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

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

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

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

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

It has now been more than 20 years since the first magnetic resonance (MR) images of the human heart were described. Although few would deny that the quality of cardiovascular MR (CMR) images has dramatically improved since that time, it has nevertheless been only in the last 5 to 10 years that large numbers of dedicated CMR clinical services have opened in the United States. Although organizations such as the Society for Cardiovascular Magnetic Resonance (SCMR; www.scmr .org) are working toward scan standardization, there is often variability among CMR centers as to what actually comprises a “routine clinical CMR.” In this setting, a chapter that attempts to describe all of the possible permutations of pulse sequences, scan protocols, and interpretation methods can quickly become unwieldy. This chapter focuses on the 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

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CHAPTER 2

700 CMR procedures (quarterly)

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

800

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

Congenital evaluation (7%)

CMR EXAM MENU (“SCANS”)

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

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

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

both the short and long axis cardiac views lie at arbitrary angles with respect to scanner coordinates and are 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

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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

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

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

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

sequence, is its high SNR, fast speed, and excellent blood/ myocardium CNR that greatly facilitates the identification of the endocardial border. Typically, a stack of multiple closely spaced (contiguous or with 1 to 2 mm gaps) short axis slices (6 to 8 mm thick) is acquired to provide full coverage of the left and right ventricles. Short axis views, perpendicular to the long axis views, can be planned on long axis scout images, as described earlier. In addition, cine images can be obtained in multiple long axis views, such as the two chamber, three chamber, or four chamber orientations. A key consideration for cine imaging is that current CMR scanner hardware is not adequate to acquire cine images with high spatial resolution and high temporal 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°

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

k-space slice 1

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

k-space slice 2 Heartbeat 1 Heartbeat 2

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Trigger

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

green segment is always acquired immediately after the R wave and is used to reconstruct the first movie frame, the orange segment is acquired next and becomes the 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.

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ECG Trigger Acquired in segments:

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30 segments acquired each heartbeat

Each segment contains 10 lines

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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

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The resulting heterogeneity in blood flow between regions supplied by coronary arteries with stenoses (little or no flow increase) and territories supplied by normal epicardial arteries (multifold increased flow) is observed as a 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

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2 min

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Saline flush Breath holding Image acquisition

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

contrast agent concentration, and therefore 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

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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

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

the images of Figure 2 26 because they would not be as crisp otherwise. For CE MRA, the most commonly used pulse sequence is a fast three dimensional spoiled GRE. This is a variant of the GRE technique that was discussed previously (see Fig. 2 19). The need for three dimensional data requires application of an additional gradient to encode the additional dimension. This gradient is played in the slice encoding direction. 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

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

Figure 2 29 summarizes how one would approach CMR scanning of a patient referred for the evaluation of aortic disease. In summary, although the definition of “routine clinical CMR” continues to evolve at the DCMRC and other 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.

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

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

This chapter is concerned with the basic principles and 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

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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

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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

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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

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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

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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

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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

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

Multiple image slices at once

Parallel MRI

Multiple k-space “slices” at once

RF coils

Rotations y ra Xitt em

Gradient steps

er

BASIC PRINCIPLES OF CARDIOVASCULAR MAGNETIC RESONANCE

MDCT

Imaged subject Imaged subject Coil 1

Coil 2

Coil 3 k=0 k=2

Detector rows

k=4

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

of each modality shown at the top and a more detailed 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.)

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comparable to that of the unaccelerated approach.109 This decreases the total scan time and reduces the susceptibility to patient motion and respiratory drift. Short breath hold 3D coronary artery CMR is challenging because of the 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.

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of the coronary arteries or any other typical cardiac view by subsequent reformatting. The same scanner and array have also been used for accelerated real time cardiac imaging115 and for rapid volumetric contrast enhanced MRA at 12 to 16 fold accelerations.88 The trend toward still larger 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.

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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

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in the subendocardial layer.18,19 The spatial resolution of conventional imaging modalities, such as positron emission tomography (PET) and single photon emission computed tomography (SPECT), was insufficient to detect blood flow deficits limited to the subendocardial layer. More specifically, the sensitivity of a myocardial perfusion imaging technique is directly related to its spatial resolution and the ability to discern transmural variations of flow.18,20 The use of CMR offers the unique possibility of 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

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

and the images are read out with several radiofrequency excitations/echo trains.23 This results in a 30% to 40% speed advantage compared with conventional GRE sequences and is particularly useful at field strengths of 1.5 T or lower. (At 3.0 T and higher, any increase in TE beyond the minimum allowed by the gradient system nearly inevitably gives rise to susceptibility artifacts, in 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

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Figure 4 2 Comparison of signal time curves for a region of interest in the left ventricle (LV) obtained in a canine with quantitative T1 imaging (T1 FARM) and T1 weighted, saturation recovery prepared fast gradient echo (TR/TE 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

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

WATER EXCHANGE AND ITS EFFECTS ON MYOCARDIAL CONTRAST ENHANCEMENT With 1H magnetic resonance imaging, the presence of a 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 1