Grainger & Allison's Diagnostic Radiology, 5th Edition

Contributors

†

: deceased

Andy Adam MBBS (Hons), FRCP, FRCS, FRCR, FFRRCSI (Hon)

Professor of Interventional Radiology Department of Radiology St Thomas’ Hospital King’s College London London, UK E. Jane Adam MBBS, MRCP, FRCR Consultant Radiologist Department of Radiology St George’s Hospital London, UK Judith E. Adams MBBS, FRCR, FRCP Chair, Diagnostic Radiology Imaging Science and Biomedical Engineering University of Manchester Honorary Consultant Radiologist Royal Infirmary Manchester, UK David J. Allison BSc, MD, MRCS, LRCP, MBBS, DMRD, FRCR, FRCP

Emeritus Professor of Imaging Imperial College London, UK Sandra Allison MD Assistant Professor of Radiology Director, Radiology Residency Program Director, Ultrasound Georgetown University Hospital Washington DC, USA Philip Anslow FRCR Consultant Neuroradiologist Department of Radiology Radcliffe Infirmary Oxford, UK Susan M. Ascher MD Georgetown University Medical Center Washington DC, USA Zelena A. Aziz MD, MRCP, FRCR Consultant Radiologist Department of Radiology London Chest Hospital London, UK

Clive I. Bartram FRCS, FRCP, FRCR Emeritus Consultant, St Mark’s Hospital and Honorary Professor of Gastrointestinal Radiology Faculty of Medicine Imperial College London, UK Philip P. W. Bearcroft FRCP, FRCR Consultant Radiologist Department of Radiology Cambridge University Hospitals NHS Foundation Trust Addenbrooke’s Hospital Cambridge, UK Anna-Maria Belli DMRD, FRCR Consultant Vascular Radiologist and Reader in Radiology Department of Radiology St George’s Hospital London, UK Anthony R. Berendt BM, BCh, FRCP Consultant Physician-in-Charge Bone Infection Unit Nuffield Orthopaedic Centre Oxford, UK Lol Berman FRCP, FRCR University Department of Radiology Addenbrooke’s Hospital Cambridge, UK Martin J. K. Blomley†

Gisele Brasil Caseiras PhD Research Fellow Insitute of Neurology University College London London, UK Jackie E. Brown BDS, MSc, FDSRCP, DDRRCR Consultant Oral and Maxillofacial Radiologist Kings College London Dental Institute Guy’s Dental Hospital London, UK Dina F. Caroline MD, PhD Professor Emerita Department of Radiology Temple University Hospital Philadelphia Pennsylvania, USA Silvia D. Chang MD, FRCP(C) Assistant Professor University of British Columbia Department of Radiology Vancouver General Hospital Vancouver, Canada W. K. ‘Kling’ Chong BMedSci, MD, MRCP, FRCR Consultant Neuroradiologist Department of Radiology Great Ormond Street Hospital for Children NHS Trust London, UK Bairbre Connolly MBBCh, BAO, FRCSI, MCH, FFRRCSI, FRCP(C)

Carol A. Boles MD Associate Professor of Radiology Associate, Surgical Sciences Orthopedic Surgery Wake Forest University North Carolina, USA Jamshed B. Bomanji MBBS, MSc, PhD Consultant in Nuclear Medicine UCLH Trust Middlesex Hospital London, UK

Medical Director and Division Head of Image Guided Therapy Pediatric Interventional Radiologist Assistant Professor, University of Toronto Department of Diagnostic Imaging The Hospital for Sick Children Toronto, Canada Susan J. Copley MBBS, MD, MRCP, FRCR Consultant Radiologist and Honorary Senior Lecturer Radiology Department Hammersmith Hospital London, UK

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CONTRIBUTORS

David O. Cosgrove MA, MSc, FRCP, FRCR Emeritus Professor Imaging Sciences Department Faculty of Medicine Imperial College Hammersmith Hospital London, UK

Robert J. Eckersley PhD Research Associate Imaging Sciences Department Faculty of Medicine Imperial College Hammersmith Hospital London, UK

Philip C. Goodman MD Professor of Radiology Chief, Thoracic Imaging Division Department of Radiology Duke University Medical Center Durham North Carolina, USA

Nigel Cowan PGDipLATHE, FRCP Oxford, UK

Andrew J. Evans MRCP, FRCR Consultant Radiologist Nottingham Breast Institute Nottingham City Hospital Nottingham, UK

Isky Gordon FRCR, FRCP Professor of Paediatric Imaging Institute of Child Health London, UK

Justin J. Cross MRCP, FRCR Consultant Neuroradiologist Department of Radiology Addenbrooke’s Hospital Cambridge, UK Paras Dalal BSc, MRCP, FRCR Research Fellow in Thoracic Imaging Department of Radiology Royal Brompton Hospital London, UK Maria Daskalogiannaki MD Registrar in Radiology Department of Radiology University Hospital of Heraklion Crete, Greece A. Mark Davies FRCR Consultant Radiologist Royal Orthopaedic Hospital Birmingham, UK Adrian K. Dixon MD, FRCR, FRCP, FRCS, FMedSci, FFRRCSI (Hon), FRANZCR (Hon)

Professor of Radiology Department of Radiology Addenbrooke’s Hospital University of Cambridge Cambridge, UK Rose de Bruyn DMRD, FRCR Consultant Radiologist Department of Radiology Great Ormond Street Hospital for Sick Children NHS Trust London, UK Claudio Defilippi MD Consultant Radiologist Department of Radiology OIRM - S. Anna Hospital Turin, Italy Sujal R. Desai MD, MRCP, FRCR Consultant Radiologist Department of Radiology King’s College Hospital London, UK

Jane Evanson BSc, MBBS, MRCP, FRCR Consultant Neuroradiologist, Barts & The London Hospital NHS Trust The Royal London Hospital London, UK Laura Fender BMedSci, BMBS, MRCP, FRCR Consultant Radiologist Nottingham University Hospital Nottingham, UK Alan H. Freeman MBBS, FRCR Consultant Radiologist Department of Radiology Addenbrooke’s Hospital Cambridge, UK Julia Gates MD Assistant Professor of Radiology Department of Radiology Tufts University School of Medicine Springfield Massachusetts, USA Robert N. Gibson MBBS, MD, FRANZCR, DDU Professor of Radiology Department of Radiology University of Melbourne Royal Melbourne Hospital Victoria, Australia Raymond J. Godwin MA, MB, Bchir, FRCP, FRCR Consultant Radiologist Department of Radiology West Suffolk Hospital Suffolk, UK Karen Goldstone BSc, MSc, Csci, FIPEM Radiation Protection Advisor Acting Head of Department of Medical Physics and Clinical Engineering East Anglian Regional Radiation Protection Service (EARRPS) Addenbrooke’s Hospital Cambridge, UK

Nicholas Gourtsoyiannis FRCR (Hon) Professor of Radiology University of Crete Faculty of Medicine Heraklion, Crete Greece Andrew J. Grainger BM, BS, MRCP, FRCR Consultant Radiologist Chapel Allerton Orthopaedic Centre Leeds Teaching Hospitals Leeds, UK Ronald G. Grainger MB ChB(Hons), MD, FRCP, DMRD, FRCR, FACR(Hon), FRACR(Hon)

Professor of Diagnostic Radiology (Emeritus) University of Sheffield Honorary Consultant Radiologist Royal Hallamshire Hospital and Northern General Hospital Sheffield, UK Philippe Grenier FRCR (Hon) Professor of Radiology Service de Radiologie Polyvalente Diagnostique et Interventionnelle Hôpital Pitié-Salpêtrière Paris, France Roxana S. Gunny BS, BSc, MRCP, FRCR Consulant Neuroradiologist Department of Radiology Great Ormond Street Hospital for Children NHS Trust London, UK Christine M. Hall MBBS, DMRD, FRCR MD Professor of Paediatric Radiology Great Ormond Street Hospital for Children NHS Trust London, UK David M. Hansell MD, FRCP, FRCR, LRSM Professor of Thoracic Imaging Department of Radiology Royal Brompton Hospital London, UK

CONTRIBUTORS

George G. Hartnell FRCR, FRCP Director of Cardiovascular and Interventional Radiology Department of Radiology Baystate Medical Center Springfield Professor of Radiology Tufts University Medical School Boston Massachusetts, USA Hedvig Hricak MD, Dr. Med, SC, Dr. h.c, FRCR (Hon) Chairman, Department of Radiology Carroll and Milton Petrie Chair Professor of Radiology, Weill Medical College of Cornell University Memorial Sloan-Kettering Cancer Center New York, USA James E. Jackson MRCP, FRCR Consultant Radiologist Department of Imaging Hammersmith Hospital London, UK H. Rolf Jäger FRCR, MD Reader in Neuroradiology Institute of Neurology University College London Honorary Consultant Neuroradiologist The National Hospital for Neurology and Neurosurgery and University College Hospital London, UK Jonathan J. James BMBS, FRCR Consultant Radiologist Nottingham Breast Institute Nottingham City Hospital Nottingham, UK Renee M. Kendzierski DO Assistant Professor of Radiology Department of Radiology Temple University Hospital Philadelphia Pennsylvania, USA Dow-Mu Koh MRCP, FRCP Senior Lecturer and Honorary Consultant Department of Radiology Royal Marsden Hospital Sutton, UK Isla Lang MBChB, MRCP, FRCR Consultant Paediatric Radiologist Sheffield Children’s Hospital Sheffield, UK

Adrian K. P. Lim MD, FRCR Consultant Radiologist and Senior Lecturer Imaging Sciences Department Faculty of Medicine Imperial College Hammersmith Hospital London, UK

Stuart E. Mirvis MD, FACR Professor of Radiology Department of Radiology University of Maryland School of Medicine Baltimore Maryland, USA

David J. Lomas MA, MB, BChir, FRCR, FRCP Professor of Clinical Magnetic Resonance Imaging Department of Radiology Addenbrooke’s Hospital Cambridge, UK

Sameh K. Morcos FRCS, FFRRCSI, FRCR Professor of Diagnostic Imaging University of Sheffield Consultant Radiologist Department of Diagnostic Imaging Northern General Hospital Sheffield, UK

Sharyn L. S. MacDonald MBChB, FRANZCR Consultant Radiologist Department of Radiology Christchurch Hospital Christchurch, New Zealand David MacVicar MA, MRCP, FRCP, FRCR Consultant Radiologist Department of Diagnostic Radiology Royal Marsden Hospital Sutton, UK Adrian Manhire BSc, MBBS, FRCP, FRCR Consultant Radiologist Nottingham City Hospital Nottingham, UK Tarik F. Massoud MA, MD, PhD, FRCR University Lecturer and Honorary Consultant in Neuroradiology University Department of Radiology University of Cambridge School of Clinical Medicine Addenbrooke’s Hospital Cambridge, UK Kieran McHugh FRCPI, DCH, FRCR Department of Radiology Great Ormond Street Hospital for Sick Children NHS Trust London, UK James Meaney FRCR, FFRRCSI Director of MRI St James’s Hospital Dublin, Ireland Hylton B. Meire FRCR, DRCOG, DMRD Consultant Radiologist (Retired) King’s College Hospital London, UK Kenneth A. Miles MD, FRCR, MSc, FRCP Clinical Imaging Sciences Centre Brighton and Sussex Medical School University of Sussex Falmar, Brighton, UK

Robert A. Morgan MBChB, MRCP, FRCR Consultant Radiologist Department of Radiology St George’s Hospital London, UK Iain Morrison MBBS, MRCP, FRCR Consultant Radiologist Radiology Department Kent and Canterbury Hospital Canterbury, UK Nestor L. Müller MD, PhD, FRCPC Professor and Chairman Department of Radiology University of British Columbia Head and Medical Director Department of Radiology Vancouver General Hospital Vancouver, Canada Graham Munneke MRCP, FRCR Consultant in Interventional Radiology Department of Radiology St. George’s Hospital London, UK Alison D. Murray MB ChB (Hons), FRCR, FRCP Senior Lecturer in Radiology Department of Radiology University of Aberdeen Aberdeen, UK Richard A. Nakielny FRCR Honorary Clinical Lecturer Directorate of Medical Imaging & Medical Physics Royal Hallamshire Hospital Sheffield, UK Hrudaya Nath MD Professor of Radiology Department of Radiology University of Alabama Hospitals Birmingham Alabama, USA

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xii

CONTRIBUTORS

Tony Nicholson MSc, FRCR Consultant Vascular Radiologist Department of Clinical Radiology Leeds Teaching Hospitals Leeds, UK Amaka C. Offiah BSc, MBBS, MRCP, FRCR, PhD Consultant Radiologist (Academic) Great Ormond Street Hospital for Children NHS Trust London, UK Simon Padley BSc, MBBS, FRCP, FRCR Consultant Radiologist Department of Radiology Chelsea and Westminster Hospital London, UK Martyn N. J. Paley PhD, FInstP Professor of MR Physics Academic Radiology University of Sheffield Sheffield, UK Nickolas Papanikolaou PhD Biomedical Engineer Department of Radiology University Hospital of Heraklion Crete, Greece Jai Patel MBChB, MRCP, FRCR Consultant Vascular Radiologist Department of Clinical Radiology St James’s University Hospital Leeds, UK Anne Paterson MBBS, MRCP, FRCR, FFR RCSI Consultant Paediatric Radiologist Radiology Department Royal Belfast Hospital for Sick Children Belfast, UK Praveen Peddu MRCS, FRCR Specialist Registrar in Radiology Department of Radiology King’s College Hospital London, UK A. Michael Peters BSc, MD, MSc, MRCP, MRCPath, FRCR

Professor of Nuclear Medicine Brighton and Sussex Medical School University of Sussex Brighton, UK William H. Ramsden BM, FRCR Consultant Paediatric Radiologist Department of Clinical Radiology St James’s University Hospital Leeds, UK

Sheila Rankin FRCR Consultant Radiologist Department of Radiology Guy’s and St. Thomas’ Foundation Trust London, UK

John Rout BDS, FDSRCS, MDentSc, DDRRCR, FRCR Consultant Oral and Maxillofacial Radiologist Birmingham Dental Hospital Birmingham, UK

Padma Rao MBBS, BSc, MRCP, FRCR,

Michael B. Rubens MB, DMRD, FRCR Consultant Radiologist Department of Radiology Royal Brompton Hospital London, UK

FRANZCR

Consultant Paediatric Radiologist Royal Children’s Hospital Parkville Melbourne Victoria, Australia Christine Reek BSc, FRCR Consultant Radiologist Department of Radiology Greenfield Hospital Leicester, UK John H. Reynolds DMRD, FRCR, MMedSci Consultant Radiologist Birmingham Heartlands Hospital Birmingham, UK Rodney H. Reznek FRANZCR (Hon), FRCP, FRCR Professor of Diagnostic Imaging The Centre for Cancer Imaging St Bartholomew’s Hospital and The London Queen Mary’s School of Medicine and Dentistry London, UK Philip M. Rich BSc, FRCS, FRCR Consultant Neuroradiologist Department of Neuroradiology Atkinson Morley Wing St George’s Hospital London, UK Andrea Rockall MD, BS, BSc, MRCP, FRCP Department of Radiology St Bartholomew’s Hospital London, UK Giles Roditi FRCP, FRCR Consultant Radiologist Department of Radiology Glasgow Royal Infirmary Glasgow, UK

Asif Saifuddin BSc (Hons), MBChB, MRCP, FRCR Consultant Radiologist Department of Radiology Royal National Orthopaedic Hospital NHS Trust Stanmore, UK Evis Sala MD, PhD, FRCR University Lecturer in Oncology Imaging University Department of Radiology Addenbrooke’s Hospital Cambridge, UK Caron Sandhu FRCR Consultant Radiologist Department of Radiology Guy’s and St. Thomas’ Hospital London, UK Dawn Saunders MD, MRCP, FRCR Consultant Neuroradiologist Department of Radiology Great Ormond Street Hospital for Children NHS Trust London, UK Daniel J. Scoffings MRCP, FRCR Specialist Registrar in Neuroradiology Addenbrooke’s Hospital Cambridge, UK Djilda Segerman MA, MSc, MIPEM Head of Nuclear Medicine Physics Department of Medical Physics Brighton and Sussex University Hospitals NHS Trust Brighton, UK

Lee F. Rogers MD Clinical Professor of Radiology Department of Radiology University of Arizona Health Services Tucson Arizona, USA

Kathirkama Shanmuganathan MD Associate Professor of Radiology Department of Radiology University of Maryland School of Medicine Baltimore Maryland, USA

Giles Rottenberg FRCR Consultant Radiologist Department of Radiology Guy’s and St. Thomas’ Foundation Trust London, UK

Ashley S. Shaw MRCP, FRCR Consultant Radiologist Department of Radiology Addenbrooke’s Hospital Cambridge, UK

CONTRIBUTORS

Mihra S. Taljanovic MD, MA Associate Professor of Clinical Radiology and Clinical Orthopedic Surgery Head - Musculoskeletal Imaging Section Department of Radiology Tucson Arizona, USA

Satinder P. Singh MD, FCCP Associate Professor of Radiology Director Cardiac CT Director, Combined Cardiopulmonary and Abdominal Fellowship Chief of Cardiopulmonary Radiology Department of Radiology University of Alabama Hospitals Birmingham Alabama, USA

FRCR

S. Aslam A. Sohaib MRCP, FRCR Radiology Department Royal Marsden Hospital London, UK

Consultant Radiologist Department of Clinical Radiology Great Ormond Street Hospital for Children NHS Trust London, UK

Alan Sprigg MBChB, DCH, DRCOG, DMRD, FRCR, FRCPCH

Consultant Paediatric Radiologist Sheffield Children’s Hospital Sheffield, UK John M. Stevens MBBS, DRACR, FRCR Lyshom Department of Neuroradiology Radiology Department The National Hospital for Neurology and Neurosurgery London, UK Dennis J. Stoker MB, FRCP, FRCS, FRCR Emeritus Consultant Radiologist Henley-on-Thames, UK Nicola H. Strickland BM BCh, MA (Hons),

Andrew M. Taylor BA (Hons), BM BCh, MRCP

Stuart Taylor BSc, MD, MRCP, FRCR Consultant Radiologist Department of Intestinal Imaging St Marks Hospital Northwick Park Harrow, UK Henrik S. Thomsen MD Professor and Chairman Department of Diagnostic Radiology Copenhagen University Hospital Herlev, Denmark Paolo Toma MD Radiologist-in-Chief Radiology Department G. Gaslini Institute Genoa, Italy

(Oxon), FRCP, FRCR

Consultant Radiologist Imaging Department Hammersmith Hospitals NHS Trust Honorary Senior Lecturer Imperial College London, UK Louise E. Sweeney MBBCH, BAO, DCH, DMRD, FRCR, FFR, RCSI

Consultant Paediatric Radiologist Radiology Department Royal Belfast Hospital for Sick Children Belfast, UK

Peter Twining FRCR, BSc, BS MB Consultant Radiologist Nottingham University Hospital Nottingham, UK John A. Verschakelen MD, PhD Professor of Chest Radiology Department of Radiology University Hospitals Gasthuisberg Leuven, Belgium

Sarah J. Vinnicombe BSc, MRCP, FRCR Consultant Radiologist Department of Radiology St Bartholomew’s Hospital London, UK Gustav K. von Schulthess MD, PhD Professor and Director Department of Radiology University Hospital Zurich, Switzerland Iain D. Wilkinson BSc, MSc, PhD, CSci, ARCP, FIPEM

Reader in Magnetic Resonance & Consultant Clinical Scientist Academic Radiology University of Sheffield and Sheffield Teaching Hospitals NHS Foundation Trust Sheffield, UK A. Robin M. Wilson FRCR, FRCP(E) Consultant Radiologist King’s College Hospital and Guy’s and St Thomas’ Foundation Trusts London, UK David J. Wilson MBBS, BSc, FRCP, FRCR Consultant Musculoskeletal Radiologist Nuffield Orthopaedic Centre and University of Oxford Oxford, UK Stuart J. Yates MSci, MSc, CSci, MIPEM Principal Physicist Department of Medical Physics & Clinical Engineering Cambridge University Hospitals NHS Foundation Trust Cambridge, UK

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Preface

We hope that this 5th edition of Diagnostic Radiology will continue to build on the original vision of Professors Grainger and Allison who, back in the early 1980s, saw the need for a ‘bible’ for doctors studying for postgraduate examinations in radiology, and to provide a bench book for reporting and reference. The success of the first four editions, which were extremely well received by an increasingly international readership, speaks for the realisation of their dream. Few could have predicted at that stage the extraordinary growth of radiology, or its increasing importance within all aspects of modern medicine. The unprecedented expansion in the imaging repertoire, together with the trend for increasing subspecialisation, have led to changes in training and in the methods used for teaching and learning. This book has had to evolve to reflect these changes, adapting to the perceived needs of those facing postgraduate examinations and also to all radiologists who wish to have an up-to-date basic general textbook for ready reference and illustration. An attempt to cover every subject in detail would have resulted in a huge book that would have been very difficult to use.We have chosen to concentrate on those subjects that most radiologists need to know well, and to pay special attention to the needs of trainee radiologists preparing for examinations. Because training throughout Europe is moving towards a three

year basic course followed by two years of training in selected subspecialties, the factual examination in the UK has moved to an earlier stage in training with less emphasis on some of the diagnostic rarities so beloved by examiners of old. The curriculum is now somewhat less comprehensive and the reduction in size of this 5th Edition reflects that – down from three volumes to two. In this electronic age there are many databases of images available on the internet, with accompanying text. Nevertheless, we believe that well structured textbooks remain an essential part of medical education and practice as they present information in a format that facilitates learning, guiding the reader through an unfamiliar field. We are convinced that Diagnostic Radiology will remain a valuable resource for many years to come. We are again extremely grateful to the distinguished international cast of authors who have all worked hard to deliver fresh and up-to-date material. We are also most grateful to Michael Houston, Gavin Smith and Nora Naughton for their professional skills and publishing expertise and to Jeremy Rabouhans for invaluable help with proof reading. We could never have done it without them! Andy Adam Adrian Dixon

Acknowledgements

This edition could not have occured without the large amount of work done by all the contributing authors and their colleagues. However, the vision and overall planning of Michael Houston at Elsevier have been fundamental in bringing the book to fruition. So, too, has the meticulous gathering and editing of material by Gavin Smith. Finally, the skilful copyediting, and other tasks provided by Nora Naughton, and her remarkable team, must not be forgotten; without them the Editors simply could not have managed. All of these col-

leagues remained remarkably cheerful throughout and kept strong heads even when chapters were late, images missing, and all the other hiccups that can hinder progress in a project of this kind. At a local level, all the Editors would like to thank their various secretaries, technicians and colleagues who have helped proofread, collect material and made various other invaluable contributions. The Editors

CHAPTER

Picture Archiving and Communication Systems (PACS) and Digital Radiology

1

Nicola H. Strickland

The role of PACS Advantages and disadvantages of PACS • Advantages of PACS • Disadvantages of PACS Planning for PACS • The PACS project team • Tendering for a PACS • The PACS contract • Economic considerations • Purchasing versus leasing a PACS • PRE-PACS workflow and preparation • Acronyms: DICOM, HL7 and IHE PACS project implementation • Implementation of digital image acquisition prior to PACS • Integration of PACS ‘Value-added’ PACS Modern PACS architecture • PACS networks • Storage requirements and solutions for PACS PACS workstations • Monitor quality on PACS workstations Conference room design Graphic user interface • Soft-copy tools Software concepts • Prefetching

Compression • Lossless • Lossy • Need for compression Digitization policy Teleradiology • Review of images from home • Teleradiology linkage between two or more hospitals for joint MDTMs • Outsourcing of imaging examinations for reporting • Tele-education Security • User-specific log-in and password • Monitor screen savers • Workstation time-outs • Encryption • Audit trails • Firewalls • Teleradiology security PACS training Quality assurance • Plain radiography • PACS workstations PACS housekeeping • Storage commitment • Modality performed procedure step Current and future directions of PACS

THE ROLE OF PACS A picture archiving and communication system (PACS) aims to replace conventional analogue film and paper clinical request forms and reports with a completely computerized electronic network whereby digital images are viewed on monitors in

conjunction with the clinical details of the patient and the associated radiological report displayed in electronic format. Clearly PACS must replace the functions of traditional X-ray film; i.e. image acquisition, storage, transportation and display.

4

SECTION 1

• IMAGING TECHNIQUES AND GENERAL ISSUES

Were these the only roles of a PACS, it would be an extremely complex and expensive means of replacing traditional film. A PACS must improve upon a film-based system, preferably in a cost-neutral manner. The major added value of a PACS is efficiency of data management. True efficiency benefits can only be realized once a PACS is at least hospital-wide, since any more limited installation means running two systems in parallel, i.e. it entails continuing to produce conventional film and moving it around the hospital, as well as the cost of installing and maintaining a PACS. Thus, even if funds are limited initially, it is advisable at least to aim and plan for growing the PACS installation into a hospital-wide system ultimately. This means decid-

ing upon a potential time scale in which the hospital-wide PACS is to be achieved, and deciding upon a scaleable PACS architecture. Ideally the whole hospital infrastructure should be adapted at the outset so that a hospital-wide PACS can be accommodated at a later stage. This includes providing an uninterruptable power supply (UPS), allowing sufficient cabling space in floors and ceilings, and adapting the air conditioning system for PACS. The hospital information technology (IT) network is likely to need upgrading to enable large amounts of image data to be transported, and it is advisable to install multiple PACS ‘drops’ (workstation outlets) so that more PACS workstations can easily be added at a later date.

ADVANTAGES AND DISADVANTAGES OF PACS Although the concept of PACS has now been in existence for over 20 years, advances in computer hardware technology only enabled it to become a realistic clinical entity in the 1990s. PACS installations are now rising exponentially worldwide, and although most institutions with PACS have achieved a completely filmless working environment, few function in a paperless mode. PACS has proved itself over the last 10 years in the clinical environment, however, and it would now be unthinkable not to implement PACS when installing a new imaging department, or a new hospital.

