 exci tation 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 effec tive 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 arte rial transit of the contrast agent. To partially capture the speed advantage of EPI and to minimize T2* and flow related artifacts, one reduces the length of the echo trains

250 200

100 LV RV

50

Signal intensity (a.u.)

150

Tissue 0

0

10

20

30

(IV contrast injection)

40

50

60

70

80

90

Time (sec)

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 par ticular, 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 inten sity 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 orienta tions. 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 reso lution.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

4 MYOCARDIAL PERFUSION IMAGING THEORY

300

LV blood pool

1000 800 Signal (a.u.)

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

PARALLEL IMAGING AND ACCELERATION TECHNIQUES FOR PERFUSION CARDIOVASCULAR MAGNETIC RESONANCE

COMPARISON OF LV CURVES WITH 0.1 MMOL/KG BOLUS OF GD-DTPA

T1-FARM Fast gradient echo

600 400 200 0 –200 0

10

20

30

40

50

Time (sec)

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 2.4/1.2 msec; flip angle 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 tech niques, 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 resolu tion than is available with nuclear imaging techniques. An illustration is shown in Figure 4 3 from studies in an experimental animal model.

RV

Parallel imaging refers to the use of independent receiver chan nels for the parallel acquisition of data from multiple elements of a receiver coil array (see Chapter 3). Parallel imaging tech niques 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 encod ing. By measuring the local coil profiles, one can reduce the number of phase encoding steps and during image recon struction replace the missing phase encodings with informa tion 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 encod ing direction. In practice, and in particular for imaging techni ques 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

ANT

RV

CMR (1st pass)

FDG-PET

TTC-stained

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 con trast agent is detected indirectly through the change in the 1 H relaxation recovery rates, T1 and T2. The effect of a con trast agent on T1 and T2 within a single uniform closed space can be described by its relaxivity, r1,2, using the follow ing equation: T1;2 ¼ r1;2  ½CR. Here, [CR] denotes the con centration 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 pre dominant 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 cap illary barrier and the cytolemmal membrane, if sufficiently fast, allows 1H spins to sample different environments dur ing 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 vascu lar 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 intrin sic relaxation rates. The limiting cases of no exchange and fast exchange pro vide 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 overesti mation 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, inter mediate 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 esti mate 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. Never theless, 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 inter pretation 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 myo cardial perfusion that are not based on the use of an exoge nous CMR contrast agent, but instead exploit endogenous contrast mechanisms related to blood flow and blood Cardiovascular Magnetic Resonance 61

4 MYOCARDIAL PERFUSION IMAGING THEORY

Another remarkable advance for fast CMR perfusion imag ing is the development of sparse sampling techniques for dynamic imaging applications. An example is the unaliasing by Fourier encoding the overlaps using the temporal dimen sion (UNFOLD) technique, which involves a trade off of tem poral 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 comple mentary 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 acqui sition 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) con trast,45,46 and magnetization transfer contrast.47,48 These techniques have to be considered technically more challeng ing 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 assump tion 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 trans ported 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 sat uration that reflects regional oxygen supply and demand (see Chapter 42). Deoxyhemoglobin is paramagnetic, whereas oxyhemoglobin is only diamagnetic. This means that deoxy hemoglobin 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 trans fer 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 myo cardium 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 eval uated based on the level of differential signal enhancement in the myocardium. With a T1 weighted imaging tech nique, myocardial segments showing reduced signal inten sity enhancement during the first pass, relative to other myocardial segments, are interpreted as hypoperfused. This type of qualitative judgment of signal enhancement differ ences 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 microvas cular 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. Abso lute 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 rela tive to the enhancement in the blood pool, both measured on the same linear scale. In fact, for the purpose of quantify ing perfusion, the tissue is often viewed as a linear, station ary 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 myo cardial perfusion imaging, including absolute quantifica tion of myocardial blood flow.53 With CMR, a quantitative analysis of contrast enhancement is based on signal inten sity 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 spill over effects. Figure 4 4 shows an example of a CMR perfu sion 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 inter est 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

(1)

0

F represents the flow blood (e.g., in units of milliliters per minute per gram of tissue). Zierler showed that if the transport within a tissue region of interest is assumed to be linear and stationary, then the response to an arbitrary arterial input is given by the convolution integral of this input with the tissue impulse response57: ðt qðtÞ ¼ cin ðt  tÞ  RðtÞdt

(2)

0

The impulse response is denoted by R(t). The response in a tissue region to an impulse input, the tissue impulse response, has the useful property that its initial amplitude equals the myocardial blood flow through that region. This can be derived by considering the special case in which the

Myocardial SI (a.u.)

RV

100

50 0.70 mL/g/min (anteriorseptal)

0

A

10

20

30

40

50

60

50

60

Time (sec)

RV

LV SI (a.u.)

600 400 200 0

B

10

C

20

30

40

Time (sec)

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 com partmental 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 deconvolu tion 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 defini tion of the impulse input, no additional tracer enters the tissue region after the initial delta function input. Further more, 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 approxi mation of the shape of the impulse response of an intravas cular 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 approxi mated 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 compari son to blood flow measurements with labeled micro spheres59,60 and also by comparison to invasive coronary flow reserve measurements,42 with similar regional perfu sion 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 tis sue. 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

4 MYOCARDIAL PERFUSION IMAGING THEORY

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 para meters need to be adjusted for a best fit of the measured resi due 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 quanti fication of physiologic changes observed indirectly with

BLOOD-TISSUE EXCHANGE UNIT

Blood plasma flow (Fp)

Axial direction (x)

PScap

Plasma volume = Vp

Interstitial space Parenchymal cell

A

Signal intensity (a.u.)

50 LV input

40 30

4 mL/g/min

20 10 1 mL/g/min

0

0

10

20

30

B

40

50

60

70

Time (sec) 200

Signal intensity (a.u.)

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

well mixed within each compartment, meaning that the con trast 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 dif ferential equations that account for the variation of tracer con centrations within each tissue region as a function of time and at least one spatial variable.62,63 The number of parameters in

5 mL/g/min 150

Fp ~ impulse amplitude

4 mL/g/min

100

2.7 mL/g/min 2 mL/g/min

50 1 mL/g/min 0 0

C

10

20

30

40

50

60

70

Time (sec)

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 condi tions, 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 con straint in the limited dynamic range of contrast enhancement, which results in signal saturation at higher contrast concen trations, 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 precon trast 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 con trast enhancement of the delay time after a saturation prepa ration. 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 dos age, each optimized to measure contrast enhancement in the myocardium and the blood pool, respectively, but with other wise identical injection protocols and under stable hemody namic 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 mini mal signal saturation.65,66 Such pulse sequence techniques with arterial input scout impose a small time penalty, but oth erwise 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 satu ration,67 as long as the calibration curves are monotonically increasing as a function of R1. An example of calibration based saturation correction appears in Figure 4 6C.

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

1000

TD = 10

800

2000 1000 0 0

200 400 600 800 1000 1200 1400

B 400

Observed CE (%)

TD = 300 ms 1000

200 0 0

A

3000

TD = 100

600

CE (%)

Contrast enhancement (CE) (%)

1200

Expected CE (%)

MODEL-BASED CALIBRATION CURVE

2

4

6

8

10

500 0

Measured LV CE Corrected LV CE

500

R1 (s)

0

C

10

20

30

40

50

Image #

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, Model based 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 per fusion 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, hyperpo larized contrast media68 could be one possible approach because it is relatively easy to selectively quench the hyper polarized 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 avoid ing 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 reso lution requirements during stress compared with rest. 3. Spatial resolution. Mild reductions of myocardial blood flow as a result of an epicardial lesion manifest them selves 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 tempo ral resolution. 4. The CMR sequence for contrast residue detection. T1 weighted 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 arti facts 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 arti fact, 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 per fusion 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 recom mended, although higher factors may work well when SSFP techniques are used. 7. Analysis and modeling. The rate of contrast enhance ment 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 enhance ment, 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 myo cardial 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 lim itations that may be limited to subendocardium or that result in only moderate blood flow reductions. For these applica tions, quantitative CMR methods have been successfully developed and validated to provide a true measure of myo cardial blood flow.

1. Stollfuss JC, Kosa I, Ziegler S, Schwaiger M, et al. Quantitative assessment of myocardial perfusion: is it of clinical relevance? Q J Nucl Med. 1996;40(1):76–84. 2. Schwitter J, Wacker CM, Wilke N, et al. Magnetic resonance Imaging for Myocardial Perfusion Assessment in Coronary artery disease Trial (MR-IMPACT II): a phase III multicenter, multivendor trial comparing perfusion cardiac magnetic resonance versus single photon emission computed tomography for the detection of coronary artery disease. Circulation. 2006;114(18):II–806. 3. Gould KL, Kirkeeide RL, Buchi M. Coronary flow reserve as a physiologic measure of stenosis severity. J Am Coll Cardiol. 1990;15(2): 459–474. 4. Gould KL, Lipscomb K. Effects of coronary stenosis on coronary flow reserve and resistance. Am J Cardiol. 1974;34:48–55. 5. Wilson RF, Wyche K, Christensen BV, et al. Effects of adenosine on human coronary arterial circulation. Circulation. 1990;82(5): 1595–1606. 6. Wilson RF. Assessment of the human coronary circulation using a Doppler catheter. Am J Cardiol. 1991;67(14):44D–56D. 7. Pijls NH, De Bruyne B, Peels K, et al. Measurement of fractional flow reserve to assess the functional severity of coronary-artery stenoses [see comments]. N Engl J Med. 1996;334(26):1703–1708. 8. Panting JR, Gatehouse PD, Yang GZ, et al. Abnormal subendocardial perfusion in cardiac syndrome X detected by cardiovascular magnetic resonance imaging. N Engl J Med. 2002;346(25):1948–1953. 9. Wilke N, Jerosch-Herold M, Wang Y, et al. Myocardial perfusion reserve: assessment with multisection, quantitative, first-pass MR imaging. Radiology. 1997;204(2):373–384. 10. Wang L, Jerosch-Herold M, Jacobs DR, et al. Coronary risk factors and myocardial perfusion in asymptomatic adults: the Multi-Ethnic Study of Atherosclerosis (MESA). J Am Coll Cardiol. 2006;47(3):565–572. 11. Wang L, Jerosch-Herold M, Jacobs DR, et al. Coronary artery calcification and myocardial perfusion in asymptomatic adults the MESA (Multi-Ethnic Study of Atherosclerosis). J Am Coll Cardiol. 2006;48 (5):1018–1026. 12. Rosen BD, Lima JA, Nasir K, et al. Lower myocardial perfusion reserve is associated with decreased regional left ventricular function in asymptomatic participants of the multi-ethnic study of atherosclerosis. Circulation. 2006;114(4):289–297. 13. Rosen BD, Edvardsen T, Lai S, et al. Left ventricular concentric remodeling is associated with decreased global and regional systolic function: the Multi-Ethnic Study of Atherosclerosis. Circulation. 2005; 112(7):984–991. 14. Decking UK, Schrader J. Spatial heterogeneity of myocardial perfusion and metabolism. Basic Res Cardiol. 1998;93(6):439–445. 15. Arai AE, Grauer SE, Anselone CG, et al. Metabolic adaptation to a gradual reduction in myocardial blood flow. Circulation. 1995;92 (2):244–252. 16. Bache RJ, Schwartz JS. Effect of perfusion pressure distal to a coronary stenosis on transmural myocardial blood flow. Circulation. 1982;65 (5):928–935. 17. Canty IJ, Piotrowski B, Klocke F. Relative reductions in subendocardial and full-thickness flow during pharmacologic vasodilation as an index of functional stenosis seventy. J Am Coll Cardiol. 1990;15:182A. 18. 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. 19. Muehling OM, Panse P, Jerosch-Herold M, et al. Cardiac magnetic resonance perfusion imaging identifies transplant arteriopathy by a reduced endomyocardial resting perfusion. J Heart Lung Transplant. 2005;24(8):1122–1123. 20. Wagner A, Mahrholdt H, Holly JA, 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(9355):374–379. 21. Bolli R, Zhu W-X, Thornby JI, et al. Time course and determinants of recovery of function after reversible ischemia in conscious dogs. Am J Physiol. 1988;254:H102–H114. 22. Tarantini G, Cacciavillani L, Corbetti F, et al. Duration of ischemia is a major determinant of transmurality and severe microvascular obstruction after primary angioplasty: a study performed with

23. 24. 25. 26. 27. 28. 29. 30. 31. 32. 33. 34. 35. 36.

37.

38.

39. 40.

41. 42. 43. 44. 45.

contrast-enhanced magnetic resonance. J Am Coll Cardiol. 2005;46 (7): 1229–1235. Epstein FH, Wolff SD, Arai AE. Segmented k-space fast cardiac imaging using an echo-train readout. Magn Reson Med. 1999;41(3): 609–613. Li W, Storey P, Chen Q, et al. Dark flow artifacts with steady-state free precession cine MR technique: causes and implications for cardiac MR imaging. Radiology. 2004;230(2):569–575. Hunold P, Maderwald S, Eggebrecht H, et al. Steady-state free precession sequences in myocardial first-pass perfusion MR imaging: comparison with TurboFLASH imaging. Eur Radiol. 2004;14(3):409–416. Chen Z, Prato FS, McKenzie CA. T1 Fast Acquisition Relaxation Mapping (T1-FARM): An Optimised Reconstruction. IEEE Trans Med Imag. 1998;17:155–160. Bellamy DD, Pereira RS, McKenzie CA, et al. Gd-DTPA bolus tracking in the myocardium using T1 fast acquisition relaxation mapping (T1 FARM). Magn Reson Med. 2001;46(3):555–564. Pruessmann KP, Weiger M, Scheidegger MB, Boesiger P. SENSE: sensitivity encoding for fast MRI. Magn Reson Med. 1999;42(5): 952–962. Pruessmann KP, Weiger M, Boesiger P. Sensitivity encoded cardiac MRI. J Cardiovasc Magn Reson. 2001;3(1):1–9. Sodickson DK, Manning WJ. Simultaneous acquisition of spatial harmonics (SMASH): fast imaging with radiofrequency coil arrays. Magn Reson Med. 1997;38(4):591–603. Niendorf T, Sodickson DK. Parallel imaging in cardiovascular MRI: methods and applications. NMR Biomed. 2006;19(3):325–341. Madore B. Using UNFOLD to remove artifacts in parallel imaging and in partial-Fourier imaging. Magn Reson Med. 2002;48(3):493–501. Madore B. UNFOLD-SENSE: a parallel MRI method with self-calibration and artifact suppression. Magn Reson Med. 2004;52(2):310–320. Di Bella EV, Wu YJ, Alexander AL, et al. Comparison of temporal filtering methods for dynamic contrast MRI myocardial perfusion studies. Magn Reson Med. 2003;49(5):895–902. Hansen MS, Battes C, Tsao J, et al. Accelerated dynamic Fourier velocity encoding by exploiting velocity-spatio-temporal correlations. MAGMA. 2004;17(2):86–94. Tsao J, Kozerke S, Boesiger P, Pruessmann KP. Optimizing spatiotemporal sampling for k-t BLAST and k-t SENSE: application to highresolution real-time cardiac steady-state free precession. Magn Reson Med. 2005;53(6):1372–1382. Donahue KM, Weisskoff RM, Chester 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(6):858–867. Wiart M, Carme S, Mai W, et al. In vivo quantification of regional myocardial blood flow: validity of the fast-exchange approximation for intravascular T1 contrast agent and long inversion time. Magn Reson Med. 2006;56(2):340–347. Donahue KM, Burstein D, Manning WJ, Gray ML, et al. Studies of GdDTPA relaxivity and proton exchange rates in tissue. Magn Reson Med. 1994;32:66–76. Wacker CM, Wiesmann F, Bock M, et al. Determination of regional blood volume and intra-extracapillary water exchange in human myocardium using Feruglose: first clinical results in patients with coronary artery disease. Magn Reson Med. 2002;47(5):1013–1016. Donahue KM, Weisskoff RM, Bogdonoff AA, et al. Measurement of vascular volume fraction and water permeability with intravascular contrast agents. In: Third Scientific Meeting. Nice: SMR; 1995. Jerosch-Herold M, Wilke N, Stillman AE, Wilson RF. MR quantification of the myocardial perfusion reserve with a Fermi function model for constrained deconvolution. Med Phys. 1998;25(1):73–84. Li X, Huang W, Yankeelov TE, et al. Shutter–speed analysis of contrast reagent bolus–tracking data: preliminary observations in benign and malignant breast disease. Magn Reson Med. 2005;53(3):724–729. Fidler F, Wacker CM, Dueren C, et al. Myocardial perfusion measurements by spin-labeling under different vasodynamic states. J Cardiovasc Magn Reson. 2004;6(2):509–516. Fieno DS, Shea SM, Li Y, et al. Myocardial perfusion imaging based on the blood oxygen level-dependent effect using T2-prepared steady-state free-precession magnetic resonance imaging. Circulation. 2004;110 (10):1284–1290.

Cardiovascular Magnetic Resonance 67

4 MYOCARDIAL PERFUSION IMAGING THEORY

References

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

46. Klocke FJ, Li D. Testing coronary flow reserve without a provocative stress: a “BOLD” approach. J Am Coll Cardiol. 2003;41(5):841–842. 47. Tsekos NV, Wilke N, Xu Y, et al. Magnetization transfer contrast in assessment of myocardial perfusion. Radiology. 1993;189(P):166. 48. Wolff SD, Balaban RS. Magnetization transfer contrast (MTC) and tissue water proton relaxation in vivo. Magn Reson Med. 1989;10:135–144. 49. Detre JA, Zhang W, Roberts DA, et al. Tissue specific perfusion imaging using arterial spin labeling. NMR Biomed. 1994;7(1–2):75–82. 50. Wacker CM, Fidler F, Dueren C, et al. Quantitative assessment of myocardial perfusion with a spin-labeling technique: preliminary results in patients with coronary artery disease. J Magn Reson Imaging. 2003;18(5):555–560. 51. Zhang H, Shea SM, Park V, et al. Accurate myocardial T1 measurements: toward quantification of myocardial blood flow with arterial spin labeling. Magn Reson Med. 2005;53(5):1135–1142. 52. Bauer WR, Nadler W, Bock M, et al. The relationship between the BOLD-induced T(2) and T(2)(*): a theoretical approach for the vasculature of myocardium. Magn Reson Med. 1999;42(6):1004–1010. 53. Schwaiger M. Myocardial perfusion imaging with PET. J Nucl Med. 1994;35(4):693–698. 54. Jerosch-Herold M, Swingen C, Seethamraju RT. Myocardial blood flow quantification with MRI by model-independent deconvolution. Med Phys. 2002;29(5):886–897. 55. Zierler K. Indicator dilution methods for measuring blood flow, volume, and other properties of biological systems: a brief history and memoir. Ann Biomed Eng. 2000;28(8):836–848. 56. Zierler KL. Theoretical basis of indicator-dilution methods for measuring flow and volume. Circ Res. 1962;10:393–407. 57. Zierler KL. Equations for measuring blood flow by external monitoring of radioisotopes. Circ Res. 1965;16:309–321. 58. Axel L. Tissue mean transit time from dynamic computed tomography by a simple deconvolution technique. Invest Radiol. 1983;18(1):94–99. 59. Christian TF, Rettman DW, Aletras AH, et al. Absolute myocardial perfusion in canines measured by using dual-bolus first-pass MR imaging. Radiology. 2004;232(3):677–684.

68 Cardiovascular Magnetic Resonance

60. Muhling OM, Wang Y, Panse P, et al. Transmyocardial laser revascularization preserves regional myocardial perfusion: an MRI first pass perfusion study. Cardiovasc Res. 2003;57(1):63–70. 61. Hsu LY, Rhoads KL, Holly JE, et al. Quantitative myocardial perfusion analysis with a dual-bolus contrast-enhanced first-pass MRI technique in humans. J Magn Reson Imaging. 2006;23:315–322. 62. Bassingthwaighte JB, Raymond GR, Chan JIS. Principles of tracer kinetics. In: Zaret BL, Beller GA, eds. Nuclear Cardiology: State of the Art and Future Directions. St. Louis: Mosby-Year Book; 1993:3–23. 63. Kroll K, Wilke N, Jerosch-Herold M, et al. Accuracy of modeling of regional myocardial flows from residue functions of an intravascular indicator. Am J Physiol (Heart Circ Physiol). 1996;40:H1643–H1655. 64. Jerosch-Herold M, Wilke N, Wang Y, et al. Direct comparison of an intravascular and an extracellular contrast agent for quantification of myocardial perfusion. Int J Card Imaging. 1999;15(6):453–464. 65. Ko¨stler H, Ritter C, Baunach D, et al. Determination of the arterial input function in high dose radial myocardial perfusion imaging. In: Scientific Sessions of the International Society of Magnetic Resonance in Medicine. Miami Beach, FL: ISMRM; 2005. 66. Gatehouse PD, Elkington AG, Ablitt NA, et al. Accurate assessment of the arterial input function during high-dose myocardial perfusion cardiovascular magnetic resonance. J Magn Reson Imaging. 2004;20 (1):39–45. 67. Cernicanu A, Axel L. Theory-based signal calibration with singlepoint T1 measurements for first-pass quantitative perfusion MRI studies. Acad Radiol. 2006;13(6):686–693. 68. Johansson E, Mansson S, Wirestam R, et al. Cerebral perfusion assessment by bolus tracking using hyperpolarized 13C. Magn Reson Med. 2004;51(3):464–472. 69. 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.