ADVANTAGES OF PACS There are a number of powerful advantages accruing from a hospital-wide PACS: • Once correctly acquired onto the PACS, no image can ever be lost or misfiled and is always available when needed.This is a major benefit considering that in many hospitals up to 20 per cent of conventional films are missing at the time they are clinically required. In addition to the convenience of always having the appropriate image available when it is wanted for viewing, no patient is re-irradiated simply because a previous key study has been lost. • PACS facilitates the easy comparison of a patient’s current and historical examinations, and of examinations performed on the same body part using different imaging techniques, which is desirable in the interests of optimum clinical practice, and always possible since none of the relevant comparative images is missing1. • All images remain accessible from the PACS archives day and night, every day of the year. • Simultaneous multilocation viewing of the same image is possible on any workstation connected to the PACS network, whereas hard-copy film can manifestly only be in one place at any one time. • Image retrieval is infinitely quicker from PACS than it is using conventional film where someone physically has to go and fetch the film packet2. • The benefits of a computerized system mean that all images correctly and permanently reside under the appropriate

•

•

•

•

imaging study, remain in their correct orientation, and are automatically chronologically ordered. Database searches for a particular patient or study are rapidly effected. Computerized data can easily be duplicated, i.e. ‘backed up’ as a precaution against loss, and cheaply stored in a distant location if desired for disaster recovery purposes. Viewing of images on monitors allows numerous postprocessing soft-copy manipulations: a range of different window width and level settings can be applied to CT images, for example, within a fraction of a second, whereas previously further sheets of film had to be printed with the appropriate settings for soft tissue, lung or bone as desired3,4. There are some direct cost savings due to PACS: there is no longer a film budget, film packet cost or chemical processing. (The ability to print film need not be retained provided a CD burner is linked to the PACS to burn imaging studies for transfer of patients outside the institution, or in the event of a PACS failure or planned downtime). Ancillary staff – filing clerks and darkroom technicians – are no longer needed. The most frequently cited benefit of PACS by nonradiological clinicians is the very substantial time saving incurred by their never having to search for or retrieve films. This time saving represents a considerable indirect financial benefit of a hospital-wide PACS, and should certainly be costed when a business case for PACS is being made. After a hospital-wide PACS has been installed, old films can progressively be removed from the film filing room (film store) starting with the oldest first.The timing of this manoeuvre will depend upon local and national policy, but it has been shown5 that most radiological comparisons are made with studies obtained during the preceding 6 months, although this may be longer in institutions with a large oncological or paediatric practice. Maintenance of a film store, with its associated lighting, heating and cleaning, and the value of the physical space itself, is particularly costly in hospitals located in cities where the space is at a premium. Even off-site film storage at a cheaper location is worth introducing as a cost-saving measure in such institutions in the early years after PACS

CHAPTER 1

•

PICTURE ARCHIVING AND COMMUNICATION SYSTEMS (PACS) AND DIGITAL RADIOLOGY

has been installed, before the film store can be dispensed with altogether. (In the UK it is the radiological report, not the images that are deemed to be the legal document, with certain restrictions for paediatric and educationally challenged patients.) • The installation of a PACS infrastructure in an institution (a local area network [LAN]) sets the stage for the introduction of teleradiology over a wide area network (WAN) if desired. Teleradiology offers the potential for improvements in efficiency, for example in geographically remote areas by centralizing a reporting service, or increasing the referrals to a particular institution. A number of perceived potential benefits of PACS have not been substantiated, or not consistently demonstrated, by audit studies. These include the possibility of a reduced hospital inpatient stay or a greater throughput of outpatients6. It is hardly surprising that such benefits cannot be attributed to PACS since there are so many other variable factors that influence these issues. Also the type of study required to prove any given putative benefit of PACS is fraught with practical difficulties. The classical ‘before’ and ‘after’ study comparing the pre-PACS with the post-PACS era is inevitably complicated by the numerous other concurrent environmental, technological (and often political) changes that have taken place in the interim. These include, for example, changes in the medical personnel and in the type of clinical practice pursued in the hospital under study. Studies comparing a PACS institution with a nonPACS institution may be similarly flawed by the difficulty in adequately correcting for other inherent differences between the two institutions, which may or may not be related to the presence of a hospital-wide PACS7.

DISADVANTAGES OF PACS The advantages of PACS described earlier must be set against its potential disadvantages: • Even though the costs of hardware and storage media continue to reduce in price, PACS remains an expensive technology. Most estimates suggest that a PACS installation should aim at becoming cost-neutral in less than 5 years. Some would argue that PACS should be viewed in the same way as any new imaging technique, and as such it represents an advance in health care management and should not be assessed merely in terms of cost–benefit. • The technological complexity of a PACS and the absolute dependency of a hospital on the PACS once it becomes filmless require a dedicated maintenance programme for the PACS and a carefully devised plan detailing how to supply a minimal clinical service should the whole PACS fail for a significant period of time. This inevitably means that there will be a requirement for new or retrained hospital personnel specializing in computer engineering/information technology, as well as a vendor-provided maintenance service, and these costs must be set against the savings made in respect of less highly paid clerical staff and darkroom technicians. • Once a hospital-wide PACS is in operation and film has been withdrawn, there is no ‘fall-back position’. The hospital is no longer equipped to run a film-based service. This is a daunting prospect that may act as an initial psychological deterrent to embarking upon a large-scale PACS project. • Changing from a hard-copy to a soft-copy imaging environment will raise many issues necessitating a change of work patterns involving: the training of the users, maintenance of the system, action to be taken in the event of a PACS failure and the institution of specific quality assurance protocols (see later).

PLANNING FOR PACS THE PACS PROJECT TEAM The planning and installation of a PACS requires the coordinated input of a multidisciplinary team that might beneficially include representatives from the IT department, computer scientists, physicists, nonradiologist clinicians, radiologists, radiographers, nurses, hospital management, the hospital financial manager and, ultimately, a representative from the chosen PACS vendor. This emphasizes that PACS must have ‘buy in’ from the whole hospital/health care enterprise, and is not a ‘radiological toy’. The project team needs a leader with sufficient time to commit to the project8. This leader need not be a radiologist, or indeed a clinician, but must have a comprehensive practical grasp of the clinical workings of a hospital environment as well as an understanding of basic IT issues and the requirements of imaging.

TENDERING FOR A PACS When tendering for a PACS, a detailed request for tender/ proposal (RFT/RFP) will need to be drawn up9,10. This

document should define the clinical and logistical requirements of the PACS, rather than merely stipulating technical specifications. It is up to the radiologists and other clinicians to specify the current and projected future performance requirements that the PACS must fulfill (including the issues of PACS maintenance, uptime and data migration) and it is the role of the vendors to specify how these clinical requirements are to be met technically.

THE PACS CONTRACT Every PACS project should be based on a firm contract between the institution and the vendor, defining the responsibilities of each in detail. The PACS contract serves as a legal document and requires careful wording. Any subsequent changes or additions to the contract should be made as formal addenda ‘change control notices’ (CCNs) to the original contract, to preserve the legal integrity of the document. This also serves as an audit trail, keeping the contract up to date, and simplifying the situation for both purchaser and vendor.

5

6

SECTION 1

• IMAGING TECHNIQUES AND GENERAL ISSUES

It is important to define what is meant by ‘the life of the system’. This is often taken as being 8–10 years after the completion date.The majority of PACS installations are now based on some form of leasing agreement, rather than an outright capital purchase of the hardware and software. The date at which the maintenance contract will commence, its cost and its terms, all need careful definition. Responsibility must be defined for migration of the PACS archive data at the end of the PACS contract, or in the event of premature termination of the contract. It is advisable to define exactly what is meant by an ‘update’ to the system (generally a minor software release that merely corrects bugs in the system), and what is meant by an ‘upgrade’ (which constitutes a new software installation comprising major new features giving enhanced functionality), whether these are to be included in the purchase price/maintenance agreement, and at what frequency these will occur. An important consideration in negotiation with a vendor is the terms under which the delivered PACS hardware will be replaced if upgrades are released that it cannot support. It is vital to define which party is ultimately responsible for the functioning of the interfaces to the various pieces of imaging equipment, and to other hospital computer systems including the hospital and radiological information systems (HIS and RIS), electronic patient record (EPR), speech recognition dictation system, and electronic remote requesting (order communications) system, if these are already extant in the institution.

ECONOMIC CONSIDERATIONS The main reason for putting in a PACS (ultimately a hospital-wide or larger PACS) is to improve the efficiency of data handling throughout the whole of that health care environment. Much of the benefit will be experienced outside the imaging department itself, which is why PACS must be hospital driven (and funded), not radiology driven. PACS should be part of the hospital-wide information management and technology strategy, with the aim of achieving cost neutrality over about 3 years.

PURCHASING VERSUS LEASING A PACS The options for PACS procurement are broadly 2-fold11: 1 Capital purchase, using the hospital’s capital allowance, or 2 Some form of leasing arrangement. The fundamental difference between the two is that with capital purchase the hospital owns the assets (the PACS), whereas with any leased arrangement the hospital procures a service that contractually provides agreed functional outcomes (the output-based specification [OBS]). The vast majority of PACS installations nowadays are on a leasing scheme with a managed service provided by the PACS vendor. A leasing arrangement means that all the PACS equipment (hardware and software) is provided as a service from

a private company (with a maintenance contract). Nothing is owned by the hospital, i.e. the private company retains all the assets. The advantage of a leasing arrangement is that it allows hospitals with no hope of ever having the large capital sum necessary to purchase a hospital-wide PACS to make a quantum leap in technology to move to a PACS solution, and it transfers the risk to the company. It has to be appreciated, as with renting a house, that the hospital would be left with nothing if it were to stop paying the lease. At the end of the contractual term, the hospital does not own anything. However, in these circumstances the hospital is usually given an option to buy the installed assets (the PACS hardware and software) at ‘a fair market value’, as negotiated with the provider. In a leased service the PACS company provides the PACS hardware and software necessary to meet the workload and performance requirements stipulated by the hospital, both at the time of leasing and in the future. For example, such requirements might include the need to perform 400 000 imaging examinations per year, with a short-term storage of a year, and a display time for these examinations (first image to screen) of 3 s or less.The PACS company is responsible for maintaining a PACS with an agreed level of technology throughout the contractual term. The technological risk rests with the provider. The risk covers the following three areas: • Utilization of the system • Availability of the system to users • Future planning and implementation of new technology into the hospital’s system. The risk is transferred to the service provider by linking their revenues to agreed performance targets for the above. In addition the hospital should be provided with a guaranteed programme of equipment replacement and a guarantee to keep pace with technology. Leasing arrangements vary from about 8–30 years, with some planned ‘built-in’ equipment replacement ‘hardware swap-out’ at 5-yearly intervals, or 3–4-yearly technological reviews with ‘technology refreshments’ of software and/or hardware as may be necessary to provide greater productivity or capability. This technology refreshment may be a chargeable issue depending upon the contract and/or negotiations. Companies generally expect to recoup their capital outlay in approximately 3–5 years in a leased arrangement, but this will of course depend upon the financial model being used.

PRE-PACS WORKFLOW AND PREPARATION A number of careful and detailed preliminary pre-PACS studies of the workload and workflow pattern within the hospital/ health care facility will need to be conducted before the clinical requirements to be met by the PACS can be specified12.This includes documenting how many imaging examinations in each modality are performed annually, the average (and maximum) number of images per examination for each modality, and making a prediction of the rate of growth of the imaging workload.This prediction will be influenced by the expectation,

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for example, of acquiring a new 64-slice (or greater) multidetector CT machine, the intention to start performing high image acquisition studies in MR such as cardiac or breast imaging and so forth. The predicted image storage requirement should always be an overestimate to allow for unexpected demand. It is also advisable to know which outpatient clinics are performed when, and how many film packets are pulled for each, to give some estimate of the network traffic to be expected. It is also important to define the role of the imaging department in other imaging-related activities, such as the radiological steps involved in conducting multidisciplinary team meetings (MDTMs) and radiological presentations at staff rounds, and undergraduate and postgraduate teaching sessions. PACS must be able to fulfill all these functions. The Imaging Directorate would be wise to know before PACS is installed, exactly how many imaging studies are never reported (for various reasons such as: the film packets are never returned to the department for reporting), the time between image examination acquisition and dictation of a report by a radiologist for each investigation and the time delay (if any) between dictation of the report and the availability of the verified report to other clinicians. These workflow deficiencies need to be addressed and corrected prior to the installation of a PACS, since the PACS will not ameliorate the situation but instead will expose these issues by making such information available in computerized form throughout the PACS institution. Seamless integration with other IT systems, and with imaging acquisition devices, is absolutely vital for a PACS to function successfully, and every modern PACS depends upon adherence to DICOM and HL7 standards and IHE (see next paragraph for an explanation of each) workflow processes to achieve this integration13. Before installing PACS, old equipment and old data information systems will need to be upgraded to a minimum level of DICOM and HL7 compliance, respectively. Often it is cheaper to replace such products with modern versions, rather than to pay to upgrade them. A full inventory must be made of the equipment and IT systems to be connected to the PACS with a precise description of the level (if any) of DICOM or HL7 compliance supported, before a PACS project can be properly planned and costed. The full DICOM conformance statement of every piece of DICOM-compliant equipment (e.g. a computed tomography [CT] scanner, an ultrasound [US] machine, a workstation etc.) is available on the Internet. It is generally the users’

responsibility to list the make and model of all the equipment possessed by the hospital/health care facility, and the PACS vendor’s responsibility to look up and interpret the DICOM conformance statements and to make it clear which pieces of equipment will need upgrading/replacing, and to explain the connectivity limitations if such DICOM upgrades are not undertaken. One of the major causes of interoperability problems when linking equipment from different vendors to a PACS is the use of ‘private DICOM attributes’ by many vendors, which, in simple terms, means that information stored in these particular private DICOM fields is not available to be shared with other apparatus, manufactured by a different vendor, which may be linked to it on a network. This leads to a loss of functionality on the recipient apparatus, for example not being able to post-process scanner images received on a workstation from a different vendor.

ACRONYMS: DICOM, HL7 AND IHE DICOM stands for digital communication in medicine and refers to a worldwide multipart standard to which all modern imaging equipment and PACS must adhere, and has now been extended to other disciplines, including cardiology, endoscopy, and ophthalmology. The DICOM conformance statement of every piece of modern imaging equipment is obligatorily available on the Internet, and the description of the various DICOM attributes possessed by each appliance will predict its connectivity with another piece of equipment. All apparatus is described as being a ‘user’ or a ‘provider’ of services such as storage.The DICOM standard is continuously under development, but each new part of the DICOM standard has to be backwardly compatible with the current DICOM standard. HL7 stands for health level 7 and refers to a worldwide standard for data information systems such as HIS and RIS. It is a less rigorous standard than DICOM. HL7 messages from data information systems are conveyed to DICOM apparatus (including PACS) by a ‘broker’, which acts as an integrating and translation device. IHE stands for integrating the health care enterprise and is not a standard, but a comprehensive workflow descriptor of how processes, such as reporting, for example, are achieved in imaging. Its use of integration profiles eliminates the need to reconcile the details of HL7 messages and DICOM conformance statements among multiple vendors. It is now being expanded outside the discipline of imaging.

PACS PROJECT IMPLEMENTATION The PACS project should be treated as a major IT project and divided into stages (milestones) with specific dates set for the completion of each stage. It may be useful in the UK to base the project management on the PRINCE guidelines published by the HMSO14 for major IT equipment contracts. Only after each consecutive milestone is satisfactorily completed and assessed is the next milestone embarked upon.

Each milestone requires conformance testing and clinical acceptance15. These are distinct entities: conformance testing should be carried out independently by hospital employees as well as by the vendor company itself. Payment is best deferred until after clinical acceptance at each stage. Final payment is only made after agreed satisfactory functioning of the system under loaded conditions, in a clinical setting.

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A realistic definition of the expected dates of completion of each milestone (including installation dates), with penalty clauses written into the contract to come into effect if these dates are overrun, protects the institution from major delays since penalties can be extracted from the vendor if such delays occur.

IMPLEMENTATION OF DIGITAL IMAGE ACQUISITION PRIOR TO PACS Plain radiography (chest, abdominal and skeletal plain images) still constitute the majority (usually 60 per cent) of imaging examinations in most general radiology departments. These examinations therefore need to be acquired in a digital format in order to be transferred to PACS. This represents a considerable challenge since conventional plain film work is still such a significant part of the total departmental workload. The other imaging investigations (computed tomography [CT], magnetic resonance [MR], nuclear medicine [NM], positron emission tomography [PET], US, digital subtraction angiography, and fluoroscopy) are either already digital in nature at acquisition or can be rendered so by screen capture/frame grabbing from the acquisition device, and can thus be easily transferred to PACS. There are three basic means of rendering plain radiographic images digital: 1 Digitizing conventional analogue film 2 Photostimulable phosphor plate technology, commonly known as computed radiography (CR). 3 Direct digital radiography. There is considerable new technology associated with acquiring plain radiographic images digitally, whichever method is chosen. In situations other than the opening of a new hospital/health care facility, it is circumspect not to introduce this new technology concurrently with a PACS to avoid the risk of ‘technology overload’ for the users. Once the bulk of the imaging studies (i.e. the plain radiography) is being acquired and stored digitally, the introduction of PACS and the complete withdrawal of film is a less daunting task. The overlap period, in which a digital archive is being built up whilst continuing to distribute film outside the imaging department, should be kept to a minimum for financial reasons: it is obviously expensive to run hard- and soft-copy systems concurrently. A period of <3 months is a reasonable timescale for the phasing in of PACS after digital plain radiography has been introduced. The digital archive acquired during this time is sufficient for the majority of comparisons with previous plain images4, and obviates the need for a great deal of retrospective digitization of conventional film.

Digitization of analogue film This is a time-consuming process that requires relatively skilled operators who are able to differentiate the various imaging modalities and body parts and orientate the films correctly for digitization and storage on PACS. Digitization must take place in a nondusty environment. A digitized image retains any intrinsic limitations present in the analogue film: areas that

were underexposed on the original film remain underexposed when digitized; areas that were overexposed on the original film would normally yield more information when viewed over a bright light, but tend just to look black when digitized. Digitization on a large scale is not a practical option and combines the expense of a film budget with the cost of digitization and maintenance of a soft-copy archive. Digitization may be a viable solution in a very-low-throughput teleradiology service, and is necessary in a PACS environment to allow hardcopy films acquired at a nonPACS institution to be entered onto the PACS for comparison with subsequent digital images acquired at the PACS institution. Since digitization is a costly and labour-intensive process, yielding digital images of suboptimal quality, a sensible minimalist digitization policy should be adopted when PACS is installed. A reasonable compromise is only to digitize one previous ‘key’ image in patients for whom it is regarded as clinically highly desirable to have a comparison image available long term on the PACS archive (for example, oncology patients on chemotherapy, or patients with slow-growing tumours or chronic disease). Some institutions choose not to digitize any previous films, but undertake to make past hardcopy imaging available on request (from off-site storage) for over a year. It is clear that unselective ‘blind’ digitization of batches of previous films is a waste of time and money considering how few previous imaging studies are ever subsequently reviewed4.

Photostimulable phosphor plate technology/ computed radiography This is the commonest means of acquiring plain radiographic images in a digital format. Photostimulable phosphor plates replace the conventional film/screen combination in a cassette and can be used with standard X-ray equipment (although not for fluoroscopy). X-ray photons falling on the phosphor plate produce a latent image by promoting electrons to metastable traps of higher energy level. When the plate is ‘read out’ in raster fashion by a scanning laser beam in the plate reader, these electrons gain sufficient energy to fall back to their basal energy level, and in so doing, emit photons of visible light that is converted to a final digital signal (Fig. 1.1). Phosphor plates have a flatter, more linear, optical density versus radiation exposure response curve than the classic sigmoid-shaped curve of the conventional film/screen combination (Fig. 1.2). This technology thus gives a greater latitude of exposure, which is useful under challenging radiographic circumstances, such as in the intensive care unit, and allows structures of widely differing radiographic density to be well visualized on the same image, such as the area behind the cardiac silhouette and the lung vessels, or the soft tissues of the neck and the cervico-thoracic junction. This feature allows a reduction, compared with film, in the number of repeat images taken due to errors in selecting the correct exposure. The spatial resolution of CR is lower (approximately 2½ line-pairs per mm) than that of film (approximately 5 line-pairs per mm), but its superior contrast resolution more than compensates for its limited spatial resolution.

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Xray -photons Exposure

Caesium Iodide

Photostimulable phosphor plate

Light Selenium

Laser

Latent image

Luminescence

Read out

Amorphous silicon panel (photodiode/transistor array)

Amorphous silicon panel (photodiode/transistor array)

Electrons

Electrons

Read out electronics

Read out electronics

Digital signal

Digital signal

Figure 1.1 Diagrammatic representation of digital image acquisition by photostimulable plate technology (CR).

Phosphor plates are repeatedly reusable for a thousand or more exposures, although they do have a finite life after which they start to crack and need replacement (and this is a cost that must be allowed for in the departmental budget). CR apparatus is cheaper than direct digital radiography (DR) equipment, especially for mobile radiographic units.

Direct digital radiography Current direct DR apparatus (other than in dental work) uses amorphous silicon with a digital detector. Different systems on the market use different detectors to capture the X-ray photons: caesium iodide or rare earth scintillators, which emit light photons, or selenium, which produces an electrical signal (i.e. a direct digital signal) (Fig. 1.3).The amorphous silicon is used to make plates consisting of an array of photodiodes/transistors, each acting as a single pixel to amplify the incoming light or electrical signal and give a digital electronic output. When it became technically possible to construct flat-panel detectors large enough, they started to be used in plain radiography.The detector plates are fragile, heavy and fixed, and have only very recently been adapted for use in mobile direct DR apparatus (which remains much more expensive than the phosphor Film/screen 3.0 CR or DR

Dark

2.0 Linear density range Lungs 1.0 Heart

Light 0.0 0.0

Xray-photons

X-rays

1.0

2.0 log Relative radiation exposure

3.0

Figure 1.2 Optical density versus radiation exposure response curves. The phosphor plate has a more linear response than the sigmoidshaped curve of the film/screen combination. Photostimulable phosphor plates have a wider latitude than film/screen.

Figure 1.3 direct DR.

Diagrammatic representation of image acquisition by

plate alternative).Thus at present there is a continuing concurrent need for CR using the much cheaper phosphor plates for mobile radiography and to provide redundancy. The spatial resolution of direct DR equals that of film and it has better detective quantum efficiency (DQE) – giving a guaranteed possible dose reduction per exposure – and better contrast resolution. Unlike photostimulable phosphor plates, a direct digital signal is produced and there is no manual processing of plates involved. The current higher cost of direct DR apparatus still favours its use mainly in mammography and in very-high-throughput radiology units16.

INTEGRATION OF PACS The seamless integration of the PACS with existing image acquisition systems and with other hospital computer systems is the most important and challenging aspect of any PACS project.This is why the person/vendor with ultimate responsibility for the various planned interfaces must be clearly identified and involved in the PACS contract.

HIS–RIS–PACS integration It is crucial to have seamless integration between PACS and RIS9 and HIS. No PACS installation should be contemplated unless there are already integrated HIS and RIS systems within the institution, or they are being installed simultaneously with the PACS. It is most unwise to implement a PACS before a HIS/RIS since the PACS database will stand alone, isolated, and the retrospective merger of databases is time consuming and expensive. It could be argued that it is unnecessary to integrate HIS with PACS (even if this is done indirectly via the RIS) (Fig. 1.4). However it is only the HIS, not the RIS, that is aware of all the patients known to the hospital/health care enterprise irrespective of whether they have previously had imaging studies. If such a patient attends for an imaging study, his or her demographic details need to be automatically transferred from the HIS to RIS and PACS. Furthermore it is only the HIS that

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HIS HIS

RIS

RIS PACS

Figure 1.5

PACS

Bi-directional integration between PACS, HIS and RIS.

Figure 1.4 Unidirectional interfaces between HIS, RIS and PACS. Here HIS is only connected to PACS via RIS. Scanner

is aware of the patients who have been scheduled to attend particular outpatient clinics, and this information can be used directly by the PACS to prefetch the imaging studies for these patients from a long-term archive prior to their attendance at the outpatient clinic. There are three main reasons in favour of having bidirectional integration between these three systems (Fig. 1.5): • There is input of demographic data only once, minimizing human error. • Any update to a patient’s demographic data on HIS, RIS or PACS is propagated to all systems automatically, ensuring that all systems contain accurate information. • Any scheduling or status data is distributed to all systems automatically, providing all systems with advanced notice of events and allowing them to prepare to play their roles in these events. A good example of this is the concept of pre-fetching, whereby pertinent historical examinations are fetched to the PACS short-term storage unit in advance of a patient’s attendance at an outpatient clinic, hospital admission or imaging examination; these data are known to the HIS and/or RIS and passed to the PACS prior to the event concerned.

Radiology information system The RIS may be a stand-alone computer platform, or may be a module of the HIS12. It stores information specific to the radiology department, including the registration of patients attending only the imaging department for an imaging investigation, but not otherwise under the care of the hospital (e.g. general practitioner referrals); schedules examinations, including date, time and room allocation; may track the patient’s progress through the imaging department; and stores radiological reports and produces statistics concerning imaging workflow. A modern RIS will incorporate some DICOM features such as DICOM modality worklist (DMWL), (Fig. 1.6).

RIS Single data entry

Figure 1.6

DICOM modality worklist.

whereby the demographic data of a scheduled list of patients is downloaded directly to the imaging apparatus (e.g. CT system) due to perform the studies (provided the imaging modality also supports DMWL). This obviates the need to retype the demographic details of patients on the imaging equipment or the PACS, and thereby eliminates work and human typographical errors, which would otherwise cause the images not to be recognized when later sent to a mismatched entry for that patient on PACS. A reference number (an ‘accession number’) is assigned by the RIS to each study and communicated to the imaging room as part of the DMWL to obtain a one-to-one correspondence between the DICOM data set and the RIS-related information.

Hospital information system or electronic patient/ medical record The HIS (or in some health care institutions, the EPR) is the most important computer system in a hospital. It constitutes the patient master index that stores demographic (textural) data on all patients known to the hospital, whether or not they have ever had any imaging procedures performed. In addition to this demographic data, e.g. patient name, identification number and date of birth, the HIS also records admission and discharge dates, outpatient appointments, clinician(s) responsible for the patient’s care and so forth. Other modules of the HIS document laboratory data, usually including haematology, biochemistry and histopathological test results. Ultimately a true EPR should contain all the case notes for a patient in digital format, but most systems are a long way from achieving this goal.