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

Left ventricular (LV) diastolic function has been recog nized 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 abnor mal 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 con sidered 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) tech niques, three dimensional (3D) data collection, spiral and steady state free precession (SSFP) imaging, and higher magnetic field strength, the prolonged tag per sistence permits easier access to diastole. Together with rapid state of the art software analysis tools, quantifi cation 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 counter clockwise 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 accompa nied by a rapid untwisting at the apex, whereas the vol ume 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 rel atively poor internal structure because of the absence of structural landmarks, making quantification of parameters, such as rotation, stress, and strain, limited.4,5 Several inva sive 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 transplan tation 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 land marks and their irregular geometric distribution. Further more, 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 struc ture of the myocardium. However, a CMR myocardial tag ging 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

5 CARDIOVASCULAR MAGNETIC RESONANCE TAGGING ASSESSMENT OF LEFT VENTRICULAR DIASTOLIC FUNCTION

CHAPTER 5

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

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 reli ably 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 acqui sitions, 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 modula tion 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 pro longs 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 mag netization 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 projec tion 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 infor mation and a second part (responsible for the fading of the tags) built up as a function of time. For CSPAMM, this sec ond component is also suppressed by the subtraction pro cedure. 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 con trast 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 myo cardium 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 min imum. 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 appro priate, 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 informa tion 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

t=405 msec

t=442 msec

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 blur ring in the images, spectral spatial RF excitations15 are a stringent requirement.14 Together with echoplanar and spi ral 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 seg mented 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 assess ment of myocardial motion based on 3D CSPAMM lattice tagging was reported by Ryf and colleagues.18 Simulta neously, the same authors proposed an extension to a pre viously 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 coa ched 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, sophis ticated algorithms that reduce quantification of tagged time series of cardiac images to seconds are now readily available,19 and variants have been successfully imple mented.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 myo cardium (Fig. 5 4) are defined and motion specific para meters (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

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

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 res olution 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 api cal 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 rota tion 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. Untwist ing 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 architec ture, 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 dysfunc tion 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 end systolic 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 imag ing, the number of RF excitations for imaging can be mini mized. 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 algo rithm (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 differ ence 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 with out 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 quantifi cation of diastolic function remains to be more fully investi gated compared with gold standard techniques. Currently, the CMR technique is not widely accessible to clinical cardiol ogists 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 accessi bility 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 fol lowing 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 sub jects’ 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.

Cardiovascular Magnetic Resonance 75

5 CARDIOVASCULAR MAGNETIC RESONANCE TAGGING ASSESSMENT OF LEFT VENTRICULAR DIASTOLIC FUNCTION

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.

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

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 reso nance (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 advance ments in molecular imaging agents, including compounds that detect pH changes, enzymatic activity, specific biomol ecules 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 indi cations 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 chap ter 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 relaxa tion, 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 tis sue, 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 conven tional units for r1 are mM-1s-1 (per millimolar per second, sometimes Lmol-1s-1). Thus, the slope of 1/T1 as a func tion 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 depen dent 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 nonlin ear. 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 con trast 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 ver sus 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 neg ative contrast can also be medically useful, and certain con trast 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 1200 msec at 1.5 T) and short T1 tissue (dashed line, T1 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, gadopente tate 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] 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 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. Chem ically, 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., gadopente tate 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 pro files, 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 (neu tral) 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 injec tion 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 docu mented safety benefit in using nonionic CMR contrast agents. One benefit of the nonionic compounds is the ability to formu late them at high concentration (1 M)10 without drastically

increasing osmolality or viscosity (see Table 6 1). These high concentration 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 draw back 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 capa ble 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 steady state 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 develop ment. 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 (dis cussed later). Gd BOPTA (gadobenate; see Fig. 6 4) has weak affinity15,16 for albumin (10% bound), which leads to a mod est 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 frac tion 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 frac tion. As the concentration of the contrast agent begins to stabilize (at 0.5 mM) the fraction bound will become con stant. The observed relaxivity will depend on the fraction bound; unlike ECF agents, T1 change in plasma is not lin early 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 steady state 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 gado linium covalently linked to macromolecules, such as poly lysine, 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 glob ular. 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 reticu loendothelial 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 relax ation properties (smaller r2/r1) and can be used for bright blood 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 con trast 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 catalyti cally 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 mag netic 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 differ ent 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 repre sents the number of water molecules bound to the metal ion (typ ically 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 effec tively 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 mat ter 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 mecha nism. 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 to proton distance, tc is the correlation time, and oH is the hydrogen Larmor frequency. The product S(Sþ1) is proportional to the mag netic moment. All other factors being equal, the higher the mag netic moment, the more efficient the relaxation. This is why Gd3þ (S ¼ 7/2) is preferred to copper (Cu2þ, S ¼ 1/2). The dipo lar effect depends on the distance between the ion and the hydro gen 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 rota tional diffusion of the molecule, tR, or the electronic T1 of the metal ion, T1e. In the MR experiment, nuclei are excited by apply ing 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 ther mal 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 Lar mor 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 pro tein targets, then their relaxivity would be increased on binding because the contrast agent would take on the rotational character istics of the protein (receptor induced magnetization enhance ment). 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 relax ivity 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 relax ivity. 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 con taining 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 config urations (half filled f and d shells, respectively), and as a result, electronic relaxation is relatively slow. Other ions, such as dyspro sium (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 magni tude 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 encoun tered 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 compo nent to relaxivity11 and the correlation times that govern outer sphere 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 slow tumbling 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 mole cules, 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 magne tism persists outside the external magnetic field. A weaker form of this is superparamagnetism, in which small particles of iron ox ides with aligned spins occur in a magnetic field. Because the par ticles 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 submi cron 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 admi nistered intravenously. The difference in particle size determines their pharmacokinetic behavior. There are no inner sphere water mole cules 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 fea ture 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 magneti zation 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 aggrega tion 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 relax ivity 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 suscep tibility effects and the SPIOs can act as T2* relaxation agents. This has important consequences when considering the effects of con trast 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 Gado mer or albumin bound MS 325, have several fold higher relaxivities than the ECF agents. As Figure 6 7 suggests, the slow tumbling com pounds 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 10 nsec, solid line), typical of albumin binding, gives high r1 that decreases with increasing field and high r2; intermediate correlation time (tR 1 nsec, short dashed line) shows relaxivity maximum pushed out to a higher field; short correlation time (tR 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 com plex. 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 MS 325 in buffer solution and in serum albumin solution. 2. Compartmentalization of the metal complex in tissue. Generally, tissue water is compartmentalized into intra vascular, interstitial (fluid space between cells and capil laries), and intracellular space, constituting roughly 5%, 15%, and 80% of total water, respectively. Cellular orga nelles 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 encoun tering 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 reduc tion in apparent T2. The result of the first two effects is that the simple relax ivity 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 concentra tion 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 [gadofosve set], 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 aque ous 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 microvascula ture; most of the water in the tissue does not have the opportunity to be relaxed by the Gd within TR of an imag ing 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 com partments 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 phe nomenon, 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 com partments 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 compart ments 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 compart ments. 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 micro scopic 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 compart ment 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 bio logic water mobility on contrast enhanced scans.37 Biologic compartmentalization also results in susceptibil ity contrast. The contrast agent causes microscopic field inhomogeneities, sometimes called “mesoscopic” inhomo geneities.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 dis tributes to extracellular space in the same manner as Gd DTPA. The higher relaxivity can result in greater conspicu ity and higher reader confidence for detecting lesions.39 Similarly, Gd BOPTA has shown efficacy in contrast enhanced 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 characteriza tion 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 Mag netics, 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 applica tions and is approved for use in Europe, Canada, and Austra lia, 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 imag ing 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 characteriza tion 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 con trast 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 com ponent of thrombus. These molecules contain a peptide sequence specific for fibrin that has been linked to several gadolinium che lates. 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 throm bus visualization with the experimental fibrin targeted agent EP 2104R 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 functiona lized 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 angio genesis.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 biodistribu tion 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 trans port 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 pathol ogies with elevated collagen levels. The Massachusetts General Hospital group developed a gado linium 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 suscep tibility (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 gadolinium based agent, EP 2104R. (A) Double oblique slice from bright blood 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 alter ing 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 parti cles as targeted contrast agents. Commercial USPIOs passively accumulate in macrophages. This property has been exploited to provide negative contrast in macrophage rich atherosclerotic pla ques. There appears to be a positive correlation between macro phage 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 mesen chymal 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 therapeu tic intervention of mouse embryonic stem cells could be followed by visualizing the stem cells and following their effect on LV func tion 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 apopto tic cardiomyocytes in a mouse model of transient left anterior des cending 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 prepara tion). This view has changed with the apparent link between Gd contrast and nephrogenic systemic fibrosis (NSF). A very rare, but potentially devastating condition affect ing patients with chronic kidney disease, NSF has an esti mated prevalence of more than 10% among patients receiving Gd who are on hemodialysis.93 Although present ing primarily with skin thickening, tethering, hyperpig mentation, and disabling joint flexion contractures, patients with NSF can also have multiorgan/systemic fibro sis, 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 glomer ular 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 che lator, 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 thermody namic 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) > Gd HPDO3A (gadoteridol)  Gd DO3A butrol (gadobutrol)  MS 325 (gadofosveset)  Gd BOPTA (gadobenate)  Gd DTPA (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 Gd DOTA (gadoterate), Gd HPDO3A (gadoteridol), and Gd DO3A butrol (gadobutrol) were the most inert with respect to Gd loss. The compound that released Gd the most rap idly was Gd DTPA BMA (gadoversetamide) under the con ditions 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 impor tant 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], Gd BOPTA [gadobenate], Gd DTPA BMA [gadoversetamide], Gd DTPA BMEA [gadodiamide), Gd DO3A butrol [gadobu trol], and Gd BOPTA [gadobenate]) appeared indistinguish able 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 approxi mately 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 hyperpolar ized 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 approxi mately 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 excel lent sensitivity with no background (high signal to noise ratio). Challenges include the distribution and availability of the hyper polarization equipment and imaging hardware compatibility for imaging nonhydrogen nuclei (not available on all clinical scan ners). 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 threat ening 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 DTPA BMEA (gadodiamide) often show spuriously low serum cal cium 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 com pounds. 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 signifi cant number of new agents, including blood pool and tis sue 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 param eter, 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 applica tions. 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

16.

17. 18. 19. 20. 21.

22.

23. 24. 25.

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.

Cardiovascular Magnetic Resonance 89

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.

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

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.

90 Cardiovascular Magnetic Resonance

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 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 stud ies.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 over view 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 satura tion 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 mate rial (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 of flight 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 time of 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 measure ment 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 associ ates,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 transfor mation 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 gra dient 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 gradi ent waveforms in the direction of desired flow measure ment. The difference in the waveforms is calculated to produce a well defined velocity related phase difference between the two images. A phase reconstruction is pro duced for each of the images, which are then subtracted pixel by pixel to produce the final velocity map. This pro cess 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 veloc ity 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 ambig uous 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 col leagues24 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 col umn 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 over all accuracy of the method is to adjust the velocity sensi tivity 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

7 BLOOD FLOW VELOCITY ASSESSMENT

45 150 msec

AA

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

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.

A

B

C

D

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 sug gested by Bittoun and associates.32 The method is a combina tion 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, particu larly 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 grad ually 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 cor rected precisely and automatically in software, with no user intervention required.38 A second common reason for back ground 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 ran dom phase or show the phase of a weak ghost that may not be visible on the magnitude image with normal bright ness 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 pix els 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 catheteriza tion.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 ventricu lar 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 lung volume 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 para meters, 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 cover age methods, such as interleaved spirals, will be important. Ultra fast flow imaging techniques have also been devel oped, 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

Magnitude Velocity

600

Aortic flow

400

Diaphragm position

20 Expiration

10 0

200

(mm)

(mm)/sec

–10

SVC flow

Inspiration

–20 –30

0 1

4

7

10

13

16

19

22

Cardiac cycles

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

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

spiral imaging,50,51 or by imaging only one spatial dimen sion.52 A compromise generally has to be made in temporal or spatial resolution and probably also in the signal to noise 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).

0

168 cm/sec

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 repre sentation would be to use the cine velocity images to calcu late 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 informa tion 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

66 (msec)

A

162 (msec)

B

Pressure

7 BLOOD FLOW VELOCITY ASSESSMENT

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.

Flow Pattern

258 (msec)

C

354 (msec)

D

450 (msec)

E

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, pos sibly because of the reduced compliance. To fully understand, visualize, and measure flow para meters in blood vessels, a method of acquisition is required that measures velocity in three dimensions and three direc tions over time, with high spatial and temporal resolution. Even with the fastest gradient systems available today, this pre sents 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 accept able 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

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

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.

98 Cardiovascular Magnetic Resonance

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.

Cardiovascular Magnetic Resonance 99

7 BLOOD FLOW VELOCITY ASSESSMENT

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.

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

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 imag ing methods. First, CMR does not use ionizing radiation. Second, the radiofrequency (RF) radiation penetrates bony structures without attenuation through relaxation para meters. Third, CMR gives additional diagnostic information about tissue characteristics. Finally, CMR provides three dimensional 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, electrocar diographic (ECG) setup, patient monitoring, and contrain dications 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 uncom fortable. During CMR, communication with the patient may be difficult because of interfering noise from the gradi ent coils. On the other hand, CMR is entirely noninvasive. Overall, the safety issues during CMR that may pose poten tial safety concerns2 can be summarized as follows:

100 Cardiovascular Magnetic Resonance

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, includ ing RF power deposition concerns 5. Auditory considerations from noise from the gradients 6. Safety considerations concerning superconducting mag net 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 differ ent 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 fibrilla tion threshold.4 In one of the investigations, however, car diac 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 fer romagnetic 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 ferro magnetic 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 exten sively reviewed and compiled.8 However, there are com mon 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 hazard ous. 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 typi cally attracted to the magnet, and this is one of the prob lems of imaging pacemakers. The second type of physical concern deals with magnet ically 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 clo sure 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 tis sue. Biologic effects of induced currents can be caused either by power deposition by the induced currents (ther mal 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

8 SPECIAL CONSIDERATIONS FOR CARDIOVASCULAR MAGNETIC RESONANCE: SAFETY, ELECTROCARDIOGRAPHIC SETUP, MONITORING, AND CONTRAINDICATIONS

Ferromagnetism

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

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 ven tricular 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 impor tant 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 cur rents 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 gra dients have the potential to cause either thermal or nonther mal 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 prox imity 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 med icine 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 estab lished, both on theoretical and experimental grounds. Application of this law, in combination with Maxwell’s law, yields an equation called the fundamental law of magne tostimulation, which has the hyperbolic form of a strength duration 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 prac tice, some precautions are necessary. First and most impor tantly, 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, circum stances 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 pel vis) 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 con traindicated in patients with pacemakers and ICDs, although recent developments are promising (see the dis cussion 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 sur gery, 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 head phones 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 ongo ing, 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 tem perature 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 seda tion with, for example, 2 to 5 mg intravenous (IV) diaze pam.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 pre fer 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 appoint ment 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 gadolin ium currently on the market is extremely good. Gadopente tate 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 avail able 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 func tion. (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

8 SPECIAL CONSIDERATIONS FOR CARDIOVASCULAR MAGNETIC RESONANCE: SAFETY, ELECTROCARDIOGRAPHIC SETUP, MONITORING, AND CONTRAINDICATIONS

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 cru cial 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 bradycar dia. In particular during adenosine infusion, atrioventricu lar heart block may develop in a small percentage of patients (0.7% to 2.8%), although this is usually asymp tomatic 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 dipyr idamole 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. Dobu tamine (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 sup raventricular tachycardia and (nonsustained) ventricular tachycardia, are seen. Dobutamine can be safely adminis tered 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 scan ner table is ideal. A cardiac arrest trolley must be main tained 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 echo cardiography. Ferrous metal, which is present in most mon itoring equipment, can distort the magnetic field, and such an item has the potential to become a projectile. In addi tion, monitoring wires that are attached to the patient, leav ing 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. There fore, specific solutions to these problems have been designed. Commercially available CMR compatible monitoring equi pment, including that used to measure ECG, blood pres sure, 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 con nected 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 sig nal for routine imaging, and each manufacturer has devel oped its own solution to the problems posed. Fiberoptic transmission of ECG signals for gating is now common place, 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 inter ference,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 lat est 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 elec trodes 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 interfer ence. 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 record ing 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 ferro magnetic and nonferromagnetic materials in CMR, for exam ple, 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 ferromagne tism, 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

8 SPECIAL CONSIDERATIONS FOR CARDIOVASCULAR MAGNETIC RESONANCE: SAFETY, ELECTROCARDIOGRAPHIC SETUP, MONITORING, AND CONTRAINDICATIONS

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 identi fied 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 cardio graphy 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 rea sons, 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 evalu ated 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 func tioning 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. How ever, 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.

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

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 establish ing 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 appro priate advice from a neurosurgical center is obtained. There have been cases of bleeding in the eye in patients with pre vious 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 risk benefit 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 pac ing 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 spe cific 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 angio plasty.50–53 Because of the increased indications for intra coronary stents, the population of patients with these stents in situ who need to undergo CMR for various diag nostic 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 vascu lar 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 inter val 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 revascu larization 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 after ward.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 associ ates.46 The authors found no significant increase in temper ature 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 influ ence 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 mini mal, 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 mil lion patients in the United States having a permanent pace maker in 2002 and more than 370,000 having an implant in 2003. Many centers consider MR absolutely contraindi cated in these patients, and none of the pacemakers or ICDs has been approved by the FDA for CMR examina tions. 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 character ized 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 com plex. In general, three hazardous MR interactions with pacemakers and ICDs should be considered. First, the static magnetic fields exert mechanical forces on the ferro magnetic components of the devices, including the pace maker 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 cur rently 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 phan toms.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. Tempera ture increases of up to 63.1 C were observed. Furthermore, no pacemaker malfunctions were observed in the asynchro nous mode. Inhibition or rapid pacing was observed dur ing 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 signifi cant adverse effects of CMR on DDD pacemakers.74 All units paced normally in the static magnetic field, but dur ing 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 config urations was studied, but none of the patients experienced any torque or heat sensation. Four non pacemaker depen dent 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 examina tions, 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 pro grammed 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 chan ges 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 fre quency of the applied imaging pulses. This could poten tially 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 state ments to be made about their suitability for CMR. Prefera bly, pacemaker dependent patients should not be scanned, but if needed, it has been suggested to program the pace makers in the asynchronous mode (VOO or DOO).83 Cardiovascular Magnetic Resonance 107

8 SPECIAL CONSIDERATIONS FOR CARDIOVASCULAR MAGNETIC RESONANCE: SAFETY, ELECTROCARDIOGRAPHIC SETUP, MONITORING, AND CONTRAINDICATIONS

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. Pacemaker dependent 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 exami nation.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 twitch ing. However, clinical scanners operate below the thresh old 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 pace makers and ICDs should not undergo CMR because of the unquantifiable risks. According to the recently pub lished 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 demon strated. 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 pros theses, 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.

Cardiovascular Magnetic Resonance 109

8 SPECIAL CONSIDERATIONS FOR CARDIOVASCULAR MAGNETIC RESONANCE: SAFETY, ELECTROCARDIOGRAPHIC SETUP, MONITORING, AND CONTRAINDICATIONS

References

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

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.

110 Cardiovascular Magnetic Resonance

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 ques tions 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 with out limitations to “acoustic windows” or overlapping struc tures. 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 dark blood spin echo,34 and 3D contrast enhanced magnetic reso nance 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 youn ger 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 diag nosis 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 exami nation must accommodate these interventions. Thus, the physician must be knowledgeable about the pre and post operative 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 hypo plastic (e.g., hypoplastic left heart syndrome). In the evalu ation 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. Assess ment of ventricular function in ventricles with uncom mon 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 consider ation. Even in older preteens or teens, cooperation may be problematic (e.g., breath holding). Although many intra vascular 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 portabil ity of CMR is disadvantageous, especially for the critically ill infant for whom transportation to the CMR suite is cum bersome. 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 con structed without regard to electrocardiogram (ECG) trigger ing 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

9 SPECIAL CONSIDERATIONS: CARDIOVASCULAR MAGNETIC RESONANCE IN INFANTS AND CHILDREN

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 dis cussed are generic. The major manufacturers of CMR equipment have similar sequences with proprietary nomen clature (see Appendix I.).

GENERAL PROTOCOL FOR CARDIOVASCULAR MAGNETIC RESONANCE IN CONGENITAL HEART DISEASE The CMR procedure in patients with CHD is not standar dized 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 func tional 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 connec tions are suspected (e.g., connection below the diaphragm).

Ao

RA

RA-PA

RPA

Anatomy Morphology

RV

LPA

C

B

A

RA

LV

D

E

F

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) acqui sitions 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 mea sure 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. “Retro spectively” acquired velocity mapping (compared with pro spectively 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 junc tion 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 pulmo nary 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 anat omy 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

9 SPECIAL CONSIDERATIONS: CARDIOVASCULAR MAGNETIC RESONANCE IN INFANTS AND CHILDREN

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 per formed 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 large volume 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 coarcta tion) 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 con straints 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

B

C

D

E

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 bright blood steady state free precession dataset. AAo, ascending aorta; DAo, descending aorta; LV, left ventricle.

F

AAo DAo

LV

G

114 Cardiovascular Magnetic Resonance

PDA

AAo

MPA

DAo

DAo

A

C

B

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.

RV LV RV

RV

LV

A

C

B RVOC AAo

LV

RV LV

D

E

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.

Cardiovascular Magnetic Resonance 115

9 SPECIAL CONSIDERATIONS: CARDIOVASCULAR MAGNETIC RESONANCE IN INFANTS AND CHILDREN

TAo

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.

A

B

C

D

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] ver sions), 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 techni ques 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 nonin vasively into “cubes of magnetization” and local defor mation 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 tag ging 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.

RPA MPA MPA

LA

C

LV

B RPV

C LV

B

A

E

C AAo

AAo

AAo C C

RV RV

LV

C

DAo

D

F

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 accu mulates 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 mid 1980s,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 addi tion, 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 imag ing, coronary artery CMR may be used to image patients with CHD who have coronary manipulation (e.g., trans position of the great arteries after arterial switch45), native disease (e.g., anomalous left coronary artery from the pulmonary artery), or acquired disease (e.g., coro nary aneurysm as a result of Kawasaki disease). Both tar geted 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 proto cols 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 symp toms they cause, and at what age they occur, but also Cardiovascular Magnetic Resonance 117

9 SPECIAL CONSIDERATIONS: CARDIOVASCULAR MAGNETIC RESONANCE IN INFANTS AND CHILDREN

LPA

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

F

F

A

B

D

E

C

F

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 proce dures with CMR are a protocol in themselves and typ ically 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 uncer tainty as to its utility.53–55 The CMR protocol includes T1 weighted imaging, cine for RV volume and func tion, 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 reso lution 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

RV

LV

RV LV

C

LAD

B

A

LCx

E

D

RCA

LAD

RCA Ao

PA

LCA LCx

LCA

A

B

C

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 four chamber 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 repeti tion 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 simul taneously. 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, impor tant to recognize that the number of calculated phases should not exceed two times the measured phases (essen tially 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, real time CMR in which all of the lines of k space are acquired in one heartbeat because of respiratory motion or arrhyth mia. 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 anes thesia during which the anesthesiologist can suspend respi ration. 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 com promise, those who have failed deep sedation, and those situations in which deep sedation would need to be mark edly 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 mon itor 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 addi tional 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 gado linium 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 gadolin ium is in the structure of interest during imaging. In chil dren, because of the quick circulatory time, intravenously administered gadolinium will reach the systemic circula tion (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 ves sels are surgically transposed to arise normally and the cor onary arteries are implanted above the native pulmonary (neoaortic) valve. The pulmonary arteries are typically “draped” over the ascending aorta in the Lecompte maneu ver (Figs. 9 11 to 9 13), which may result in left PA stenosis. It is advantageous for the surgeon to perform a

LPA

RPA

AAo

DAo

B A LPA

RPA

RVOT

RVOT

C

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 9 11C showing a narrowed right ventricular outflow tract (RVOT). This is a diastolic frame, so no turbulence is seen. RPA, right pulmonary artery.

D

LPA

RVOT

RPA RPA

A

Lecompte maneuver rather than making the right PA course tortuous by bringing it directly posterior and then under neath 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

B

assess for stenoses, which generally can occur at the take off of the left PA or in the right PA between the superior vena cava (SVC) and the aorta. Cines of the RV and LV out flow 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.

RPA LPA

RVOT Ao

RPA

B

A

RPA Ao

RVOT

RVOT

LPA

C

D

quantitate ventricular performance and to search for residual VSD. Data suggest that including RV trabeculations and pap illary 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 Gad olinium 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 obstruc tion 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 coro nary 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 ven tricles are linked in such a way that separation of the circu lations 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 asso ciated 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, physiol ogy, 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 assess ment. Anomalous venous structures, such as a decom pressing 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 func tion 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 assess ment of the distal aortic arch for obstruction and the aor tic to pulmonary anastomosis, visualization of the aortic to pulmonary shunt (typically, a right Blalock Taussig shunt) or the right ventricle to PA shunt (i.e., Sano proce dure) is important. This is generally done with double inversion dark blood imaging (or CE MRA) because the turbulence across this structure by cine generally causes

nAo

nPA

nAo

nPA

C

B

A

D LPA Inn v

LPA

SVC SVC SVC B RPA

E

F

G

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 aor tic 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, hypo plasia, 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 hypoplas tic left heart syndrome; see Fig. 9 15); ventricular out flow tract obstruction (especially in patients with a bulboventricular foramen; see Fig. 9 16); aortic pulmo nary collateral vessels; anomalous venous structures; and pulmonary or systemic venous obstruction (espe cially 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 vol ume and mass, stroke volume, cardiac index, and atrio ventricular 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, two chamber (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).

RV

A

C

B

D

signal loss. Qp/Qs is obtained by velocity mapping in the ascending aorta (over each semilunar valve) and by veloc ity 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 hemi Fontan procedure was performed, imaging should assess whether any leak was present from the SVC to PA anasto mosis into the atrium. Collated flow is assessed with veloc ity maps in the pulmonary veins. After the Fontan procedure, the most important structure to image is the entire systemic venous pathway (Fontan baf fle) for obstruction or clot (see Fig. 9 16). Visualization of 124 Cardiovascular Magnetic Resonance

fenestration flow (using cine) and assessment of this struc ture 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 deter mine 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 des cending aorta. It may have a number of associated abnorm alities, 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,

PA

LV

A

E

RVOC

B

C

VENC=3 m/s

D

VENC=5 m/s

F

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 to descending 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 impor tant 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 bicus pid aortic valve (see Fig. 9 6). Aortic stenosis or regurgita tion 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 regurgi tant fraction and peak velocity is obtained for aortic regurgi tation and stenosis, respectively. In addition, through plane velocity mapping just below the coarctation and at the dia phragm 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 myo cardial 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

9 SPECIAL CONSIDERATIONS: CARDIOVASCULAR MAGNETIC RESONANCE IN INFANTS AND CHILDREN

Ao

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 year old child with a typical coarctation of the proximal descending aorta (DAo) (arrow). D and E, Shaded surface display from a three dimensional contrast enhanced magnetic resonance angiography showing extensive collateral vessels. The coarctation is readily seen (arrow). AAo, ascending aorta.

DAo

AAo

A

B

AAo

DAo

C

E

D

A

B

C

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, adoles cents, 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.

Cardiovascular Magnetic Resonance 127

9 SPECIAL CONSIDERATIONS: CARDIOVASCULAR MAGNETIC RESONANCE IN INFANTS AND CHILDREN

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

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

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.

128 Cardiovascular Magnetic Resonance

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 influ encing 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 non breath 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 mis registration effects, and it may also vary during the breath hold period itself,2 resulting in image blurring and arti facts. In addition, the scan parameters are limited by the need to perform imaging within the duration of a comfort able breath hold period, and for a number of patients, this period may be very short. An alternative to breath holding is to monitor respira tory 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 techni ques being strongly dependent on the accuracy of the method of motion assessment. During normal tidal respira tion, 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 determin ing diaphragm motion during respiration. The navigator echo is the signal from a column of material oriented per pendicular to the direction of the motion to be monitored. On Fourier transformation, this signal results in a well defined edge of the moving structure. The navigator echoes may be interleaved with the imaging sequence and conse quently 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 devel opments, considers the various choices that have been studied in the implementation of navigators, and discusses their importance. There are many variables in the applica tion 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 breath ing methods. The first of these uses the navigator informa tion 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 mea surement 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 nav igator 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 effi ciency 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 dia phragm displacements that occurred during data acquisi tions 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

10 USE OF NAVIGATOR ECHOES IN CARDIOVASCULAR MAGNETIC RESONANCE AND FACTORS AFFECTING THEIR IMPLEMENTATION

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 aver aging of data to improve the signal to noise ratio, increas ing 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 arti facts 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 tech nique has been shown to produce good results with reason able 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 sub jects, 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 over sampled, typically by a factor of five, and then sorted retro spectively 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 tech nique that combines the benefits of phase ordering with an automatic window selection that enables the highly effi cient 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 pro spective 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 navi gator echo. With the spin echo technique, a spin echo signal is gen erated from the column of material formed by the intersec tion 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 well defined 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 two dimensional (2D) RF pulse to excite a column of approxi mately 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.)

Cardiovascular Magnetic Resonance 131

10 USE OF NAVIGATOR ECHOES IN CARDIOVASCULAR MAGNETIC RESONANCE AND FACTORS AFFECTING THEIR IMPLEMENTATION

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 narrow est 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 techni ques, 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 win dow of 5 mm or less is required, and this generally results in relatively poor scan efficiency. In addition, as noted ear lier, many subjects and patients undergo a drift in dia phragm 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 prede fined acceptance window.13 With this method, one com plete 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 off set 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 posi tion. 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 accept reject 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 accept reject 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 mea sure the position of the diaphragm. However, the motion of the heart is not straightforward, and only the inferior bor der 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 cor onary 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 rela tively 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 fol lowing 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 expira tory breath hold scans before the coronary imaging proto col. 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 improv ing scan efficiency. The need for a subject specific correc tion 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 subject specific 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 follow ing should be performed with subject specific factors deter mined from free breathing acquisitions. An additional or alternative approach to the real time slice following described earlier is to use a postprocessing adap tive motion correction technique4 to correct an image retro spectively for movement occurring during data acquisition. This technique, which can be used to correct a 2D acquisi tion for in plane displacement or a 3D acquisition for in plane 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 indi vidual 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 spi ral29 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 col umn. The dome of the right hemi diaphragm is higher than that of the left, and the two move coherently with respira tion, 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 later ally 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 vary ing 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 ante rior 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 echo measured 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 lon ger 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 diffi cult 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 anterior posterior 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 direc tions 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 anterior posterior measurements, made over a period of approxi mately 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 effi ciency 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 poten tial 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 com puting architecture of the scanner being used (discussed Cardiovascular Magnetic Resonance 133

10 USE OF NAVIGATOR ECHOES IN CARDIOVASCULAR MAGNETIC RESONANCE AND FACTORS AFFECTING THEIR IMPLEMENTATION

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 subject specific 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 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 hemi diaphragm, (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 regu larly 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 acquisi tions 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

10 USE OF NAVIGATOR ECHOES IN CARDIOVASCULAR MAGNETIC RESONANCE AND FACTORS AFFECTING THEIR IMPLEMENTATION

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. Partic ularly if prenavigators only are being used, the longer this interval, the greater the potential for errors caused by respi ratory 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 multi slice imaging and also provide some potential for estimat ing internavigator respiratory motion. The potential prob lem with this is that the navigator signal to noise ratio could be reduced and this may affect the accuracy of navi gator 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 read out. However, the precision of the measurement is depen dent to a large extent on the signal to noise ratio of the navigator measurement. The most important factor affect ing 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 imag ing area of interest and the region of navigator detection. On the other hand, if phased array coils are used, it is pos sible to position one coil specifically for navigator detec tion, possibly over the region of the right diaphragm. Another important factor in the precision of the measure ment 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 sec tion 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 detec tion, correlation, or least squares fit algorithms. For rapid tracking (repetition time < 100 msec) or narrower col umns, 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 How ever, most navigator techniques acquire only one or two navigators per cardiac cycle, and in such cases, the signal to noise ratios are usually relatively high and edge detec tion algorithms are generally adequate.

MORE RECENT APPROACHES Other Forms of Navigators As has been mentioned, there are problems with the con ventional 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 can not be implemented simply and efficiently to give a mea sure of this motion in 3D.37 A number of ingenious alternative approaches have been described. In 2003, Nguyen and colleagues38 described a method that selec tively excited the epicardial fat, followed by a rapid read out scheme that instantaneously gave three 1D images of its position in the x, y, and z directions. Tested on six sub jects, the method showed a slight improvement over con ventional 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 measure ment 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 fur ther or used. Subsequently, Stehning and associates used radial imaging for “self navigation.” They developed an inter leaved 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 cor onary scans. Hardy and colleagues41 used cross correla tion of low resolution real time 2D spiral coronary artery images to accept or reject images for averaging. This adaptive averaging technique was extended to high resolution segmented acquisitions by cross correlating subi mages reconstructed from individual data segments. Breath ing autocorrection with spiral interleaves (BACSPIN) is a similar technique42 that involves the acquisition of a multi slice 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 diffi culty 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 one dimensional (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 transla tional 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 low resolution 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 col leagues46 acquired images at a number of breath hold posi tions 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 preac quired during free breathing. An affine model of the respi ratory related cardiac motion was then extracted from these images, and this was steered by real time navigators to con trol 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 10 mm 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 com puters. 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 con trol and adjustment of parameters associated with the scan ning 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 com puter, 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 con trol 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. There fore, 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 respira tion, 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 imag ing.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.

Cardiovascular Magnetic Resonance 137

10 USE OF NAVIGATOR ECHOES IN CARDIOVASCULAR MAGNETIC RESONANCE AND FACTORS AFFECTING THEIR IMPLEMENTATION

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 cor rection that has been applied to spiral imaging is to acquire a low resolution 3D dataset on the fat resonance immedi ately before a high resolution interleaf on the water reso nance. 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 dark blood coronary vessel wall imaging, where high quality images have been obtained without the need for a gating window.43

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

References 1. Atkinson DJ, Edelman RR. Cineangiography of the heart in a single breath hold with a segmented turboFLASH sequence. Radiology. 1991;178:357–360. 2. Holland AE, Goldfarb JW, Edelman RR. Diaphragmatic and cardiac motion during suspended breathing: preliminary experience and implications for breath-holding. Radiology. 1998;209:483–489. 3. Wang Y, Riederer SJ, Ehman RL. Respiratory motion of the heart: kinematics and the implications for the spatial resolution in coronary imaging. Magn Reson Med. 1995;33:713–719. 4. Ehman RL, Felmlee JP. Adaptive technique for high-definition MR imaging of moving structures. Radiology. 1989;173:255–263. 5. Liu YL, Riederer SJ, Rossman PJ, Grimm RC, Debbins JP, Ehman RL. A monitoring, feedback, and triggering system for reproducible breath-hold MR imaging. Magn Reson Med. 1993;30:507–511. 6. 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. 7. Wang Y, Grimm RC, Rossman PJ, et al. angiography in multiple breath-holds using a respiratory feedback monitor. Magn Reson Med. 1995;34:11–16. 8. Keegan J, Gatehouse PD, Taylor AM, Yang GZ, Jhooti P, Firmin DN. Coronary artery imaging on a mobile 0.5Tesla scanner: implementation of real-time navigator-echo controlled segmented k-space FLASH and interleaved spiral sequences. Magn Reson Med. 1999; 41:392–399. 9. Taylor AM, Jhooti P, Wiesmann FW, 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. 10. Lenz GW, Haacke EM, White RD. Retrospective cardiac gating: a review of technical aspects and future directions. Magn Reson Imaging. 1989;7:445–455. 11. Hofman MB, Paschal CB, Li D, Haacke EM, van Rossum AC, Sprenger M. MRI of coronary arteries: 2D breath-hold vs. 3D respiratory-gated acquisition. J Comput Assist Tomogr. 1995;19:56–62. 12. Oshinski JN, Hofland L, Mukundan S, Dixon WT, Parks WJ, Pettigrew RI. Two-dimensional coronary MR angiography without breath-holding. Radiology. 1996;201:737–743. 13. Sachs TS, Meyer CH, Irarrazabal P, Hu BS, Nishimura DG, Macovski A. The diminishing variance algorithm for real-time reduction of motion artifacts in MRI. Magn Reson Med. 1995;34:412–422. 14. Maki JH, Prince MR, Londy FJ, Chenevert TL. The effects of time varying intravascular signal intensity and k-space acquisition order on three-dimensional MR angiography image quality. J Magn Reson Imaging. 1996;6:642–651. 15. 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. 16. Sinkus R, Bornert P. Motion pattern adapted real-time respiratory gating. Magn Reson Med. 1999;41:148–155. 17. Sinkus R, Bo¨rnert P. Extension of real-time MR gating to cope with changes in motion pattern: making MR gating autarkic. In: Proceedings of the Sixth Scientific Meeting of ISMRM, Sydney. 1998:2127. 18. Jhooti P, Gatehouse PD, Keegan J, Bunce HH, Taylor AM, Firmin DN. Phase ordering with automatic window selection (PAWS): a novel motion-resistant technique for 3D coronary imaging. Magn Reson Med. 2000;43:470–480. 19. Hackenbroch M, Nehrke K, Gieseke J, et al. 3D motion adapted gating (3D MAG): a new navigator gated 3D coronary MR-angiography. Eur Radiol. 2005;1598–16606. 20. Langreck H, Schnackenburg B, Nehrke K, et al. MR coronary artery imaging with 3D motion adapted gating (MAG) in comparison to a standard prospective navigator technique. J Cardiovasc Magn Reson. 2005;7:793–797. 21. Pauly J, Nishimura D, Macovski A. A k-space analysis of small tipangle excitation. J Magn Reson. 1989;81:43–56. 22. McConnell MV, Khasigawala VC, Savord BJ, et al. Prospective adaptive navigator correction for breath-hold MR coronary angiography. Magn Reson Med. 1997;37:148–152. 23. Danias PG, McConnell MV, Khasigawal VC, Chuang ML, Edelman RR, Manning WJ. Prospective navigator correction of image position for coronary MR angiography. Radiology. 1997;203:733–736.

138 Cardiovascular Magnetic Resonance

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.

Cardiovascular Magnetic Resonance 139

10 USE OF NAVIGATOR ECHOES IN CARDIOVASCULAR MAGNETIC RESONANCE AND FACTORS AFFECTING THEIR IMPLEMENTATION

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 arti facts. Standard CMR planes have evolved from other imag ing modalities, including body computed tomography (CT) imaging, echocardiography, and X ray contrast angiogra phy, and consistent nomenclature across imaging mod alities 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 spe cific 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 func tion and are familiar from other cardiac imaging techni ques. 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 ver tical 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 differ ent 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 char acterization, the spin echo images typically permit the dif ferentiation of fat (white) from muscle (intermediate gray). Black regions in spin echo CMR studies may repre sent 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 struc ture of interest. Presaturation bands can be added to remove specific artifacts. Other preparatory prepulses can be applied to emphasize or deemphasize the signal contri bution of specific tissues. For example, in the evaluation of arrhythmogenic right ventricular cardiomyopathy, high spatial 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 vascu lar 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 out flow tract and are therefore helpful in the evaluation of patients with congenital heart disease and RV cardiomyop athy. 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

C

11 NORMAL CARDIAC ANATOMY, ORIENTATION, AND FUNCTION

A

B

D

E

F

G H Figure 11 1 For legend see next page.

Cardiovascular Magnetic Resonance 141

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

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 intra cardiac 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 advan tage of CMR is the ability to precisely position the long axis planes to avoid the foreshortening that can occur in con trast 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 ori entation to obtain views of the atria can be useful for asses sing 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 encom pass 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 two chamber 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. 11 1D), with the short axis planes perpendicular to an imagi nary 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 sys tolic 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 tar geted 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. Sev eral 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 tra beculae carneae and crista terminalis and does not rep resent an atrial mass.7
Lipomatous hypertrophy of the atrial septum (Fig. 11 3B). 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 pericar dium 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 interpre tation. These artifacts relate primarily to several features of CMR. The acquisition time is often relatively long in com parison 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

F

E

Right ventricular wall

Inferior vena cava

Right atrium Coronary sinus

G Figure 11 2

Coronary sinus

H

For legend see page 145.

Cardiovascular Magnetic Resonance 143

11 NORMAL CARDIAC ANATOMY, ORIENTATION, AND FUNCTION

Ascending aorta

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

Ascending aorta

Right pulmonary artery

Pericardium

Left atrium

RVOT Pericardium Epicardial fat

Right atrium

Right ventricular wall

J

I

Transverse aortic arch

Ascending aorta

K

L

RCA origin Aortic valve leaflets PDA

M

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

R

S

T

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 experi ence 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

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

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 real time imaging approaches, CMR requires gating to the electro cardiogram (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 implemen ted 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 rela tive 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 pro vide 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 pat terns. Strategies to reduce respiratory artifact include the use of sustained breath hold, presaturation of the high intensity 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 pres ent 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. Fig ure 11 4F shows the artifact related to a bioprosthetic aor tic 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 car diac function, and CMR is often considered as the gold standard for the noninvasive evaluation of cardiac func tion,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 accu rately identify end systole. Historically, X ray left ventricu lography 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 trans ducer but typically is at least 30 frames/sec (temporal reso lution 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 view sharing 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 fre quent 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 car diac cycle, the image data are generally acquired over mul tiple 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. How ever, 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) measure ments as in echocardiography, but with CMR, more accu rate 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 for mulas) 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 sum mation 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 polyno mial approximation for numerical integration;26 failure to distinguish between the mathematical and medical defini tions of Simpson’s rule leads to confusion between clini cians and engineers or medical physicists. The term modified Simpson’s rule is sometimes used in the echocardiog raphy 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 qual itatively (“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 Asso ciation1 model for qualitative assessment of wall thickening. Wall thickening can also be determined quanti tatively by using centerline or other methods, and commer cial software is available for such analyses. Myocardial tagging methods28–30 can be used to quantify myocardial contractility using strain without the need to identify endo cardial or epicardial borders explicitly. However, software for tracking deformation of the tag lines over time none theless requires human interaction, prolonging analysis times. The harmonic phase (HARP) technique was pro posed to obviate the human postprocessing issue by com puting 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, dis placement 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

11 NORMAL CARDIAC ANATOMY, ORIENTATION, AND FUNCTION

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 fre quency 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 dis placed in the phase encoding direction. As shown in Fig ure 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 sat uration 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 17 segment model of the LV (2). Ant, anterior; Ap, apical; As, anteroseptal; Inf; inferior; Lat, lateral; Sept, septal; Bs, basal.

Apex Ap-Sept

Ap-Lat

Mid-Sept MidLat Bs-Lat

Bs-Sept

MidSept

MidLat

Mid-Inf

Inf/Lat

Ap-Sept

Ap-Ant

Ap-Inf

Ap-Lat

Apex Ap-Inf Mid-Inf Bs-Inf

A

Ap-Ant MidAnt Bs-Ant

B

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 read ily applied to the RV. Studies in ventricular casts have shown an excellent correlation between CMR and displace ment 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 sub tracting the end systolic volume from the end diastolic vol ume. 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 deter mined 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 assess ment 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 multiply ing 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 myocar dial 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 correla tion 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 lin ear measurements and cubed power geometric formulas, such as the echocardiographic Penn formula,43 generally overestimate volumetric mass even when the linear mea surements 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 sys tematically 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 Dal las Heart Study populations. Table 11 2 shows cine SSFP based 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 differ ent ethnic groups. Higher field strength CMR systems (i.e., 3 Tesla[T]) offer signal to noise advantages over “conventional” 1.5 Tesla systems and are well established for neurologic imag ing. 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

63)

Men 95th Percentile

115 36 155 0.69 0.70 (DHS) 66 21 89 58 54 (DHS) 18 78

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

Women (n

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

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

Women (n

367)

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

11 NORMAL CARDIAC ANATOMY, ORIENTATION, AND FUNCTION

mass with SSFP CMR is smaller than CMR gradient echo mass. This has implications for what constitutes normal range 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

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

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 pre served 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 optimiza tion and understanding of how to identify and correct arti facts 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 measure ments of velocity and vessel flow without the constraints of Doppler methods. Although there is potential for arti facts, CMR allows accurate and reproducible measurement of vessel flow in vivo.37,59–61 The approach for the mea surement of aortic flow is straightforward. The coronal scout image is used, and the imaging plane is placed per pendicular 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 wrap around. The reconstructed images are typically presented as a set of magnitude images that are used to determine the cross sectional 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 differ ence 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 techni ques similar to those for aortic flow. This is particularly valuable in evaluation of patients with left to right intracar diac shunting to determine the ratio of pulmonary to sys temic 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 dif ference 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 mor phology, motion, competence, and effects on ventricular function. Imaging cardiac valve morphology poses signifi cant 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 pathol ogy 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 ste nosed 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 LVOT view 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, mea surements 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 tur bulence (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. Regur gitant jets can be underestimated or overlooked entirely based on visual assessment alone. Quantitative flow techni ques 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 Inter rogation 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 regurgita tion 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 contraindica tion for CMR74 except in the case of probable valve dehis cence,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 pro duce 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.

References 1. 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;105:539–542. 2. El-Khoury GY, Bergman RA, Montgomery WJ. Sectional Anatomy by MRI. New York: Churchill Livingstone; 1995. 3. 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.