‘VALUE-ADDED’ PACS Once hospitals have PACS installed, it will gradually become clear to them that having a PACS, whereby imaging studies and their reports are electronically stored and transmitted, is only the first step in achieving a fully integrated, seamless, electronic system of health care workflow. Becoming filmless is not enough: all health care systems are a long way off the aim of also becoming paperless. In order fully to achieve the latter, there can be no paper request forms, and no paper copies of reports.

Paper request forms can be integrated into the PACS environment using a work-around solution, such as digitizing the paper request forms prior to acquisition of the imaging study, or using a bar code hand-held laser reader on the patient label on the request form at the time of imaging and reporting to locate the electronic entry on RIS, PACS and the speech recognition system. Such processes have a number of drawbacks, do not produce a streamlined workflow and, most importantly, they do not eliminate the existence of a paper request. It is only by

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ceasing to be dependent upon the physical existence of a paper request that health care institutions can function as virtual hospitals, and become location independent with respect to the sites of generation of the imaging study request, the acquisition of its imaging study, and the generation of its report17. To eliminate paper requests, imaging studies must be booked using a remote electronic requesting system (sometimes known as ‘order communications [order comms]’). The image study request details must be entered electronically at the time of booking (either by directly typing them into the system or by using speech recognition to enter them). The booking of imaging studies is rather more demanding of such electronic requesting systems than the booking of other investigative studies, such as haematology or biochemistry profiles or histopathology examinations. Radiologists in the UK are legally bound by IRMER (ionizing regulations in medicine regulations), whereby studies using ionizing radiation must be clinically justified and therefore the clinicians’ requests for such imaging studies must be vetted by a radiologist to ensure they are indeed clinically indicated. An electronic requesting system for radiological studies must allow this vetting step to take place electronically, with an audit trail of the person who changed an imaging request and the reason for the change; the details of the new procedure scheduled must be made known to the referrer and his/her team, and to the patient. The generation of the radiology reports should ideally be via an electronic speech recognition system fully integrated with both the RIS and the PACS. In this way when an imaging study is displayed on the PACS monitors for reporting, the integration means that both the RIS and the speech recognition system will automatically ‘know’ which imaging study is being reported. The speech recognition system immediately converts the dictated report into text, which is corrected there and then by the radiologist author, and is electronically signed. This report is permanently stored in the RIS (the location where radiological reports are conventionally stored long term) and is visible on the PACS and in the HIS/EPR. Speech recognition adds considerable value to the PACS since imaging workflow efficiency is markedly improved, by eliminating any delays between dictation and typing, and between typing and verification by the radiologist. Once a report is dictated

and signed by the radiologist it is instantaneously available to any authorized user accessing the study on the PACS. The ‘imaging loop’ should be closed by installing electronic feedback of results, whereby the relevant clinician(s) are electronically shown the reports on imaging studies they have requested on their patients. Returned results therefore have to be grouped by referring clinician, or preferably, by referring clinical firm (team). There needs to be a means whereby the reporting radiologist can, at the time of dictating the report, electronically add an ‘alert’ flag to a report revealing an unexpected or urgent clinical finding to draw it urgently to the attention of the referring clinician and his/her team. The imaging loop is closed by the clinician (or his/her team) responsible for the health care of the patient reading the report on the imaging study and taking responsibility for it. There needs to be a robust audit trail to record who read the report, and, more importantly, who took responsibility for understanding and acting upon it to avoid imaging reports being overlooked, and also for medico-legal purposes should an important imaging result be missed. It is not adequate to rely upon communication of urgent results by telephone in case (as often happens nowadays) the relevant clinician is no longer on duty or is unable to be contacted for some reason. Paper reports sent by conventional mail or fax can also be unreliable and there is no confirmation that they have reached their destination and been read by the relevant clinician. Thus the electronic communication of results system must incorporate a two-step receipt procedure whereby the reader must actively click twice (e.g. in two different check boxes), once to acknowledge reading and again to indicate understanding of (i.e. taking responsibility for) the result of the imaging procedure. These steps must be traceable by the system’s audit trail. If this form of verification is not followed then an electronic result may appear to have been read (for example, as with an email programme, by changing from bold to normal type font) whereas in fact it has merely been opened, or opened by a medical student or junior doctor who will not take responsibility for dealing with the consequences of the report findings. Unfortunately these electronic requirements for closing the imaging loop have yet to be implemented on current commercial system software.

MODERN PACS ARCHITECTURE Today’s high-speed computers and networks mean the hardware architecture is no longer an issue; it is the logical architecture that matters. For large PACS, everybody builds essentially the same hardware architecture today (multiple servers on fast networks); the only real difference is whether the systems present a logically centralized view or a distributed (autorouting) view to the user.

PACS NETWORKS The network is the critical link for all the individual components in a PACS. Its performance is a major factor in determining the ability of a PACS to support the examination volume of the hospital/health care institution.

Most PACS products are industry-standard networks, usually consisting of a high-speed network, often called a backbone, which spans the entire institution, with branches that connect individual components or groups of components to the backbone. Networks can be differentiated in several ways, including cable type, speed, communication protocols and topology. Data are transmitted over the physical network using a protocol.The performance of the PACS, i.e. its response time (judged by the radiologist as the speed with which an image appears on the PACS monitor after the display command is activated on the workstation) is determined by the cable quality and speed, and by the protocols used for carrying imaging data.

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It is the quality and speed (bandwidth) of the enterprisewide network that determines the rapidity with which images are displayed on the PACS workstation and web browser computer monitors. The quality of the PACS hardware and software is immaterial if image delivery is rate limited by an inadequate portion of the network. Radiological images contain very large amounts of data: a single plain chest radiograph contains from 8–10 megabytes of data, and there are considerably more data in the very large (around 1000–2000 individual images) thin-section CT and MR cross-sectional image datasets often now acquired for a single study. Networks nowadays run standard ethernet protocol, with TCP/IP (Transmission Control Protocol/Internet Protocol) of the Internet. At least 100 megabit ethernet is needed to transmit radiological images and some institutions are now running 1 gigabit ethernet throughout the health care enterprise. The core of many PACSs still run propriety network protocols to achieve greater speed of transmission. Modern hospitals have a switched network rather than a hub network because a switched network allows the whole bandwidth (e.g. 100 Mbps ethernet) to be available to each switch in the network (whereas in a hub network the entire hub accommodates the total bandwidth (e.g. the 100 Mbps) and all the attached items have to share a portion of that bandwidth).

STORAGE REQUIREMENTS AND SOLUTIONS FOR PACS The PACS requires two basic different types of storage for: • Image data • Indexing data.

PACS image data storage The PACS requires a very large amount of storage for the images (which includes associated objects such as presentation states), but the requirements are fundamentally different from the types of storage used by other large data users such as banks. The peculiar features of PACS image storage are: • There is an ever-increasing need for more storage space • The data once written are rarely (if ever) changed • Images tend only to be accessed a few times during their entire life and normally by only one user at a time • Required disk access speed is not actually very high, as enduser access is more limited by network speed than by disk limitations.

PACS indexing data storage The PACS has more ‘conventional’ requirements for its indexing data storage where: • The total volume of data is tiny by modern standards (a few gigabytes at most) • Data are constantly being updated and changed • Concurrent access by multiple users (for queries) is required.

Historically, the costs of image data storage were the main determinant of the cost of a PACS, but nowadays, they should only be a minor component, as disk costs have fallen so dramatically, and PACS software costs are much higher than its hardware costs (although vendor mark-up for hardware can skew this ratio).

Storage volumes and types In the early days of PACS, storage volumes were mainly determined by the size and number of CR images, but multi-slice CT and large MR studies are now overtaking CR, and required data volumes are still increasing, such that even after applying lossless compression, an average district general hospital will need about 2 Tbyte and a major teaching hospital about 5 Tbyte per year18. There have been many types of storage used over the years for PACS to handle such large volumes, including optical disks (magneto-optical disks, MOD or compact disk/digital versatile disk, CD/DVD), tape-drives and jukeboxes containing both of these. All these mechanical devices, with their attendant reliability issues, are now obsolete, and the only sensible method of storing images for normal use is Redundant Array of Inexpensive Disks (RAID) – a system that spreads data across many disks, providing the following advantages: • The data are still accessible if one (or, in some configurations, more than one) drive fails • The speed of the system as a whole is faster than the individual disks, because they work in parallel. RAID itself can be used in multiple different configurations and ways of connecting the PACS server, with widely different pricing, but the main categories are: 1 Directly attached to the server. This is cheap and simple, but expansion then requires the whole system to be stopped and restarted, and expansion is also limited. 2 Attached to a network using file servers (network attached storage [NAS]). This is mature standard technology that offers an excellent choice of products as fully supported ‘off-the-shelf ’ hardware, to which any competent server should be able to connect.

Storage area networks SANs are basically RAID, but connected to their servers using special connections (normally optical fibre). They provide extensive facilities for management, backup, flexible use of space and sharing of disk space across multiple servers, but this flexibility comes at a price, as they are much more expensive than equivalent space bought in the form of file servers, normally being 5–10 times as expensive per Gbyte.

Proprietary storage systems It is possible to replace the above ‘open’ and standard connection methods by proprietary equivalents, which are claimed to have unique special features, but costs are high and there is then ‘lock-in’ to the particular vendor concerned.

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Mirrored storage It is possible to have duplicate copies of the data held on two or more devices, preferably in physically separate locations. In theory this is possible for any of the above basic types, but in practice it is difficult to do for directly attached storage, and for SANs and proprietary methods it can be prohibitively expensive, so it is most commonly applied to NAS devices.

Backup Whilst it is useful to have all the institution’s PACS data immediately available on RAID, there is always the risk that it could be lost, through any of the following: • Disk failure (RAID should be robust, if the fault is spotted and the disk replaced before another fails) • Other hardware failure (RAID is not immune from this) • Fire, flood and malicious physical damage • Software errors and viruses (which can even affect mirrored RAID). Therefore, disaster recovery backup onto media that are not on-line and therefore not accessible to viruses etc. is important, and tape is perfect for this. Note that there is no requirement for a jukebox, merely a good fire- and flood-proof safe to isolate and protect the tapes! Backup of PACS image data is very different from the backup needed in most industries, due to the ‘static’ nature of the images, and therefore it is important that the backup strategy is properly designed around this pattern. In particular: • Regular backing up of the whole set of images stored is a waste of time (since once acquired, the images are not going to change) • Each image probably only needs to be on two tapes • Those tapes should be ‘refreshed’ every year or two depending on the manufacturer’s stated data life • Conventional ‘tape rotation’ must be avoided, as it could cause good (original) data to be over-written by errors.

Compression implications for storage It is important to check that an appropriate compression strategy is used, as properly managed lossless (reversible) compression with the same compression used over the network can: • Reduce required storage space by about 60 per cent • Decrease image access times (as less data need to be read off the disk and moved over the network).

These benefits have no significant downsides, and money spent on uncompressed storage space is simply wasted. There are academic arguments in favour of moderate lossy (irreversible) compression, which could probably reduce longterm storage space by a further 75–85 per cent (90–95 per cent overall) but this is a more contentious subject, and is not as clear-cut as the incontrovertible arguments in favour of lossless compression.

Optimal PACS storage Bearing in mind all the points mentioned earlier, there are two very different requirements for PACS data storage: 1 Indexing data. These are the sort of data well understood by industry, requiring the highest speed, most robust and manageable disk available, combined with regular complete backups.The cost per Gbyte is obviously high, but the actual space required is very small. 2 Image data. These data do not require fast, over-managed, repeatedly backed-up storage – merely large quantities of ‘cheap and cheerful’ storage – preferably duplicated RAID for which network-attached storage or similar is ideal. Furthermore, by buying only what is needed as required, prices decline over time in line with commodity storage advances.

Commercial considerations Clearly, the arguments given earlier can only be put into practice if it is possible to specify the type of storage used for a PACS, and this is not possible if decisions have been made elsewhere precluding such choice, or if it is not possible to buy the ‘core PACS’ and the storage independently. At least one of the major PACS vendors has however been known to allow such independent purchase in some areas of the world, enabling the purchase of PACS software and integration expertise from those best able to supply it (the PACS companies) whilst at the same time, buying the storage itself from those best able to supply that component – the commercial storage solution providers. Nevertheless, if the PACS storage is not purchased (at a premium) from the PACS vendor, there must be sufficient highly competent in-house expertise available to manage this storage independently from the PACS vendor, and the PACS contract will need to make special reference to the responsibility for maintaining and managing the PACS storage separately from the core PACS software and integration.

PACS WORKSTATIONS MONITOR QUALITY ON PACS WORKSTATIONS Fitness for purpose ‘Diagnostic quality’ monitors means that the monitor used for viewing the digital image should be capable of showing the image at its maximum spatial and contrast resolution (i.e. the spatial and contrast resolution at which the image was acquired). The spatial and contrast resolution of images from the various imaging modalities differs, for example: plain

radiography (such as a digitally acquired chest X-ray) has a higher intrinsic data content than a single image from a CT, MR, or US scan, and therefore would need a better quality monitor to show all the data it contains. The key issue in every part of a PACS specification, including the question of monitor quality, is fitness for purpose. The following questions have to be answered, with respect to image display monitors: ‘Why am I looking at the image in this particular location?’ Is it to make a final diagnostic report

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(as issued by a radiologist)? Is it to identify urgently treatable pathology before the image has been seen and reported by a radiologist (as might be the case for a clinician on the intensive care ward, or in the Accident and Emergency [A&E] department), trying to diagnose a pneumothorax or pneumonia)? Is it as a reminder of the imaging findings on this patient by reviewing his/her imaging studies, generally also with access to the definitive radiological report (as might be the case in the outpatient clinic or on a ward round)? The monitor requirements in all these scenarios differ. It is perfectly acceptable for a nonradiological clinician to make a diagnosis from an image that is not displayed at its maximum resolution, and to go on to treat the disease displayed (e.g. a pneumonia or a fracture) before the image is seen and reported by a radiologist viewing the image on a fullresolution monitor. Fitness for purpose is the key here: it would not be appropriate for the clinician to base patient management upon such an image viewed solely upon a mobile phone screen, but conversely most lesions requiring urgent treatment can be well seen on a 1 megapixel monitor, with access to a soft-copy magnification tool, and do not have to be seen on a 3 megapixel monitor to be detected19. It is also important to bear in mind that many subtle radiological signs seen on plain radiographs (such as early interstitial lung disease, or subperiosteal bone resorption) relate to diagnoses that are generally not nowadays made clinically by the plain radiographic appearance. In the cases cited, the early interstitial lung disease would be suspected clinically and on lung function tests and confirmed radiologically by high-resolution CT; the subperiosteal resorption would occur in gross hyperparathyroidism already diagnosed biochemically. In any case most subtle plain radiological appearances do not require immediate treatment before the images can be further reviewed on a higher resolution monitor, and other investigations using different imaging modalities and various clinical tests to confirm the diagnoses, will be carried out before any treatment is commenced. A digital chest radiograph acquired using either CR (phosphor plate technology), or direct digital technology, displayed at full size at its acquisition resolution, requires a monitor with the following minimum specification: • 3 megapixels (for spatial resolution) • 700 cd/m2 brightness (high luminance) • Greyscale, not colour • 256 linear-sized levels of greyscale (DICOM part 14 compliant display), usually of 10 bit depth (for contrast resolution). Diagnostic quality monitors of the specification cited earlier are only necessary for viewing ‘plain radiography’ (such as a chest radiograph) i.e. CR, or direct DR, at their full spatial and contrast resolution. Other images (e.g. CT, MR, US, NM, angiography or fluoroscopy) will be displayed at their maximum acquisition resolution on a 1 megapixel monitor (that is ~ 1200 × 1600 pixel monitor with a 19-inch diagonal). Digital mammograms, however, require a 5 megapixel monitor to be displayed in their entirety at full resolution.

A portion of a digital chest radiograph (or other plain radiograph) can be displayed at its full acquisition spatial resolution by use of a pan-zoom or a magnifying glass soft copy tool on a lower resolution monitor (such as a 1 megapixel monitor), provided the soft copy tool truly does show more pixels within its magnified area and does not just ‘stretch out’ the existing pixels already displayed on the 1 megapixel display (Fig. 1.7). When comparing plain radiography images (such as chest radiographs), it is essential to be able to see at least the current and a previous radiograph simultaneously. Only a dualmonitor workstation will enable two chest radiographs to be simultaneously compared at full size. A single monitor will enable two chest radiographs to be compared simultaneously, but they will inevitably be of reduced size, and this is not satisfactory in circumstances where the operator is trying to detect subtle sequential changes in pathology (such as resolving/worsening respiratory distress syndrome in the intensive care setting).

Ambient lighting The ambient lighting conditions in the area where the image monitor is placed need to be optimal for image viewing, and the importance of this cannot be overemphasized. This means that complete blackout should be possible, with dark (black or navy) blackout blinds fitted to the windows, and a variable level of low ambient lighting should be achievable with dimmer switches fitted to the room lighting. The workstation monitors should not be placed opposite a window, to avoid daylight reflecting from the screen. On the wards, the workstations should, if possible, be sited in a recess off the main ward so that independent dimming of the lights is possible in the recessed area. Flat-screen monitors are preferable to cathode ray tube screens since they do not reflect light, or show parallax (but do have to be viewed directly en face), and they take up less desktop space. Monitor resolution less than image resolution

True magnification

Pixel stretching

Figure 1.7 Diagrammatic illustration of true magnification. The monitor shows an image at less than its full resolution. The middle monitor illustrates a true magnification tool where more image data (more pixels) are displayed within the area of the magnification tool. The right-hand monitor shows a ‘magnification tool’ that merely expands the displayed data (stretches the pixels) and does not display more image data.

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It is clearly pointless installing high-quality monitors without optimizing the ambient lighting for image viewing. It is important to appreciate that high monitor luminance and good low-light viewing conditions are at least as important as (and some would argue, more important than) the absolute screen resolution of the monitor.

Location and type of PACS display devices The introduction of a PACS inevitably imposes some changes on the location where images are viewed around the hospital. One of the disadvantages of film was that it could only be viewed by one user, or group of users all in the same location, at any one time. PACS allows multilocation viewing simultaneously of the same image at the same time by numerous users all at different PACS display devices (dedicated PACS workstations and/or PACS web browser computers).

‘PACS workstations’ versus ‘web browsers’ The only difference between a so-called ‘PACS workstation’ and a PACS ‘web browser’ is their software. It has nothing to do with the resolution of the viewing monitors used, since any resolution of hardware monitor can be used to show any type of software package. Unfortunately the terminology has led to considerable confusion and misunderstanding because, apart from the software differences, there are often other associated differences, depending upon: • The manufacturer’s products • The type of hardware used for the PACS workstation and that used to run the web browser software • The network used to connect the PACS workstations and the PACS web browser to the on-line image storage device • The type and degree of compression (if any) applied to images displayed on the PACS workstation and on the PACS web browser. In general, the term ‘PACS workstation’ refers to a highspecification computer that runs proprietary sophisticated software dedicated to image diagnosis and review, whereas a ‘PACS web browser’ depends upon a public domain web browser (usually Internet Explorer) to display images for review purposes, and can be loaded on any off-the-shelf computer. Any computer running PACS web browser software can be used to review PACS images and thus could be called a ‘PACS workstation,’ but since it is an off-the-shelf computer it can also perform numerous other functions (e.g. making slides, writing documents, accessing the Internet etc.) depending upon which software packages have been loaded onto it, and to which networks it is connected. The PACS images viewed using a web browser are not necessarily inferior to those viewed on a dedicated PACS workstation. The resolution of the images depends upon: 1 The quality (resolution, size and luminance) of the monitor(s) fitted to the computer upon which the web browser software is running 2 Whether the images are compressed 3 The ambient lighting conditions where the monitor is viewed.

Thus the uncompressed images viewed on a PACS web browser could be of the highest diagnostic quality if a 3 megapixel (or even a 5 megapixel) monitor were fitted to the computer. An important, relatively new hardware development is the manufacture of dual-monitor computers (with the appropriate circuitry to run from one central processing unit (CPU), one graphics card and use one keyboard). If PACS web browser software were run on such a computer, then two chest radiographs from the same patient could be compared simultaneously, at large size, one on each monitor. The potential ‘drawback’ of a PACS web browser in some circumstances is the relative simplicity of its software compared with that found on a proprietary PACS workstation dedicated to image display for diagnostic or comparative review purposes. Again this issue is one of fitness for purpose, since in the majority of clinical settings such sophisticated software is unnecessary and never used, and the web browser software is entirely adequate for displaying the images (and reports) of a particular patient, and for comparing different studies on the same patient. Thus web browser software is entirely appropriate and sufficient for viewing patients’ images in the ward or clinic setting. The monitors chosen for the computer upon which the PACS web browser is to be run should also be fit for purpose, so that if it is deemed necessary that the images to be viewed in that location should be displayed at a resolution fit for primary diagnosis (i.e. the same resolution as would be seen by a radiologist reporting in the Imaging department), then very-high-resolution, highluminance, large (possibly dual) monitors will have to be chosen, and adjustable low ambient lighting and blackout facilities will have to be installed.

Requirements for sophisticated PACS software The indications for using the sophisticated (expensive) proprietary PACS workstation software (rather than the simpler web browser software) are as follows: 1 For the radiologist performing daily lists of primary diagnostic reporting and issuing definitive reports. The sophisticated software allows the radiologist seamlessly (by a single mouse click/key stroke) swiftly to work through a reporting worklist of imaging studies on patients whose names are not known. This software is particularly useful for reporting plain radiographs where there is a rapid throughput of work, and the radiologist should not be wasting time typing in the name of every patient whose images are to be reported. The current images of the next patient in the worklist to be reported are automatically displayed on the PACS workstation monitor, in conjunction with a configurable number of previous relevant imaging studies (of the same modality and body part) in sequential reverse chronological order. This automatic display of current and past images in the required format across the monitors is known as default display protocols (DDPs) or hanging protocols. When the report is complete, the software will mark the study dictated, thereby changing its status on the PACS. None of this software is necessary for the nonradiologist clinician, who is not involved in formal radiological report generation,

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does not have to work rapidly through a reporting worklist, and who is usually seeking the PACS images of a particular patient whose name and/or other demographic details are known to them. 2 For PACS database searching not based upon specific patient name or other demographic details. This is particularly useful to radiologists who often need to find a particular imaging study, or examples of certain imaging studies, which showed specific features, but they frequently do not know/ cannot remember the patient name or other demographic details. Currently web browser software generally does not allow searches of this nature, for example finding all cardiac MR examinations performed in January 2007. Nonradiological clinicians are almost always searching for images and reports on individual specific patients for clinical review, whose demographic data (be that name, date of birth, hospital number, or all of these) are known to them. These are exactly the data used by web browsers to search for specific imaging studies and their reports. 3 For any activity that benefits from the use of multipatient or single patient folders. Folders are useful for compiling and conducting MDTMs or other clinico-radiological conferences. Sophisticated PACS software (generally not available on web browsers) is required to add and remove patient studies or images to and from folders.Thus an MDTM folder of all the specific studies to be reviewed at a specific MDTM can be compiled electronically on PACS in advance of the meeting. This allows the radiologist who will be responsible for commenting upon the radiology in the MDTM to review the relevant imaging studies of the selected patients in advance of the meeting, with the benefit of the DDPs for automated display of the latest relevant preceding images. During the MDTM the conference folder can be progressively emptied of the studies already discussed, whilst some may be retained: for example, if they are to be discussed again the following week after additional imaging or other tests have been carried out, or if an addendum needs to be added to a report. Folders are useful for direct display from PACS (without having to make up a separate PowerPoint slide presentation) at hospital staff/grand rounds. This is especially true if the software allows anonymity with removal of the patient demographic details in the header bar across the top of the screen, so the folder can be displayed in a potentially public setting without breech of patient confidentiality. Multipatient folders are of benefit in the academic and research settings.They can be used to filter images directly from PACS into a particular group of images also stored and able to be displayed on PACS. Examples include images of patients involved in a particular research study, images selected for a book chapter or lecture, or teaching cases for specific tutorials. In summary therefore, sophisticated dedicated PACS software needs to be located in areas with the above requirements for use, i.e.: • The imaging department for diagnostic reporting • Seminar and conference rooms where teaching, MDTMs and staff rounds take place.

It is only in the imaging department (or any other area where final diagnostic image report generation takes place) that highresolution monitors need to be installed.

Location of PACS web browser computers Physicians’ ward rounds often follow a format of discussing all the patients first, with their case notes and X-rays, in a seminar room, followed by a ward visit of all the patients.This working practice can easily be adapted to PACS by holding the discussion phase of the ward round in a seminar room where there is access to a PACS web browser. Surgical, intensive care and physicians’ business ward rounds are often conducted at the bedside where the clinical charts are located, sometimes with the notes trolley being pushed round with the ward round, and any relevant X-rays being looked at ‘on the fly’, on the nearest viewing box (or, regrettably, by holding the radiograph up to the window!). Unless PACS viewing facilities are provided at the bedside, these ambulant ward rounds will have to detour to view the patients’ images on the nearest web browser as each patient is visited. Citing the ward-based PACS web browser computers is important: they should be easily accessible from the bedside, but in an alcove or area of the room where there are blackout facilities, lighting that can be dimmed, and no reflected light. Bedside viewing of the PACS images is possible, but is generally an expensive option. There are battery-powered mobile PACS trolleys/carts, rechargeable from mains electricity. They need a long battery life to last the entire ward round, and pose a security problem since theft of a mobile computer is much easier than a standard, fixed computer. One option for bedside viewing of PACS images is to make them available on the TV monitor located at each patient’s bedside for watching regular television (usually on a fee-paying basis). However this may lead to patient complaints if the ward round timing happens to coincide with a patient’s favourite TV programme, especially when the patient is paying for the privilege of watching it at the bedside! Hand-held personal digital assistants (PDAs) is another potential solution. Such wireless devices can be used to view a patient’s images at the bedside as a miniature prompt reminder of the imaging study in question. Obviously this assumes that the hospital is running a wireless network (unusual in the UK) and PDAs are considered far too expensive to be provided by the NHS! In each operating theatre web browser access to PACS is necessary. In theatres, use of wall- or ceiling-mounted computer monitors has been advocated by some, whilst others favour mobile carts. Wall- or ceiling-mounted monitors are more cumbersome to adjust and repair if they malfunction, and mobile cart computers raise issues of possible transfer of infection from one theatre to another if moved between theatres. The most important consideration is to ensure acceptable viewing monitor(s) (19 in) are installed, with a relatively long logout time and no screensaver, so that the images remain displayed for a sufficiently long time period and do not keep ‘disappearing’ during the surgical operation. Images viewed in theatre will already (hopefully!) have been studied in detail elsewhere by the operating surgeon, and

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their display in theatre is as an ‘aide-mémoire’ for review purposes only. Clearly the bright ambient lighting necessary in the operating theatre will render the image viewing conditions suboptimal anyway, so the monitors chosen for the PACS web browser computer do not have to be of particularly high specification. In the outpatient clinic access to a PACS web browser will be necessary in each clinic room for the majority of clinics. Since outpatient clinic rooms are often small, desktop integration of the PACS with the hospital information system/

patient archive system (HIS/PAS) on one computer (preferably using the same login for each system) is essential to allow sufficient desktop space. This is an additional reason for seamless integration between hospital IT systems. In addition there is the obvious need to be able to click directly from the imaging study to its report, and vice versa. Ideally it should be possible to view the patient’s other electronic data (e.g. blood test results) with the same user login, on the same computer, and without having to launch another software package from scratch.

CONFERENCE ROOM DESIGN Good overhead display of PACS images is essential in rooms used for MDTMs, staff rounds, teaching and other presentations.The overhead displays need not be of high resolution since they are for demonstration purposes only. The PACS workstation monitor(s) should be slaved to overhead projector(s) for projection on the wall, so that the whole audience can see the images well and are involved in the conference. It is not satisfactory (even with a small group) to attempt to run conferences from the PACS monitors themselves since visibility, even in the second row, is not good enough. Colour landscape-style monitors (maximum dimension in the horizontal plane) are preferable to portrait-style monitors (maximum dimension in the vertical plane) for linking to an overhead projector, since the latter were developed for the film industry and their technology is designed for colour landscape projection. Such projectors only synchronize well to the refresh rates of landscape-style monitors. Although a PACS web browser can be used to run conferences, more sophisticated PACS software is highly preferable for the reasons cited earlier. For example, PACS web browsers generally do not permit the use of software such as conference folders into which the examinations to be reviewed at the conference can be placed in advance of the conference. When

a PACS web browser is used, every patient to be reviewed has to be called up manually in real time by typing in their demographic details. This is time consuming, and does not allow easy pre-conference review of the images by the radiologist. The PACS requirements for a hospital staff round or teaching presentation differ from those of an MDTM since staff rounds and teaching merely require demonstration of the salient features in an image and, unlike MDTMs, have no diagnostic review purpose. Thus a single PACS monitor is adequate and can be of low specification. The PACS software permitting the use of PACS folders is, however, highly desirable for staff round and teaching presentations. Conversely, in an MDTM the radiologist should have two PACS workstation monitors (even if only one monitor is projected for the audience to see) to allow comparison of full-size current and previous chest radiographs simultaneously, or to compare and contrast different imaging modality studies simultaneously on the two monitors. Obviously the considerations of good blackout facilities and room lighting that can be dimmed also apply to the conference room. The relatively low luminance of the projected overhead image demands almost full blackout, and certainly there must at least be the facility fully to switch off the lights at the front of the room nearest the overhead screen(s).

GRAPHIC USER INTERFACE Whatever the specification of the PACS workstation, it is to be expected that its graphic user interface (GUI) should be intuitive and user friendly. The GUI is the screen layout and computer tools presented to the user.The ability to manipulate and post-process soft-copy images on monitors represents a major change in working practice for radiologists and other clinicians used to working with hard-copy film and now migrating to working with soft-copy images on PACS monitors20.

SOFT-COPY TOOLS All PACS vendors now offer a variety of soft-copy tools for image display and manipulation. User preference for the tools varies and many users now like the option to be able to configure their own customized toolbar/tool palette for their most frequently used soft-copy tools, some of which will be modality dependent.

Contrast width and level The contrast window width and level tool is the most frequently used soft-copy tool, and is often allocated to one of the mouse buttons for easy accessibility. Alteration of the contrast width and level allows soft tissue and bony structures to be well seen on a single exposure. For example, it often permits the left lower lobe to be assessed behind the left cardiac silhouette, whereas this information is not available on a relatively underexposed hard-copy chest radiograph. There is thus an increase in the amount of information that can be extracted from an image, and this is particularly noticeable for plain radiography. In addition to the contrast tool for continuous manipulation of greyscale width and level, there is generally a facility to enter personalized contrast width and level presets so that these can be rapidly applied to displayed images. This is particularly useful for instantly flipping from soft tissue to lung to

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bone windows on CT images (whereas previously different window settings had to be printed onto film by the radiographers).

CT density values A pixel tool will measure Hounsfield unit values on CT images. A region of interest (ROI) tool will give a mean density value.

Greyscale inversion A greyscale inversion tool causes video inversion of the image from the more conventional white-on-black presentation to black on white and vice versa. Many users prefer to use the black-on-white presentation for viewing barium work, angiography, hepatobiliary fluoroscopy and for identifying the tips of indwelling lines and catheters on plain radiography.

Stack/cine mode The ability to demonstrate images in stack-mode display leads quickly to a change in practice when viewing cross-sectional images (and angiographic images) on PACS monitors21. The stacking facility allows images to be demonstrated either on an automatic cine (useful in cardiac MR, dynamic US or MR contrast studies), or in a manual-scroll mode whereby the user can scroll up and down through a stack of images at a speed directly controlled by the speed of mouse movement (Fig. 1.8). All PACS vendors now provide software to link image stacks so that they run synchronously after matching for corresponding anatomical level. This is particularly useful for comparing new and old studies in the follow up of oncology patients, and for examining a chest CT simultaneously on lung and mediastinal window settings. An additional useful GUI tool is the ability to cross-link series in different anatomical planes (or scout views). This is mostly used in MR simultaneously to pinpoint a particular structure in, say, the transverse, coronal and sagittal planes. It is the exploitation of stack-mode display on a PACS monitor that allows diagnostic workstations with only two monitors (or even one monitor) still to be suitable for reporting cross-sectional imaging modalities. This viewing mode is so useful that it is desirable to have stack-mode display available

on the web browser review stations as well as on the diagnostic workstations.

Image orientation Image rotation and a left–right inversion tool will flip the orientation of an image (although if the radiographers have performed their quality assurance correctly, this should very rarely be necessary)22. It can also be of use in CT or MR procedures in which the patient has been examined in a lateral or prone position.

Magnification Magnifying glass functions include both the ability to magnify a segment of the image (preferably with a variable size to this magnified segment and a continuously variable degree of magnification within it) and the provision of a pan zoom tool that allows magnification of the entire image with a roam facility. Not all PACS vendors provide both tools, but each does have its individual uses. The magnification glass tool enables a segment of the image to be magnified, whilst retaining an overall view of the entire image, at its original size. The pan zoom tool enlarges the entire image, and is most useful for enlarging extremities (digits) and small paediatric images. If the magnification tool displays the original data at full resolution (Fig. 1.7), its presence can obviate the need to have 3 megapixel display monitors on workstations used for occasional diagnostic plain radiographic reporting, which greatly reduces the cost of the diagnostic workstations.

Annotation Annotation tools for teaching and labelling purposes include text, a scaleable arrow, the ability to paste measurement and density value data directly onto the image (and store it there as part of the overlay), and a freehand draw tool that is a useful adjunct for real-time teaching purposes. A closed-loop freehand draw tool will provide a cross-sectional area measurement of an irregular shape and is particularly useful for drawing round the tumour mass on radiotherapy planning scans.

Significant images It is highly desirable to have a means of designating selected images in an examination as ‘significant’ (‘key’ or ‘flagged’). The recall and display and/or printing of only these selected images can be useful under certain circumstances, such as for: • Nonspecialist clinicians • For demonstration purposes • As an aide-mémoire when referring to a historical examination.

Measurement Ti l e

Sta c k / cine

Figure 1.8 Diagram illustrating ‘tile’-mode display (left) and ‘stack-’ or ‘cine’-mode display (right). (Reproduced from Strickland N H, Allison D J 1995. Default display arrangements of images on PAC monitors. Br J Radiol 68:252–260 with permission).

Measurement functions include tools for making linear, area and angle measurements. There is also a requirement for certain computed measurements for particular specialties. In orthopedics the scoliotic curve is often assessed using the Lippman-Cobb angle. Dedicated orthopedic templating software packages are available to facilitate the planning

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required to predict the requisite joint prosthesis to be inserted at surgery.

Format The image format changes the number (upcount) of tile images displayed per monitor, and determines whether the examination will be displayed in tile or stack mode.

Linked cursors A linked cursor tool (dual user arrows) allows interactive conferencing between two distant users on different PACS workstations, communicating either by telephone, audio conferencing or by typing into dialogue boxes on their workstations. The two arrow cursors have a physically distinct appearance and one is assigned to each of the two workstations conferencing with one another. This feature is particularly important in the teleconference setting.

White masking/shuttering It is desirable to have a means of eliminating the glare from the white nonexposed area surrounding images acquired on phosphor plates or by fluoroscopic devices. This can currently be achieved by manual shuttering with a black frame, in some cases at the review workstation of the acquisition device, in others on the PACS workstation itself. This is a suboptimal solution because it is time consuming and it generally only allows rectangular shuttering frames to be applied. Such frames are unhelpful when an image, such as of the tibia and fibula, has been acquired in a diagonal orientation across the imaging plate, or when multiple exposures have been made on the same plate. The process ought to be achieved electronically such that the edge of the exposed area is automatically recognized and the white surround masked.

Save option The ‘save’ command should be a password-based user privilege. It is undesirable to have images saved on the PACS with inap-

propriate settings and therefore the save privilege is often made available only to radiologists and radiographers, who, with their specialist knowledge of how an image should appear, ought to be sufficiently responsible to leave an image so that it looks acceptable for viewing by the next PACS user. The save privilege may be made available to other users only in their private folders. An alternative approach is to have the image reappear as last saved by that specific user according to log in, and to default to how it was saved by the radiographer who first verified its appearance on PACS if not previously accessed by that user. It is sometimes useful to be able to display the original version of the image, as it was acquired onto PACS before any soft-copy manipulations were performed on it by other users. The problem with having this version of the image appear as the default is that frequently the plain radiography images will have required subsequent adjustment for orientation by the radiographer prior to verification on PACS, and will thus appear the wrong way up or inverted.

Some other desirable features An ‘undo’ command is useful for reversing the last configurable number of commands (as in a normal word processing package). It is useful to have a pull-down menu allowing easy access to the last viewed examinations so that the user can navigate back to them quickly and easily. It is important to be able to display the clinical details or report associated with any active examination by a one-step procedure. An ‘abort’ button is an extremely useful feature if a lengthy search or command has been instituted in error. The display of a progress bar in addition to a time cursor on the screen is desirable psychologically for the user, and indicates how far the task is towards completion. It also allows the user to distinguish whether the workstation has in fact ‘frozen/crashed/hung’ or whether the command is merely taking a long time to implement.

SOFTWARE CONCEPTS Sophisticated PACS software is essential to the radiologist (and to a lesser extent the nonradiological clinician in certain settings, as discussed in the section on workstations) to allow the PACS to be used efficiently in the clinical environment. It is the PACS software that is the discriminatory feature determining whether or not hospital-wide functioning using a PACS will be sustainable in a busy clinical environment23. Four main software concepts are considered in this section: 1 Prefetching 2 Worklists and folders 3 Default display protocols (DDPs) or hanging protocols 4 Reporting/dictation macros or worklists.

PREFETCHING Prefetching is the process whereby previous images on a patient are automatically retrieved from the long-term archive

onto the short-term storage device, either for review when the patient is due to attend the hospital as an in- or outpatient, or prior to the acquisition and viewing of a new imaging examination on that patient. Now that the cost of RAID storage has reduced, local ‘on line’ storage devices are often huge, with effectively all imaging studies kept there, immediately accessible to PACS users. Under such circumstances there is no call for prefetching of any imaging studies from a long-term archive. However, where numerous PACS facilities are sending their imaging studies to shared long-term archives (as in the English Connecting for Health countrywide PACS project, where there are five large long-term PACS cluster stores serving the whole country), the software concept of prefetching is again becoming a very important issue. Prefetching is commonly sufficiently sophisticated to take the form of an ‘intelligent’ prefetch. This means that only a

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be accessed in different ways (e.g. by body part or by referring clinician) according to the clinical setting. Depending upon the nature of the workflow in a particular institution the precise requirements for various worklists and folders will differ, and the pertinent software needs to be tailored and made site specific. One example is reporting worklists: some sites may prefer to divide up their plain radiographic work into body part specific worklists to be reported by specialized radiologists – chest radiographs, skeletal, skull and spine etc.; others may have radiologists who are happy to work from one general plain radiography reporting worklist. Direct integration of both HIS and RIS to PACS means that any ‘tag’ (or header field) in these systems can be used to construct a customized worklist or folder on PACS. (If only the RIS is directly interfaced to PACS then HIS fields cannot be used to construct PACS worklists or folders).

configurable number of past examinations, from the same modality and/or the same body part as that currently being imaged or reviewed, are retrieved from the long-term archive. It is usually only examinations of the same body part (and often of the same technique) that will be of relevance in making a diagnostic comparison, and there is no point in unnecessarily overloading the network and PACS short-term storage device with data irrelevant to the current clinical problem. Prefetching is essential to reduce the number of ad hoc fetches of examinations to a minimum and is usually scheduled to take place during the night before the next booked patient attendance or imaging study, since traffic on the network is slackest during the night.This is important because retrieval of examinations from the long-term storage (whatever its nature) takes much longer than the almost instantaneous display from short-term storage, and therefore needs to be kept to a minimum during clinical practice. Current hardware enables prefetched images to appear on the workstation monitor in <3 s, this being the time between the execution of the command to display an examination and the appearance of the first image of that examination on the monitor. Any subsequent re-display of this examination will be even quicker since it will already be in the random access memory (RAM) of the individual workstation. This contrasts markedly with the time taken to fetch an examination from the long-term archive, which is generally 2–3 min if that archive is local or even much longer (5–10 min) if it is far away on a wide-area network. Prefetching algorithms rely upon connections with the HIS and RIS to trigger prefetches for inpatients, outpatients, patients booked to have radiological examinations without a hospital attendance (referred from general practitioners), and patients due to be reviewed at MDTMs24. All these events (except the MDTM patient list) are known to the HIS and/or RIS, since this is where the booking for such events takes place, and this information needs to be transmitted to the PACS to trigger the prefetch. This is one of the key reasons why integration between the HIS, RIS and PACS databases is so crucial. If the HIS and RIS are governed by a master patient index, then it is this master patient index that communicates with the PACS database. It can be useful to have the PACS itself able to trigger prefetching; an example of this requirement is when patients have to be reviewed at a MDTM. Many such patients may not be current in- or outpatients and may not have had recent imaging examinations performed in the hospital, so there will be no trigger outside the PACS for the studies relating to these patients to be prefetched prior to the conference.

These are software algorithms for ordering and displaying the images of current and a configurable number of past related examinations in a predetermined arrangement on the workstation monitors, so that they appear as the radiologist and/or other clinician wishes to view them25–27. DDPs display only previous imaging studies of the same body part (and usually the same modality) for comparison with the current study, and they display these studies in the correct chronological order by default. All these features save the user time since the images do not have to be rearranged on the monitors prior to reporting or viewing. DDPs can now frequently be made user specific (based on login) and user configurable, and may be configured to suit different modalities or different body parts. Most PACS software allows specialized DDPs to be constructed, for example for cardiac MR, and for 3D reconstructions associated with cross-sectional imaging.The introduction of multi-slice spiral CT, large MR angiograms and dynamic MR contrast studies (e.g. in the breast) produce vast numbers of images at acquisition. Some centres consider it more appropriate to store these images on a local archive directly linked to the acquisition device, and only to send definitively to the PACS some selected reconstructed images for review. Some PACS vendors sell workstations with a third monitor, which is a colour, relatively low-resolution monitor (approximately 1 megapixel) linked to their high resolution (3 megapixel, greyscale) diagnostic monitors, on which a post-processing package, such as multiplanar reformatting, can be run simultaneously for cross-sectional imaging studies displayed on the diagnostic PACS monitors.

Worklists and folders

Reporting/dictation macros/worklists

These software concepts are a means of subdividing the vast amount of imaging data acquired onto the PACS into clinically relevant, easily accessible subgroups. Although ‘worklists’ and ‘folders’ are generally identical from the engineering viewpoint, worklists usually refer to ‘work that has to be done’ (such as lists of patients’ examinations that are to be acquired or reported), whereas folders are designated groupings of examinations that have generally been reported but need to

These terms refer to software mechanisms that provide the reporting radiologist with a means of working through a reporting worklist (equivalent to the old ‘reporting pile’ with film) in a manner that is efficient and thus time saving. This is achieved by minimizing the number of commands necessary to accomplish various necessary steps. For example, a single keystroke or mouse click might be expected to display the next examination in the reporting worklist with a predeter-

Default display protocols or hanging protocols

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mined number of previous related examinations, generally of the same body part and modality, in the correct chronological order and format according to the appropriate (or chosen) DDP. The same single keystroke or mouse click should also save the changes made to the displayed examination, if any, after dictation, and mark that examination reported when closed, by changing its status and thereby removing it from the worklist. The reporting macro will then automatically display the next unreported examination in the worklist, with its relevant previous examinations for comparison and so forth. Automated reporting/dictation macros or worklists are essential to the radiologist for rapid, efficient plain radiographic reporting work. This software can only work if the pertinent clinical details associated with the imaging study request appear automatically in electronic format with the imaging study displayed on PACS. Ideally this should be achieved by the integration of remote electronic requesting (order communications) with the PACS and the RIS; however a less satisfactory work-around can be achieved either by typing the clinical requesting details onto PACS or RIS, or by scanning in the paper request form. A highly desirable feature of some reporting macros is that the software automatically skips displaying the next exam in the reporting worklist if it has already been opened in reporting macro mode by another radiologist, working at a different workstation. This allows numerous radiologists to work from the same reporting worklist, at physically distant locations within the same or different institutions (as in a virtual hospital environment), and protects them from simultaneous dictation of the same exam. A useful (but rarely present) refinement of such reporting macros is to have an intermediate status assigned to examina-

tions if they have been reported by a junior radiologist and require checking at a later time. Such an intermediate status will temporarily protect these examinations from being subsequently reported by a more senior radiologist, and will allow their easy retrieval for checking after dictation by the junior.

Proprietary software It is important to realize that there is no standard for PACS software analogous to the DICOM standard for PACS and imaging hardware. All PACS software is thus proprietary, and vendor specific. This being said, one can ‘mix and match’ the hardware from different vendors when constructing a PACS, but a PACS workstation from one particular vendor will only display images according to the software programs that it is running. Some progress in this area has been made due to Clinical Context Object Workgroup (CCOW), which is a HL7 standard protocol designed to enable disparate applications to synchronize in real-time, and at the user-interface level. It is vendor independent and allows applications to present information at the desktop in a unified way. PACS databases are also not covered by the DICOM standard and so the only way for a workstation made by one particular vendor to obtain an examination from another vendor’s database is for that workstation to send a query/ retrieve DICOM format message to the database supplied by the different vendor. Such query/retrieve DICOM messages, although effective in being able to accomplish the task of displaying the desired images, will inevitably be slower than a proprietary query/retrieve between a workstation and a PACS database supplied by the same vendor. At present there is, unfortunately, no prospect of a universal software standard being developed.

COMPRESSION There are two major categories of image data compression28: lossless (reversible) and lossy (irreversible).

to be employed without the irreversible data loss being visibly detectable in the image.

LOSSLESS

NEED FOR COMPRESSION

Lossless compression techniques can currently achieve a maximum compression ratio of 3:1. As the name implies, lossless compression reversibly compresses the data so that no data are permanently lost from the image, which is restored to its original state when decompressed.

The main requirement for image data compression arises in connection with teleradiology. Compressed (reduced amounts of ) data can be transmitted more quickly, and therefore more cheaply, over long distance transmission lines. There is currently no British legislation regarding whether or not radiological images may be irreversibly compressed and used for diagnosis or archived in a compressed form. The British Royal College of Radiologists’12 recommend that if a definitive diagnostic report is to be made on an image then it should be viewed in an uncompressed form, although it may be subsequently archived in a lossily compressed form. Thus only lossless compression would be acceptable for an image that is to be definitively diagnostically reported on the receiving workstation. It has been argued in some countries (e.g. Germany) that only the ‘original’ (uncompressed) image should be archived. This argument is spurious, however, in that there is no clear definition of what constitutes the

LOSSY Lossy compression techniques irreversibly compress the image and data are permanently lost, although the aim is that this should not be detectable to the observer when the image is subsequently viewed29. Several types of compression algorithm exist, for example joint photographic expert group (JPEG) compression, and wavelet compression. The latter is currently employed by many companies in their web browser technology since it permits higher compression ratios (of up to 60:1)

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‘original’ image. Even the images as first acquired, be they CT or on phosphor imaging plate, have already been modified in some way, by the application of various image-processing algorithms. Now that long-term archiving media are becoming cheaper and of higher density (i.e. greater capacity) the need irreversibly to compress imaging data on a local area network (LAN) has diminished. The lossless compression of images being transferred from a short-term storage device to diagnostic workstations is of benefit, however, since even a 3:1 lossless compression ratio will noticeably increase the speed of transmission and thereby reduce the lag time between requesting the display of an image and seeing it appear on the PACS workstation monitors. Lossy compression of images transmitted to PC review workstations significantly reduces their transmission time30.

Many web browsers operate intelligent ‘streaming’ of data transmission such that reversibly compressed data are transmitted and the degree of decompression is tailored to fit the maximum display resolution of the monitors of each workstation/computer, thereby saving on unnecessary decompression time overhead, allowing the images to be displayed faster on the monitors. Progressive image decompression of all the images in the study takes place in the background whilst the first image(s) are displayed for viewing. A display time of 3 s between the keyboard display command and the appearance of the first image of the study on the display monitor is required for effective clinical workflow. In the diagnostic teleradiology setting it is useful to be able to view the data initially in a lossily compressed form (so that it arrives quickly) and subsequently in an uncompressed form to check that nothing that may be of diagnostic relevance has been missed.

DIGITIZATION POLICY Digitization is a laborious and time-consuming process, and needs to be performed by someone who is able to recognize the difference between the various imaging techniques (such as distinguishing CT from MR) and different body parts (for example someone who is able to differentiate films belonging to a barium meal, a small bowel enema and a barium enema). It therefore needs to be performed by personnel with some radiographic training, generally radiographers, and not filing clerks. Every study needs to have an entry sent from the RIS to the PACS prior to digitization; and needs to have the report, which resides in the RIS, made visible via the PACS. Digitized film will always be inferior to the original film, or to direct digital acquisition, however good the quality of the digitizer. Institutions differ in their approach to digitization of historical hard copy films when PACS is installed. It would be a total waste of resources to digitize large numbers of historical films, especially since so few historical films are ever referred to again. Many institutions digitize no previous films, but continue to send film packets to the outpatient clinics, and keep their film store on site, for a period of at least 3 months, and undertake rapidly to retrieve any requested hard copy for a further more prolonged period of, say, 2 years, when the film store is moved to an off-site storage facility. Some PACS sites (for example, Hammersmith Hospital in 1995/1996) allowed nonradiological clinicians to select films from a single previous imaging study for digitization prior to PACS ‘going live’. Such a film(s), deemed crucial for future reference, can be selected by the clinician from a particular patient’s film packet prior to refilling, after the patient has attended the outpatient clinic, or been discharged from hospital. (The examination for digitization was placed in a temporary plastic outer film jacket that also contained the patient’s normal cardboard film packet.) An additional policy, which is very effective, is to start building up a soft-copy PACS archive of digital images (from CR and DR, as well as the more traditional digital modalities of CT, MR etc.), some time before the institution withdraws film and goes live with PACS.This means that there will already be a substantial number of historical digital images on the PACS

for comparison when the institution becomes filmless. Clearly it is expensive to run this overlap period for too long, since any cost benefits of PACS can only start to be realized when film is withdrawn, and a period of approximately 3 months would be a reasonable compromise. Modalities that are already acquiring digital images (such as CT, MR many US and NM machines etc.) can easily send any configurable number of historical studies to PACS from the local modality storage medium. It is important that the reports of these examinations should also accompany the images and be visible on PACS. Every PACS institution will need to have the facility to digitize film and to upload CDs of digital images (and preferably also of their reports) onto PACS in order to deal with the images of patients referred to the hospital from elsewhere, so that these historical images can be easily viewed in conjunction with the subsequently acquired imaging studies performed on the patient when under the care of the PACS hospital. This requirement will clearly be greater for tertiary referral centres than for district general hospitals. In England the amount of digitization/uploading of digital images should reduce as the whole of England’s hospitals have a PACS installed and are able to send their images to the central cluster stores. It will always be necessary for the institution to have a means of burning imaging studies, with their reports, onto a CD for sending with patients returning from referral from abroad (outside England). This CD burner should automatically burn the images onto the CD in DICOM format, with a self-launching DICOM viewer, and also automatically burn the report associated with the imaging study. PACS vendors generally sell their own CD burners, but it is possible to use cheaper freeware versions available on the Internet, such as that supplied by efilm. In addition it is useful to be able to read CDs brought by the patient in the outpatient clinic or in the MDT room. Thus the computers in clinic rooms etc. (those also running the PACS web browser) should have a CD drive and should have sufficient RAM to be able to run a large cross-sectional imaging study in stack mode. If the CD files are copied

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to the hard drive of the computer the imaging studies will obviously run faster than trying to play them directly from the CD. After reviewing the images on the CD it may be decided that they are not of relevance to the current problem and therefore do not require uploading to the institution’s PACS. Such CDs containing images from a referring hospital should have been burnt in DICOM format with a DICOM viewer that self-launches when the CD is opened.

Unfortunately this is not always the case: sometimes the DICOM format of the images is incomplete, sometimes no DICOM viewer is included on the CD, and sometimes the images may be sent in JPEG, or other, format31. All these problems can make the CD unreadable on a normal PC, and thus it is advisable to have DICOM and JPEG reader software programs installed on these PCs to attempt to read such CDs.