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

11 NORMAL CARDIAC ANATOMY, ORIENTATION, AND FUNCTION

E

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

7. Jarvinen VM, Kupari MM, Hekali PF, Poutanen VP. Right atrial MR imaging studies of cadaveric atrial casts and comparison with right and left atrial volumes and function in healthy subjects. Radiology. 1994;191:137–142. 8. Mortele KJ, Mergo PJ, Williams WF. Lipomatous hypertrophy of the atrial septum: diagnosis with fat suppressed MR imaging. J Magn Reson Imaging. 1998;8:1172–1174. 9. Shirani J, Roberts WC. Clinical, electrocardiographic and morphologic features of massive fatty deposits (“lipomatous hypertrophy”) in the atrial septum. J Am Coll Cardiol. 1993;22:226–238. 10. Noma M, Kikuchi Y. Images in cardiovascular medicine: lipomatous hypertrophy of the atrial septum. Circulation. 1999;100:684. 11. 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;42:361–370. 12. Hundley WG, Meshack BM, Willett DL, et al. Comparison of quantitation of left ventricular volumes, ejection fraction and cardiac output in patients with atrial fibrillation by cine magnetic resonance imaging: versus invasive measurements. Am J Cardiol. 1996;78:1119–1123. 13. Danias PG, Stuber M, Botnar RM, et al. Relationship between motion of coronary arteries and diaphragm during free breathing: lessons from real-time MR imaging. AJR Am J Roentgenol. 1999;172:1061–1065. 14. Taylor AM, Jhooti P, Wiesmann F, et al. Magnetic resonance navigatorecho monitoring of temporal changes in diaphragm position: implications for magnetic resonance coronary angiography. J Magn Reson Imaging. 1997;7:629–636. 15. Botnar RM, Stuber M, Danias PG, et al. Improved coronary artery definition with T2-weighted, free-breathing, three-dimensional coronary MRA. Circulation. 1999;99:3139–3148. 16. Bellenger NG, Gatehouse PD, Rajappan K, et al. 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. 17. Babcock EE, Brateman L, Weinreb JC, et al. Edge artifacts in MR images: chemical shift effect. J Comput Assist Tomogr. 1985;9:252–257. 18. Lotan CS, Cranney GB, Doyle M, Pohost GM. Fatshift artifact simulating aortic dissection on MR images. AJR Am J Roentgenol. 1989;152: 385–386. 19. Mor-Avi V, Jenkins C, Kuhl HP, et al. Real-time 3-dimensional echocardiographic quantification of left ventricular volumes: multicenter study for validation with magnetic resonance imaging and investigation of sources of error. J Am Coll Cardiol Img. 2008;1:413–423. 20. Foo TK, Bernstein MA, Aisen AM, et al. 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. 21. Heidemann RM, Ozsarlak O, Parizel PM, et al. A brief review of parallel magnetic resonance imaging. Eur Radiol. 2003;13:2323–2327. 22. Yang PC, Kerr AB, Liu AC, et al. New real-time interactive cardiac magnetic resonance imaging system complements echocardiography. J Am Coll Cardiol. 1998;32:2049–2056. 23. Dodge HT, Sandler H, Ballew DW, et al. The use of biplane angiocardiography for the measurement of left ventricular volume in man. Am Heart J. 1960;60:762–776. 24. Sandler H, Dodge HT. The use of single plane angiocardiograms for the calculation of left ventricular volume in man. Am Heart J. 1968;75:325–334. 25. Chuang ML, Hibberd MG, Beaudin RA, 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. 26. Kaplan W. Advanced Calculus. New York: Addison-Wesley; 2002. 27. Strohm O, Schulz-Menger J, Pilz B, et al. Measurement of left ventricular dimensions and function in patients with dilated cardiomyopathy. J Magn Reson Imaging. 2001;13:367–371. 28. 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. 29. Axel L, Dougherty L. MR imaging of motion with spatial modulation of magnetization. Radiology. 1989;171:841–845. 30. Fischer SE, McKinnon GC, Maier SE, Boesiger P. Improved myocardial tagging contrast. Magn Reson Med. 1993;30:191–200. 31. 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.

156 Cardiovascular Magnetic Resonance

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.

Cardiovascular Magnetic Resonance 157

11 NORMAL CARDIAC ANATOMY, ORIENTATION, AND FUNCTION

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.

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

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 sev eral 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 compre hensive CMR studies has been hindered by long image acqui sition 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 complemen tary 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 pro tocols 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 man agement of most patients with CAD: cardiac morphology, global and regional myocardial function, myocardial ische mia, viability status of dysfunctional myocardium, and the presence of flow limiting coronary stenoses. CMR can pro vide 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 ven tricles.16,17 Cine imaging forms an essential component of any CMR study in CAD. With the introduction of steady state free precession (SSFP) pulse sequences, data acquisi tion times have been reduced to around 5 minutes for acqui sition 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 acquisi tion.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: contrast enhanced myocardial perfusion and dobutamine stress MR (DSMR). The place of myocardial perfusion imaging in a compre hensive 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 min utes 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 min utes 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 recov ery 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 car diac 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 pro vides 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 uni versal comprehensive CMR protocol cannot be prescribed. The protocols suggested below should therefore be seen as general guidance providing a framework that requires adjust ments 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 proxi mal coronary stenosis.4,16,32 Acquisition times are more variable than for other CMR methods. Free breathing tech niques 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 time consuming. 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 min utes 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 navi gator gated MR coronary angiography into a compre hensive 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 ische mia 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 com parison, DCMR was found to be more accurate than stress perfusion CMR to detect ischemia.34 However, for compre hensive CMR protocols, most authors have favored adeno sine 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 theoreti cally 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 compli cations than inotropic stress, which is of particular rele vance in the MR environment.35,36 Third, the administration protocol for vasodilators is less time consuming 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 imag ing 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 lit erature, 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 con trast 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. Twenty three patients completed the protocol in a mean scan time of 46
5 minutes. In 2000, Sensky and coworkers presented a comprehen sive 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 consec utive patients presenting to the emergency room with chest pain.39 The CMR protocol commenced with perfusion imag ing 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 per fusion 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

1 (tagging) 1 2 2 1 1 1 1

3 3 1 1 2 2 3

Stress Perfusion

LGE

Low Dose DCMR

Coronary CMR 2 (coronary flow)

4 4 3 3 2

2 3 5 4 4 4 3

3 2

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 com bined 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 perfu sion study in preparation for LGE. Chiu and coworkers studied patients with acute coronary syndromes, complet ing 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 com menced with resting perfusion, which was followed by cine imaging, a targeted coronary MRA of the right coro nary 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 coro nary 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 coronar ies) to predict the presence of CAD on subsequent coro nary 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 coro nary 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 signifi cantly increased the predictive power of a logistic regres sion 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 combi nation 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 accu racy. 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 approxi mately 40 minutes. There is early evidence that the com bined analysis of these methods improves the diagnostic accuracy and prognostic value over the analysis of individ ual CMR methods. Resting perfusion imaging may be added at little additional time expense to aid in the identification of perfusion artifacts and to allow semiquan titative 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 NSTE MI (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 pre dict recovery of dysfunctional segments at 6 months was assessed. CMR was completed in 50
10 minutes. Low dose 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 sensi tivity of 97% with a sensitivity of 91%, while the combina tion 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, agree ment between the two methods was not good in the seg ments 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 seg ments 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 pre sented by Bodi and coworkers, who assessed four CMR derived 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 dif ferent 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 multi variate analysis, LGE had the highest predictive power, followed by DCMR. On the basis of the multivariate anal ysis, the authors generated a comprehensive score com bining 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 nontrans mural 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 meth ods to predict the likelihood of functional recovery after myocardial infarction. In summary, despite some controversy, the current evi dence 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 mod ified. For example, if the scan is requested for the noninva sive screening of patients at risk of CAD, a high negative predictive value will be paramount. A combined CMR anal ysis 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 myocar dium will benefit from revascularization, a higher positive predictive value may be more relevant, and the study should be interpreted as positive only if several measure ments 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. Coro nary MRA at present is less accurate than perfusion imag ing 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 avail able, 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 via ble, 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 compre hensive 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 com pleted 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 qual ity 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 proto cols and to establish the optimal combination of CMR methods in different clinical scenarios. Standardized proto cols for image acquisition and analysis need to be devel oped and validated in large scale clinical trials. Finally, the cost effectiveness of comprehensive CMR in compari son with established and other evolving management stra tegies needs to be determined.

References 1. Kramer CM. Integrated approach to ischemic heart disease: the onestop shop. Cardiol Clin. 1998;16:267–276. 2. Poon M, Fuster V, Fayad Z. Cardiac magnetic resonance imaging: a “one-stop-shop” evaluation of myocardial dysfunction. Curr Opin Cardiol. 2002;17(6):663–670. 3. Al-Saadi N, Nagel E, Gross M, et al. Noninvasive detection of myocardial ischemia from perfusion reserve based on cardiovascular magnetic resonance. Circulation. 2000;101:1379–1383. 4. Danias PG, Roussakis A, Ioannidis JPA. Diagnostic performance of coronary magnetic resonance angiography as compared against conventional X-ray angiography a meta-analysis. J Am Coll Cardiol. 2004;44:1867–1876. 5. Kim RJ, Wu E, Allen R, et al. The use of contrast-enhanced magnetic resonance imaging to identify reversible myocardial dysfunction. N Engl J Med. 2000;343:1445–1453. 6. 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. 7. Klein C, Nekolla SG, Bengel FM, et al. Assessment of myocardial viability with contrast-enhanced magnetic resonance imaging: comparison with positron emission tomography. Circulation. 2002; 105:162–167. 8. Nagel E, Klein C, Paetsch I, et al. Magnetic resonance perfusion measurements for the noninvasive detection of coronary artery disease. Circulation. 2003;108:432–437. 9. 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. 10. 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;103:2230–2235. 11. Selvanayagam JB, Kardos A, Francis JM, et al. Value of delayedenhancement cardiovascular magnetic resonance imaging in predicting myocardial viability after surgical revascularization. Circulation. 2004;110:1535–1541. 12. 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. 13. Barkhausen J, Ruehm SG, Goyen M, et al. MR evaluation of ventricular function: true fast imaging with steady-state precession versus fast low-angle shot cine MR imaging: feasibility study. Radiology. 2001; 219:264–269. 14. Plein S, Bloomer TN, Ridgway JP, Jones TR, Bainbridge GJ, Sivananthan MU. Steady-state free precession magnetic resonance imaging of the heart: comparison with segmented k-space gradientecho imaging. J Magn Reson Imaging. 2001;14(3):230–236. 15. Plein S, Smith WH, Ridgway JP, et al. Measurements of left ventricular dimensions using real-time acquisition in cardiac magnetic resonance imaging: comparison with conventional gradient echo imaging. MAGMA. 2001;13:101–108. 16. Myerson SG, Bellenger NG, Pennell DJ. Assessment of left ventricular mass by cardiovascular magnetic resonance. Hypertension. 2002;39:750–755. 17. Niwa K, Uchishiba M, Aotsuka H, et al. Measurement of ventricular volumes by cine magnetic resonance imaging in complex congenital heart disease with morphologically abnormal ventricles. Am Heart J. 1996;131:567–575. 18. Greil GF, Germann S, Kozerke S, et al. Assessment of left ventricular volumes and mass with fast 3D cine steady-state free precession k-t

168 Cardiovascular Magnetic Resonance

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.

Cardiovascular Magnetic Resonance 169

12 COMPREHENSIVE CARDIOVASCULAR MAGNETIC RESONANCE IN CORONARY ARTERY DISEASE

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.

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

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 avail able 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 innova tion 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 distor tions,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 conven tional 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 com mercial 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 tech niques 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 quantita tively 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 bor der definition facilitated automated or semiautomated con touring.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 phys ical 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 ana tomic 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 shim ming 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

ECG amplitude (mV)

2.5 3.0 T

2

1.5 T

1.5

1.0 T

1

0.0 T

0.5 T-Wave 0 0

500

1000

1500

2000

Time (msec)

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 fad ing results in a higher temporal resolution and more accu rate identification of the tagged lines, which facilitates identification of subtle changes in myocardial motion dur ing the cardiac cycle. In Figure 13 3, three CSPAMM18 line tagged cine frames out of a series with a temporal res olution of 50 frames per R R interval are displayed. These images were acquired by using spiral imaging (8 msec spi ral readouts) in conjunction with spectral spatial excitation for fat suppression and ramped flip angles for constant tag ging CNR. Data have confirmed the increase of CNR at 3 T for myocardial tagging techniques.13,14,19 Adenosine myo cardial stress tagging has been implemented at 3 T in com bination 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 paral lel to the image plane, combined with out of plane phase encoding gradients in the perpendicular section select direction.21 Implementation of strain encoding on higher field systems may increase the spatial and temporal resolu tion 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

13 HIGH FIELD CARDIOVASCULAR MAGNETIC RESONANCE

3

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

End-Diastole

End-Systole

Late Diastole

A

B

C

D

E

F

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 coro nary artery disease.23 However, visualization of the more dis tal 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 seg ments. 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 sur face receive coil. High resolution imaging was performed by using a three dimensional (3D) segmented k space gradi ent 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 dur ing free breathing.27 To minimize the sensitivity of the navi gator 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 sup press contamination of the navigator by spurious signal ori ginating 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 imag ing 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 dia stolic trigger delay28 was used to minimize intrinsic myocar dial 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

LV

Ao RV LM

B

A

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 sur rounding 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 coro nary 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 sup pression fast gradient echo sequences32 provided high spa tial resolution for the delineation of coronary arteries with good interobserver variability for quantitative vessel sharp ness 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 sig nificantly 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

A

Adiabatic T2 Prep

B

Cardiovascular Magnetic Resonance 173

13 HIGH FIELD CARDIOVASCULAR MAGNETIC RESONANCE

PA LAD

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

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 measure ment but also the differentiation between plaque compo nents 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 loca tion 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 pre pulse with the inversion time (TI)41 adapted for the pro longed 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 pre pulse, 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 in plane 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 suppres sion 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 gradi ent echo technique with ECG triggering and navigator gat ing. These results have been confirmed by others using a single breath hold43 and using navigator gating.43,44 Part icularly the use of adiabatic pulses for both inversion and reinversion44 has been reported to be robust to off resonance 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 extra cellular 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 impedi ments 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 alterna tive to contrast perfusion for assessment of ischemia, noncon trast 3 T stress with myocardial blood oxygenation level dependent imaging compares favorably with positron emission tomography (PET).48

RCA wall

A

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 inhomo geneities 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. Vol ume tracking allowed the suppression of tissue outside the planned voxel and provided the acquisition of high quality spectra during free breathing.12

Parallel Imaging Parallel imaging techniques such as simultaneous acquisi tion of spatial harmonics51 or sensitivity encoding for fast MRI (SENSE)52 are particularly interesting at higher field strengths and for selected applications. While the theoreti cally 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 abbre viated scanning times using parallel imaging with an accel eration 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 pro longed 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 coro nary 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 tech niques with prolonged signal readout (spiral techniques, EPI techniques).2 However, by using parallel imaging, echo trains or readout times can be shortened, which may sup port optimized image quality using such techniques at higher field strength. Sodickson and colleagues also demonstrated that parallel imaging may be employed ben eficially for first pass contrast agent bolus injection techni ques.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 method ology, 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

13 HIGH FIELD CARDIOVASCULAR MAGNETIC RESONANCE

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. Ongo ing 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 resolu tion, 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 steady state magnetization with subsequent relative loss in SNR, this effect can in turn also be used to improve the efficacy of myo cardial 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 per fusion imaging. Together with navigator technology, it may be possible to implement cardiac spectroscopy during free breathing58 while the enhanced SNR will support spectro scopic data acquisition in relatively small and well defined volumes. If extrinsic and intrinsic myocardial motion is suffi ciently constrained, diffusion imaging together with fiber tracking may allow for a more detailed insight into fiber archi tecture and the orientation of myocytes.59

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

After further refinement of imaging sequences and adap tation of hardware components (larger coil arrays), the spe cific 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.

176 Cardiovascular Magnetic Resonance

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.

Cardiovascular Magnetic Resonance 177

13 HIGH FIELD CARDIOVASCULAR MAGNETIC RESONANCE

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 func tion is a fundamental aim of noninvasive cardiac imaging. It forms the foundation upon which much of the assessment and management of myocardial dysfunction, ischemia, via bility, 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 seman tics, 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 develop ing 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 impor tant 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 improv ing 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 indi vidual patient’s management, serial studies need to be per formed 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 nonin vasive, 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 depen dent.9 The quantification of ventricular function is limited by a priori geometric assumptions that may provide a rea sonable assessment in the normal ventricle but are less reli able 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) echocardi ography, with the ability of imaging the heart in tomo graphic views, considerably improved the accuracy of LV volume measurement, providing the opportunity to derive the cardiac function from cardiac volumes by the area length and Simpson’s method of discs (Fig. 14 2).12 This, however, relies on good visualization of the entire endocar dial border, which is frequently not possible, although sev eral 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

14 ASSESSMENT OF CARDIAC FUNCTION

CHAPTER 14

ISCHEMIC HEART DISEASE

CW RV IVS ESD

LV

EDD

PVW

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 expe rience.16 Recently, there have been two important advances in echocardiography. On one hand, contrast echocardiogra phy 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) echocardiogra phy has emerged as a more accurate and reproducible approach to LV quantitation by removing the need for geo metric assumptions about the LV shape.20 Volumetric analy sis 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

RV

A

RA LA

Figure 14 2 Echo measurement by Simpson’s method of discs. This represents a more accurate measurement than M mode or area length 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, echocar diography 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 inade quate discrimination of the endocardial border. Overall, the 182 Cardiovascular Magnetic Resonance

Nuclear cardiology with radionuclide ventriculography is commonly used to measure ventricular function by measur ing 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 can not be obtained. The use of gated perfusion single photon emission computed tomography (SPECT) has allowed the development of 3D solutions to global and regional ventric ular 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 per fusion assessment,25 but it is not being performed solely to assess ventricular function. This technique has good repro ducibility,26 and ventricular volumes are reported as reli able,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 echocardi ography is more accurate than gated SPECT for estimation of LV remodeling.30 Recently, ECG gated F 18 fluorodeoxyglu cose positron emission tomography (PET) has been shown to be reasonably accurate for measurement of cardiac func tion,31 but this technique is time consuming and is performed only in viability studies. For these nuclear cardiology techni ques, the need for repeated radionuclide doses in follow up studies is problematic, especially for research, in which radia tion exposure must be justified in a milieu of competing tech nologies 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 accu rate 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, quali tative assessment of function, similar to that of echocardi ography. The main advantage of CMR, however, lies in its quantitative accuracy and reproducibility. There are two main methods for measuring the end diastolic and end systolic volumes. The earliest semiquanti tative 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 bor der can be traced, as well as the length (L) from the apex to the mitral annulus. The volume (V) for systole and dias tole is then easily calculated: V ¼ 0.85A2/L. A more accu rate 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 analy sis,38 it suffers from the same limitations that echocardiog raphy 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 contigu ous 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.

A

B

C

D

Cardiovascular Magnetic Resonance 183

14 ASSESSMENT OF CARDIAC FUNCTION

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 cineventriculo graphy,34 SPECT,35 and cine magnetic resonance imaging.36 Postprocessing tools allow fast and semiautomatic determina tion of LV function parameters from MDCT data in analogy to known CMR evaluation approaches. It cannot be consid ered a first line modality, owing to the exposure to radiation, the use of potentially nephrotoxic contrast, and the low tem poral 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 func tion is lower than that with CMR, with an interobserver varia bility of 4.2% for end diastolic volume, 5.4% for end systolic volume, and 4.1% for LVEF.37

ISCHEMIC HEART DISEASE

ED

ES

LV End Diastole LV End Systole LV Stroke Volume LV Ejection Fraction LV Mass LV Mass Index

219 mL 107 mL 112 mL 51% 243 g 125 g/m2

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 dimen sions to be measured simultaneously.

Short axis A3

A1

A2

Diastole

Systole

1 2 3 4 5 6 7

8

9

10

11 12 13

184 Cardiovascular Magnetic Resonance

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 arti facts and makes the cines independent of inflow enhance ment and based on the ratio of T2 to T1. This results in substantially improved blood myocardium contrast, espe cially 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 charac teristics, 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 differ ences 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 nowa days, further developments in this direction are limited, owing to physiologic constraints such as the risk of periph eral 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 imag ing 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 (five fold to eightfold acceleration) or improved spatial resolu tion 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 repro ducible in normal as well as abnormal ventricles.57–61 Much validation work was done by using conventional non breath hold cine CMR on older scanners, however, and currently, breath hold sequences are routinely employed in

200

CMR LV volume (mL)

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 reduc ing 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 develop ment, 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 pulmo nary disease.52 Global RV function is difficult to assess adequately by echocardiography, while radionuclide ventri culography 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 dis cussed in greater detail in Chapters 11 and 28.

150

100

50

R = 0.997

0 0

50

100

150

200

LV cast volume (mL) Figure 14 6 Validation of left ventricular (LV) volume measurements by CMR. CMR derived volume (vertical scale) is compared with the displacement volume of the left ventricular casts from human postmortem hearts. There is excellent accuracy, and the regression line slope is close to identity. Source: Data from Rehr RB, Malloy CR, Filipchuk NG, et al. Left ventricular volumes measured by MR imaging. Radiology. 1985;156:717 719. Cardiovascular Magnetic Resonance 185

14 ASSESSMENT OF CARDIAC FUNCTION

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 con siderable 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.

190

110 Cadaver RV mass (g)

LVSV (mL)

100 90 80 70 60

170

150

R = 0.94

R = 0.97

50 130

40 40

50

60

70

80

90

100

110

120

RVSV (mL)

250

200

150

100

R = 0.99

50 50

100

150

200

250

CMR LV mass (g) Figure 14 8 Validation of left ventricular (LV) mass with comparison of CMR derived mass against human cadaver hearts. Source: Data from Katz J, Millikem MC, Stray Gunderson J, Buja LM, Parkey RW, Mitchell JH, Peshock RM. Estimation of human myocardial mass with MR imaging. Radiology. 1988;169:495 498.

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

130

150

170

190

CMR RV mass (g)

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.