TELERADIOLOGY Teleradiology is radiology at a distance (from the Greek, tele, meaning distance). It has more stringent requirements than other types of telemedicine because of the vast amount of data in radiological images and, if a definitive primary diagnosis is to be made at the receiving site, the requirement is to see this data at its original resolution. Teleradiology has many forms, but the most commonly used types of teleradiology are: 1 Review of images from home when on call 2 Teleradiology linkage between two or more hospitals for joint MDTs 3 Out-sourcing of imaging examinations by understaffed hospitals for reporting elsewhere by a private company 4 Tele-education whereby a digital image teaching library is accessed remotely by users. Again, the key here is that the hardware and software should be ‘fit for purpose’.

REVIEW OF IMAGES FROM HOME This is to allow the on-call radiologist or physician/surgeon to review plain radiographs (such as the intensive care chest Xray) or emergency CT to decide upon immediate medical or surgical management. It is not intended to provide the quality of image upon which the final definitive diagnostic report will be issued. These types of management issues generally depend upon observing fairly gross pathology, such as lobar lung collapse, free intraperitoneal gas, or intracranial haemorrhage.Veryhigh-resolution monitors are therefore not required, and a single 17 in PC monitor will suffice. It is, however, important to have a timely transfer of images (it would be medically negligent to wait for more than 2 or 3 min to receive a CT study showing a dissection of the aorta, for example), and for this a reliable broadband connection and streamed image compression (tailored to the resolution of the receiving PC monitor) will be necessary. Use of a virtual private network (VPN) minimizes the amount of data encryption necessary, which reduces the time overhead required for the encryption and decryption process.

TELERADIOLOGY LINKAGE BETWEEN TWO OR MORE HOSPITALS FOR JOINT MDTMS The complexity of the hardware requirements will vary depending upon the nature of the teleconference, for example, one site may always be the dominant site transmitting to vari-

ous satellite sites, or bidirectional links may be required so that either site can send or receive data. If one site wishes to send images whilst the other simultaneously transmits microscopy data for example, then two independent lines of transmission will be needed. Most clinicians prefer to use their screens for receiving as much medical data as possible, rather than reserving one for a photographic image of their colleagues in the other location(s), which is a concept video conferencing companies, used to installing equipment for business teleconferencing, often have difficulty in understanding. Large overhead projection onto a wall screen is preferable to smaller plasma screens that are not well seen from the back of a conference room. Some PACS software products include linked or sitespecific cursors to be used simultaneously on images during a teleconference. If a company different from the PACS vendor is used to install the teleradiology conferencing equipment, then the same stringent contract requirements for equipment service, maintenance (including a helpline facility and dial in troubleshooting), uptime and updates/upgrades should be incorporated into the contract signed as is used for the PACS project itself or for any medical IT project.

OUT-SOURCING OF IMAGING EXAMINATIONS FOR REPORTING This service constitutes diagnostic primary reporting remote from the site where the imaging examinations were acquired.This demands exactly the same hardware and software requirements (including access to the patients’ previous imaging examinations with their reports) as if the diagnostic reporting were being carried out in the primary hospital contracting out the service.

TELE-EDUCATION The hardware requirements for remote access of a digital image teaching library are minimal. The user’s home PC will be adequate. Software for such teaching libraries will be of variable sophistication. The better teaching modules will allow the user to perform basic manipulation of the images, including altering contrast width and level, magnification, viewing in stack mode. Flexible use of the teaching material is desirable, with the possibility for the user to review cases on one particular topic, or to work through a random selection of cases in ‘exam mode’, and to use bookmarking in order to be able to return to the same place in the teaching programme after an interruption, and so forth.

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SECURITY Any PACS and teleradiology installation must incorporate security measures to protect patient confidentiality32. These operate at several levels in a PACS: • User-specific log in and password • Monitor screen savers • Workstation time-outs • Encryption • Audit trails • Firewalls.

USER-SPECIFIC LOG-IN AND PASSWORD Users of the PACS should be registered with the system administrator and provided with a personal log-in and password. It is becoming more common for passwords to have to contain special characters and numbers as well as letters. Some institutions implement an enforced password change every few months in order to maintain an updated list of currently approved users.

MONITOR SCREEN SAVERS Screen savers on PACS workstations guard against the risk of a patient’s data being inadvertently left displayed on an unattended workstation.

AUDIT TRAILS A PACS, like any other multi-user computerized system, should have an audit trail mechanism that logs to a specific user all actions they carry out on the system.This audit trail can be used to trace any inappropriate data viewing, or any action that could potentially pose a risk to the security of the system and the integrity of its data. Such actions on a PACS include the deletion of images and the merger of patients. There are legitimate reasons for occasionally wishing to carry out such procedures, usually to correct human error, for example, the image(s) of one patient may erroneously have been acquired into the wrong patient’s folder; the same patient may inadvertently be entered onto the PACS twice with different identification numbers or different name spellings.The privilege to make such changes on the PACS is generally restricted to the system administrator(s).

FIREWALLS A firewall is a configuration that protects a computer system from unauthorized access from outside the institution. It only allows access from specified Internet protocol (IP) addresses and ports using certain protocols. It will protect a PACS from access by unauthorized users.

TELERADIOLOGY SECURITY WORKSTATION TIME-OUTS An automatic time-out feature for every workstation minimizes the risk of workstations in use being left unattended after a user has logged in. Different time-out settings can be applied to the various workstations throughout the hospital based upon their intensity of clinical use.

ENCRYPTION Encryption is not generally deemed necessary for an intranetbased PACS, but is essential when teleradiology, with data transmission over public networks, is being employed. There are numerous different types and levels of encryptions. Commonly encryption involves private and public encryption keys (codes) to encrypt the data and to access it once encrypted.

When patient data are transmitted over public networks, as with teleradiology, security measures must be undertaken to ensure confidentiality. Thus an encryption program must be part of any teleradiology installation to codify the textural data. Generally the images per se cannot be recognized as patient specific and therefore image data need not be encrypted, unless there is identifiable textural patient information burnt into the image itself, as occurs with screen-captured images or images digitized from hard-copy film. Many would argue that it is safer and simpler to encrypt both image and textural data sent over a public network. Alternatively a VPN can be used for teleradiology instead, obviating the need for encryption and thus saving on the time overhead taken for encryption and decryption of large volumes of data. A VPN is accessed by designated users via a randomly generated code on a digital keypad.This code changes for every log-on session, and allows the user to be connected to the network.

PACS TRAINING PACS training is part of the PACS installation and is carried out in the first instance by the PACS vendor, as will be specified in the PACS contract. PACS training (as for any IT installation) is provided on a cascade basis, so that the vendors’ applications specialists train key personnel, who are then responsible for cascading that training to other members of staff. It is important to keep a formal permanent record of

the names of these ‘core trainers’ who attended the training sessions, so that they can later be held responsible for training other members of their teams. One of the challenges in any hospital is to disseminate information to its staff. The times and locations of the training sessions need to be well publicized, with sufficient choice of times to allow the maximum number of users to attend. Since

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attendance at such training sessions is often given a low priority by busy clinicians, it can be very valuable to schedule the training sessions just prior to the PACS go-live date (or PACS upgrade). It is also very good for applications specialists to be present in key locations, such as the outpatient clinics and the A & E department, immediately after the PACS (or major PACS upgrade) is introduced, to help users in the clinical setting when they have no choice but to use the PACS. On-line computerbased training and help menus are useful adjuncts to training. Basic PACS training and information on how to obtain a PACS password must be included in new hospital doctors’ induction day instructions and be available on the hospital intranet.

The training required by a nonradiological clinician to use the PACS is relatively modest: merely to learn how to call up the required imaging study, display it with any relevant historical imaging study, manipulate the image with respect to simple soft-copy tools including changing the contrast settings and magnification, and display the report. Most clinicians also like to be able to display large cross-sectional imaging studies (CT and MR) in stack mode and to scroll through them manually. Clinicians also need to be taught how to add imaging studies to folders so they can populate the MDTM folders with the imaging studies they wish to have reviewed in advance of the meetings.

QUALITY ASSURANCE A thorough programme of quality assurance must be instigated to maintain a PACS. This will generally be undertaken by in-house specialists, such as radiation physicists and IT personnel, as well as input from the PACS vendor as part of the maintenance contract. Technical support via a modem connection is not sufficient. The tasks involved in a quality assurance programme for a PACS hospital differ substantially from those previously required in a conventional-film-based hospital.

PLAIN RADIOGRAPHY The acquisition of plain radiographic images in digital format so they can be transferred to the PACS still constitutes the major workload in the imaging departments of most general hospitals (see ‘Implementation of Digital Image Acquisition Prior to PACS’). The quality assurance of these images, which are still largely acquired by CR, is an important task integral to the PACS project. Quality assurance is a topic that, unfortunately, has been largely ignored by PACS and CR manufacturers. The radiographic technique required for CR and direct DR differs in a number of respects from the procedure required with conventional film, for example, the correct body part algorithm must be applied to the image. In order effectively to run a quality control program it is essential that ‘rejected’ unsatisfactory images should not be lost from the system. This can be avoided if a mechanism is introduced whereby such inadequate images are sent to PACS by default and then, by ‘rejection’ on the PACS, made invisible to the clinical user, but visible to the physicist who is performing the quality control analysis. The same mandatory send to PACS of all unsupervised fluoroscopic and angiographic images acquired by junior radiologists in training allows their work to be assessed for quality assurance purposes. It is helpful to make certain fields on the CR device mandatory, including: • Entry of the radiographer’s identification, so that quality control issues can be discussed with the appropriate person • A record of the exposure factors (mAs and kVp) used on each image, since the wide latitude of CR and DR will ‘salvage’ an overexposed image that would have looked

black on film, but will appear acceptable on digital imaging after appropriate contrast adjustment, resulting in even gross overexposure passing undetected.

PACS WORKSTATIONS Regular quality assurance of the monitors on diagnostic PACS workstations ensures they are functioning optimally33. Their brightness, fidelity of greyscale range, and number of dead pixels (on flat-panel monitors) need surveillance. An image should have the same greyscale appearance whatever the device upon which it is displayed, including monitors of differing qualities and makes, film, CDs, and even paper prints. DICOM definition of greyscale display conformance greatly facilities this task by making it amenable to automation. Even trivial quality assurance tasks are important and monitors need regular cleaning to remove finger marks and dust, and keyboards also need cleaning. Laser mice do not suffer from mechanical failure seen in mice containing roller balls.

PACS failure The day-to-day maintenance, percentage uptime, fault-fixing times, and the software and hardware upgrades and ‘refreshes’ of the dedicated PACS workstations and the web browser computers must form part of the PACS contract. Responsibility for the dedicated PACS workstations and their software normally lies with the PACS vendor company, whereas the hospital IT department is usually responsible for the web browser computers, their software and antivirus programs, and for the PACS and hospital networks. The amount of out-ofhours cover is predicated by cost. Often this emergency cover is channelled via the in-house IT helpline. It is vital that these IT helpdesk workers understand that a PACS breakdown has dire clinical consequences (culminating in the total breakdown of all clinical practice), due to the fundamental dependence of all-modern clinical practice upon imaging. Any PACS problem must be given the highest priority by the IT helpdesk, with immediate escalation, unlike a reported email system failure, for example. It is essential to have a robust protocol in place to deal with major potential failures of the integrated PACS, including knowing what to do in the event of failure of:

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The RIS Sending from the modality to PACS The PACS itself, or part of the PACS The network Any other IT systems that may be integrated with the PACS, such as remote electronic requesting, speech recognition and the EPR.

Every major PACS upgrade will currently be associated with planned PACS downtime (although in future it is likely that even large upgrades will be mainly conducted in the background, thereby not disturbing the user), and these upgrades need careful advance planning and hospital-wide coordination34.There has to be a mechanism (by printing film or burning CDs for example) of maintaining an emergency imaging service during the PACS downtime.

PACS HOUSEKEEPING PACS does not run itself and needs daily housekeeping to keep the database tidy. This task is best performed by a dedicated team to provide redundancy (for sickness and holiday leave), although only one full-time equivalent (FTE) is necessary per PACS site producing 80–100 000 examinations per year. Radiographers are best suited to this task (after special training) since they are already familiar with radiological workflow, and understand the difference between the various imaging modalities and body parts imaged. A team of four such radiographers for one FTE post avoids the risk of boredom arising if they were to perform solely PACS housekeeping duties for much longer than a week at a time, and enables them also to maintain their radiography skills by working on a rota basis. PACS housekeeping is an ongoing task that will last throughout the life of the PACS and it is essential that provision be made for such posts at the time of PACS implementation. If the appointment of such posts is delayed, huge backlogs of PACS housekeeping tasks develop, which are difficult to clear. Housekeeping tasks can be summarized according to the resultant problems visualized on the PACS, for each of which there may be several different causes: 1 There is a study entry on PACS but no images 2 The study entry appears as ‘unspecified’ (or similar nomenclature) on PACS 3 The study has a dictated status on PACS but there is no associated report visible on the PACS 4 The study has a nondictated status on PACS and no report 5 The study has a verified state on PACS but needs to be reopened to add more images to it 6 The study is unretrievable from the long-term archive 7 There are CR/direct DR errors 8 Burning imaging studies and their reports to CDs and uploading imaging studies from external CDs/digitizing external films. There will be a large amount of extra housekeeping work generated every time there is an upgrade to the PACS or to any of the IT systems integrated to PACS (such as RIS, speech recognition or HIS/EPR).

1. There is a study entry on PACS but no images Possible causes • The patient was expected for the study at a set time (booked appointment on RIS) but did not attend (DNA), or defaulted from the study before images could be acquired

for some other reason, such as claustrophobia in the MR scanner. The RIS entry automatically generated an order entry on PACS, but no images were generated to go into the study. This situation requires deletion of the RIS entry and of the PACS entry (although it is advisable to keep a permanent record somewhere [usually on the RIS] of the fact that the patient DNA). • The modality on which the images were acquired is not connected to the PACS, such as a mobile US machine printing paper images, or a mobile fluoroscopy unit used in the operating theatres, and therefore the images taken could not be transferred to PACS. However since the study was performed, a report should be generated, and it is often valuable to have such reports accessible on the PACS, despite the lack of images, for later comparative purposes when subsequent studies are performed that are acquired onto the PACS. Clearly such genuine studies, with zero images on PACS but a generated or pending report, should not be deleted.

2. The study entry appears as ‘unspecified’ (or similar nomenclature) on PACS Main possible causes • No RIS entry was made for the study: the demographic details of the patient were entered directly on the imaging modality and the images were then sent to PACS. • The link between RIS and PACS is down so there is no entry on PACS to receive the sent images. • The DICOM modality worklist between RIS and the modality is down (or has not been implemented). The demographic details of the patient therefore have to be entered again, manually, directly on the modality. Also due to typographical errors, there is a mismatch of demographical data sent to PACS from the modality acquiring the images and that from the RIS entry sent to PACS. • Extra images are sent to PACS for an exam that has already been closed on PACS, for example, post-micturition US images, or CT/MR reformatted images. This is not a problem for equipment where DICOM modality performed procedure step (MPPS) has been implemented, whereby individual steps of an imaging study can be closed without closing the whole study. • The imaging plate (or image acquisition) was not assigned to any patient, i.e. no name was supplied. This could arise, for example, if different CR plates containing images from

CHAPTER 1

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PICTURE ARCHIVING AND COMMUNICATION SYSTEMS (PACS) AND DIGITAL RADIOLOGY

the same examination on the same patient were put through different CR readers, if these readers are not networked and so do not communicate with one another. The problem in any of these scenarios is that the PACS is regarding these ‘unspecified’ images as belonging to a study on a separate patient. This means that none of the previous imaging studies on the patient is available for comparison with the unspecified examination, which is not only undesirable but could be medically dangerous. The first bullet scenario should never arise, since it must be made compulsory policy for a RIS entry to be made first before any imaging examination is carried out (and this must apply in the emergency situation and out of hours too). The solution for the next three bullets is to merge the ‘unspecified’ images into the correct PACS folder. In the case of the fifth bullet, no demographic data were supplied at all yet the images were still sent to PACS. Under these circumstances it is essential for the correct patient to be identified, and to reassign the image(s) to the correct patient demographic data on the CR plate reader, and then for the ‘no name supplied’ images to be merged into one folder, bearing the correct demographic details, on PACS.

3. The study has a dictated status on PACS but there is no associated report visible on the PACS Possible causes • Human error: the study is erroneously marked dictated on PACS, by a keystroke or mouse click, by a radiologist (or other reporter). • The dictation is ‘lost’. This can be due to the loss or corruption of a tape (if cassette tapes are being used to record verbal dictation), or to malfunction of a speech recognition or digital dictation system. Action required to remedy this problem is to decide upon a date range to be investigated (choosing an end-point where the report would have been expected to reach PACS), and to check whether a report exists on the RIS or speech recognition/digital dictation system. If it does, then this report needs to be resent to RIS and/or PACS and the failure in this pathway corrected. If no report exists, then the status of the examination needs to be restored to ‘unreported’ status on PACS so that it will reappear in an unreported worklist for a report to be dictated by a radiologist. If the examination is of a ‘special’ type, such as an angiogram, a specific radiologist may need to be alerted and so made aware that this ‘recycled’ study exists and requires a report. It may be necessary to resend the study entry from RIS to PACS, depending upon the type of integration implemented between PACS, RIS and the speech recognition system (if in use). If there is no backward connection from PACS to RIS and/or the speech recognition system, merely changing the status of the study to ‘unreported’ on PACS will not affect its status on RIS or on the speech recognition system. These systems will still view the study as having a reported status, and will therefore block the study from being re-reported.

4. The study has a nondictated status on PACS and no report This refers to studies on PACS that have been ‘missed’ and not reported. In some PACS where there is a relatively small short-term archive, when such studies become old they ‘fall off ’ the short-term archive, their status changes from ‘unreported’ to some different intermediate status, and they are then only present on the long-term archive, but have still not been reported. There are various reasons for examinations to be missed in reporting, including: • The study was performed out of hours on call by a junior radiologist. A verbal report and/or hand-written report in the case notes was issued, but no definitive report was issued awaiting discussion with a consultant radiologist (this never took place). • If cross-sectional imaging work is allocated to radiologists on a sessional basis, a study may be missed at hand-over between, say, morning and afternoon sessions, with each of the radiologists responsible thinking the other is going to report the study. • A particular type of imaging examination may be unpopular to report either because it requires specialist expertise, such as a temporo-mandibular joint MR, or because it is tedious and time consuming, such as a bone age study. Such studies may be left for prolonged periods of time in the hope that they will be reported by another colleague! The solution is as in item (3) above in addition to recalling it to the short-term PACS archive.

5. The study has a verified state on PACS but needs to be reopened to add more images to it This problem arises simply due to the current lack of implementation of DICOM storage commitment (SC) and DICOM MPPS on the modality, the PACS, or commonly both (see later for a description of SC and MPPS). A study may be closed on PACS by a radiographer verifying that all the images have been transferred to PACS, or by a radiologist (or other reporter) reporting the study and marking it ‘dictated’. At a later stage it is realized that further images need to be added to the study. This may arise in a number of circumstances, for example: • Post-processing reconstructions were performed on a specialist workstation after the study was completed. • Dedicated radial head views were being performed at the request of a different radiologist from the one reporting the study, who did not realize that further views had been requested. • The study was reviewed by the reporting radiologist working on a specialist dedicated workstation (e.g. for crosssectional imaging studies), who then reported the study on PACS without opening it, not realizing that not all the images had yet arrived on PACS. • Delayed images were later acquired on an IV urogram. In these circumstances the study status needs to be ‘re-opened’ on PACS, i.e. its status needs to be changed on PACS to

27

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• IMAGING TECHNIQUES AND GENERAL ISSUES

‘unverified’ (or the equivalent nomenclature), and the additional images need to be sent to the study folder on PACS and the verified/reported status of the study reinstated. Obviously if a radiologist has already reported the study without being aware that additional images exist, they must be alerted to the presence of these extra images in case the images alter the validity of the report already issued, requiring a report addendum to be added.

6. The study is unretrievable from the long-term archive If a study cannot be fetched from the long-term store back onto the local short-term archive this usually requires investigation by a PACS engineer, who needs to be alerted. Robust mechanisms will need to be set in place to deal with this problem when linked regional cluster stores are in widespread use nationally (as planned for the Connecting for Health, CfH project in England).

7. Computed radiography/direct digital radiography problems These are various and could reasonably be classified as radiographic and radiographic equipment problems rather than PACS problems, but since plain radiography forms such an integral part of any PACS, they are best dealt with by the PACS housekeeping team. One such example, when the imaging plate (or image acquisition) is not assigned to any patient, i.e. no name was supplied, has already been detailed earlier in item 2. Other examples include: • An image is assigned to the wrong exam of the same, or a different patient. • The wrong processing algorithm is used on the image, which might cause the image to look too grainy or poorly defined. • The imaging plate is cracked, scratched or dirty. The first bullet scenario is an emergency situation for the PACS housekeeper if an image is assigned to the wrong patient, since it could potentially have medically dangerous consequences. The immediate visibility of PACS images to users throughout the health care institution theoretically means that the wrong management could be started on the wrong patient (e.g. a chest drain placed in a patient who did not have a tension pneumothorax), before the fault was rectified by reassigning the image to the correct patient. The very first housekeeping step when an image has been assigned to the wrong exam is to write and save a warning annotation on the image (such as ‘Error: beware this image does not belong to this patient’), which will be immediately visible as an alert to all PACS users. This incorrectly assigned image must then be deleted on PACS. Then, on the CR reader (or other modality), the image must be un-assigned from the wrong patient/study and re-assigned to the correct patient/study and resent to PACS. When the wrong processing algorithm has been used, the suboptimal appearance of the image cannot be corrected merely by changing the contrast window width and level.The

raw data needs to be reprocessed on the CR or DR console using the correct algorithm, and resent to PACS such that the new image resides in the correct PACS study folder. Cracked, scratched or dirty imaging plates need to be identified and replaced or cleaned as appropriate.

8. Burning imaging studies and their reports to CDs and uploading imaging studies from external CDs/ digitizing external films Transfer of imaging studies to institutions outside England (not linked by a shared long-term archive) requires the burning of the images and their associated report onto a CD for physical transfer. Private patients are entitled to a copy of their imaging study and report for future reference. The images should be burnt to the CD in DICOM format with a self-launching DICOM viewer so that the imaging studies can be easily viewed on the recipient’s PC, and can be uploaded to another PACS without having to convert their format. There is now no place for printing images to film, which is more expensive than a single CD and is not amenable to image manipulation (such as changing contrast window width and level, or magnification). CDs (or films) containing imaging studies from outside hospitals/health care centres may be accepted for uploading (or digitizing) onto the institution’s PACS, if a patient is being considered/has been accepted for continuing care at that institution. All such studies will require a RIS entry, and, ideally, a study report, associated with the imaging study transferred to the PACS for review. The amount of uploading of imaging studies from external CDs is escalating exponentially as more institutions implement PACS. (In England it is hoped that adequate implementation of the cluster stores, the Englandwide National spine and the N3 network should alleviate this problem.) Unfortunately many institutions are burning CDs without a DICOM viewer, or with images in non-DICOM format that pose problems for the recipients in reading them and transferring them to their own PACS. The PACS housekeepers should be responsible for trying to resolve these issues in a timely manner, since no one else has been delegated to take on this unexpected extra work.

STORAGE COMMITMENT (Fig. 1.9) SC is a relatively recent DICOM service and a major advance for PACS. It provides for the first time a way for the modality and the PACS to confirm that images sent by the modality are safely stored in the PACS. The PACS sends a message back n images sent Modality

PACS

n images sent

Figure 1.9

DICOM storage commitment.

CHAPTER 1

•

PICTURE ARCHIVING AND COMMUNICATION SYSTEMS (PACS) AND DIGITAL RADIOLOGY

to the sending modality when it has successfully received and committed to storage the expected ‘n’ number of images. On receipt of this message from the PACS, the modality operator (e.g. a CT superintendent) will know that all the images in a study have been successfully acquired onto PACS, and it is now safe to delete those images from the acquisition modality (the CT scanner in this case) without risk of any unarchived images being ablated.

MODALITY PERFORMED PROCEDURE STEP (Fig. 1.10) In MPPS, the modality sends a message to the RIS (and to the PACS as well) whenever a defined part of an imaging acquisi-

PACS RIS Modality

Figure 1.10

DICOM modality performed procedure step.

tion process has been accomplished, so that the RIS and the PACS can progress the case on to the next step (which may be to store the case in the PACS, or to move the case to the reporting stage).

CURRENT AND FUTURE DIRECTIONS OF PACS PACS has now proved itself in sustained filmless clinical practice in numerous sites around the world. The acceptance and implementation of the DICOM standard for imaging equipment and PACS, and of the HL7 standard for RIS and HIS, and the introduction of the EPR, has promoted easier integration of systems from different vendors. The introduction of fast standard (nonproprietary) protocols and the application of web browser technology have enabled PACS installations to become open systems using commercial offthe-shelf hardware, which in turn is reducing the overall cost of PACS. Now that PACS is accepted into routine clinical practice in modern medical institutions, it is the ‘added value’ of PACS in improving clinical workflow efficiency that must take priority. Seamless integration of PACS with other electronic systems in the health care enterprise implies a single log in, using a single workstation, with easy ‘point and click’ navigation between the various software packages. PACS is merely one part of the ultimate total EPR where electronic requesting and booking, speech recognition, electronic feedback and acknowledgement of results, digital recording of the entire case notes, electronic pharmacy, stocktaking and ordering etc. all form a part of the whole. Radiology has led the way in developing PACS to its current high standard and ubiquitous deployment. Other disciplines using clinical images (such as cardiology, obstetric US, endoscopy and bronchoscopy, ophthalmology, dermatology, histopathology etc.) must show the same rigor in workflow process to achieve the PACS equivalent, which requires as priorities: (A) an information system like RIS to book the studies, send the demographic data to PACS and the image acquisition device, and store the reports; and (B) extensive IT support and active housekeeping on a daily on-going basis. There is no technical reason why all clinical images and other clinical digital records should not be integrated into the elusive EPR: it is merely the process and discipline of implementation that is still currently lacking outside the domain of radiology.

ACKNOWLEDGEMENT The author thanks Dr David Harvey for writing the subsection on ‘Storage Requirements and Solutions for PACS’.

REFERENCES 1. Reiner B I, Siegel E L, Hooper F, Protopapas Z 1998 Impact of filmless imaging on the frequency of clinician review of radiology images. J Digit Imaging 11(suppl 1):149–150 2. Redfern R O, Kundel H L, Polansky M, Langlotz C P, Horii S C, Lanken P N 2000 A picture archival and communication system shortens delays in obtaining radiographic information in a medical intensive care unit. Crit Care Med 28:1006–1013 3. Reiner B I, Siegel E L, Hooper F J, Glasser D 1998 Effect of film-based versus filmless operation on the productivity of CT technologists. Radiology 207:481–485 4. Reiner B I, Siegel E L, Hooper F J, Pomerantz S, Dahike A, Rallis D 2001 Radiologists’ productivity in the interpretation of CT scans: a comparison of PACS with conventional film. Am J Roentgenol 176:861–864 5. Siegel E L, Reiner B I 1998 Recommendations for image prefetch or film digitization strategy based on an analysis of an historic radiology image database. J Digit Imaging 11:94–99 6. Bryan S, Weatherburn G, Buxton M, Watkins J, Keen J, Muris N 1999 Evaluation of a hospital picture archiving and communication system. J Health Serv Res Policy 4:204–209 7. Mariani C, Tronchi A, Oncini L, Pirani O, Murri R 2006 Analysis of the Xray work flow in two diagnostic imaging departments with and without a RIS/PACS system. J Digit Imaging 19(suppl 1):18–28 8. Strickland N H 1998 Role of a project leader implementing a hospitalwide PACS. In: Horrii S C, Blaine J G (eds) Proceedings of SPIE Medical Imaging 3339: PACS design and evaluation: engineering and clinical issues, pp 476–478 9. Pilling J 1999 Problems facing the radiologist tendering for a hospital wide PACS system. Eur J Radiol 32:101–105 10. Perry J H 2000 PACS: A PACS RFP Toolkit, www.dejarnevtte.com/ efinegan/pacsprim.htm 11. Reiner B, Siegel E, McKay P 2000 Adoption of alternative financing strategies to increase the diffusion of picture archiving and communication systems into the radiology marketplace. J Digit Imaging 13(suppl 1):108–113 12. The Royal College of Radiologists 2007 Guide to information technology in radiology. BFCR (in preparation). Royal College of Radiologists, London

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13. Boochever S S 2004 HIS/RIS/PACS integration: getting to the gold standard. Radiol Manage 26:16–24 14. CCTA 1998 Managing successful projects with PRINCE (PRINCE 2 reference manual). Central Computing and Telecommunications Agency, London 15. Lewis T E, Horton M C, Kinsey T V, Shelton P D 1999 Acceptance testing of integrated picture archiving and communications systems. J Digit Imaging 12(suppl 1):163–165 16. Andriole K P 2002 Productivity and cost assessment of computed radiography, digital radiography, and screen-film for outpatient chest examinations. J Digit Imaging 15:161–169 17. Graschew G, Roelofs T A, Rakowsky S, Schlag P M, Heinzireiter P, Kranzlmuller D, Volkert J 2006 New trends in the virtualization of hospitals – tools for global e-Health. Stud Health Technol Inform 121:168–175 18. Meenan C, Daly B, Toland C, Nagy P 2006 Use of a thin-section archive and enterprise 3D software for long-term storage of thin-slice CT data sets. J Digit Imaging 19(suppl 1):84–88 19. Brem M H, Bohner C, Brenning A et al 2006 Evaluation of low-cost computer monitors for the detection of cervical spine injuries in the emergency room: an observer confidence-based study. Emerg Med J 23:850–853 20. Strickland N H 2000 PACS (picture archiving communications systems): filmless radiology. Arch Dis Childhood 83:82–86 21. Mathie A G, Strickland N H 1997 Comparison between ‘stack’ and ‘tile’ mode CT image display on PACS. Radiology 203:207–209 22. Luo H, Luo J 2006 Robust online orientation correction for radiographs in PACS environments. IEEE Trans Med Imaging 25:1370–1379 23. Strickland N H 1997 Implications of a filmless hospital for a radiological service. In: Proceedings of the 5th International Conference on Image Management and Communication (IMAC) ‘97. Seoul, Korea. Proceedings of the PakMedTek Symposium 24. Strickland N H, Deshaies M J, Reynolds R A, Turner J E, Allison D J 1997 Short term storage allocation in a filmless hospital. In: Horii S C, Blaine J (eds) Proceedings of SPIE Medical Imaging 3035: PACS design and evaluation: engineering and clinical issues, pp 509–514

25. Strickland N H, Allison D J 1995 Default display arrangements of images on PACS monitors. Br J Radiol 68:252–260 26. Strickland N H, Allison D J, Reynolds R A, Crowell J, Schuurmans M D, Keller S A 1995 Organization of intensive therapy unit images on PACS monitors. In: Jost R G, Dwyer S J (eds) Proceedings of SPIE Medical Imaging 2435: PACS design and evaluation: engineering and clinical issues, pp 278–285 27. Strickland N H, Allison D J, Deshaies M J 1997 Design for the optimal arrangement of magnetic resonance images on PACS monitors. In: Kim Y (ed.) Proceedings of SPIE Medical Imaging 3031: Image display, pp 432–439 28. Okkalides D 1998 Assessment of commercial compression algorithms, of the lossy DCT and lossless types, applied to diagnostic digital image files. Comput Med Imaging Graph 22:25–30 29. Toney M O, Dominguez R, Dao H N, Simmons G 1997 The effect of lossy discrete cosine transform compression on subtle bone fractures. J Digit Imaging 10:169–173 30. Bennett W F, Spigos D G, Tzalonikou M T, Terrell J E, Augustyn M A 1999 Web-based viewing of picture archiving and communications systems images – Part II: the effect of compression on speed of transmission. J Digit Imaging 12(suppl 1):116–118 31. van Ooijen P M, Roosjen R, de Blecourt M J, van Dam R, Broekema A, Oudkerk M. 2006 Evaluation of the use of CD-ROM upload into the PACS or institutional web server. J Digit Imaging 19(suppl 1):72–77 32. Doupi P, Ruotsalainen P, Pohjonen H 2005 Implementing interoperable secure health information systems. Stud Health Technol Inform 115:187–214 33. Crespi A, Bonsignore F, Paruccini N, Macchi I 2006 Grayscale calibration and quality assurance of diagnostic monitors in a PACS system. Radiol Med (Torino) 111:863–875 34. Strickland N H, Allison D J, Beam K W et al 1998 Implementation of a major hardware and software upgrade in a hospital-wide PACS during clinical use. In: Horii S C, Blaine J G (eds) Proceedings of SPIE Medical Imaging 3339: PACS design and evaluation: engineering and clinical issues, pp 496–499

CHAPTER

Intravascular Contrast Media for Radiology, CT and MRI

2

Ronald G. Grainger, Henrik S. Thomsen, Sameh K. Morcos, Dow-Mu Koh, Giles Roditi

INTRAVASCULAR RADIOLOGICAL IODINATED CONTRAST MEDIA Ionic monomers (high osmolar contrast media [HOCM]) Non-ionic monomers (low osmolar contrast media [LOCM]) Ionic dimers Non-ionic dimers Quantity of iodine required Informed consent Adverse reactions to radiological contrast media Types of adverse reaction Complications Principles of investigating patients at increased risk Contrast media related to specific clinical areas Renal Tract Nervous system Cardiovascular System Contrast media, blood coagulation and the endothelium Prevention of adverse reaction to RCM Ethical, financial and cost–benefit considerations High- versus low-osmolar contrast media Conclusion MRI CONTRAST MEDIA Extracellular MRI contrast agents Pharmacokinetics Clinical use of extracellular MRI contrast agents

Safety of extracellular MRI contrast agents Contrast medium-induced nephropathy Nephrogenic systemic fibrosis (NSF) Prevention of adverse reactions Use of extracellular MRI contrast agents in special groups of patients Interaction between extracellular MRI contrast agents with clinical tests Use of extracellular MRI contrast agents for radiographic examinations CONTRAST-ENHANCED MR IMAGING OF THE LIVER AND PANCREAS Non-specific extracellular gadolinium chelates Arterial phase Portovenous phase Interstitial phase Hepatocyte-selective gadolinium chelates Non-gadolinium contrast media CONTRAST-ENHANCED MR IMAGING OF THE PANCREAS Gadolinium contrast-enhanced imaging Mn-DPDP-enhanced MR imaging MRI CONTRAST AGENTS AND TECHNIQUES FOR CONTRASTENHANCED MR ANGIOGRAPHY Concepts of contrast medium administration Peripheral run-off and whole-body CE-MRA Blood pool contrast agents

INTRAVASCULAR RADIOLOGICAL IODINATED CONTRAST MEDIA Ronald G. Grainger Iodine (atomic weight 127) is the only element that has proved satisfactory for general use as an intravascular radiological contrast medium (RCM).The iodine provides the radio-opacity: the other elements of the RCM molecule provide no radio-opacity but act as carriers of the iodine, greatly increasing the solubility and markedly reducing the toxicity of the total molecule.

The problem has always been how to package the iodine so that it may be delivered safely into very sensitive arterial systems (e.g. brain, heart, kidney) in the very large amounts required to produce adequate film-screen radio-opacity1 (see later). Organic carriers of iodine will probably remain the basis of all intravascular RCM for the foreseeable future. Since the 1950s,

32

SECTION 1

• IMAGING TECHNIQUES AND GENERAL ISSUES

Structural formulae of contrast media

Iodine atoms +

Cation

particles

Osmolality osmols/kg H2O 280 mg I2/ml

— Coo

A. Ionic monomer (HOCM) Conventional contrast medium salt e.g. sodium iothalamate (Conray)

3:2

1500

3:1

600

6:2

560

6:1

300

I

I R5

R3 I

B. Nonionic monomer (LOCM) Substituted amide e.g. iohexol (Omnipaque), iopamidol (Niopam)

CONH-R I

I R5

R3 I Cation

+ —

C. Ionic dimer (LOCM) Mono-acid dimeric salt e.g. sodium ioxaglate (Hexabrix)

R

Coo I

I R

R I

D. Nonionic dimer (LOCM) e.g. iodixanol (Visipaque) iotrolan (Isovist)

I

I

I

R

R I

I

I

I

R

IONIC MONOMERS (HIGH OSMOLAR CONTRAST MEDIA [HOCM])

R I

there are four chemical varieties of iodinated RCM in clinical use (Fig. 2.1,Tables 2.1, 2.2). All four are tri-iodo benzene ring derivatives with three atoms of iodine at 2,4,6 positions (in monomers) and six atoms of iodine per molecule of the ring anion (in dimers); they are very hydrophilic, have low lipid solubility, low toxicity, low binding affinities for protein, receptors or membranes, and have molecular weights less than 20002 (Table 2.1). On intravascular injection, all of these four products are distributed rapidly because of high capillary permeability into the extravascular, extracellular space (except in the central nervous system [CNS]). They do not enter the interior of blood cells or tissue cells and they are rapidly excreted, with over 90 per cent being eliminated by glomerular filtration by the kidneys within 12 h. None of these four chemical types of RCM have marked pharmacological actions. This is highly desirable as they are used purely as imaging and not therapeutic agents.

I

Figure 2.1 Structural formulae of the four different types of iodinated intravascular contrast media: osmolalities and ratio of iodine atoms to osmotically active particles.

All ionic monomers are salts with sodium or meglumine (Nmethyl glucamine) as the non-radio-opaque cation and a radioopaque tri-iodinated fully substituted benzoic acid ring as the anion. These anions include diatrizoate (Urografin, Hypaque), iothalamate (Conray), ioxithalamate, metrizoate – the first two

Table 2.1 IODINATED CONTRAST MEDIA: CHARACTERISTICS, IODINE CONTENT AT 300 mosmols kg-1 WATER (PHYSIOLOGICAL) (MODIFIED AFTER MORRIS 19932) Classification iodine per particle ratio

Molecular weight

Iodine content mg ml−1 at 300 mosmol kg−1 water

Osmolality mosmol kg−1 water at 300 mgl ml−1

Classification of RCM generic names

Iodine atoms per molecule

Osmotic particles per molecule

Ionic monomer HOCM

3

2

3:2

600–800

70

1500–1700

3

1

3:1

700–800

150

600–700

6

2

3:1

1269

150

560

6

1

6:1

1550–1626

300

300

Diatrizoate Iothalamate Metrizoate Ioxithalamate Non-ionic monomer LOCM Iohexol Iomeron Iopamidol Iopromide Ioversol Ioxilan Xenetix Ionic dimer LOCM Ioxaglate Non-ionic dimer LOCM isotonic Iodixanol Iotrolan

CHAPTER 2

being by far the most frequently used. Each molecule completely dissociates in water solution into two ions – one non-radioopaque cation and one tri-iodinated radio-opaque anion, giving an iodine:particle ratio of 3:2 (1.5) (Table 2.1 and Figure 2.1). They are very hypertonic – 1600 mosmols kg−1 water at 300 mg iodine kg−1 compared to physiological osmolality of 300 mosmol kg−1 water.

NON-IONIC MONOMERS (LOW OSMOLAR CONTRAST MEDIA [LOCM]) Almén introduced the concept of non-ionic (and therefore low osmolar) contrast media in 1969 with metrizamide. The 2nd generation non-ionic monomers in 2006 include iohexol, iopamidol, iopromide, ioversol, ioxilan, etc. (Tables 2.1 and 2.2) and are much more stable, much more soluble, much less toxic and much less expensive than metrizamide. None of these molecules dissociate in solution. They are tri-iodinated non-ionizing compounds and therefore in solution they provide three atoms of iodine to one osmotically active particle (the entire molecule), producing an iodine: particle ratio of 3:1 (Tables 2.1, 2.2, Fig. 2.1). They have less than half of the osmolality of HOCM. They have an osmolality of about 600 mosmol kg−1 water at a concentration of 300 mg I /ml-1 compared with the physiological osmolality of 300 mosmols.

IONIC DIMERS Ioxaglate is the only compound in this group. It is a mixture of the sodium and meglumine salts of a monoacidic double benzene ring with each benzene ring having three atoms of iodine at C2, 4, 6 positions. The total molecule therefore contains six atoms of iodine and in solution each molecule dissociates into one radio-opaque hexa-iodinated anion and one non-radio-opaque cation (sodium and/or meglumine). Ioxaglate therefore has an iodine:particle ratio of 6:2 or 3:1 (Tables 2.1, 2.2, Fig. 2.1).

Table 2.2

•

INTRAVASCULAR CONTRAST MEDIA FOR RADIOLOGY, CT AND MRI

NON-IONIC DIMERS Iotrolan (Schering) and iodixanol (Nycomed) are both examples of non-ionizing chemicals. They do not ionize or dissociate in solution. Each molecule contains two non-ionizing tri-iodinated benzene rings linked by a bridge (Tables 2.1, 2.2, Fig. 2.1). Each molecule therefore provides in solution six atoms of iodine for one molecule, i.e. an iodine:particle ratio of 6:1. Non-ionic dimers are physiologically isotonic (300 mosmol kg−1 water) in solution at 300 mg I/ml-1. Iotrolan has been recently voluntarily withdrawn for intravascular use because of unexpected delayed reactions – usually with allergic and skin manifestations that are rarely serious but seem to be more severe and frequent in the Japanese. These delayed adverse reactions may also apply to other dimers. There are recent but unconfirmed reports that non-ionic dimers may be less nephrotoxic than non-ionic monomers3,4 (see renal toxicity later).

QUANTITY OF IODINE REQUIRED Very large amounts of iodine are necessary because of the low sensitivity inherent in conventional photographic film-screen radiography. This low sensitivity to iodine is not an inherent property of the X-ray beam but is due to the low sensitivity of the photographic process in distinguishing minor differences in radio-opacity. Alternative nonphotographic X-ray detection recording systems, such as photodetectors with electronic amplification, like those used in CT scanning and in digital subtraction angiography (DSA), are very much more sensitive to minor differences in iodine concentration as demonstrated by 95 per cent tumour blush in all cerebral tumours on CT after IV injection of RCM compared to about 20 per cent blush after intra-arterial injection using photographic film-screen radiography.

REPRESENTATIVE LOW-OSMOLAR MEDIA (LOCM)

Generic name

Registered trade name

Iodine mg ml−1 water

Osmolality mosmols kg−1 water

Viscosity cps at 3.7°C

320

580

7.5

Ionic LOCM Dimer Meglumine ioxaglate 32.9% (Sodium ioxaglate 19.65%)

Hexabrix 320 Non-ionic monomer

Iopamidol 61%

Niopam 300

300

616

4.7

Ioversol 64%

Optiray 300

300

645

5.5

Iopromide 62.3%

Ultravist 300

300

620

4.6

Iobitridol 65.8%

Xenetix 300

300

695

6.0

Iohexol 64.6%

Omnipaque 300

300

640

6.1

Non-ionic dimer Iodixanol 65%

Visipaque 320

320

290

11.4

Iotrolan 64%

Isovist 300

300

320

8.1

33

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The average daily physiological turnover of iodine is 0.0001 g. The total iodine content in the body (mainly in the thyroid) is 0.01 g. The requirement for a 2-, 3- or 4vessel angiogram (with conventional film-screen photographic recording) may be 70 g iodine, delivered in, say, 30 min into the arterial supply trees of very sensitive organs, for example, the brain, kidney, or heart. This is 700 000 times the daily body turnover of iodine and it is testimony to the low toxicity of RCM (either HOCM or LOCM) that this huge quantity of iodine can be accepted into the arterial tree of the most sensitive organs of the body with very rare evidence of severe toxicity. In complicated cases, these massive quantities may have to be increased by 100 per cent or even more, involving more than 200 g of iodine (400++ g of RCM) being injected intravascularly for film-screen radiographic angiography, probably the largest quantity of any drug used in the whole of clinical medicine.

INFORMED CONSENT Informed consent is essential for any invasive procedure (angiography, angioplasty, vascular embolization, biopsy, etc.), but probably is not essential in many countries for uncomplicated procedures (e.g. intravenous urography [IVU]) in a reasonably fit patient, as it may alarm the pvatient to learn that there is a risk (however small) of a severe or fatal ADR. In some countries, e.g. the USA, informed consent may well be necessary, even in these patients. Like every other drug, contrast agents should be administered only when there are clear and defensible clinical indications and when the prospects of benefit to the patient outweigh the risks and discomfort that may occur. With modern contrast agents (either HOCM or LOCM), it is more likely that a complication occurring during a radiological arterial procedure will be due to the consequences of the instrumentation (e.g. vessel damage, thromboembolism) rather than to the contrast agent, if the latter is used in an appropriate dose and manner.

ADVERSE REACTIONS TO RADIOLOGICAL CONTRAST MEDIA6–12 An excellent (1999) review of adverse reactions, their mechanism, prophylaxis and treatment is presented by Sidhu and Dawson6. In clinical medicine, every technical procedure and every drug given to a patient by whatever route (oral, rectal, IV, IA, intrathecal) may cause adverse drug reactions (ADRs), which are sometimes fatal. Radiographic procedures and drugs are no exceptions. It is inevitable that even when administered appropriately, RCM will very occasionally result in a severe or even fatal ADR, sometimes leading to litigation. It has become increasingly difficult to defend inappropriate indications, inadequate resuscitation facilities and inappropriate experience, especially in small imaging departments, in either private or public hospitals.

TYPES OF ADVERSE REACTION Adverse reactions to RCM may be divided into: 1 Idiosyncratic anaphylactoid reactions 2 Non-idiosyncratic reactions 3 Combined 1 and 2 reactions.

Idiosyncratic anaphylactoid reactions These ADRs are the most dreaded and most serious and fatal complications of RCM injection as they occur without warning, cannot be reliably predicted and are not preventable in the present state of our knowledge. These reactions usually (85 per cent) begin either during or immediately after the injection of contrast medium. ADRs are more frequent in patients who have had a previous adverse reaction to contrast medium, asthmatics, allergic and atopic patients, patients with impaired cardiovascular and renal systems, diabetics (Table 2.3), and patients on beta-adrenergic blockers and possibly on nonsteroidal analgesics. Idiosyncratic reactions vary greatly in severity and clinically may closely mimic true anaphylactic and allergic reactions, although there is no confirmatory evidence of a true antigen IGE antibody immunological reaction to contrast media. Some possible mechanisms5,6,10,11 involved in causing these reactions are: 1 Inhibition of enzymes such as cholinesterase (which de-activates and hydrolyses acetylcholine), resulting in increased concentration of acetylcholine; this may result in symptoms of vagal overstimulation, e.g. cardiovascular collapse, bradycardia, bronchospasm. 2 Release of vasoactive substances such as histamine, serotonin or bradykinin may result in vasomotor collapse. 3 Activation of physiological cascade systems including the complement activation system, the kinin system with bradykinin release, the coagulation system inducing intravascular coagulation and the fibrinolytic system causing lysis of fibrin and blood clots. 4 The immune system disturbances. 5 Anxiety, apprehension and fear of the radiological procedure probably plays a significant part in adverse reactions by activating a hypothalamic reaction resulting in cardiovascular and respiratory collapse and even death. It is probable that many of these mechanisms are inter-related and that a self-perpetuating vicious circle is established, resulting in severe and perhaps fatal cardiovascular collapse. Unlike the chemotoxic and hyperosmolar reactions, the earlier mentioned anaphylactoid reactions are not dose dependent and death has been known to occur following a 1 ml IV test dose, or after the full dose of RCM has been given after a negative test dose.

Non-idiosyncratic reactions Unlike the idiosyncratic reactions, these non-idiosyncratic reactions are dose dependent and therefore relate to the chemical composition, osmolality and concentration of contrast medium and the volume, speed and multiplicity of the injection.

CHAPTER 2

Table 2.3 ADVERSE DRUG REACTIONS TO IV INJECTION OF HOCM (IONIC) AND LOCM (NON-IONIC) MONOMERIC CONTRAST MEDIA (MODIFIED FROM KATAYAMA ET AL 19907) Ionic HOCM

Patients

Non-ionic LOCM

169 284

Total number of patients

163 363

12.66%

Total adverse reactions

3.13%

0.22%

Severe ADR

0.04%

(1)

Deaths (uncertain cause) (1)

(1)

All ADR% ionic HOCM

Symptoms

All ADR% non-ionic monomeric LOCM

4.58

Nausea

1.04

2.29

Heat

0.92

2.97

Vomiting

0.36

3.16

Urticaria

0.47

1.12

Flushing

0.16

0.40

Vascular pain

0.05

0.58

Coughing

0.15

0.17

Dyspnoea

0.04

Ionic HOCM, severe ADR (%)

Clinical history

Non-ionic LOCM, severe ADR (%)

1.88

Asthma

0.23

0.18

Without history of allergy

0.03

0.53

With history of allergy

0.10

0.13

Without history of previous ADR to RCM

0.03

0.73

With history of previous ADR to RCM

0.18

0.53

Heart disease

0.10

0.20

Renal disease

0.04

0.22

Diabetes mellitus

0.05

1 Chemotoxic reactions. Chemotoxic adverse reactions are probably due to toxicity to the contrast medium anion rather than to its iodine content, as the iodine is very firmly bound to the benzene ring. Some patients with a low iodine intake (often geographical), have an increased incidence of endemic goitre, are very sensitive to an increased intake of iodide, and may develop thyrotoxicosis after injection of RCM contrast agents. Presumably in these patients this is due to the very small amount of free iodide ions, which is known to be present in contrast medium solutions. Special care should be taken with patients who are thyrotoxic or goitrous. Chemotoxic side effects include cardiac, neurological and renal toxicity as well as vascular manifestations. Cationic toxicity may result from changes (either greater or lesser) in sodium or calcium ion concentration induced by injection of large amounts of RCM. 2 Hyperosmolar reactions. The very high osmolality of high concentrations of HOCM (ratio 3:2) is 5–8 times

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INTRAVASCULAR CONTRAST MEDIA FOR RADIOLOGY, CT AND MRI

the physiological osmolality (300 mosmol kg−1 water) of every cell in the body (Table 2.1, Fig. 2.1). It is not surprising therefore that adverse reactions may occur due to the marked hyperosmolality that develops when large volumes of concentrated HOCM are injected into the circulation. The degree of hyperosmolality is much reduced if non-ionic ratio 3:1 (LOCM) and (even more) if ratio 6:1 products isotonic non-ionic dimeric LOCM are injected instead of injecting HOCM of ratio 3:2 (Tables 2.1, 2.2). The adverse reactions due in part to hyperosmolality of the contrast medium include: 1 Erythrocyte damage due to the loss of intra-erythrocytic water, leading to rigidification and reduction of the vital essential deformability of red cells, which is necessary for the erythrocytes (7 µm diameter) to pass through capillaries (diameter 2–4 µm). 2 Endothelial damage. All RCMs, especially HOCM injure the endothelium by virtue of their chemotoxicity, ionicity and osmolality. The normal endothelium permits passage of molecules less than 50 000 molecular weight (HOCM or LOCM [even dimers] have molecular weights less than 20 000) by freely diffusing through intercellular pores and by transcytosis. Thus all RCMs are readily passed through the endothelium into the tissue fluid except in the CNS (blood–brain barrier – see later). Damage to the capillary wall of the injected arterial territory (perhaps by RCM) leads to its increased permeability and the passage of toxic substances from the blood into the extravascular tissue fluid that bathes the tissue cells. In addition to its semi-permeability membrane function, the endothelial cells of the capillaries also constitute a highly significant and very sensitive pharmacological tissue and they are stimulated to produce several active substances such as Factor XII, arachidonic acid, phospholipids, prostacyclins, angiotensin-converting enzymes, nitric oxide, pro-coagulants and anti-coagulants etc., which have considerable effects in vasoregulation and coagulation6,10. 3 Blood–brain barrier damage is a specific example of endothelial damage. The central nervous system is the only tissue in the body that is normally protected from the passage of contrast medium molecules through the capillary wall, because the capillaries of the central nervous system are relatively impervious and have very tight intercellular connections. If this capillary blood–brain barrier is damaged, the molecules or ions of RCM may pass through the capillary wall into the extravascular neurological tissue fluid and exert a direct toxic effect on the delicate nerve tissue cells. A pathological increase of permeability of the blood–brain barrier occurs in several conditions, e.g. many cerebral and meningeal tumours, cerebral infarcts and in some infections. This is the basis of the very useful diagnostic sign of tumour staining, which is particularly well seen on CT, even after IV injection of contrast medium.