Cadaver LV mass (g)

ISCHEMIC HEART DISEASE

120

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 illu strated by considering the sample size required for a drug trial designed to show a 10 g decrease in LV mass with anti hypertensive 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% dif ference 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 mea sures 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

A

Semelka normal

Bogaert normal

Interobserver % variability

14 12 10 8 6 4 2

B no oga rm er al t

Be no llen rm ge a r Be l lle LV ng H er Be l D leng C e M r

Se no me rm lka a Se l m LV elk H a Se m D el C ka M

0

B

14

Interstudy % variability

12 10 8 6 4 2 0

C

Bellenger normal

Bellenger DCM

EDV ESV

Bogaert normal

Semelka DCM

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). Fig ure 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 num ber of phases acquired are related through the repetition time (TR) of the sequence. Modern CMR scanners with fas ter 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 com promise 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 com promise. 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 gat ing, 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 par tial 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 stud ies 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 signal to noise ratio and image quality. 3D imaging may even tually allow more partitions with thinner slices.  Analysis of the short axis slices is relatively straightfor ward, provided that the quality of the images is reason able (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 sys tolic 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 informa tion from the heart besides volumetry alone. For example, flow in the major vessels is easily measured with great accu racy 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 determin ing 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 (17 (42 (28 (58

130 (90 43 (25 87 (59 67 (58 107 (71 76 25 51 40 63

(59 (16 (39 (30 (48

99) 38) 66) 40) 91)

171) 62) 115) 76) 143) 93) 34) 63) 49) 77)

79 26 53 34 74

(62 (15 (41 (28 (57

126 (86 41 (22 85 (57 68 (59 108 (72 73 24 50 40 63

(56 (14 (37 (31 (48

97) 37) 65) 41) 91)

166) 59) 113) 77) 144) 90) 33) 62) 49) 78)

77 25 52 34 73

(60 (14 (40 (28 (57

122 (82 39 (20 83 (56 69 (60 109 (73 70 22 48 40 63

(53 (13 (36 (31 (48

95) 36) 64) 41) 90)

162) 57) 111) 78) 145) 87) 31) 60) 49) 78)

75 24 51 34 73

(58 (13 (39 (28 (56

118 (77 36 (18 81 (54 69 (60 110 (74 67 21 47 40 63

(50 (12 (34 (31 (49

93) 35) 63) 41) 89)

158) 55) 109) 78) 146) 84) 30) 59) 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 atrioven tricular 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 commu nication) 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 follow up 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 print out 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 (16 (39 (25 (21

108) 45) 71) 41) 48)

130 (87 46 (20 84 (58 65 (53 49 (28

172) 72) 109) 77) 70)

76 27 49 38 29

(57 (14 (37 (28 (19

94) 40) 61) 49) 39)

82 28 54 34 33

(59 (13 (38 (26 (20

105) 42) 70) 42) 46)

124 (81 42 (15 82 (56 67 (55 47 (26

166) 68) 108) 79) 68)

72 24 48 39 27

(53 (11 (36 (29 (17

90) 37) 60) 49) 37)

79 25 53 35 32

(56 (11 (37 (27 (19

101) 40) 69) 42) 45)

117 (75 37 (11 80 (55 69 (57 45 (24

160) 63) 106) 81) 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 variabil ity 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 cir cular flow pattern turning back toward the septum and outflow tract, persisting through diastole.77 With 3D auto mated 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 pro duced 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 echocar diography, 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 Stud ies using real time CMR are now being published; these too show that CMR is superior to dobutamine stress echocardi ography in patients with limited acoustic access.84,85 Several methods based on cine CMR have been sug gested 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 con traction, 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 per formance is produced. Tagging is discussed in more detail in Chapter 5.

+95% CI Mean –95% CI

20

30

40

50

60

70

80

LV end-systolic volume/BSA - males (mL/m2)

LV end-diastolic volume/BSA - males (mL/m2)

120 110 100 90 80 70 60 50 40

50 +95% CI 40 30 20

Mean –95% CI

10 0 20

30

40

70

Mean

60

–95% CI

LV mass/BSA males (g/m2)

LV ejection fraction - males (%)

90 +95% CI

50 40 20

30

40

50

60

60

70

80

Age (years)

Age (years)

80

50

70

120 110 100 90 80 70 60 50 40

80

+95% CI Mean –95% CI

20

30

40

Age (years)

50

60

70

80

Age (years)

120 110 100 90 80 70 60 50 40

+95% CI Mean –95% CI

20

30

40

50

60

70

80

LV end-systolic volume/BSA - females (mL/m2)

LV end-diastolic volume/BSA - females (mL/m2)

FEMALES 50 40

+95% CI

30

Mean

20

–95% CI

10 0 20

30

Age (years)

70

Mean

60

–95% CI

LV mass/BSA females (g/m2)

LV ejection fraction - females (%)

+95% CI

50 40 20

30

40

50

60

Age (years)

50

60

70

80

70

80

Age (years)

90 80

40

70

80

120 110 100 90 80 70 60 50 40

+95% CI Mean –95% CI 20

30

40

50

60

Age (years)

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

+95% CI Mean –95% CI

200 100 0 20

30

40

50

60

70

400 300 200 100

+95% CI Mean –95% CI

0 20

80

30

40

LV early/active peak filling rate ratio - males

50

60

70

80

70

80

Age (years)

Age (years) 8 7 6 5 4 3 2 1 0

+95% CI

Mean

–95% CI 20

30

40

50

60

70

80

Age (years)

LV active peak filling rate/BSA - females (mL/s/m2)

FEMALES LV early peak filling rate/BSA - females (mL/s/m2)

700 600 500

+95% CI Mean

400 300

–95% CI

200 100 0 20

30

40

50

60

70

80

400 300 +95% CI 200 Mean 100

–95% CI

0 20

30

Age (years) LV early/active peak filling rate ratio - females

ISCHEMIC HEART DISEASE

MALES

8 7 6 5 4 3 2 1 0

40

50

60

Age (years) +95% CI

Mean

–95% CI 20

30

40

50

60

70

80

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 car diac 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

14 ASSESSMENT OF CARDIAC FUNCTION

THE FUTURE

ISCHEMIC HEART DISEASE

35.

36. 37. 38.

39. 40. 41. 42.

43.

44.

45.

46. 47. 48. 49. 50. 51.

52. 53.

54.

55. 56.

patients with coronary artery disease. Am J Roentgenol. 2002;179: 1545–1550. Yamamuro M, Tadamuro E, Kubo S, et al. Cardiac functional analysis with multidetector row CT and segmental reconstruction algorithm: comparison with echocardiography, SPECT and MR imaging. Radiology. 2005;234:381–390. Shuijf JD, Bax JJ, Salm LP, et al. Noninvasive coronary imaging and assessment of left ventricular function using 16-slice computed tomography. Am J Cardiol. 2005;95:571–574. Mahnken AH, Heuzler D, Klotz E, et al. Determination of cardiac output with multislice spiral computed tomography: a validation study. Invest Radiol. 2004;39:451–454. Sievers B, Brandts B, Franken U, Trappe HJ. Single and biplane TrueFISP cardiovascular magnetic resonance for rapid evaluation of left ventricular volumes and ejection fraction. J Cardiovasc Magn Reson. 2004;6:593–600. Longmore DB, Klipstein RH, Underwood SR, et al. Dimensional accuracy of magnetic resonance in studies of the heart. Lancet. 1985;1:1360–1362. Rehr RB, Malloy CR, Filipchuk NG, et al. Left ventricular volumes measured by MR imaging. Radiology. 1985;156:717–719. Buser PT, Auffermann W, Holt WW, et al. Noninvasive evaluation of global left ventricular function with use of cine nuclear magnetic resonance. J Am Coll Cardiol. 1989;13:1294–1300. 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. 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. 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. 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. Fang W, Pereles FS, Bundy J, et al. Evaluating left ventricular function using real-time TrueFISP: a comparison with conventional MR techniques abstract. J Cardiovasc Magn Reson. 1999;1:310. 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. Tsao J, Boesiger P, Pruessmann KP. k-t BLAST and k-t SENSE: dynamic MRI with high frame rate exploiting spatiotemporal correlations. Magn Reson Med. 2003;50:1031–1042. Madore B, Hoge WS, Kwong R. Extension of the UNFOLD method to include free breathing. Magn Reson Med. 2006;55:352–362. 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. Gorcsan 3rd J, Murali S, Counihan PJ, Mandarino WA, Kormos RL. Right ventricular performance and contractile reserve in patients with severe heart failure: assessment by pressure-area relations and association with outcome. Circulation. 1996;94:3190–3197. Sandoval J, Bauerle O, Palomar A, et al. Survival in primary pulmonary hypertension: validation of a prognostic equation. Circulation. 1994;89:1733–1744. Rominger MB, Bachmann GF, Pabst W, Rau WS. Right ventricular volumes and ejection fraction with fast cine MR imaging in breathhold technique: applicability, normal values from 52 volunteers, and evaluation of 325 adult cardiac patients. J Magn Reson Imaging. 1999;10:908–918. Lorenz CH, Walker ES, Graham 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(suppl II): 233–239. Katz J, Millikem MC, Stray-Gunderson J, et al. Estimation of human myocardial mass with MR imaging. Radiology. 1988;169:495–498. 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.

194 Cardiovascular Magnetic Resonance

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.

Cardiovascular Magnetic Resonance 195

14 ASSESSMENT OF CARDIAC FUNCTION

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 mea sure 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, myocar dial 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 syn thetic beta 1 selective catecholamine agonist that binds to myocyte beta 1 receptors and increases intracellular calcium concentrations and myocardial contraction.8 In 1984, dobuta mine 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 methyltrans ferase, 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 steno sis, 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 incapacita tion, pulmonary disease, or chronic deconditioning.15 Other benefits of dobutamine include its tolerance for peripheral vein infusion and its effectiveness in both perfu sion 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 contrain dicated in patients with an obstructive cardiomyopathy, severe aortic stenosis, second and third degree heart block, sudden death syndrome, tachyarrhythmias, and pre vious episodes of dobutamine hypersensitivity. Although dobutamine exhibits positive inotropic activ ity, 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 recep tors. 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 incre ments of 0.5 mg up to a total of 2 mg. Side effects of atro pine 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 uti lizing DSE and DCMR (Table 15 1).21–25 Major complica tions among more than 5000 DSE tests included death, ventricular fibrillation, sustained ventricular tachycardia, complete atrioventricular block, acute myocardial infarc tion, cardiac rupture/ventricular septal defect, transient ischemic attack, and severe hypotension. Minor complica tions 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 discon tinue 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 ejec tion 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 test ing and emergency medications (Table 15 2) for treat ment of potential cardiac complications need to be readily available to improve the probability of a success ful 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, fol lowed 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 atro pine 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. Myo cardial ischemia is defined as an increase in wall motion score during testing or the persistence of severe hypokin esis 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 unex pected neurologic findings.21 DSE is useful for assessing risk of future myocardial infarction, death, and the need for coronary interven tion.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 cardiotho racic surgery, obstructive airway disease, or advanced age.17,43,44 In addition, there is a high interobserver varia bility 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 myocar dium 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.

Cardiovascular Magnetic Resonance 197

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. Electrocar diogram (ECG) interpretation within the magnetic for ST segment changes is difficult, owing to ECG distortion from the magnetohydrodynamic effect of pulsatile blood. There fore, a resting supine 12 lead ECG is first performed out side of the magnet room. Intravenous access is then established, and a phased array cardiac surface coil, bra chial 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 contin uously monitor the patient’s heart rhythm, respiratory rate, oxygen saturation, and blood pressure. The protocol is sim ilar 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 api cal 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 cal culate 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 develop ment of ischemia, as evidenced by a new wall motion abnor mality (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 12 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 two chamber 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 X ray coronary angiography. Study subjects received dobuta mine 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 dobuta mine 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 collea gues 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 cor rectly identified in the left anterior descending (LAD), 62% in the left circumflex (LCX), and 78% in the right cor onary 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 collea gues.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 quanti tative58 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%). Sen sitivity for detection of single , double , and triple vessel disease was 75%, 80%, and 100%, respectively. In the quantitative portion of the study, 39 patients with previ ously 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 prese lected 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 triple vessel 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 four chamber 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 quan titative and qualitative methods. Each of the early studies was relatively small, was per formed in a single institution, and enrolled participants in whom the investigators were aware of the extent of coro nary 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 prema turely 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 per form DCMR in patients with suspected CAD. In 1999, Nagel and colleagues published the first study using high dose DCMR (up to 40 mg/kg/min) and atropine. Compari son 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 seg ment 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 sensitiv ity 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 left anterior descending coronary artery; peach bars left circumflex coronary artery, blue bars 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 narrow ings. From these data, it was concluded that DCMR demon strates 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 col leagues 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 sig nificant 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

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**

Specificity 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 myo cardial segments not visualized with DSE, and 90% had eight or more. Continuous cines were obtained in three api cal (horizontal, vertical, and apical long axis) views and three LV short axis (base, mid, and apex) views during 5 minute 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 pro vides 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 revas cularization and were difficult to assess with DSE. The sensi tivity 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 com plication 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 identi fying 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 con tractile reserve compared DCMR with DSE and radionuclide studies. Development of tagging has better enabled quantita tive 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 impor tant determinants of potential recovery after vascularization. Dendale and colleagues were among the first to explore whether low dose DCMR represents a reasonable alterna tive64 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 collea gues 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 meta bolic 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 com pared with positron emission tomography (PET) with 18F fluorodeoxyglucose, 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 revasculari zation. There was an excellent correlation between LV end systolic 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 demon strate 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 con siderably. To determine the contribution of residual subepi cardial viable myocardium to global LV function, Bogaert and colleagues studied 12 patients with single vessel disease after successful reperfusion after their first transmural ante rior 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. Conse quently, 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 reper fusion, 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 exam ined percent intramyocardial circumferential segment shortening in a group of 20 patients after their first re perfused 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 dysfunc tional 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 func tion 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 revasculari zation. 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 revas cularization (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 com monly 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 estab lished. 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 multivar iate 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 cor onary 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 dobu tamine 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 preopera tive cardiovascular risk in cardiac and noncardiac surgery patients. Rerkpattanapipat and colleagues studied a group of 102 patients for preoperative risk evaluation for noncar diac surgery.73 Congestive heart failure, myocardial infarc tion, 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 con gestive heart failure (N ¼ 1) during surgery or the postop erative 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 tech nique, 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 tag ging at three LV short axis levels. Patients with new wall motion abnormalities underwent X ray coronary angiogra phy. 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 quali tative 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 poten tial assessment of LV diastolic function. Paetsch and collea gues 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 tag ging 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 quan titative analysis of the data typically takes less than 3 min utes. 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 mecha nism 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

15 WALL MOTION STRESS CARDIOVASCULAR MAGNETIC RESONANCE: ISCHEMIA, VIABILITY, AND PROGNOSIS

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.

ISCHEMIC HEART DISEASE

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 semiau tomatic method for tracking three dimensional (3D) car diac 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 set tings, build the mesh, and track the 3D images in approxi mately 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 cor onary vasodilation, producing up to fivefold increases in coronary artery blood flow. Dipyridamole raises the inter stitial levels of adenosine by inhibiting adenosine

Perfusion had the highest accuracy values for coronary ste noses 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 speci ficity. Both studies demonstrated that adenosine and dipyr idamole 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 consec utive patients who had perfusion CMR images acquired in three LV short axis views at rest and at peak dobutamine atropine stress.89a Combined perfusion and wall motion methodology improved sensitivity for CAD (70% diame ter stenosis) from 85% to 91% but provided lower specific ity (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 arti facts associated with relatively long image acquisition times. Faster CMR sequences have become available that allow for rapid image acquisition after exercise stress. Roest and col leagues studied 16 healthy subjects using bicycle exercise in the supine position on a magnetic resonance compatible bicycle ergometer (Fig. 15 14).90 Respiratory motion arti facts were minimized by brief end expiratory breath holding for all LV short axis images. Stroke volume and ejection frac tion increased in both ventricles, while the end systolic vol ume of both ventricles decreased. End diastolic volume remained the same. Similar results were obtained for GRE and with standard echo planar CMR. This study demon strates that exercise CMR can be used to assess physiologic changes in both the left and right ventricles. While other studies have examined supine bicycle ergo metry stress, Rerkpattanapipat and coworkers demon strated 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

15 WALL MOTION STRESS CARDIOVASCULAR MAGNETIC RESONANCE: ISCHEMIA, VIABILITY, AND PROGNOSIS

breakdown via the adenosine deaminase pathway, and it inhibits the facilitated uptake of adenosine into cells. Nor mally, there is also a mild rise in heart rate and fall in blood pressure. Because caffeine and theophylline block adeno sine binding and can significantly reduce adenosine’s vaso dilatory 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 dizzi ness, 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 adminis tered 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 co workers 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 dys function was 67%.88 Not surprisingly, there was an appar ent 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%, respec tively. 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 dipyrida mole. Wall motion and the perfusion reserve index exhib ited 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.

ISCHEMIC HEART DISEASE

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 clin ical utility of exercise induced LV wall motion abnormal ities 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 contractil ity 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, qualita tive and quantitative assessments of LV wall motion and strain are possible. While not yet perfected for routine clinical use, these approaches may soon provide real time 3D determinations of regional and global LV function during DCMR.

References 1. Master AM, Oppenheimer ET. A simple exercise tolerance test for circulatory efficiency with standard tables for normal individuals. Am J Med Sci. 1929;177(2):223–243. 2. Sechtem U, Sommerhoff BA, Markiewicz W, et al. 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(1):145–151. 3. Pattynama PM, de Roos A. MR evaluation of myocardial ischemia and infarction. Magn Reson Imaging. 1995;7(4):218–231. 4. 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(1):134–140. 5. Nagel E, Al-Saadi N, Fleck E. Cardiovascular magnetic resonance: myocardial perfusion. Herz. 2000;25(4):409–416. 6. 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. 7. Tuttle RR, Mills J. Dobutamine: development of a new catecholamine to selectively increase cardiac contractility. Circ Res. 1975;36(1):185–196. 8. Baig MW. Pharmacologic perfusion imaging: who needs it and why? Post grad Med. 1992;91(1):185–187, 190–192.

210 Cardiovascular Magnetic Resonance

9. Rabinovitch MA, Kalff V, Chan W, et al. The effect of dobutamine on exercise performance in patients with symptomatic ischemic heart disease. Am Heart J. 1984;107(1):81–85. 10. Brunton LL, Lazo JS, Parker KL, et al. Goodman & Gilman’s The Pharmacological Basis of Therapeutics, 11th ed New York: McGraw-Hill, 2001. 11. Pennell DJ, Underwood SR. The cardiovascular effects of dobutamine assessed by magnetic resonance imaging. Post Grad Med J. 1991;67 (Suppl 1):S1–8; discussion S8–9. 12. Iskandrian AS, Verani MS, Heo J. Pharmacologic stress testing: mechanism of action, hemodynamic responses, and results in detection of coronary artery disease. J Nucl Cardiol. 1994;1(1):94–111. 13. Kugiyama K, Inobe Y, Ohgushi M, et al. Comparison of coronary hemodynamics during infusions of dobutamine and adenosine in patients with angina pectoris. Jpn Circ J. 1998;62(1):1–6. 14. Warltier DC, Zyvoloski M, Gross GJ, Hardman HF, Brooks HL. Redistribution of myocardial blood flow distal to a dynamic coronary arterial stenosis by sympathomimetic amines: comparison of dopamine, dobutamine and isoproterenol. Am J Cardiol. 1981;48(2): 269–279. 15. Dagianti A, Agati, L, Sciomer S, et al. Stress echocardiography: comparison of exercise, dipyridamole and dobutamine in detecting and predicting the Extent of Coronary Artery Disease. J Am Coll Cardiol. July; 26(1):18–25 (8).