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4 Vasodilatation of the arteriolar and capillary bed of the injected artery is closely related to the hyperosmolality of the contrast medium solution. It may lead within a few seconds to a rapid increase in blood flow to several times the physiological flow rate. This is associated with a local uncomfortable feeling of warmth, heat, discomfort, or severe pain during peripheral arteriography. Systemic generalized vasodilatation following injection of large quantities of hyperosmolar contrast medium may lead to severe systemic hypotension, peripheral venous pooling and a diminished venous return to the heart, resulting in cardiovascular failure that may be extremely severe and even fatal. 5 Hypervolaemia results from the osmotic attraction of extravascular tissue fluid into the circulation induced by the hyperosmolality of RCM. This may increase the blood volume by about 10 per cent within a few minutes, thus reducing haemoglobin concentration and increasing the demand on cardiac output and stress on the left ventricle. 6 Cardiac depression with systemic hypotension, diminished venous cardiac return, myocardial ischaemia and the direct depression of myocardial contractility, particularly after coronary angiography. All of these hyperosmolar reactions are closely related to the hyperosmolality of RCM and therefore depend on the dose, concentration and volume of RCM injected and are consequently considerably reduced by substituting LOCM for the very hypertonic HOCM.

Vasomotor reactions These vasomotor ADRs may occur following either idiosyncratic (not dose-dependent) or non-idiosyncratic (dosedependent) reactions induced by RCM, or may develop independent of other ADRs. There may be severe hypotension, tachycardia or bradycardia with marked apprehension, anxiety, sweating and disturbance, and depression of myocardial contraction with a markedly reduced cardiac output, leading to unconsciousness and cardiopulmonary collapse and possibly death. Vaso-vagal reactions may also occur, characterized by bradycardia, whereas the vasomotor collapse states mentioned earlier are usually accompanied by tachycardia.

COMPLICATIONS Minor reactions7 Flushing, nausea, arm pain, pruritus, mild urticaria, vomiting and headache7 (Table 2.3) occur in 5–10 per cent of patients injected with HOCM but less than 1 per cent when LOCM is injected. These are usually mild in severity, of short duration, self-limiting and generally require no specific treatment other than reassurance and restoration of the patient’s confidence. Occasionally an oral antihistamine (e.g. chlorpheniramine [Piriton] 4 mg) for urticaria, a tranquillizer (e.g. diazepam [Valium] 5 mg) or a mild analgesic (e.g. aspirin 300 mg or paracetamol 500 mg) may be helpful. Arm pain may be due to a small unrecognized extravasation, but venous spasm on entirely intravascular injection may also

occur, especially with HOCM. There may be local oedema, very rarely resulting in cellulitis, compartment syndrome, or tissue necrosis. A major problem may arise if there is a large extravasation such as in powered injections of RCM for dynamic CT. A major recent advance is an automatic device to electrically stop the pump5. The electrical skin impedence changes if there is significant extravasation. Electrodes over the site of injection are connected by cable to the power injector, which is automatically immediately cut off if significant extravasation occurs. This is an important safety device that has been adopted by several manufacturers of powered injectors, and permits higher IV flow rates with considerably increased safety5.

Intermediate reactions More serious degrees of the above generalized symptoms, including moderate degrees of hypotension and bronchospasm occur in 0.5–2 per cent of IV injections of HOCM and is probably reduced to one-quarter of that frequency following IV LOCM7 (Table 2.3). These intermediate reactions disturb both the patient and the doctor but do not cause alarm. They usually respond readily to appropriate therapy – reassurance, chlorpheniramine (Piriton) (4–10 mg orally, intramuscularly or intravenously) for urticaria or angioneurotic oedema, diazepam (5 mg) for anxiety, salbutamol β2 agonist inhalation for bronchospasm, hydrocortisone 100–500 mg IM or IV, adrenaline 0.1–0.5 ml of 1/1000 solution intramuscularly or subcutaneously (children 0.01 ml kg−1 body weight up to 3 mg max.) for more severe bronchospasm – repeated if necessary.

Severe life-threatening reactions These include severe manifestations of all of the above symptoms and occur in about 0.2 per cent after HOCM and about 0.04 per cent following LOCM7 (Table 2.3). Severe reactions include convulsions, unconsciousness, laryngeal oedema, severe bronchospasm, pulmonary oedema, severe cardiac dysrhythmias and cardiac arrest, cardiovascular and respiratory collapse.Treatment is urgent: the patient is very alarmed and the doctor is seriously worried about the patient’s possible demise. The responsible clinician (usually the radiologist) must be well trained in emergency treatment as these severe ADRs are unpredictable in occurrence and in response to treatment. Treatment6,10–12 must be urgently and expertly administered according to a pre-arranged well-practised schedule. Additional medical (including anaesthetic) assistance should be immediately available. A crash trolley complete with oxygen cylinders, flow valves, suction equipment, oral airways, nasal tubes, endotracheal tubes, face masks, ambu-type bag, stethoscope, sphygmomanometer, syringes, needles, cannulae and tracheotomy set is essential. A suction machine, electrocardiogram (ECG), blood pressure monitor and DC defibrillator must also be readily available. There must be a comprehensive selection of drugs and IV fluids necessary for resuscitation and

CHAPTER 2

the means and the personnel for their safe administration must be readily available. The airway must be secured and oxygen, artificial respiration, external cardiac massage and electrical DC defibrillation must be administered as and when required. IV fluid infusion (normal saline, lactated Ringer's solution) through an indwelling IV catheter is essential to restore blood volume and to administer IV drugs, such as: • a powerful diuretic such as frusemide (Lasix) 20–40 mg IV slowly or IM for pulmonary oedema • diazepam and barbiturates for convulsions • adrenaline (see later) • salbutamol (β2 agonist metered dose inhaler) • hydrocortisone or methyl prednisolone (100–1000 mg) • aminophylline (very slowly, 250–500 mg) intravenously for intense bronchospasm • chlorpheniramine for allergic or anaphylactic symptoms • vasopressors, e.g. noradrenaline (or metaraminol [Aramine] 0.5–5 mg slow IV infusion) • dihydroxyphenylaline (or dopamine) infusion (2.5–5 µg kg−1 min−1) for hypotension with monitoring of the blood pressure Sodium bicarbonate infusion should be administered to correct any acidosis, and atropine 0.6–1.0 mg IV or IM repeated if necessary (children 0.02 mg kg-1 repeated if necessary to 2 mg total) is used for vasovagal reactions, bradycardia and cardiac failure. Adrenaline (recently prepared solution and in date) is much more rapidly effective than steroids: it is the main therapeutic agent and should always be readily available for it is the essential treatment. In severe anaphylactoid reactions it should be administered as soon as possible and may be life-saving. Adrenaline (epinephrine) (0.3–0.5 ml, 1/1000 solution [children 0.01 ml kg−1 body weight] by deep SC or IM injection repeated at 10–20-minute intervals) provides the most rapid and reliable relief for bronchospasm, angioneurotic oedema and other anaphylactoid symptoms. Adrenaline has occasionally precipitated ventricular fibrillation when given by IV injection, but in very severe anaphylactoid reactions, even this route should be considered, injecting 1–5 ml, 1/10 000 solution, very slowly, preferably under ECG control. Corticosteroids take several hours to become fully effective, even when given intravenously, but they should be administered immediately and intravenously in severe anaphylactoid reactions.

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INTRAVASCULAR CONTRAST MEDIA FOR RADIOLOGY, CT AND MRI

cized (reasonably) on methodological grounds, they are very useful contributions to the international literature and knowledge of RCM.They both found that deaths due to LOCM are no fewer than to HOCM6. A very interesting analysis of Deaths Related to Iodinated Contrast Radiological Media (ICRM) reported to the United States Federal Drugs Administration between 1979 and 1994 (estimated 170 million IRCM studies) was published in 19979 and concluded that despite the increasing availability and usage of the safer LOCMs in the USA in 1978–1994, the data do not show a decrease in RCM-related deaths to LOCM as opposed to HOCM injection.

Factors significantly predisposing to adverse drug reactions (Table 2.4) All ADRs are significantly more frequent with HOCM (12.66 per cent) than with LOCM (3.13 per cent)7 (Table 2.3). Previous adverse reactions to RCM and a history of asthma are the two most dangerous adverse predisposing factors.These and others include: 1 Previous adverse reaction. A previous adverse reaction to contrast medium increases the risk factor of a reaction to a second LOCM injection (severe ADR 0.18 per cent)7 (Table 2.3). 2 History of asthma or bronchospasm. Patients with a previous history of bronchospasm, however mild, have an increased risk of an adverse reaction to RCM (0.23 per cent severe ADR, Table 2.3). Subsequent reactions to a second RCM injection may be very severe and even fatal. 3 History of allergy or atopy. Patients with known allergy, particularly to foodstuffs and perhaps especially to iodides have an increased incidence of an ADR (0.10 per cent severe ADR)7 (Table 2.3). 4 Cardiac disease. The risk of severe ADRs in severe cardiac disease is 0.10 per cent severe ADR following LOCM7 (Table 2.3). A high sodium load caused by a large volume of HOCM containing sodium salts should be avoided, as this may exacerbate cardiac decompensation due to hypervolaemia and myocardial depression.

Table 2.4 PATIENTS CONSIDERED TO BE AT GREATER THAN USUAL RISK OF A SEVERE ADVERSE REACTION TO IODINATED RADIOLOGICAL CONTRAST AGENTS 1. Patients with a previous ADR to RCM (excluding mild flushing, nausea)

Deaths

2. Asthmatics

The earlier mentioned severe adverse reactions may not respond even to expert, rapidly administered treatment and the patient may die. Death may however be immediate and without warning. It usually results from intractable cardiopulmonary collapse, pulmonary oedema, or intense bronchospasm. The death rate following IV injection of contrast medium is not known with accuracy. One death in 40 000 IV injections of HOCM is often quoted as a medium figure. Both the Katayama7 (in Japan) and Cashman (in Australia) papers were unblinded large surveys, but despite being criti-

3. Allergic and atopic patients 4. Cardiac patients with decompensation, unstable arrhythmia, recent myocardial infarction 5. Renal patients in failure, diabetic nephropathy, on metformin 6. Feeble infants and aged patients 7. Patients with a severe general debility 8. Very nervous, anxious patients 9. Patients with various metabolic and haematological disorders 10. Thyrotoxic: goitrous patients

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5 Renal disease6. Mild temporary renal dysfunction may occur in around 5 per cent; serious renal impairment occurs in about 0.1 per cent7 (Table 2.3). The most important safeguard against RCM-induced renal failure is adequate hydration, given orally or intravenously. 6 Dehydration. All patients must be well hydrated (if necessary by IV) before and after contrast medium is injected. The previous practice to dehydrate the patient deliberately for IVU (in order to produce an improved image of the renal calices) is very dangerous and must be abandoned in all patients. 7 Haematological and metabolic conditions. Patients with sickle cell anaemia may react adversely, particularly to HOCM, which may cause an increase in sickling, with potentially serious or even fatal complications following selective cerebral and coronary angiography. 8 Anxiety and apprehension. Anxiety and apprehension definitely predispose to adverse reactions. 9 Other conditions. Manifest hyperthyroidism is a contraindication to RCM administration. Neonates and aged and infirm patients are probably at increased risk because of their associated morbidities. 10 Delayed adverse reactions. Serious adverse reactions usually begin during or within 15 min of IV injections of RCM. In the last few years it has become apparent that some patients (perhaps as much as 10–20 per cent) experience reactions that may be delayed for a few days.These delayed reactions are usually uncomfortable, time-limited and not dangerous. Patients with a thrombotic tendency such as severe polycythaemia may have an increased risk, especially after arterial injections of non-ionic LOCM. HOCM inhibits thrombosis more than LOCM. Patients with myelomatosis may develop renal failure after contrast medium injection, with proteinaceous casts and precipitation of Bence-Jones protein and Tamm-Horsfall protein possibly causing renal tubular obstruction. These reactions are now much less frequent because LOCMs are much less toxic and less protein binding, and the previous practice to dehydrate the patient before IV urography is now completely reversed. Patients with phaeochromocytoma may develop a hypertensive crisis during angiography or IV urography as a reaction to the contrast agent, to abdominal compression or to manipulation of the tumour. Whenever phaeochromocytoma is suspected, these patients should be prepared with preliminary alpha- and beta-adrenergic blockade before angiography. Patients with pre-existing renal disease, particularly the renal complications of diabetes mellitus, have an increased risk of an adverse reaction to RCM, which may precipitate renal failure. LOCM is preferred especially if the patient has pre-existing renal impairment6,10–12. Recent evidence3 suggests that iso-osmolar dimer (iodixanol) LOCM may be even less nephrotoxic than monomeric LOCM (iohexol), which has twice the physiological osmolality. In a multicentred trial of diabetics with pre-existing renal

impairment (serum creatinine 115–308 mmol-1) 129 patients underwent angiography3. There was a significant increase in contrast-induced nephrotoxicity with increased serum creatinine in 26.2 per cent of the iohexol (non-ionic monomer) group compared to 3.1 per cent in the iodixanol (non-ionic dimer) group. This important finding is being strongly challenged partially because of the large size of the dimer molecule (which may itself impose a renal problem) and the possibility of an increase in delayed ADRs. In 2005 an in vitro study considered the cytotoxic adverse effects on renal tubular cells of non-ionic dimers (iodixanol and iotrolan) compared with monomer LOCM (ioversol and iomeprol). This study demonstrated significantly more toxic effects of the dimers when compared on an equi-molar concentration with the monomers4. The authors conclude that their in vitro research indicates considerable caution in accepting the advantages of the dimers reported in the previous paper3. Metformin (phenoformin), an oral biguanide hyperglycaemic therapy for diabetes mellitus Type 2, is primarily excreted by the kidney. In patients with severe renal impairment, intravascular accumulation of the biguanide may occur after RCM, and this may precipitate biguanide lactic acidosis (vomiting, diarrhoea, somnolence) – a potentially fatal complication. RCM-related lactic acidosis is extremely rare in diabetics receiving metformin if the patient has normal renal function before RCM. It is therefore recommended12 that metformin need only be discontinued if there is pre-existing renal impairment and it may be recommenced 48 h later after re-assurance of good renal function after RCM. These delayed reactions are more common in women: arm pain (7 per cent), delayed rash (5 per cent), flu-like symptoms (headache, skin lesions, salivary gland swelling – ‘iodide mumps’ (9 per cent) probably due to ‘iodism’) are the more important late manifestations. Presumably the reaction is idiosyncratic and not dose dependent. A few cases (including deaths) of delayed vasculitis, disseminated lupus erythematosis and Stevens–Johnson syndrome have been reported after injection of LOCM. Delayed reactions in general may be more frequent with LOCM than with HOCM and with dimers compared to monomers. Iotrolan – a non-ionic dimer – has been voluntarily withdrawn from intravascular clinical usage because of a worryingly high incidence of delayed adverse reactions, mainly reported in Japan. It is possible that similar delayed ADRs may occur after other similar dimeric LOCMs.

PRINCIPLES OF INVESTIGATING PATIENTS AT INCREASED RISK Manifest active hyperthyroidism is an absolute contra-indication to RCM administration. The following suggestions should be noted: 1 Despite the above extensive list of those patients who have a known predisposition to adverse reactions, many com-

CHAPTER 2

2

3

4

5

6

7

8 9

10 11

plications of contrast medium injections are unpredicted and unpredictable. The radiologist must always weigh the possible advantages to the patient of the diagnostic procedure against the very small but ever-present risk of a severe adverse reaction (and even death) to the contrast medium and the technique of its administration. In many patients deemed to be at increased risk of an adverse reaction, the RCM procedure may be replaced by other imaging investigations, such as ultrasound (US), computed tomography (CT) or magnetic resonance imaging (MRI), often not requiring injections of any contrast agent but producing equal or improved diagnostic information. The referring clinician should be consulted and advised if the patient is thought to be at increased risk. The patient should be informed, consulted and reassured as necessary. The discussion and result should be recorded in the patient's notes. Although a routine test dose is generally not necessary, it is advisable to perform a 1 ml or smaller test dose, either by IV or mucosal testing in patients considered to be at increased risk of an ADR. The smallest dose of RCM that will give reliable, comprehensive diagnostic results should be used. This is especially important in patients with pre-existing renal impairment or diabetes, those who are poorly hydrated, infants and those who may require further RCM procedures. All patients receiving RCM, particularly if there is previous renal impairment and those patients receiving large quantities of RCM, must be very well hydrated (if necessary intravenously) before and after the injection in order to excrete the RCM as completely and quickly as possible. Do not leave the patient unattended for the first few min after the injection. Premedication is strongly advised for previous adverse reactors, especially in asthmatic and allergic patients, who should be given pre-injection and post procedure 24 h of oral corticosteroids and possibly antihistamines. In all patients considered to be at increased risk, LOCM should be administered in preference to HOCM. Before administering the injection, the emergency trolley, alarm system and availability of experienced assistance must be confirmed.

CONTRAST MEDIA RELATED TO SPECIFIC CLINICAL AREAS RENAL TRACT6,10–12 There is no doubt that high doses of contrast media impair renal function, usually peaking at 3–5 d, causing a decrease in urine output and an increase in serum creatinine and urea levels, decreased creatinine clearance and reduced glomerular filtration rate (GFR). In some patients this may proceed to renal failure with anuria, uraemia and death. Renal dialysis (either intravas-

•

INTRAVASCULAR CONTRAST MEDIA FOR RADIOLOGY, CT AND MRI

cular or peritoneal) is very effective and may be life-saving as an alternative method of excreting RCM and uraemic metabolic products. These severe adverse reactions are very unlikely to occur if the patient is well hydrated and has normal renal function before the RCM injection. Particularly important adverse factors are pre-existing renal failure and oliguria, diabetic nephropathy, nephrotoxic drugs, patients who are not well hydrated and patients who are liable to be injected with very high doses of RCM for multiple sequential examinations repeated within a few days. Alternative imaging procedures must always be considered. The usual recommended dose for IV urography in the normally well-hydrated adult with normal renal function is 15–25 g iodine; this dose may be increased provided the patients are well hydrated (by IV normal saline if necessary before, during and after RCM injection). A maximum of about 70 g of iodine (1 g iodine kg-1 body weight in adults) is generally advisable even in patients with good renal function, but considerably larger quantities (up to 200 g or even 300 g of iodine, i.e. up to 600 g of RCM) may be required, particularly in difficult angiographic and interventional procedures. Nephrotoxic drugs (diuretics, antibiotics, non-steroidal anti-inflammatory drugs, cytotoxic therapy) may enhance the adverse effects of RCM. Diabetic patients receiving metformin should be given special attention (see earlier). After early uncertainty, it is now established that HOCMs are more nephrotoxic, and LOCMs are preferred for all patients considered to be at increased risk, especially with diabetic nephropathy. There is increasing evidence that RCM (particularly HOCM and large doses) may induce damage to the tubules in the renal medulla and reduce intra-medullary blood flow in patients with acute calculus renal colic. It is not yet established whether monomer LOCM or dimer LOCM have the lower renal toxicity (see earlier)3,4.

NERVOUS SYSTEM LOCMs are much more comfortable for cerebral arteriography in the conscious patient and are always preferred. Cerebral RCM photographic arteriography is being strongly challenged and partially displaced by CT and MR angiography (MRA).

Cerebral angiography Adverse reactions to RCM include dilatation of the external carotid arterial territory causing facial pain and heat. Damage to the blood–brain barrier may cause dangerous cerebral oedema, bradycardia and hypotension.

Myelography Myelography is being rapidly replaced by MRI (and to a lesser extent by CT) and very few myelograms are now performed in well-equipped imaging departments. If contrast medium myelography is essential, then non-ionic LOCM must be injected, for ionic contrast media (both HOCM and LOCM) are very toxic and often fatal when injected into the subarachnoid space.

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Myelography with oily products, such as iophendylate (Myodil, pantopaque – standard products in the 1940s–1980s) have been completely abandoned as it is now known that severe chronic adhesive arachnoiditis may follow their injection, sometimes causing considerable pain, disability and compression of the lumbosacral nerve roots with severe neurological deficit.

CARDIOVASCULAR SYSTEM Peripheral Arteriography The usual iodine concentration required for conventional film–screen angiography is about 300 mg I ml−1 contrast medium. Conventional HOCM (1500 mosmol kg−1 water) has been completely displaced for peripheral arteriography by LOCM (600–700 mosmol kg−1 water), because the latter causes much less warmth, discomfort, pain and movement. LOCM (Table 2.3) permits almost painless angiography in all territories, usually eliminating discomfort, movement and the need for general anaesthesia.

Aortography Injections of LOCM at the 300–400 mg ml−1 iodine concentration are greatly preferred as they cause much less discomfort and vasodilatation. Coronary angiography HOCMs (e.g. Urografin 76 per cent) with physiological levels of sodium and which do not bind avidly to serum calcium (related to buffer agents) had a good reputation for selective coronary angiography, but LOCMs are even safer for they cause less marked haemodynamic, myocardial and physiological changes and depression. Coronary angioplasty (PTCA) and stent placement During therapeutic coronary angioplasty, the intima and the arterial wall are deliberately torn and stretched by the intra-arterial balloon considerably increasing the risk of local arterial occlusion. The most important controllable factor in reducing this risk is the expertise of the operator (see later). Many investigators utilize full heparinization of the patient and LOCM injections.

Peripheral Venography Venography of the leg for possible deep vein thrombosis (DVT) is the most frequent venographic study. The procedure is performed by injecting RCM into a small vein of the foot, with the leg dependent and tourniquets restricting peripheral venous return. If the deep veins are already compromised and partly thrombosed, peripheral venography of the leg is a potentially dangerous procedure, as both deep and superficial venous return from the leg are compromised, and some cases of venous gangrene due to venous endothelial damage and thrombosis have been induced by attempted venography. LOCM is strongly advised because of its lower osmolality and its less irritant effect on the venous endothelium. Endothelial contact time should be reduced to the minimum by washing out with saline, massaging and exercising the leg immediately after satisfactory radiographs have been obtained. Termination of the injection of contrast medium must be seriously considered if the injection causes pain or if the deep veins are seen to be extensively thrombosed, for all RCM may induce thrombophlebitis. Although leg venography is still the ‘gold standard’ for imaging DVT, it is being increasingly replaced by compression ultrasound (without contrast medium).

Cardiac and Coronary Angiography Intracardiac injections LOCM injections are much preferred as they cause less disturbance of cardiac function, depression of myocardial contractility, peripheral vasodilatation, hypervolaemia, systemic hypotension and ECG changes. They are also much better tolerated subjectively than HOCM. Pulmonary angiography LOCM injections should be used for pulmonary angiography as they cause less elevation of the pulmonary artery pressure, coughing, movement and discomfort. Separate unilateral pulmonary artery injections should replace main stem pulmonary artery injection.

CONTRAST MEDIA, BLOOD COAGULATION AND THE ENDOTHELIUM6 In the 1990s it was confirmed that unlike ionic LOCM, non-ionic contrast media do not strongly inhibit thrombin formation or fibrin polymerization (with the consequent coagulation of blood), nor do they significantly inhibit platelet activity and aggregation. LOCMs do not inhibit blood coagulation as markedly as do HOCMs. Despite this disadvantage, LOCMs are now utilized for almost all angiographic and angioplasty procedures. Meticulous care must be taken to prevent catheter/guide wire thrombosis or sludging by constant perfusion of heparinized saline, continuous trans-catheter pressure recording and preventing blood entering the syringe (preferably plastic, not glass). Some angiographers routinely heparinize the patient for angiography and most angiographers do so during angioplasty. The production of many vaso-active physiological agents by the endothelium when exposed to RCM is discussed earlier6,11.

PREVENTION OF ADVERSE REACTION TO RCM There is no known premedication regimen that can be relied upon to eliminate completely the risk of a severe adverse reaction to RCM. Several strategies for prophylactic premedication have been suggested, usually involving the use of (some or all of ) corticosteroids, antihistaminics, H1 and H2 antagonists and ephedrine. The present author (Grainger) strongly recommends that in all patients considered to be at increased risk of adverse reactions (Table 2.4), alternative imaging should always be considered; premedication with corticosteroids (and possibly with antihistaminics) should be administered for the 12–24 h

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before and after injection of RCM, and that LOCM (rather than HOCM) should be used in all these patients in whom further administration of contrast medium is considered to be essential.

ETHICAL, FINANCIAL AND COST–BENEFIT CONSIDERATIONS HIGH- VERSUS LOW-OSMOLAR CONTRAST MEDIA The incidence of mild, moderate and severe adverse reactions is markedly less frequent with LOCM than with HOCM7 (Table 2.3). In the late 1990s, the only factor inhibiting the total abandonment of HOCM and its replacement by LOCM was financial, because LOCM was considerably more expensive. Recent major reduction in LOCM prices (industrial competition, expiration of patent protection, improved manufacturing techniques and user pleadings) has markedly changed this late 1990s position. The strong recommendation must be that LOCM should be used for intravascular injection in all patients, especially those at increased risk. Unfortunately there are still many impoverished countries where total conversion to LOCM is not a major medical requirement and is not financially justified, and they reasonably continue to use HOCM.

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INTRAVASCULAR CONTRAST MEDIA FOR RADIOLOGY, CT AND MRI

CONCLUSION Both HOCM and LOCM are probably the safest drugs used in clinical medicine. It is not yet proven that LOCM reduces the fatality rate compared to HOCM. There is no doubt that LOCMs produce less discomfort on IV injection, less pain on intra-arterial injection, less physiological and haemodynamic disturbance and fewer adverse reactions to contrast agents than does HOCM. Recent major reduction in cost has encouraged the use of LOCM in about 80–90+ per cent of all intravascular radiographic injections in Europe, North America, Australasia, Japan and many other countries. Because of the concern about the ill-effects of radiation exposure, it is probable that in the next few years, radiological methods will be progressively replaced by MRI using blood pool and other specific contrast agents. The Manual on Contrast Media, 4th edition, published by the American College of Radiology in 1999 and The European Society Guidelines on Contrast Media (2003) Version 3.0 are excellent practical handbooks. Two encyclopaedic textbooks (Dawson, Cosgrove and Grainger6 and Thomsen, Muller and Mattrey11) dealing exclusively with contrast agents (radiological, US and MRI) with contributions from internationally respected clinicians and industry scientists were published in 1999 and are recommended if further information is required.