38. Cerqueira Weisman NJ, Dilsizian V. Standardized myocardial segmentation and nomenclature for tomographic imaging of the heart. Circulation. 2002;105:539–542. 39. Poldermans D, Arnese M, Fioretti PM, et al. Improved cardiac risk stratification in major vascular surgery with dobutamine-atropine stress echocardiography. J Am Coll Cardiol. 1995;26:648–653. 40. Davila-Roman VG, Waggoner AD, Sicard GA, et al. Dobutamine stress echocardiography predicts surgical outcome in patients with an aortic aneurysm and peripheral vascular disease. J Am Coll Cardiol. 1993;21:957–963. 41. Kertai MD, Boersma E, Bax JJ, et al. A meta-analysis comparing the prognostic accuracy of six diagnostic tests for predicting perioperative cardiac risk in patients undergoing major vascular surgery. Heart. 2003;89:1327–1334. 42. Fiuza M, Pereirinha A, de Padua F, Lopes MG. The classic exercise test or by imaging: new applications of Doppler echocardiography. Rev Port Cardiol. 1993;12(5):483–493, 409. 43. Chaudhry FA, Tauke JT, Alessandrini RS, et al. Enhanced detection of ischemic myocardium by transesophageal dobutamine stress echocardiography: comparison with simultaneous transthoracic echocardiography. Echocardiography. 2000;17(3):241–253. 44. Fedele F, Scopinaro F, Vitarelli A, et al. Characterization of reversible myocardial dysfunction by magnetic resonance imaging. Herz. 1994;19:210–220. 45. Vignola PA, Bloch A, Kaplan AD, Walker HJ, Chiotellis PN, Myers GS. Interobserver variability in echocardiography. J Clin Ultrasound. 1977;5(4):238–242. 46. Hundley WG, Hamilton CA, Rerkpattanapipat P. Magnetic resonance imaging assessment of cardiac function. Curr Cardiol Rep. 2003; 5(1):69–74. 47. Kaufman L, Crooks L, Sheldon P, et al. The potential impact of nuclear magnetic resonance imaging on cardiovascular diagnosis. Circulation. 1983;67:251–257. 48. Shan K, Constantine G, Sivananthan M, Flamm SD. Role of cardiac magnetic resonance imaging in the assessment of myocardial viability. Circulation. 2004;109:1328–1334. 49. Darty SN, O’Neal J, Wesley-Farrington D, Davis AD, Link KM, Hundley G. Cardiovascular magnetic resonance imaging. Prog Cardiovasc Nurs. 2004;19(2):60–67. 50. Vogel M, Stern H, Bauer R, Buhlmeyer K. Comparison of magnetic resonance imaging with cross-sectional echocardiography in the assessment of left ventricular mass in children without heart disease and in aortic isthmic coarctation. Am J Cardiol. 1992;69(9):941–944. 51. Kuijpers D, van Dijkman PR, Janssen CH, et al. Risk stratification with dobutamine cardiovascular magnetic resonance in patients suspected of myocardial ischemia. Eur Radiol. 2004;14(11):2046–2052. 52. Mazur W, Brucks S, Darty SN, Rerkpattanapipat P. Magnetic resonance imaging in detection and functional assessment of coronary artery disease. Curr Atheroscler Rep. 2002;4(2):128–133. 53. Cigarroa CG, deFilippi CR, Brickner ME, Alvarez LG, Wait MA, Grayburn PA. Dobutamine stress echocardiography identifies hibernating myocardium and predicts recovery of left ventricular function after coronary revascularization. Circulation. 1993;88:430–436. 54. van Rugge FP, van der Wall EE, Bruschke AV. New developments in pharmacologic stress imaging. Am Heart J. 1992;124(2):468–485. 55. Pennell DJ, Underwood SR, Manzara CC, et al. Magnetic resonance imaging during dobutamine stress in coronary artery disease. Am J Cardiol. 1992;70(1):34–40. 56. Baer FM, Theissen P, Smolarz K, et al. Dobutamine versus dipyridamole magnetic resonance tomography: safety and sensitivity in the detection of coronary stenosis. Z Kardiol. 1993;82(8):494–503. 57. Baer FM, Voth E, Theissen P, et al. Coronary artery disease: findings with GRE MR imaging and Tc-99mm-methoxyisobutyl-isonitrile SPECT during simultaneous dobutamine stress. Radiology. 1994; (193):203–209. 58. van Rugge FP, van der Wall EE, Bruschke AV, 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 (1):127–138. 59. van Rugge FP, van der Wall EE, de Roos A, Bruschke AV. Dobutamine stress magnetic resonance imaging for detection of coronary artery disease. J Am Coll Cardiol. 1993;(22):431–439. 60. Nagel E, Lehmkuhl HB, Bocksch W, et al. Noninvasive diagnosis of ischemia-induced wall motion abnormalities with the use of high-dose

Cardiovascular Magnetic Resonance 211

15 WALL MOTION STRESS CARDIOVASCULAR MAGNETIC RESONANCE: ISCHEMIA, VIABILITY, AND PROGNOSIS

16. 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;6(5):317–326. 17. 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. 18. Pereiro Dr LMP. Atropine; Practical Procedures: Issue 6 (1996) Article 5: Page 1 of 1. 19. Fetiveau R, Lanzarini L, Poli A, et al. The side effects and safety of the echo-dobutamine test. G Ital Cardiol. 1995;25(2):193–201. 20. Krenning BJ, Geleijnse ML, Poldermans D, Roelandt JR. Methodological analysis of diagnostic dobutamine stress echocardiography studies. Echocardiography. 2004;21(8):725–736. 21. Cortigiani L, Picano E, Coletta C, et al. Echo Persantine International Cooperative (EPIC) Study Group; Echo Dobutamine International Cooperative (EDIC) Study Group: 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. 22. Mertes H, Sawada SG, Ryan T, Segar DS, Kovacs R, Foltz J, Feigenbaum H. Symptoms, adverse effects, and complications associated with dobutamine stress echocardiography: experience in 1118 patients. Circulation. 1993;88:15–19. 23. Rodriguez Garcia MA, Iglesias-Garriz I, Picano E. Evaluation of the safety of stress echocardiography in Spain and Portugal. Rev Esp Cardiol. 2001;54(8):941–948. 24. Picano E, Mathias Jr W, Pingitore A, Bigi R, Previtali M. Safety and tolerability of dobutamine-atropine stress echocardiography: a prospective, multicentre study. Echo Dobutamine International Cooperative Study Group. Lancet. 1994;344(8931):1190–1192. 25. 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. 26. Wahl A, Paetsch I, Nagel E, et al. Safety and feasibility of high-dose dobutamine-atropine stress cardiovascular magnetic resonance for diagnosis of myocardial ischemia: experience in 1000 consecutive cases. Eur Heart J. 2004;25(14):1230–1236. 27. Kuijpers D, Janssen CH, van Dijkman PR, Oudkerk M. Dobutamine stress MRI Part I. Safety and feasibility of dobutamine cardiovascular magnetic resonance in patients suspected of myocardial ischemia. Eur Radiol. 2004;14(10):1823–1828. 28. Hamilton CA, Link KM, Salido TB, Epstein FH, Hundley WG, et al. Is imaging at intermediate doses necessary during dobutamine stress magnetic resonance imaging? J Cardiovascular Magn Reson. 2001;3 (4):297–302. 29. Tighe Jr JF, Steiman DM, Vernalis MN, Taylor AJ. Observer bias in the interpretation of dobutamine stress echocardiography. Clin Cardiol. 1997;20(5):449–454. 30. Pedone C, Bax JJ, Poldermans D, et al. Long-term prognostic value of ejection fraction changes during dobutamine-atropine stress echocardiography. Coron Artery Dis. 2005;16(5):309–313. 31. Autenrieth G, Angermann C, Goss F, Bolte HD. Stress echocardiography in patients with coronary heart disease. Verh Dtsch Ges Inn Med. 1977;83:231–236. 32. Limacher MC, Quinones MA, Poliner LR, Nelson JG, Winters WL, Waggoner AD. Detection of coronary artery disease with exercise two dimensional echocardiography. Circulation. 1983;67:1211–1218. 33. Paul B, Kasliwal RR. Dobutamine stress echocardiography-methodology, clinical applications and current perspectives. J Assoc Physicians India. 2004;52:653–657. 34. Pingitore A, Picano E, Colosso MQ, et al. The atropine factor in pharmacologic stress echocardiography. J Am Coll Cardiol. 1996;27(5): 1164–1170. 35. Kamaran M, Teague SM, Finkelhor RS, Dawson N, Bahler RC. Prognostic value of dobutamine stress echocardiography in patients referred because of suspected coronary artery disease. Am J Cardiol. 1995;76(12):887–891. 36. Mathias JW, Doya EH, Duprat FR, et al. Detection of myocardial ischemia during dobutamine stress echocardiography: correlation with coronary cineangiography. Arq Bras Cardiol. 1993;60(4): 229–234. 37. Armstrong WF, Pellikka PA, Ryan T, Crouse L, Zoghbi W. Stress echocardiography: recommendations for performance and interpretation of stress echocardiography. J Am Soc Echocardiogr. 1998;11(1):98–103.

ISCHEMIC HEART DISEASE

61.

62.

63.

63a. 64.

65.

66.

67.

68.

69. 70. 71. 71a.

72.

73.

dobutamine stress magnetic resonance imaging: comparison with dobutamine stress echocardiography. Circulation. 1999;99:763–770. 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. 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. Nandalur KR, Dwamena BA, Choudhri AF, et al. Diagnostic performance of stress cardiac magnetic resonance imaging in the detection of coronary artery disease: a meta-analysis. J Am Coll Cardiol. 2007;50:1343–1353. Kelle S, Hamdan A, Schnackenburg B, et al. Dobutamine stress cardiovascular magnetic resonance at 3 Tesla. J Cardiovasc Magn Reson. 2008;10(1):44. 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(1):134–140. Saito I, Watanabe S, Masuda Y. Detection of viable myocardium by dobutamine stress tagging magnetic resonance imaging with threedimensional analysis by automatic trace method. Jpn Circ J. 2000;64 (7):487–494. 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(4):1006–1015. 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;82(9):1149–1151, A10. Bogaert J, Maes A, Rademakers FE. Functional recovery of subepicardial myocardial tissue in transmural myocardial infarction after successful reperfusion: an important contribution to the improvement of regional and global left ventricular function. Circulation. 1999;99(1):36–43. Geskin G, Kramer CM, Rogers WJ, et al. Quantitative assessment of myocardial viability after infarction by dobutamine magnetic resonance tagging. Circulation. 1998;98(3):217–223. 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;109(18):2172–2174. Hundley WG, Morgan TM, Neagle CM, Hamilton CA, Rerkpattanapipat P, Link KM. Magnetic resonance imaging determination of cardiac prognosis. Circulation. 2002;106(18):2328–2333. Wallace EL, Morgan TM, Walsh TF, et al. Dobutamine cardiac magnetic resonance predicts cardiac prognosis in women with known or suspected ischemic heart disease. J Am Coll Cardiol Img. 2009;2: 299–307. 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. Rerkpattanapipat P, Morgan TM, Neagle CM, et al. Assessment of preoperative cardiac risk with magnetic resonance imaging. Am J Cardiol. 2002;90(4):416–419.

212 Cardiovascular Magnetic Resonance

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 cardiovas cular 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 epidemi ologic and economic consequences of introducing new techniques for diagnosing coronary artery disease (CAD). The CMR technique offers an almost unbeatable versatil ity 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 pro posed 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 produc tion of nitric oxide is reduced, which promotes the recruit ment 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 produc tion of matrix degrading enzymes9 and collagen production by macrophages,10 which can cause fibrous cap destabiliza tion. 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 myocar dial infarction.12,14–17 Figure 16 1 demonstrates in approxi mately 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 com position. While direct visualization of plaques in the coro nary 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 clin ical 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 hyper emic 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

15

10

5

0 0

10

20

30

40

50

60

70

80

90 100

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 high grade 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 hemody namics 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 per formed in patients with CAD. With this technique, a correlation of CFR with percent area stenosis was demon strated; 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 ste nosed coronary arteries to hyperemic flow supplied by non stenosed 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 exam ple, 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 tech nique must guarantee a linear relationship between CMR derived 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 man datory 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 supply demand 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 cor onary artery, increased oxygen demand is not met by sup ply, 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 Con sequently, when contractile function was assessed during vasodilator induced stress, the sensitivity for detection of CAD in clinical studies was very low27 (Fig. 16 3). More over, 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 pro tons 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 perfu sion 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 orien tations 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 endog enous 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 pharma cologically induced hyperemia, oxygen content will increase in well perfused myocardium but not in myocar dium 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 sig nal change being required for a reliable stenosis detec tion.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

80

80

60

Specificity (%)

Sensitivity (%)

ISCHEMIC HEART DISEASE

Detection of coronary artery disease: >50% diameter stenosis

*

40 20

60 40 20

0

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 deoxygen ated 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 acqui sition 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 techni ques either under development or in clinical routine are designed to meet the following requirements: (1) provide high spatial resolution to permit detection of small suben docardial 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 abnor mally 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 mag netization schemes are most important. Since the first pass of contrast media during hyperemia lasts only 5 to 10 sec onds, 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 ische mia 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 perfu sion 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 tri sodium, Lantheus Medical, Billerica, MA) and poly lysine Gd compounds,45,46 differences between normally per fused 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 con centration of contrast media in the voxel but also its intra voxel distribution (homogeneous versus inhomogeneous) determines its T2* shortening effect, rendering such a T2* approach susceptible to vessel architecture (vessel ori entation, intervessel distance). Since T2* approaches are affected by these geometry factors (which are not known a priori), measurements of tissue contrast media concentra tions by T1 techniques are considered superior.

Hyperpolarized Contrast Media Conventional Gd chelates modulate the MR signal by accelerating the relaxation rates of surrounding water pro tons. 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 mag netic 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 liq uid 13C in various compounds) could be increased by a fac tor 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

100

0.50

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

0 0

A

10 Time (image #)

20

0

B

0.25

0.5

0.75

1

1-specificity

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 tra cers, which is observed after rest injection or rest reinjec tions. 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 assess ment of the optimal contrast media dose for perfusion imag ing. 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 T1 weighted sequences showed absence of a susceptibility induced 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 simi lar to radioactive tracers, where radioactivity is directly pro portional 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 avail able for a short time period also allows to study for exam ple, 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 possi ble.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 conse quently 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 multi slice 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 condi tions of magnetization [steady state free precession (SSFP) sequences] appear promising, since they preserve magneti zation 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 vol unteer 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. How ever, 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 fac tor 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 satu ration 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 corre lations 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 res olution 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 ren ders 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 min imal cardiac motion, that is, into late systole and middias tole. 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

13

SNR

10

7

4

1 0

0.2

0.4

0.6

0.8

1.0

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 satu ration 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, prep aration 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 pre contrast 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. Increas ing 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 How ever, 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 disadvanta geous, and larger clinical trials would be required to dem onstrate a potential advantage of perfusion imaging at 3.0 T. In a recent study, we implemented a novel spatiotem poral correlation approach for perfusion data (5 kt SENSE) 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 spa tial 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 visu ally differentiate hypoperfused regions from artifacts, both typically causing lower signal during first pass con ditions.2,62,71,74 However, the advantage of experience is traded for subjectivity, that is, reduced interreader reproduc ibility. Therefore, a computer assisted or automatic analysis of perfusion data would be highly desirable and could poten tially 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, ade quate reproducibility should rely not on individual experi ence 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 experi ence 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 pre defined 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 reduc tion 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 infe rior 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 semi quantitative 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 interac tion 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 quantita tive 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 por tions 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 analy sis 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 inten sity time curves have to correct for inhomogeneous coil sensitivities. For this purpose, it has been suggested to sub tract precontrast signal from the first pass signal intensi ties.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 applicabil ity 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 rela tively close correlation with perfusion data is reported in both animal79,84 and human studies,25 at least for the low flow range. Its robustness could also be demonstrated in a multicenter trial.1 This upslope parameter is relatively insen sitive 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 param eter 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 con centration 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 esti mates 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 homo geneously 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 vol ume 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, hyperpo larized 13C contrast media (see below) can act as intravas cular 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  e F=lðt tÞdt ¼ CA ðtÞ  F  e F=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  e

PS=F

Þ

(3)

where PS is the permeability surface area product (in milliters per gram per second) and Ki is the unidirectional influx con stant (in milliters per gram per second). For all these models, knowledge about the arterial input function, that is, the mea surement 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 mea surements. 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 lon ger 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 reso lution 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 con trast 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 condi tions 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 condi tions 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 unidirec tional 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 differ ences 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 rate 222 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 perfu sion.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 exper imental 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 quantifica tion45,86 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 nor mal 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 per formed 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) N

Authors

Prep Slices*

CM

Dose

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

99 99 241

SR multi SR multi SR multi

Gd DTPA Gd DTPA Gd DTPA BMA

0.05 0.10 0.10

Rest

Stress

Analysis

Reference Standard

Sens

Spec

AUC

+ + +

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%

0.80

Ado Ado Ado

Visual Upslopesubendo Visual

QCA (
50%) QCA (
50%) QCA (
50%)

93% 91% 90%

75% 78% 73%

0.90 0.91 0.87

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 Multicenter Studies Wolff et al.2 Giang et al.1 Schwitter et al.3

+ + + +

0.76 0.91 0.93 0.93 0.90}

#

+ +

16 STRESS CARDIOVASCULAR MAGNETIC RESONANCE: MYOCARDIAL PERFUSION

Card ovascu ar Magnet c 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 con trast medium first pass is most widely used to quantita tively assess perfusion data.25,33,58–61,69,78,108 From a rest stress 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 angi ography (QCA) (at a CFR threshold of 1.5).61 For an arte rial 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 periph eral administration of the same contrast media dose) yielded a similar performance with a sensitivity and speci ficity of 88% and 90%, respectively, for the detection of ste noses with 75% or more area reduction on QCA.69 When CFR was determined by perfusion CMR in women, a spec ificity 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 perfor mance of the hyperemic upslope parameter when thresh olds 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 subendo cardial 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 myo cardial perfusion in humans. However, if perfusion CMR should become the first line method in clinical routine, its robustness, reliability, and safety have to be demon strated 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) per formed 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 sensitiv ity/specificity of 91/78% and 94/71%, respectively.1 Thus, this multicenter trial confirmed the high diagnostic perfor mance 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 increas ing number of participating sites, homogeneity of data qual ity, 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 advan tages 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 MR IMPACT multicenter program (MR Imaging for Myocardial Perfusion Assessment in Coronary Artery Disease Trial),3 241 patients underwent both invasive coronary angiogra phy 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 demon strated 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 multi vessel 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 non invasive detection of severe and, thus, dangerous coronary lesions by CMR and to treat these lesions before they rup ture, 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 progres sion 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.

References 1. Giang T, Nanz D, Coulden R, et al. Detection of coronary artery disease by magnetic resonance myocardial perfusion imaging with various contrast medium doses: first European multicenter experience. Eur Heart J. 2004;25:1657–1665. 2. Wolff S, Schwitter J, Coulden R, et al. Myocardial first-pass perfusion magnetic resonance imaging: a multicenter dose-ranging study. Circulation. 2004;110:732–737. 3. Schwitter J, Wacker CM, van Rossum AC, et al. Magnetic resonance imaging for myocardial perfusion assessment in coronary artery disease trial: comparison of perfusion cardiac magnetic resonance with single photon emission computed tomography for the detection of coronary artery disease in a multicenter, multivendor, randomized trial. Eur Heart J. 2008;29:480–489. 4. Celermajer DS, Sorensen KE, Gooch VM, et al. Non-invasive detection of endothelial dysfunction in children and adults at risk of atherosclerosis. Lancet. 1992;340:1111–1115. 5. Suwaidi JA, Hamasaki S, Higano ST, et al. Long-term follow-up of patients with mild coronary artery disease and endothelial dysfunction. Circulation. 2000;101:948–954. 6. Clarkson P, Celermajer DS, Donald AE, et al. Impaired vascular reactivity in insulin-dependent diabetes mellitus is related to disease duration and low density lipoprotein cholesterol levels. J Am Coll Cardiol. 1996;28:573–579. 7. Lefer AM, Ma XL. Decreased basal nitric oxide release in hypercholesterolemia increases neutrophil adherence to rabbit coronary artery endothelium. Arterioscler Thromb. 1993;13:771–776. 8. Barker SGE, Beesley JE, Baskerville PA, Martin JF. The influence of the adventitia on the presence of smooth muscle cells and macrophages in the arterial intima. Eur J Vasc Endovasc Surg. 1995;9: 222–227. 9. Galis ZS, Sukhova GK, Lark MW, Libby P. Increased expression of matrix metalloproteinases and matrix degrading activity in vulnerable regions of human atherosclerotic plaques. J Clin Invest. 1994;94:2493–2503. 10. Libby P. Molecular bases of the acute coronary syndromes. Circulation. 1995;91:2844–2850. 11. Davies MJ. A macro and micro view of coronary vascular insult in ischemic heart disease. Circulation. 1990;82:Ii38–Ii46. 12. Bruschke AV, Kramer Jr JR, Bal ET, et al. The dynamics of progression of coronary atherosclerosis studied in 168 medically treated patients who underwent coronary arteriography three times. Am Heart J. 1989;117:296–305. 13. Wilcox JN, Smith KM, Schwartz SM, Gordon D. Localization of tissue factor in the normal vessel wall and in the atherosclerotic plaque. Proc Natl Acad Sci U S A. 1989;86:2839–2843. 226 Cardiovascular Magnetic Resonance

14. Ellis S, Alderman E, Cain K, et al. Prediction of risk of anterior myocardial infarction by lesion severity and measurement method of stenoses in the left anterior descending coronary distribution: a CASS Registry Study. J Am Coll Cardiol. 1988;11:908–916. 15. Waters D, Lesperance J, Francetich M, et al. A controlled clinical trial to assess the effect of a calcium channel blocker on the progression of coronary atherosclerosis. Circulation. 1990;82:1940–1953. 16. Ojio S, Takatsu H, Tanaka T, et al. Considerable time from the onset of plaque rupture and/or thrombi until the onset of acute myocardial infarction in humans: coronary angiographic findings within 1 week before the onset of infarction. Circulation. 2000;102:2063–2069. 17. Kemp H, Kronmal R, Vlietstra R, Frye R. Seven year survival of patients with normal or near normal coronary arteriograms: a CASS registry study. J Amer Coll Cardiol. 1986;7:479–483. 18. Naghavi M, Libby P, Falk E, et al. From vulnerable plaque to vulnerable patient: a call for new definitions and risk assessment strategies: Part I. Circulation. 2003;108:1664–1672. 19. Gould KL, Kirkeeide RL, Buchi M. Coronary flow reserve as a physiologic measure of stenosis severity. J Am Coll Cardiol. 1990;15:459–474. 20. 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: a comparison with positron emission tomography. Circulation. 2000;101:2696–2702. 21. Uren NG, Melin JA, De Bruyne B, Wijns W, Baudhuin T, Camici PG. Relation between myocardial blood flow and the severity of coronary artery stenosis. N Engl J Med. 1994;330:1782–1788. 22. Di Carli M, Czernin J, Hoh CK, et al. Relation among stenosis severity, myocardial blood flow, and flow reserve in patients with coronary artery disease. Circulation. 1995;91:1944–1951. 23. Sambuceti G, Parodi O, Marcassa C, et al. Alterations in regulation of myocardial blood flow in one-vessel coronary artery disease determined by positron emission tomography. Am J Cardiol. 1993;72: 538–543. 24. Picano E, Parodi O, Lattanzi F, et al. Assessment of anatomic and physiological severity of single-vessel coronary artery lesions by dipyridamole echocardiography: comparison with positron emission tomography and quantitative arteriography. Circulation. 1994;89: 753–761. 25. 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;103:2230–2235.

48. 49. 50. 51. 52. 53. 54. 55. 56. 57.

58.

59.

60. 61. 62. 63.

64. 65.

66. 67. 68. 69. 70. 71.

intravascular contrast agent NC100150 injection: preliminary results in eight patients. Radiology. 2001;221:822–826. Hall DA, Maus DC, Gerfen GJ, et al. Polarization-enhanced NMR spectroscopy of biomolecules in frozen solution. Science. 1997;276:930–932. Golman K, Axelsson O, Johannesson H, et al. Parahydrogen-induced polarization in imaging: subsecond 13C angiography. Magn Reson Med. 2001;46:1–5. Johansson E, Mansson S, Wirestam R, et al. Cerebral perfusion assessment by bolus tracking using hyperpolarized 13C. Magn Reson Med. 2004;51:464–472. Schwitter J. Myocardial perfusion imaging by cardiac magnetic resonance. J Nucl Cardiol. 2006;13:841–854. Edelman RR, Li W. Contrast-enhanced echo-planar MR imaging of myocardial perfusion: preliminary study in humans. Radiology. 1994;190:771–777. Schwitter J, Debatin JF, von Schulthess GK, McKinnon GC. Normal myocardial perfusion assessed with multishot echo-planar imaging. Magn Reson Med. 1997;37:140–147. 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. Schwitter J, Sakuma H, Saeed M, Wendland MF, Higgins CB. Very fast cardiac imaging. Magn Reson Imaging Clin N Am. 1996;4:419–432. Saeed M, Wendland MF, Sakuma H, et al. Coronary artery stenosis: detection with contrast-enhanced MR imaging in dogs. Radiology. 1995;196:79–84. Manning WJ, Atkinson DJ, Grossman W, Paulin S, Edelman RR. First-pass nuclear magnetic resonance imaging studies using gadolinium-DTPA in patients with coronary artery disease. J Am Coll Cardiol. 1991;18:959–965. Lauerma K, Virtanen KS, Sipila LM, Hekali P, Aronen HJ. Multislice MRI in assessment of myocardial perfusion in patients with singlevessel proximal left anterior descending coronary artery disease before and after revascularization. Circulation. 1997;96:2859–2867. Matheijssen NA, Louwerenburg HW, van Rugge F, et al. Comparison of ultrafast dipyridamole magnetic resonance imaging with dipyridamole SestaMIBI SPECT for detection of perfusion abnormalities in patients with one-vessel coronary artery disease: assessment by quantitative model fitting. Magn Reson Med. 1996;35:221–228. Eichenberger AC, Schuiki E, Kochli VD, et al. Ischemic heart disease: assessment with gadolinium-enhanced ultrafast MR imaging and dipyridamole stress. J Magn Reson Imaging. 1994;4:425–431. Al-Saadi N, Nagel E, Gross M, et al. Noninvasive detection of myocardial ischemia from perfusion reserve based on cardiovascular magnetic resonance. Circulation. 2000;101:1379–1383. 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. Klocke FJ, Simonetti OP, Judd RM, et al. Limits of detection of regional differences in vasodilated flow in viable myocardium by first-pass magnetic resonance perfusion imaging. Circulation. 2001;104:2412–2416. Schreiber WG, Schmitt M, Kalden P, et al. Dynamic contrastenhanced myocardial perfusion imaging using saturation-prepared TrueFISP. J Magn Reson Imaging. 2002;16:641–652. Wang Y, Moin K, Akinboboye O, Reichek N. Myocardial first pass perfusion: steady-state free precession versus spoiled gradient echo and segmented echo planar imaging. Magn Reson Med. 2005;54:1123–1129. Pruessmann KP, Weiger M, Scheidegger MB, Boesiger P. SENSE: sensitivity encoding for fast MRI. Magn Reson Med. 1999;42:952–962. Kellman P, Derbyshire JA, Agyeman KO, McVeigh ER, Arai AE. Extended coverage of first-pass perfusion imaging using sliceinterleaved TSENSE. Magn Reson Med. 2004;51:200–204. Plein S, Ryf S, Schwitter J, Radjenovic A, Boesiger P, Kozerke S. Dynamic contrast-enhanced myocardial perfusion MR imaging accelerated with k-t SENSE. Magn Reson Med. 2007;58:777–785. Nagel E, Klein C, Paetsch I, et al. Magnetic resonance perfusion measurements for the noninvasive detection of coronary artery disease. Circulation. 2003;108:432–437. 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. Walsh EG, Doyle M, Lawson MA, Blackwell GG, Pohost GM. Multislice first-pass myocardial perfusion imaging on a conventional clinical scanner. Magn Reson Med. 1995;34:39–47.

Cardiovascular Magnetic Resonance 227

16 STRESS CARDIOVASCULAR MAGNETIC RESONANCE: MYOCARDIAL PERFUSION

26. 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: 943–951. 27. 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 99 mTc-MIBI single photon emission computed tomography. Circulation. 1993;87:345–354. 28. Reeder SB, Atalay MK, McVeigh ER, Zerhouni EA, Forder JR. Quantitative cardiac perfusion: a noninvasive spin-labeling method that exploits coronary vessel geometry. Radiology. 1996;200:177–184. 29. Poncelet BP, Koelling TM, Schmidt CJ, et al. Measurement of human myocardial perfusion by double-gated flow alternating inversion recovery EPI. Magn Reson Med. 1999;41:510–519. 30. Wendland MF, Saeed M, Lauerma K, et al. Endogenous susceptibility contrast in myocardium during apnea measured using gradient recalled echo planar imaging. Magn Reson Med. 1993;29:273–276. 31. Belle V, Kahler E, Waller C, et al. In vivo quantitative mapping of cardiac perfusion in rats using a noninvasive MR spin-labeling method. J Magn Reson Imaging. 1998;8:1240–1245. 32. Fieno DS, Shea SM, Li Y, et al. 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. 33. Bertschinger KM, Nanz D, Buechi M, et al. Magnetic resonance myocardial first-pass perfusion imaging: parameter optimization for signal response and cardiac coverage. J Magn Reson Imaging. 2001;14: 556–562. 34. 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. 35. Reeder S, Faranesh A, Boxerman J, McVeigh E. In vivo measurement of T2* and field inhomogeneity maps in the human heart at 1.5 T. Magn Reson Med. 1998;39:988–998. 36. 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. 37. Atalay MK, Reeder SB, Zerhouni EA, Forder JR. Blood oxygenation dependence of T1 and T2 in the isolated, perfused rabbit heart at 4.7 T. Magn Reson Med. 1995;34:623–627. 38. Saeed M, Wendland MF, Yu KK, Li HT, Higgins CB. Dual effects of gadodiamide injection in depiction of the region of myocardial ischemia. J Magn Reson Imaging. 1993;3:21–29. 39. Sakuma H, O’Sullivan M, Lucas J, et al. Effect of magnetic susceptibility contrast medium on myocardial signal intensity with fast gradient-recalled echo and spin-echo MR imaging: initial experience in humans. Radiology. 1994;190:161–166. 40. Schwitter J, Saeed M, Wendland MF, et al. Assessment of myocardial function and perfusion in a canine model of non-occlusive coronary artery stenosis using fast magnetic resonance imaging. JMRI. 1998;9:101–110. 41. Schwitter J, Saeed M, Wendland MF, et al. Influence of the severity of myocardial injury on the distribution of macromolecules: extra- versus intravascular gadolinium-based MR contrast media. J Am Coll Cardiol. 1997;30:1086–1094. 42. 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. 43. Knuesel PR, Nanz D, Wyss C, et al. Characterization of dysfunctional myocardium by positron emission tomography and magnetic resonance: relation to functional outcome after revascularization. Circulation. 2003;108:1095–1100. 44. Kraitchman DL, Chin BB, Heldman AW, Solaiyappan M, Bluemke DA. MRI detection of myocardial perfusion defects due to coronary artery stenosis with MS-325. J Magn Reson Imaging. 2002;15:149–158. 45. Wilke N, Kroll K, Merkle H, et al. Regional myocardial blood volume and flow: first-pass MR imaging with polylysine-Gd-DTPA. J Magn Reson Imaging. 1995;5:227–237. 46. Kraitchman DL, Wilke N, Hexeberg E, et al. Myocardial perfusion and function in dogs with moderate coronary stenosis. Magn Reson Med. 1996;35:771–780. 47. Bjerner T, Johansson L, Ericsson A, et al. First-pass myocardial perfusion MR imaging with outer-volume suppression and the

ISCHEMIC HEART DISEASE

72. Araoz PA, Glockner JF, McGee KP, et al. 3 Telsa MR imaging provides improved contrast in first-pass myocardial perfusion imaging over a range of gadolinium doses. J Cardiovasc Magn Reson. 2005;7: 559–564. 73. Plein S, Schwitter J, Suerder D, et al. k-t SENSE-accelerated myocardial perfusion MR imaging at 3.0 Tesla: comparison with 1.5 Tesla. Radiology. 2008;249:493–500. 74. Hartnell G, Cerel A, Kamalesh M, et al. Detection of myocardial ischemia: value of combined myocardial perfusion and cineangiographic MR imaging. AJR Am J Roentgenol. 1994;163:1061–1067. 75. Holland AE, Goldfarb JW, Edelman RR. Diaphragmatic and cardiac motion during suspended breathing: preliminary experience and implications for breath-hold MR imaging. Radiology. 1998;209: 483–489. 76. McConnell MV, Khasgiwala VC, Savord BJ, et al. Prospective adaptive navigator correction for breath-hold MR coronary angiography. Magn Reson Med. 1997;37:148–152. 77. Chuang ML, Chen MH, Khasgiwala VC, et al. Adaptive correction of imaging plane position in segmented k-space cine cardiac MRI. J Magn Reson Imaging. 1997;7:811–814. 78. Keijer JT, van Rossum AC, van Eenige MJ, et al. Magnetic resonance imaging of regional myocardial perfusion in patients with singlevessel coronary artery disease: quantitative comparison with (201) Thallium-SPECT and coronary angiography. J Magn Reson Imaging. 2000;11:607–615. 79. 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. 80. Panting JR, Gatehouse PD, Yang GZ, et al. Abnormal subendocardial perfusion in cardiac syndrome X detected by cardiovascular magnetic resonance imaging. N Engl J Med. 2002;346:1948–1953. 81. Doyle M, Fuisz A, Kortright E, et al. The impact of myocardial flow reserve on the detection of coronary artery disease by perfusion imaging methods: an NHLBI WISE study. J Cardiovasc Magn Reson. 2003;5:475–485. 82. Keijer JT, van Rossum A, van Eenige M, et al. Semiquantitation of regional myocardial blood flow in normal human subjects by first-pass magnetic resonance imaging. Am Heart J. 1995;130: 893–901. 83. Lombardi M, Jones RA, Westby J, et al. Use of the mean transit time of an intravascular contrast agent as an exchange-insensitive index of myocardial perfusion. J Magn Reson Imaging. 1999;9:402–408. 84. Christian TF, Rettmann DW, Aletras AH, et al. Absolute myocardial perfusion in canines measured by using dual-bolus first-pass MR imaging. Radiology. 2004;232:677–684. 85. Kety S. The theory and application of the exchange of inert gas in lung and tissues. Pharmacol Rev. 1951;3:1–41. 86. Jerosch Herold M, Hu X, Murthy NS, Rickers C, Stillman AE. Magnetic resonance imaging of myocardial contrast enhancement with MS-325 and its relation to myocardial blood flow and the perfusion reserve. J Magn Reson Imaging. 2003;18:544–554. 87. Kroll K, Wilke N, Jerosch Herold M, et al. Modeling regional myocardial flows from residue functions of an intravascular indicator. Am J Physiol. 1996;271:H1643–H1655. 88. Jerosch-Herold M, Swingen C, Seethamraju RV. Myocardial blood flow quantification with MRI by model-independent deconvolution. Med Phys. 2002;29:886–897. 89. Wilke N, Jerosch HM, Wang Y, et al. Myocardial perfusion reserve: assessment with multisection, quantitative, first-pass MR imaging. Radiology. 1997;204:373–384. 90. Larsson HB, Fritz Hansen T, Rostrup E, et al. Myocardial perfusion modeling using MRI. Magn Reson Med. 1996;35:716–726. 91. Diesbourg L, Prato F, Wisenberg G, et al. Quantification of myocardial blood flow and extracellular volumes using a bolus injection of Gd-DTPA: kinetic modelling in canine ischemic disease. Mag Reson Med. 1992;23:239–253. 92. Tong CY, Prato FS, Wisenberg G, et al. Measurement of the extraction efficiency and distribution volume for Gd-DTPA in normal and diseased canine myocardium. Magn Reson Med. 1993;30: 337–346.

228 Cardiovascular Magnetic Resonance

93. Fritz-Hansen T, Rostrup E, Larsson HB, et al. Measurement of the arterial concentration of Gd-DTPA using MRI: a step toward quantitative perfusion imaging. Magn Reson Med. 1996;36:225–231. 94. Larsson HB, Stubgaard M, Sondergaard L, Henriksen O. In vivo quantification of the unidirectional influx constant for Gd-DTPA diffusion across the myocardial capillaries with MR imaging. J Magn Reson Imaging. 1994;4:433–440. 95. Gatehouse PD, Elkington AG, Ablitt NA, et al. Accurate assessment of the arterial input function during high-dose myocardial perfusion cardiovascular magnetic resonance. J Magn Reson Imaging. 2004; 20:39–45. 96. Wendland MF, Saeed M, Yu KK, et al. Inversion recovery EPI of bolus transit in rat myocardium using intravascular and extravascular gadolinium-based MR contrast media: dose effects on peak signal enhancement. Magn Reson Med. 1994;32:319–329. 97. Judd RM, Atalay MK, Rottman GA, Zerhouni EA. Effects of myocardial water exchange on T1 enhancement during bolus administration of MR contrast agents. Magn Reson Med. 1995;33:215–223. 98. Judd RM, Reeder SB, May Newman K. Effects of water exchange on the measurement of myocardial perfusion using paramagnetic contrast agents. Magn Reson Med. 1999;41:334–342. 99. Donahue KM, Burstein D, Manning WJ, Gray ML. Studies of GdDTPA relaxivity and proton exchange rates in tissue. Magn Reson Med. 1994;32:66–76. 100. Larsson HB, Rosenbaum S, Fritz Hansen T. Quantification of the effect of water exchange in dynamic contrast MRI perfusion measurements in the brain and heart. Magn Reson Med. 2001;46: 272–281. 101. Fritz-Hansen T, Rostrup E, Sondergaard L, et al. Capillary transfer constant of Gd-DTPA in the myocardium at rest and during vasodilation assessed by MRI. Magn Reson Med. 1998;40:922–929. 102. Cullen JH, Horsfield MA, Reek CR, et al. A myocardial perfusion reserve index in humans using first-pass contrast-enhanced magnetic resonance imaging. J Am Coll Cardiol. 1999;33:1386–1394. 103. Neyran B, Janier MF, Casali C, Revel D, Canet Soulas EP. Mapping myocardial perfusion with an intravascular MR contrast agent: robustness of deconvolution methods at various blood flows. Magn Reson Med. 2002;48:166–179. 104. Bjornerud A, Bjerner T, Johansson LO, Ahlstrom HK. Assessment of myocardial blood volume and water exchange: theoretical considerations and in vivo results. Magn Reson Med. 2003;49:828–837. 105. Dauber IM, VanBenthuysen KM, McMurtry IF, et al. Functional coronary microvascular injury evident as increased permeability due to brief ischemia and reperfusion. Circ Res. 1990;66: 986–998. 106. Svendsen JH, Bjerrum PJ, Haunso S. Myocardial capillary permeability after regional ischemia and reperfusion in the in vivo canine heart: effect of superoxide dismutase. Circ Res. 1991;68:174–184. 107. 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. 108. Keijer JT, van Rossum AC, Wilke N, et al. Magnetic resonance imaging of myocardial perfusion in single-vessel coronary artery disease: implications for transmural assessment of myocardial perfusion. J Cardiovasc Magn Reson. 2000;2:189–200. 109. Wackers FJT, Bodenheimer M, Fleiss JL, Brown M. Multicenter Study on Silent Myocardial Ischemia Investigators: factors affecting uniformity in interpretation of planar Thallium-201 imaging in a multicenter trial. J Am Coll Cardiol. 1993;21:1064–1074. 110. Schwitter J, Wacker C, Wilke N, Al-Saadi N, Hoebel N, Simor T. MRIMPACT II. Circulation 2006. Annual Scientific Meeting of the AHA, Chicago; 2006. 111. Heart Disease and Stroke Statistics: Update 2005. American Heart Association; 2005. 112. Iskander S, Iskandrian AE. Risk assessment using single-photon emission computed tomographic technetium-99m sestamibi imaging. J Am Coll Cardiol. 1998;32:57–62. 113. Nagel E, Lehmkuhl HB, Bocksch W, et al. Noninvasive diagnosis of ischemia-induced wall motion abnormalities with the use of highdose dobutamine stress MRI: comparison with dobutamine stress echocardiography. Circulation. 1999;99:763–770.

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 stra tegies. 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 differ ences that may lead to seemingly discrepant test results. The most important differences are the capability of the pharmaco logic 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 deliv ery: 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 coro nary 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 alle viating myocardial infarction and demand ischemia, respec tively. 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 regu late total coronary resistance and, thus, coronary blood flow. In healthy humans, resting myocardial oxygen consump tion 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 myocar dial 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 pharma cologic stimulation, a short description of the basic physio logic 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 trans port vessels with minimal resistance. Further distal,

CORONARY AUTOREGULATION

Cardiovascular Magnetic Resonance 229

17 COMPARISON OF PERFUSION AND WALL MOTION CARDIOVASCULAR MAGNETIC RESONANCE IMAGING

CHAPTER 17

ISCHEMIC HEART DISEASE

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 meta bolic needs of the heart is virtually impossible by further increasing oxygen extraction from incoming arterial blood. Accordingly, changes in myocardial demand mandate con comitant 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 con strict at rest, thereby increasing resistance and decreasing blood flow. During physical exercise, though, these micro vessels 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 reg ulatory mechanism and can amount to a factor of 2 to 5 with a wide range of interindividual variability. Per defini tion, 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 esti mated 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 catheteri zation 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 dila tion 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 maldistribu tion of blood flow within the respective myocardial terri tories, 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 cathe ter) at maximal hyperemia.3–5 FFR has a clear normal value of 1.0 and a well established abnormal value of less than 0.75. FFR is advan tageous 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 circula tion. FFR measurements have been well validated in 230 Cardiovascular Magnetic Resonance

comparison with other, noninvasive imaging modalities, such as stress echocardiography and single photon emis sion 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 electro cardiogram (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 phenom ena are therefore referred to as the spatiotemporal disparity between myocardial perfusion and wall motion abnormal ities and are deemed to be responsible for the observed higher sensitivity of perfusion abnormalities over dys functional 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, how ever, 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 vasodila tor 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 recep tors. 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 myocar dium. 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 cor onary artery stenoses. However, in the presence of a high grade 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 abol ished, and coronary blood flow is directly related to the driving pressure. Consequently, myocardial resistance in areas supplied by normal coronary arteries is further down regulated 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 myocar dial blood flow, it is crucial to achieve the condition of true steady state hyperemia. The accuracy of measure ments 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 compar ing 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 mal distribution, 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 recom mended 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 dose dependent: 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 sys temic 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 dobuta mine.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 experi ence a drop in systolic blood pressure at higher dobuta mine 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 pro vides a reasonable estimate of myocardial workload (a heart rate pressure product greater than 20,000 mm Hg/min usu ally 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 myo cardial blood flow and provoking myocardial perfusion abnormalities. The flow increase (twofold to threefold base line values) is reportedly less than that elicited by adeno sine. However, dobutamine stress myocardial perfusion imaging using radioactive isotopes including 201Tl, 99m TC sestamibi, and 99mTC tetrofosmin has been success fully 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 seg ments reaching a sensitivity, specificity, and diagnostic accu racy of 81%, 73%, and 77%, respectively. In this study, however, stress imaging has been done at low dose dobuta mine (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 min ute) while preserving a high spatial resolution. With the more recent development of parallel imaging techniques (e.g., sensitivity encoding, SENSE),27–29 dynamic, multi slice 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 imag ing 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 dobu tamine 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 baro receptor 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 hyper trophy, but these effects are usually transient and resolve spontaneously. Because adenosine is a respiratory stimulant primarily through activation of carotid body chemorecep tors, 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 radionu clide 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 termi nated 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 rou tine 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 (first degree 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 stan dard 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 examina tions 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 nonsus tained atrial fibrillation was provoked in 16. Only one patient suffered sustained ventricular tachycardia and underwent suc cessful 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 vir tually 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 bronch oconstriction (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 tar get heart rate defined as age predicted submaximal heart rate ([220  age]  0.85) is reached. Additional termina tion 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 beta blockers 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 neces sary 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 antian ginal 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, cal cium 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, pericar ditis 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/Ca antagonists/ 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 precau tions 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. Out side 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 instan taneously. Alternatively, with a workstation next to the scanner console, the cine loops can be automatically trans ferred and displayed in a synchronized quadscreen. For

assessment of perfusion abnormalities with CMR, the inter pretation 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 min utes), 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. Conse quently, 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 myocar dial segments according to the standardized myocardial seg mentation 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 imag ing of an additional long axis view during CMR perfusion imaging failed to show ischemia in this region.39 Further more, the analysis of only the inner three out of five to eight short axis slices resulted in a significantly improved diagnos tic 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 mor phologic 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 lim ited. 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 infu sions via a peripheral line. DOBUTAMINE For dobutamine, the likelihood of adverse effects increases with infusion duration rather than dosage. Hence, the exam iner 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 beta blocker do not occur long after the dobutamine effect is gone. Nitroglycerin might be used to treat chest pain symp tomatically, 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 coro nary vasodilator.41 Of note, with the standard adenosine regimen, 16% of patients did not reach a maximal hyper emic 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 min utes 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 single session 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 accu racy, 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 combina tion (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). Com pared 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 single session 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 dysfunc tional 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 con tractility, 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 accom plished 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 contraindica tions 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.

Cardiovascular Magnetic Resonance 239

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 per fusion 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 pathophy siologies 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 via bility 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 find ings. CMR stress testing is capable of detecting residual myocardial ischemia. Various important complications of myocardial infarction can be imaged, such as LV dysfunc tion, LV thrombus, ventricular septal defect, aneurysm, pseudoaneurysm, and valvular complications. Thus, infor mation that has prognostic value or is useful in risk stratifi cation 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 detec tion 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 per spective 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 rela tively easily imaged with steady state free precession (SSFP) cine CMR.1,2 SSFP cine CMR has been accepted as a refer ence 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 accu racy of dobutamine stress tests. Recent comparisons indicate that cine CMR performs significantly better than non contrast 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 unrecog nized 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 transtho racic 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 man age coronary artery disease (CAD) in a combined experience in 877 patients, cine CMR performed with a sensitivity rang ing 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 pre dicted maximum and patients may be experiencing angina. Cine CMR also performed well in detecting acute coro nary 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 myo cardial 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 presenta tion 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 abnormal ities such as myocarditis. On the other hand, in a patient with a prior infarction, the presence or absence of a regional wall motion abnor mality frequently does not alter the clinical approach to a patient. In the latter situation, evidences of inducible ische mia or residual viability become additional pieces of infor mation 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 syn drome 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 differ ent mechanisms. Finally, it is important to differentiate per fusion 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 extra cellular 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 dif fuses into the interstitial space but is excluded from the intracellular space in healthy tissues. Thus, viable myocar dium enhances to a degree based on the amount of gado linium 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 win dow of 5 to 30 minutes after contrast administration pro vide 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 deposi tion but a relatively small intracellular space within fibro blasts. 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 infarc tion.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 inver sion time that results in uniform nulling of normal myocar dium (see Chapter 1). Nulling normal myocardium refers to a situation whereby the signal intensity of normal myo cardium becomes uniformly very dark. Thus, any abnorm alities 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 no entry 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 meth odology,19 basic validations, and clinical testing (Table 18 1).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 gad olinium 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 histopa thology 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 deter mined by single photon emission tomography (SPECT) imag ing.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 measur ing 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 reso lution.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 ven tricular (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 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 infarc tions, even those that are not identified by other imaging meth ods. Measurements of thickness of nonenhanced myocardium compared with measurements of percent transmural enhance ment also provide prognostic accuracy for predicting improved systolic wall thickening following AMI.38 Physiologic correlations to the spatial distribution of gado linium 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 myocar dial infarction from other causes of LGE.41 When the resolu tion 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 rela tive 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 reper fusion status of the myocardium and the degree of residual coronary artery stenosis. Postinfarct perfusion status modu lates 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 ade quate to understand the relative signal intensity on LGE CMR. However, a complete understanding of contrast enhancement after AMI needs to consider whether the cor onary artery has been reperfused, whether there is a resid ual coronary artery stenosis, the amount of collateral blood flow to the ischemic bed, and the status of myocardial per fusion 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 nor mal myocardium. In fact, it is possible to detect the perfu sion 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 tran sient phenomena.

Collateral blood flow can bring gadolinium contrast into the infarcted region as long as circulating levels of gadolin ium 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 min utes, 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 obstruc tion 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, includ ing (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 well reperfused 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 distribu tion is higher than normal. Doppler flow wire measure ments 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 por tions 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 microvascu lar obstruction was associated with adverse clinical out come.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 dur ing an average 16 month follow up. Although microvascu lar 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 col leagues found that the presence of microvascular obstruc tion 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 rel atively simple to perform. The CMR relationships between postinfarct microvascu lar 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. Further more, 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 likeli hood 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 neces sary 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 man aging 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 car diac 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 predic tive value over an assessment of viability. This question is not fully resolved, but the explanation may lie in a few pos sibilities. 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 statisti cally 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 ampli fied 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 obstruc tion tends to occur in more severe infarcts with later reper fusion.54 Thus, microvascular obstruction tends to be associated with more transmural infarcts. Third, the sever ity of a residual perfusion defect associated with microvas cular obstruction may alter postinfarct healing or remodeling. Severe microvascular obstruction may be asso ciated with more thinning of the myocardium in the infarct zone29 and thus worse peak systolic strain, further impair ing regional wall thickening.55–57 Fourth, severe microvas cular 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 circumferen tial 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 seg ment 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.

ISCHEMIC HEART DISEASE

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 clini cal approach works in most patients, it does not work well in many patients, particularly in non ST elevation myocar dial infarction and patients with left bundle branch blocks. T2 weighted CMR can uniquely differentiate AMI from chronic myocardial infarction and from normal myocar dium on the basis of differences in signal intensity. In gen eral, 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 sig nal. Just as soft tissue such as skin and muscle swell after mechanical injuries, myocardium swells after acute myocar dial ischemia and infarction. The extra water in the myocar dium 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 myo cardial infarction.58 With very high sensitivity and specific ity, they were able to determine the infarct related artery and differentiate AMI from chronic myocardial infarction. Subsequently, Cury and colleagues demonstrated the com bination 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 asso ciated 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 abnor mality 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 conve nient 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, par ticularly 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 pro vide 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 anti coagulation. 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). Infarct related ventricular septal defect can be identified and charac terized 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 two chamber view (left: end systolic frame from cine CMR; right: LGE images). The arrow points to the apical thrombus. LA, left atrium; LV, left ventricle.

ISCHEMIC HEART DISEASE

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 suit able for detecting AMI and assessing many clinically impor tant 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 appropriate ness criteria.65 It is important to note that postinfarct patients must be imaged efficiently and quickly, so signifi cant CMR experience is important for high quality clinical results.

References 1. Barkhausen J, Ruehm SG, Goyen M, Buck T, Laub G, Debatin JF. MR evaluation of ventricular function: true fast imaging with steady-state precession versus fast low-angle shot cine MR imaging: feasibility study. Radiology. 2001;219:264–269. 2. Pereles FS, Kapoor V, Carr JC, Simonetti OP, Krupinski EA, Baskaran V, Finn JP. Usefulness of segmented trueFISP cardiac pulse sequence in evaluation of congenital and acquired adult cardiac abnormalities. AJR Am J Roentgenol. 2001;177:1155–1160. 3. Bellenger NG, Burgess MI, Ray SG, Lahiri A, Coats AJ, Cleland JG, Pennell DJ. 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. 4. Bellenger NG, Davies LC, Francis JM, Coats AJ, Pennell DJ. Reduction in sample size for studies of remodeling in heart failure by the use of cardiovascular magnetic resonance. J Cardiovasc Magn Reson. 2000;2:271–278. 5. Grothues F, Smith GC, Moon JC, Bellenger NG, Collins P, Klein HU, Pennell DJ. 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. 250 Cardiovascular Magnetic Resonance

6. Strohm O, Schulz-Menger J, Pilz B, Osterziel KJ, Dietz R, Friedrich MG. Measurement of left ventricular dimensions and function in patients with dilated cardiomyopathy. J Magn Reson Imaging. 2001;13:367–371. 7. Hoffmann R, von Bardeleben S, Kasprzak JD, Borges AC, ten Cate F, Firschke C, Lafitte S, Al-Saadi N, Kuntz-Hehner S, Horstick G, et al. Analysis of regional left ventricular function by cineventriculography, cardiac magnetic resonance imaging, and unenhanced and contrastenhanced echocardiography: a multicenter comparison of methods. J Am Coll Cardiol. 2006;47:121–128. 8. Hundley WG, Hamilton CA, Thomas MS, Herrington DM, Salido TB, Kitzman DW, Little WC, Link KM. 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. 9. 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. 10. 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;107:1592–1597.

28. Selvanayagam JB, Porto I, Channon K, Petersen SE, Francis JM, Neubauer S, Banning AP. Troponin elevation after percutaneous coronary intervention directly represents the extent of irreversible myocardial injury: insights from cardiovascular magnetic resonance imaging. Circulation. 2005;111:1027–1032. 29. Baks T, van Geuns RJ, Biagini E, Wielopolski P, Mollet NR, Cademartiri F, van der Giessen WJ, Krestin GP, Serruys PW, Duncker DJ. 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. 30. Gerber BL, Belge B, Legros GJ, Lim P, Poncelet A, Pasquet A, Gisellu G, Coche E, Vanoverschelde JL. Characterization of acute and chronic myocardial infarcts by multidetector computed tomography: comparison with contrast-enhanced magnetic resonance. Circulation. 2006;113:823–833. 31. Plein S, Younger JF, Sparrow P, Ridgway JP, Ball SG, Greenwood JP. Cardiovascular magnetic resonance of scar and ischemia burden early after acute ST elevation and non-ST elevation myocardial infarction. J Cardiovasc Magn Reson. 2008;10:47. 32. Fieno DS, Kim RJ, Chen EL, Lomasney JW, Klocke FJ, Judd RM. 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. 33. Klein C, Nekolla SG, Bengel FM, Momose M, Sammer A, Haas F, Schnackenburg B, Delius W, Mudra H, Wolfram D. Assessment of myocardial viability with contrast-enhanced magnetic resonance imaging: comparison with positron emission tomography. Circulation. 2002;105: 162–167. 34. Kuhl HP, Beek AM, van der Weerdt AP, Hofman MB, Visser CA, Lammertsma AA, Heussen N, Visser FC, van Rossum AC. Myocardial viability in chronic ischemic heart disease: comparison of contrastenhanced magnetic resonance imaging with (18)F-fluorodeoxyglucose positron emission tomography. J Am Coll Cardiol. 2003;41: 1341–1348. 35. 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. 36. Ibrahim T, Bulow HP, Hackl T, Hornke M, Nekolla SG, Breuer M, Schomig A, Schwaiger M. Diagnostic value of contrast-enhanced magnetic resonance imaging and single-photon emission computed tomography for detection of myocardial necrosis early after acute myocardial infarction. J Am Coll Cardiol. 2007;49:208–216. 37. Kumar A, Abdel-Aty H, Kriedemann I, Schulz-Menger J, Gross CM, Dietz R, Friedrich MG. Contrast-enhanced cardiovascular magnetic resonance imaging of right ventricular infarction. J Am Coll Cardiol. 2006;48:1969–1976. 38. Ichikawa Y, Sakuma H, Suzawa N, Kitagawa K, Makino K, Hirano T, Takeda K. Late gadolinium-enhanced magnetic resonance imaging in acute and chronic myocardial infarction. J Am Coll Cardiol. 2005;45:901–909. 39. Rehwald WG, Fieno DS, Chen EL, Kim RJ, Judd RM. Myocardial magnetic resonance imaging contrast agent concentrations after reversible and irreversible ischemic injury. Circulation. 2002;105:224–229. 40. Maskali F, Poussier S, Marie PY, Tran N, Antunes L, Olivier P, Plenat F, Maitrejean S, Zannad F, Karcher G. High-resolution simultaneous imaging of SPECT, PET, and MRI tracers on histologic sections of myocardial infarction. J Nucl Cardiol. 2005;12:229–230. 41. Reimer KA, Jennings RB. The “wavefront phenomenon” of myocardial ischemic cell death. II. Transmural progression of necrosis within the framework of ischemic bed size (myocardium at risk) and collateral flow. Lab Invest. 1979;40:633–644. 42. 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;343:1445–1453. 43. Hunold P, Schlosser T, Vogt FM, Eggebrecht H, Schmermund A, Bruder O, Schuler WO, Barkhausen J. Myocardial late enhancement in contrast-enhanced cardiac MRI: distinction between infarction scar and non-infarction-related disease. AJR Am J Roentgenol. 2005;184: 1420–1426. 44. Kim RJ, Chen EL, Lima JA, Judd RM. Myocardial Gd-DTPA kinetics determine MRI contrast enhancement and reflect the extent and severity of myocardial injury after acute reperfused infarction. Circulation. 1996;94:3318–3326. 45. Kloner RA, Ganote CE, Jennings RB. The “no-reflow” phenomenon after temporary coronary occlusion in the dog. J Clin Invest. 1974;54:1496–1508.

Cardiovascular Magnetic Resonance 251

18 ACUTE MYOCARDIAL INFARCTION: CARDIOVASCULAR MAGNETIC RESONANCE DETECTION AND CHARACTERIZATION

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.

ISCHEMIC HEART DISEASE

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.

252 Cardiovascular Magnetic Resonance

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 contrac tile function, and myocardial electrical features are modi fied as a consequence of local mechanical aberrations, systemic hemodynamic changes, alterations in the neuro hormonal milieu, paracrine and intercellular factors, and intracellular factors including modifications of gene expres sion.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 altera tions 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 asso ciated 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 obstruc tion (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 Neutro phils release serine proteases and activate matrix metallo proteases, 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 expan sion 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 sys tem 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 mus cle 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 assess ment 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 col leagues18 demonstrated that fast breath hold CMR techni ques have a better interstudy reproducibility than did echocardiography in serially assessing LV end systolic vol ume, 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

19 ACUTE MYOCARDIAL INFARCTION: VENTRICULAR REMODELING

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 col leagues19 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 statisti cally significant only for ejection fraction (EF).20 Various investigators have validated the CMR measure ment 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 ini tial 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 two fold increase in LV end diastolic volume and an almost six fold increase in LV end systolic volume. While there was no significant change in RV end diastolic volume or end sys tolic 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 signifi cantly 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 orienta tion 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 eigenvec tor of the diffusion tensor coincides with the local fiber ori entation. 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 myo cardium of control rats. However, the overall fiber orienta tion 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 Koner mann 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 follow up from 2.0
1.6 mm to 0.5
2.9 mm, the extent being dependent on infarct size, but independent of infarct loca tion. 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 short axis 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 end systolic volume was found. The increase in LV end diastolic volume was disproportionate such that the LV volume to mass 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 col leagues25 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 infarc tion. 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 des cending (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 predic tor of an increase in LV end diastolic volume index was peak creatinine kinase.26 Schroeder and colleagues27 performed serial CMR stud ies at 1 week, 13 weeks, 26 weeks, and 52 weeks post infarction in 51 patients. They found a 5% increase in LV mass index among those treated with percutaneous trans luminal 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). Contrast enhanced 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 ana lyzed functional data with a 3D wall thickening using a centerline method. Wall thickening was reduced in myo cardial 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 cir cumferential 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 nonin farcted 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 Moul ton 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 func tion 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 dys function occurs early after infarction between the infarcted and remote regions. These investigators calculated the per centage 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 post infarction 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 maxi mum (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 per centage of intramyocardial %CS was only 17
%, signifi cantly 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 improve ment 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%. How ever, 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 correla tion between regional function and regional loading condi tions 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 condi tions increased between the initial and follow up scans, while in peri infarct regions, no significant change in defor mation 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 postinfarc tion ventricular remodeling, the utility of such sequences will be the quantitative assessment of the effects of pharma cologic, 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 remo deled 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 subendo cardial 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 alter ation of myocardial high energy phosphate levels in remo deled myocardium. The extent of remodeling, as indexed by the increase in end diastolic volume and mass over time, cor related 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 dysfunc tion. The CP to ATP ratio in remodeled noninfarcted myo cardium 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 bioe nergetic abnormalities reflects the severity of LV dysfunc tion 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 con centrations 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 sig nal 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 euthana sia at 3 days and 6 weeks after infarction. MRS has been used to serially document established molec ular 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 relaxa tion have been particularly useful in the assessment of ventricular remodeling. Gadolinium based agents have been used to study MO, scar formation, and viability. Gadolinium diethylenetriamine pentaacetic acid (Gd DPTA) is an extra cellular 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 col lagenous 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 high energy 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.

ISCHEMIC HEART DISEASE

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 hypoenhance ment 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 postinfarc tion 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 per formed CMR on 25 patients at 1 week and 8 weeks post infarction. LVEF did not improve in patients with MO. Partial functional improvement was noted in segments with non transmural 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 infu sion 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 estab lished 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

Dog 2

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 post infarction. They demonstrated an increase in scar size and the MO zone during the first 48 hours. Fieno and col leagues57 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 per manently. LGE CMR was performed at 3 days, 10 days, 4 weeks, and 8 weeks to assess the circumferential and longi tudinal extents of infarction. Infarct resorption during the 8 week follow up period correlated with initial infarct size (Fig. 19 4). At the 4 week scan, the extent of infarct resorp tion 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 col leagues,28 in which patients received stenting with contempo rary 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 base line 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 car diovascular 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 pre dicted 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 char acteristic 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 pri mary 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, electrocar diographic, 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 deli neating acute infarctions. Watzinger and colleagues62 per formed 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 Gd mesoporphyron 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.

ISCHEMIC HEART DISEASE

pharmacologic agents have focused on the renin angiotensin aldosterone axis. CMR evaluation of interven tional 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 under lying 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 col leagues.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, demonstrat ing that use of an ACE I during the first 8 weeks after ante roapical infarction limited the increase in LV end diastolic volume that was noted in untreated controls.66 The limita tion 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, high dose ACE I limited the increase in LV mass and end diastolic 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 per formed 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 reduc tion 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 com pared 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 vol ume. However, LVEF did improve over the first 8 weeks in comparison to ACE I alone. The most powerful pharmaco logic 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 post infarction. At baseline, end systolic volume index was lower and LVEF was higher in the transgenic mice. The ini tial 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 recep tor (AT1R) in postinfarction remodeling, these investigators demonstrated equivalent remodeling in AT2R overexpres sing 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 overexpres sion 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 col leagues77 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 thin ning 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 remo deling post infarction. Nicorandil is an ATP sensitive potas sium 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 reper fusion 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 end diastolic 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 HMG CoA reductase inhibition in LV remodeling to determine whether the mechanism of any positive effect on remo deling was NOS dependent. Adult rats underwent left coronary ligation followed by treatment with placebo, ceri vastatin, or cerivastatin plus N methyl L arginine methyl ester (L NAME), a NOS inhibitor. Rats that were fed cer ivastatin 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 CoA reductase 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 col leagues84 utilized an ovine model of LAD occlusion to cre ate 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 con trols. 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

ISCHEMIC HEART DISEASE

function and prevention of LV dilation in a mouse infarc tion model that received stem cell transplantation. How ever, scar size did not decrease. Amado and colleagues86 used CMR with a porcine model of infarction to demon strate 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 perma nent 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 stud ied with 31P CMRS. LVEF was higher in animals treated with stem cells (p < 0.05). The ratio of endocardial phos phocreatine 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 nonin vasive evaluation of LV function. These factors have also allowed clinical trials evaluating LV remodeling to be ade quately 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 symp tom onset to intravenous enalaprilat and then oral ACE I therapy or placebo for 1 month post infarction. Twenty three of the patients underwent CMR at 1 month for evalu ation of infarct expansion, defined as the ratio of the infarct to noninfarct endocardial segment length. Other para meters 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 col leagues90 utilized CMR to assess ACE I therapy post infarction 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 treat ment 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 end systolic 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 base line 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 vari ables 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 carve dilol 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 sev eral hundred) to demonstrate quantifiable differences between treatment and control groups.95 CMR has been used to assess post infarction LV remo deling after PCI. Studies have been conducted to assess the effect of both early and late recanalization of infarct related 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 vol ume. These findings were not observed in the delayed ther apy group.96 Subsequently, Silva and colleagues97 reported

160 140

p = 0.001*

Baseline 6 months

120 p <0.001

100

p = ns

80 60 p = 0.008

40 20 0 LVEDVI

LVESVI

LVEF

LVMI

PLACEBO 160 140

Baseline 6 months

p = ns

120 100

p = ns

p = ns

80 60 40

p = ns

20 0 LVEDVI

LVESVI

LVEF

LVMI

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 regi men including aspirin, beta blocker, ACE I, and a lipid lowering 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 vol ume 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 differ ence 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 aver age 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 seg ments 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 fol low 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 improve ment 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 end systolic volume index and LVEF. CMR is a particularly useful modality for assessing PCI, not only, as was previ ously 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 postin farction LV remodeling after myocardial stem cell transplan tation. In the BOOST trial, 60 patients were randomized to autologous bone marrow cell intracoronary injection or medical therapy after successful PCI for an acute ST segment 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 progeni tor 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

ISCHEMIC HEART DISEASE

Finally, CMR has also been used in studies assessing nontraditional therapies of LV remodeling post infarction. Dubach and colleagues103 studied 25 patients with LV dys function (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. Inves tigators 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

264 Cardiovascular Magnetic Resonance

17.

18.

19.

20.

21. 22. 23. 24. 25.

26.

27.

28.

29.

30.

predictors and impact of long-term use of captopril: information from the Survival and Ventricular Enlargement (SAVE) trial. Circulation. 1997;96(10):3294–3299. 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(1):44–51. Grothues F, Smith GC, Moon JC, et al. Comparison of interstudy reproducibility of cardiovascular magnetic resonance with twodimensional echocardiography in normal subjects and in patients with heart failure or left ventricular hypertrophy. Am J Cardiol. 2002;90(1): 29–34. Matheijssen NA, Baur LH, Reiber JH, et al. Assessment of left ventricular volume and mass by cine magnetic resonance imaging in patients with anterior myocardial infarction intra-observer and inter-observer variability on contour detection. Int J Card Imaging. 1996;12(1): 11–19. 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. Kramer CM, Lima JA, Reichek N, et al. Regional differences in function within noninfarcted myocardium during left ventricular remodeling. Circulation. 1993;88(3):1279–1288. McDonald KM, Francis GS, Matthews J, Hunter D, Cohn JN. Longterm oral nitrate therapy prevents chronic ventricular remodeling in the dog. J Am Coll Cardiol. 1993;21(2):514–522. Shapiro EP, Rogers WJ, Beyar R, et al. Determination of left ventricular mass by magnetic resonance imaging in hearts deformed by acute infarction. Circulation. 1989;79(3):706–711. Grossman W, Jones D, McLaurin LP. Wall stress and patterns of hypertrophy in the human left ventricle. J Clin Invest. 1975;56(1):56–64. Konermann M, Sanner BM, Horstmann E, et al. Changes of the left ventricle after myocardial infarction: estimation with cine magnetic resonance imaging during the first six months. Clin Cardiol. 1997; 20(3):201–212. Kramer CM, Rogers WJ, Theobald TM, Power TP, Geskin G, Reichek N. Dissociation between changes in intramyocardial function and left ventricular volumes in the eight weeks after first anterior myocardial infarction. J Am Coll Cardiol. 1997;30(7):1625–1632. Schroeder AP, Houlind K, Pedersen EM, Nielsen TT, Egeblad H. Serial magnetic resonance imaging of global and regional left ventricular remodeling during 1 year after acute myocardial infarction. Cardiology. 2001;96(2):106–114. Hombach V, Grebe O, Merkle N, et al. 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(6):549–557. Wiesmann F, Frydrychowicz A, Rautenberg J, et al. Analysis of right ventricular function in healthy mice and a murine model of heart failure by in vivo MRI. Am J Physiol Heart Circ Physiol. 2002;283(3): H1065–H1071. Nahrendorf M, Hu K, Fraccarollo D, et al. Time course of right ventricular remodeling in rats with experimental myocardial infarction. Am J Physiol Heart Circ Physiol. 2003;284(1):H241–H248.

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;.

Cardiovascular Magnetic Resonance 265

19 ACUTE MYOCARDIAL INFARCTION: VENTRICULAR REMODELING

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.

ISCHEMIC HEART DISEASE

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.

266 Cardiovascular Magnetic Resonance

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 interven tions,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 revasculariza tion 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 nega tives. In other words, the sensitivity of a test is the number of diseased persons with a positive test divided by the total num ber 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 with out 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 spec ificity. In the viability literature, however, sensitivity indicates the ability of a test to identify viable myocardium, and specific ity 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 myo cardial 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 mix tures 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 sub acute 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 thin ning. Some thinning may be observed, however, depending on the degree to which the endocardially located infarct extends throughout the wall. Therefore, the finding of pre served 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, pre served end diastolic wall thickness cannot be used to distin guish between viable and nonviable myocardium. Recently, a patient was described with a small subendocar dial infarct with a thickness of 1.5 mm covered by an epicar dially 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 cir cumstances 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

20 MYOCARDIAL VIABILITY

CHAPTER 20

ISCHEMIC HEART DISEASE

include sympathomimetic agents14 or postextrasystolic poten tiation.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 respon sive myocardium will usually recover function after appropri ate revascularization.15 However, it appears that there is also some spontaneous improvement in the response to dobuta mine over the course of infarct healing after reperfused myo cardial 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 approxi mately 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 micro scopy 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 ery throcytes 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 mag netic 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 hypoper fusion 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 pro longed washin time constant. This lack of reperfusion after restoration of flow in epicardial vessel is known as the no reflow phenomenon. When the myocardium is imaged by CMR early after injection after gadolinium (early gadolin ium enhancement), no reflow zones appear dark in com parison 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 relax ation 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 nor mal 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 con trast washout.19,25 The enhancement pattern that is seen will depend on regional differences in tissue washin/washout kinet ics 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 signal to noise ratio