MRI CONTRAST MEDIA EXTRACELLULAR MRI CONTRAST AGENTS Henrik S. Thomsen and Sameh K. Morcos Extracellular MRI contrast agents are diagnostic pharmaceutical compounds containing paramagnetic metal ions that affect MR signal properties of surrounding tissues. Paramagnetic agents are mainly positive enhancers that reduce the T1 and T2 relaxation times and increase tissue signal intensity on T1-weighted MR images and have almost no effect on T2-weighted images. Initially it was thought that MRI did not need contrast media due to the excellent soft tissue contrast compared to CT. However, in 2005, up to 40 per cent of all MRI examinations worldwide were performed with contrast agents. The majority of contrast-enhanced MRI examinations are performed with extracellular, not organ-specific agents. However, organ-specific agents (e.g. liver, lymph nodes) have recently become available (see later section). Almost 25 years ago, copper (Cu2+), manganese (Mn2+) and gadolinium (Gd3+) were considered as potential paramagnetic ions to be used for MRI. However, Gd is the most powerful, with seven unpaired electrons, but unfortunately gadolinian is the most toxic of those ions and therefore it is necessary to encapsulate it by a chelate. The first agent to be introduced into clinical practice was the gadolinium diethylene triamine pentacetic acid salt (GdDTPA). The chelate DTPA was already known from its use in nuclear medicine (99mTc-DTPA). Several extracellular Gd

chelate MRI contrast agents are currently available for clinical use in addition to Gd-DTPA. Some of these agents are ionic with high osmolality; others are non-ionic with osmolality twice that of the blood (Table 2.5). All the extracellular MRI contrast agents have a high tolerability in animal studies and good relaxation properties. In this chapter the pharmacokinetics, clinical use and safety of these agents will be discussed, including possible interaction with clinical tests and their use for radiographic examinations.

PHARMACOKINETICS The pharmacokinetics of all extracellular MRI contrast agents with the exception of gadobenate dimeglumine (Gd-BOPTA) are similar to iodinated water-soluble contrast media. They have low molecular mass and after IV injection they are rapidly diffused into the interstitial extravascular space. Reverse diffusion into the intravascular compartment also occurs and a state of equilibrium between diffusion in and out is usually reached within 2 h. Gd chelates are eliminated unchanged from the intravascular compartment by passive glomerular filtration and by 24 hours >95 per cent of the injected dose is excreted in urine with normal renal function. A very small amount (<0.1 per cent) is eliminated via faeces. The biological half-life is approximately 1.5 h. Extracellular MRI contrast agents do not cross the intact specialized vascular blood–brain barrier. They do not bind to proteins or receptors and there is no intracellular penetration.

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Table 2.5 CURRENTLY AVAILABLE EXTRACELLULAR MRI CONTRAST AGENTS

Generic agent name

Gadopentate Gadoterate dimeglumine meglumine (Gd-DPTA) (Gd-DOTA)

Gadodiamide (Gd-DPTA-BMA)

Gadoteridol (Gd-HP-DO3A)

Gadobenate dimeglumine (Gd-BOPTA)

Gadobutrol (Gd-BT-DO3A)

Gadovesetamide Gd-DTPA-BMEA

Year of introduction

1988

1989

1993

1994

1998

1998

2000

Molecular structure

Liner, ionic

Cyclic, ionic

Linear, non-iconic

Cyclic, non-ionic

Linear, ionic

Cyclic, non-ionic

Linear, non-ionic

Molecular concentration

0.5 mmol/L

0.5 mmol/L

0.5 mmol/L

0.5 mmol/L

0.5 mmol/L

1 mmol/L

0.5 mmol/L

Osmolality mosmol/kg H2O

1960

1350

650

630

1970

1600

1110

Viscosity (mPas at 37°C)

2.9

2.0

1.4

1.3

5.3

4.96

2.0

*Thermodynamic stability constant

22.1

25.8

16.9

23.8

22.6

21.8

16.6

*Conditional stability

18.1

18.8

14.9

17.1

18.4

N/A

15

*Kinetic stability (dissociation T ½ *t pH 1.0)

10 min

>1 month

35 s

3h

N/A

24 h

N/A

**Amount of excess chelate (mg/ml)

0.40

None

12.00

0.23

None

N/A

28.40

*The higher the figure, the higher is the stability of the contrast agent. **The presence of excess chelate is an indirect marker of the instability of the contrast agent. The higher the figure, the higher the instability of the agent. N/A, not applicable.

Unlike all other Gd chelates, gadobenate dimeglumine has a capacity for weak and transient protein binding and is eliminated through both the renal and hepatobiliary pathways. The hepatic uptake represents 2–4 per cent of the injected dose. It behaves as a conventional extracellular contrast agent in the first minutes following IV administration and as a liver-specific agent in a later delayed phase (40–120 min after administration) when it is taken up specifically by normal functioning hepatocytes.13

CLINICAL USE OF EXTRACELLULAR MRI CONTRAST AGENTS These agents accumulate in tissues with abnormal vascularity (malignant and inflammatory lesions) and in regions where the blood–brain barrier is disrupted. Owing to their rapid equilibration in the interstitial space of both normal and tumour/ inflammatory tissues, the use of dynamic MR imaging after bolus injection makes the best use of the narrow imaging window with a transiently increased tumour/inflammatory to normal tissue contrast. Generally the recommended dose for clinical use is 0.1 mmol kg-1 body weight but up to 0.3 mmol kg-1 may be used, particularly in MRA.13–15

anaphylactoid reaction, which are observed with radiographic iodinated contrast media, may also be seen with extracellular MRI contrast agents but the incidence is much lower. The incidence of mild adverse effects is less than 5 per cent. Life-threatening reactions are very rare, with an incidence around 1:100 000. Fatal reactions may however occur but are extremely rare.13

CONTRAST MEDIUM-INDUCED NEPHROPATHY Extracellular MRI agents are more nephrotoxic than iodinated contrast media in equimolar doses. However, nephrotoxicity after contrast-enhanced MRI examinations is not common even in patients with renal disease as the dose of Gd-based contrast medium required for a diagnostic MRI examination is small in comparison to the doses of iodinated contrast media that are routinely used for CT or other radiographic examinations. However, in patients with marked reduction in renal function (particularly due to diabetes mellitus), contrast medium-induced nephropathy may occur even at the standard doses of extracellular MRI contrast agents (see later section on NSF).16

SAFETY OF EXTRACELLULAR MRI CONTRAST AGENTS

NEPHROGENIC SYSTEMIC FIBROSIS (NSF)

Extracellular MRI contrast agents are well tolerated with a low incidence of adverse effects. There are no differences in the safety of the various agents except when it comes to extravasation in soft tissue; then the osmolality is important and high osmolar agents are likely to induce more local damage. In blood, the osmotic load of all Gd-based contrast media is very low, compared to iodinated contrast media, because only a small amount of the contrast agent is required to produce a diagnostic MRI examination. Adverse reactions such as nausea/vomiting, urticaria, bronchospasm, laryngeal oedema, hypotension and generalized

Recently (2006 and 2007) several reports have suggested that the administration of extracellular gadolinium based contrast agents (Gd-CA) may trigger the development of NSF in patients with advanced renal impairment (GFR < 30 ml/ min) or on dialysis. The incidence of this condition in this group of patients is around 3–5% and the onset varies from a few days to 3 months after administration of Gd-CA. The disease is characterised by scleroderma- like skin changes that mainly affect the limbs and trunk. The induration of the skin can progress to cause flexion contracture of joints. The fibrotic changes may also affect other organs such as mus-

CHAPTER 2

cles, heart, liver and lungs. The disease can be aggressive, in some patients leading to serious physical disability or even death. The considerable majority of reported NSF complications followed the administration of non-ionic gadodiamide, but in a few patients NSF followed injection of gadopentate dimeglamine, gadovestomide and other gadolinium contrast agents.18–21 The stability of the binding of the gadolinium ion (Gd+++) within the chelate could be an important factor in the pathogenesis of NSF and may explain the strong association between gadodiamide and this condition. The stability of the Gd-chelates is influenced by the configuration of the molecule whether linear or macrocyclic and ionicity. Macrocyclic chelates offer better protection and binding to Gd+++ in comparison to the linear molecules. Replacement of carboxyl group by a less strongly coordinating group to produce a non ionic Gd-CA also weaken the binding of the chelate to Gd+++ particularly in the non ionic linear molecule. Stability measurements (Table 2.5) indicate that the least stable molecules are the non ionic linear chelates, and most stable is the ionic cyclic chelate. In vivo, transmetallation with endogenous ions leading to the release of the highly toxic gadolinium ions (Gd+++) is likely to occur with unstable Gd-CA particularly if they remain in the body for a long period, as is the case in patients with advanced renal impairment or on dialysis. Recent reports demonstrated gadolinium in the skin lesions of patients with NSF strengthening the hypothesis that free gadolinium play an important role in triggering this condition. It is advisable at this stage of our understanding to *avoid the administration of non-ionic linear chelates in patients with advanced renal impairment (GFR <30ml/min) including those on dialysis. A more stable Gd-CA such as macrocyclic Gd-chelates might prove less hazardous if contrastenhanced MRI examination is thought to be necessary in such a group of patients, including pregnant and lactating women with end-stage renal failure. There is no evidence that immediate haemodialysis after administration of Gd-CA protects the patient against NSF*.20, 21

PREVENTION OF ADVERSE REACTIONS It is not recommended to give any prophylaxis before administering extracellular MRI agents. However, if there has been a previous hypersensitivity-like reaction to a gadolinium-based compound, an alternative imaging investigation should be considered. If the administration of a Gd-based contrast agent is deemed essential, a different MRI contrast agent to the one incriminated in the previous reaction should be used and corticosteroid prophylaxis might offer some protection. There is no cross-reaction between extracellular MRI contrast agents and radiographic iodinated contrast media. Patients with multiple allergies may react to both iodinated and gadolinium-based contrast media.13

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USE OF EXTRACELLULAR MRI CONTRAST AGENTS IN SPECIAL GROUPS OF PATIENTS Patients on dialysis Extracellular MRI contrast agents are easily removed by haemodialysis or continuous ambulatory peritoneal dialysis. There is no need to schedule the MRI examination in relation to the dialysis session.

Pregnancy and lactation Mutagenic and teratogenic effects have not been described after administration of gadolinium-based contrast media.Therefore, if the MR examination is deemed necessary, an extracellular MRI contrast agent may be given to a pregnant female. Only tiny amounts of gadolinium-based contrast media given to a lactating mother reach the milk and only a minute proportion is absorbed from the baby’s gut. Therefore, stopping breast-feeding for 24 h following administration of a gadolinium-based contrast media is not warranted.13

INTERACTION BETWEEN EXTRACELLULAR MRI CONTRAST AGENTS WITH CLINICAL TESTS Extracellular MRI contrast agents may cause spurious hypocalcaemia as they can interfere with the assay method of calcium measurement. Awareness of this effect on calcium measurements by some MRI contrast agents is important in order to avoid incorrect and potentially hazardous treatment. Caution should be exercised when using colorimetric assays for angiotensin-converting enzyme, calcium, iron, magnesium, total iron binding capacity and zinc in serum samples of patients who have recently received Gd-based contrast media. Therefore, biochemical assays are better performed before contrast media injection, or delayed for at least 24 h afterwards or longer in patients with renal impairment. Urgent laboratory tests performed on specimens collected shortly after contrast media injection should be carefully assessed. Accuracy of unexpected abnormal results should be questioned and discussed with colleagues from the hospital laboratories.22

USE OF EXTRACELLULAR MRI CONTRAST AGENTS FOR RADIOGRAPHIC EXAMINATIONS Gadolinium attenuates X-rays and hence their use for radiographic examinations instead of iodinated contrast media has been advocated in patients with history of serious adverse reactions to iodinated contrast media and in patients who shall undergo imminent thyroid treatment with radioactive iodine, providing the renal function of these patients is normal. In patients with renal impairment, Gd-based contrast media should not be used for radiographic examinations as they can induce nephrotoxicity at the doses required to produce adequate radiographic enhancement. It is also important to point out that Gd-based contrast agents have not been approved for radiographic examinations or intra-arterial injection and the doses approved for MRI examinations will provide only a limited enhancement of radiographic examinations.16

* NB: The drug safety authorities in Europe have recently recommended that the non-ionic linear chelate gadodiamide agents should not be used in patients with GFR < 30 ml/min, on dialysis or those who have had or who are waiting for liver transplantation. Gd-based agents should be used in neonates and infants up to one year of age, only after careful consideration.

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CONTRAST-ENHANCED MR IMAGING OF THE LIVER AND PANCREAS Dow-Mu Koh The IV contrast media currently available for MR imaging the liver can be classified into three broad categories. The types, actions and common side effects of these contrast media are summarized in Table 2.6.

NON-SPECIFIC EXTRACELLULAR GADOLINIUM CHELATES These contrast compounds distribute freely in the extracellular space and shorten the T1-relaxation time, resulting in T1 signal enhancement. Dynamic MR imaging of the liver is performed after bolus IV contrast injection. The use of fat-saturated three-dimensional (3D) volume interpolated MR imaging (e.g. VIBE, THRIVE, FAME) allows high spatial resolution imaging of the entire liver to be acquired in a 20-second breath hold. Imaging is repeated in the arterial, portovenous and parenchymal phases of liver enhancement.

ARTERIAL PHASE In the hepatic arterial phase (typically 20–30 s after IV injection), both benign (e.g. focal nodular hyperplasia, adenoma and haemangiomas) and malignant lesions (e.g. hepatocellular carcinoma and metastases arising from neuroendocrine tumours) may appear hypervascular. Further differentiation depends on the pattern and rate of contrast medium change within the lesion. For example, haemangiomas typically dem-

Table 2.6

onstrate peripheral nodular enhancement23 (Fig. 2.2), while focal nodular hyperplasia shows intense uniform enhancement. Heterogeneous irregular enhancement is typical of hepatocellular carcinoma.

PORTOVENOUS PHASE The portovenous phase occurs at 60–90 s after IV contrast medium administration during which maximum liver parenchymal enhancement is achieved.This increases conspicuity of hypovascular (e.g. hypovascular metastases) and non-vascular (e.g. hepatic cysts) lesions. During this period, benign lesions such as adenoma and focal nodular hyperplasia may fade to isointensity with the liver. Hypervascular metastases can also appear isointense. The pseudo-capsule and internal septae associated with hepatocellular carcinoma become more prominent.

INTERSTITIAL PHASE This phase occurs from approximately 90 s to 5 min after administration of IV contrast agent. Differential washout of contrast agent from the hepatic parenchyma and focal liver lesions aids lesion characterization. In this phase, haemangiomas continue to enhance centripetally (Fig. 2.2); adenomas and focal nodular hyperplasia are typically isointense with the liver, although late enhancement of the central scar of focal nodular hyperplasia may be visible24. Small hepatocellular carcinoma fades to contrast hypointensity. Delayed enhancement of cholangiocarcinoma becomes apparent.

SUMMARY OF CONTRAST MEDIA AVAILABLE COMMERCIALLY FOR MR IMAGING OF THE LIVER

Classes Examples

Nonspecific extracellular gadolinium chelates

Hepatocyte selective gadolinium chelates

Magnevist™ Omniscan™

Multihance™ Primovist™

Teslascan™ (Manganese based)

Endorem™ (Iron oxide based)

Resovist™ (Iron oxide based)

Non-gadolinium-based contrast media

Dotarem™ Prohance™ Gadovist™ Constituents

Low molecular weight gadolinium chelates

Low molecular weight gadolinium chelates: Gd-BOPTA Gd-EOB-DTPA

Magafodipir trisodium (MnDPDP)

Superparamagnetic iron oxide particles (ferrumoxides 80–150 nm in size)

Superparamagnetic iron oxide particles (ferucarbotran 60 nm in size)

Administration

Intravenous bolus

Intravenous bolus

Intravenous infusion

Intravenous infusion

Intravenous bolus

Action

Distributes freely in the extracellular space

Initially distributes freely in extracellular space but undergoes hepatic excretion

Selective uptake by hepatocytes and excreted into bile ducts

Selective uptake by Kupffer cells

Selective uptake by Kupffer cells

Imaging

T1-weighted imaging

T1-weighted imaging

T1-weighted imaging

T2*-weighted ± T1-weighted imaging

T2*-weighted ± T1-weighted imaging

Common side effects

Rash

Rash

Nausea and abdominal discomfort

Back pain

Back pain

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INTRAVASCULAR CONTRAST MEDIA FOR RADIOLOGY, CT AND MRI

Figure 2.2 Dynamic T1-weighted MR images obtained using a 3D volume interpolated technique after IV bolus of non-specific extracellular gadolinium chelate. Imaging was performed in the (A) arterial, (B) portovenous, (C) interstitial and (D) delayed (10 min) phases of hepatic enhancement. Note the discontinuous peripheral nodular enhancement of the lesion (arrows) in segment VI of the right lobe, which increases centripetally, typical of a haemangioma. Lesion characterization is possible by evaluating the rate and pattern of contrast enhancement.

HEPATOCYTE-SELECTIVE GADOLINIUM CHELATES These are also administered intravenously as a bolus. However, a proportion of the contrast medium is selectively excreted by hepatocytes into the bile ducts. The percentage of hepatic uptake is 5 per cent for MultiHance and up to 50 per cent for Primovist. Dynamic T1-weighted MR imaging is performed in the arterial, portal and interstitial phases after injection, similar to

the nonspecific extracellular gadolinium contrast media. However, a delayed hepatic enhancement phase is observed (Fig. 2.3).This occurs at 15–30 min for Primovist and approximately 1–3 h for MultiHance. MultiHance has been used to distinguish between adenomas and focal nodular hyperplasia25. As adenomas do not contain normal biliary radicals, they appear hypointense at delayed T1weighted imaging. Metastases (nonhepatocellular) and hepatocellular carcinoma (poorly functioning hepatocytes) also frequently appear hypointense. By comparison, focal nodular

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Figure 2.3 Dynamic T1-weighted MR images using a 3D volume interpolated technique after IV bolus of hepatocyte selective gadolinium chelate (Gd-EOB-DTPA) in a normal liver. Imaging was performed in the (A) arterial, (B) portovenous and (C) interstitial phases of contrast enhancement, similar to the use of non-specific extracellular gadolinium chelates. However, a (D) delayed hepatic phase (15–30 min) was observed with intense hepatic parenchymal enhancement accompanied by clearance of contrast medium from hepatic vasculature due to hepatocyte contrast excretion. Lesions of non-hepatocellular origin (e.g. metastases) would increase in conspicuity during this phase

hyperplasia appears hyper- or isointense at delayed imaging. Primovist results in a much more intense parenchymal enhancement due to increased biliary excretion. However, its role in detection and characterization of focal liver lesions is still under investigation.

NON-GADOLINIUM CONTRAST MEDIA Unlike the gadolinium-based contrast media, dynamic imaging is not performed with the use of these MRI agents. These

contrast media are actively transported into liver and can be divided into two subcategories: 1 Hepatocyte-selective manganese-based contrast medium. Teslascan (MNDPDP) is infused intravenously and its selective uptake by hepatocytes results in intense T1 liver signal enhancement at approximately 30 min, which persists for several hours.26 There is normally complete clearance of contrast agent from the liver by 24 h. The contrast medium is also taken up by the pancreas, adrenal glands and the pituitary.

CHAPTER 2

2 Cysts, haemangiomas and metastases from non-neuroendocrine primary tumours remain hypointense at T1-weighted imaging. There is variable enhancement of focal nodular hyperplasia, adenomas and hepatocellular carcinomas.26 Metastases from neuroendocrine tumours also show enhancement by an uncertain mechanism. It has been found that delayed imaging at 24 h may be useful for detection and characterization of colorectal hepatic metastases due to delayed clearance of contrast material surrounding the metastases27 (Fig. 2.4).

Reticulo-endothelial system targeted iron-based contrast media These comprise small paramagnetic iron oxide particles that are administered intravenously. The iron accumulates within

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INTRAVASCULAR CONTRAST MEDIA FOR RADIOLOGY, CT AND MRI

Kupffer cells in the normal liver, resulting in reduction in signal intensity of the liver at T2*-weighted gradient-echo imaging. Lesions containing Kupffer cells demonstrate signal reduction whereas lesions that are Kupffer cell depleted would remain high signal. Immediately following IV administration, the contrast agent also causes T1 relaxation and early T1-weighted imaging may be performed. Iron oxide-based contrast media have been shown to increase the detection of small hepatic metastases prior to surgery.28 (Fig. 2.5). The contrast agent can also be used to demonstrate the extent of cholangiocarcinomas in the liver. However, its role in distinguishing between hepatocellular carcinoma, focal nodular hyperplasia and adenomas is less certain, as variable amounts of Kupffer cells in these lesions results in considerable overlap of their imaging appearances.

Figure 2.4 T1-weighted MR imaging of the liver obtained in two different patients with colorectal hepatic metastases at 24 h after MnDPDP contrast (selective hepatocytic) administration. Note the clearance of contrast medium from the normal liver. However (A) geographic enhancement and (B) rim enhancement can be seen adjacent to metastases believed to be related to delayed clearance as a result of regional cholestasis.

Figure 2.5 Gradient T2*-weighted MR imaging of liver obtained (A) before and (B) after paramagnetic iron oxide (Kupffer cell) contrast medium injection. The metastasis (no Kupffer cells) in the right lobe of the liver increases in conspicuity after contrast agent administration due to signal reduction in the normal liver.

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CONTRAST-ENHANCED MR IMAGING OF THE PANCREAS An MR cholangio-pancreaticogram (MRCP) is usually performed in conjunction with MR imaging of the pancreas.

GADOLINIUM CONTRAST-ENHANCED IMAGING Low molecular weight extracellular gadolinium chelates are most frequently used to characterize pancreatic lesions. Before contrast enhancement, unenhanced MR imaging should be carefully reviewed for focal signal abnormalities. In particular, fat-suppressed T1-weighted imaging is useful for identifying low signal intensity foci within the pancreas. T2-weighted imaging is helpful for visualizing high signal cystic lesions. Most solid pancreatic tumours appear hypointense or isointense on T1-weighted imaging, and isointense or mildly hyperintense on T2-weighted imaging. Following bolus contrast injection, T1-weighted MR imaging is performed. The use of a fat-suppressed 3D volume interpolated T1-weighted imaging sequence, performed in repeated breath-holds, is particularly useful for detecting small focal lesions29,30 and for assessing vascular involvement.

In the arterial phase, there is enhancement of hypervascular neuroendocrine tumours of the pancreas. Both cystic and solid hypovascular pancreatic carcinomas are also well demarcated in the arterial phase (Fig. 2.6). In the portovenous phase, hypervascular tumours usually fade to isointensity. Despite the superior contrast resolution of MR imaging, it is still controversial whether the technique is better than multi-detector CT for the detection and characterization of focal pancreatic lesions. However, there is gathering enthusiasm for using MR imaging to assess for tumour resectability.31

Mn-DPDP-ENHANCED MR IMAGING Mn-DPDP is also excreted by the normal pancreatic parenchyma. Anecdotal reports have demonstrated the potential utility of the technique in detecting pancreatic carcinoma as relative hypointense lesions against the enhancing pancreas at T1-weighted imaging4. However, no studies have shown a clear advantage over the extracellular gadolinium chelates. Furthermore, the uptake of the contrast agent (Mn-DPDP) by neuroendocrine tumours can result in reduction of contrast between the tumour and the enhancing pancreatic tissue.

Figure 2.6 Carcinoma of the pancreas (A) T2-weighted axial image showing a mildly hyperintense lesion within the pancreatic tail (arrows), which appears hypointense (arrows) on the (B) fat-suppressed T1-weighted MR image. (C) Immediately following IV extracellular gadolinium contrast administration, the mass is well seen as a hypointense mass (arrows) against the enhancing pancreatic parenchyma using a fat-suppressed T1-weighted imaging sequence.

MRI CONTRAST AGENTS AND TECHNIQUES FOR CONTRAST-ENHANCED MR ANGIOGRAPHY Giles Roditi The flow-sensitive MRA techniques of time-of-flight and phase contrast imaging first developed in the 1980s rely on maximization of signal difference between flowing blood in the vessel lumen versus static vessel wall and surrounding tissues. These have now in large part been superseded by contrast-enhanced MRA (CE-MRA) techniques, wherein gadolinium-based paramagnetic contrast agents increase the

blood pool signal by causing a profound shortening of the T1 relaxation time. A fast T1-weighted gradient-echo pulse sequence is used, specifically designed with short TR, short TE and RF spoiling to produce an image with low background signal while vascular contrast depends on T1 shortening due to the gadolinium rather than flow effects. Since CE-MRA is not susceptible to in-plane flow saturation, it

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INTRAVASCULAR CONTRAST MEDIA FOR RADIOLOGY, CT AND MRI

Blood T1(msec)

T1 (msec)

can be performed in the plane of the vessel, increasing the practicable field of view and reducing acquisition time32,33. Breath-hold techniques are desirable (i.e. a volume acquired in 30 s or less) as they improve diagnostic accuracy for branch vessels. Most modern MRI machines are now capable of imaging a clinically useful MRA volume in a comfortable breath-hold, and parallel imaging techniques are starting to allow time-resolved imaging.

•

1200

Fat T1 (270 msec)

1000

Gadofosveset Steady state (150 msec)

800 Contrast Bolus First pass 0.2 mmol/kg (50 msec)

600 400

CONCEPTS OF CONTRAST MEDIUM ADMINISTRATION Delivery of contrast agents (Gd-CA) for CE-MRA by IV injection is more complex than for ‘simple’ MRI studies and necessitates an understanding of their mode of action and kinetics to maximize contrast effect and minimize artefacts. In order to provide contrast, a sufficient concentration of gadolinium contrast agent must be achieved in the vessels of interest to reduce the T1 of blood to less than 270 ms (equivalent to the T1 of fat for a 1.5 T system). The effect of a contrast agent on a tissue is given by the equation:

200 0 0

1

2

3

4

5

6

7

8

9 10

[Gad] (mMolar) Figure 2.7 Graph of whole blood T1 value as a function of gadolinium concentration compared to T1 of fat (highest signal background). Note how T1 of blood falls from 1200 ms to below that of fat (270 ms) above 0.5 mmolar gadolinium concentration. Approximate T1 of blood during first pass contrast bolus passage (50 msec) and in steady state for Gadofosveset BPCA (150 msec) is also shown.

R' = R + rC where R is the original relaxation rate (1/relaxation time), r is the agent’s specific relaxivity rate, C its concentration and R' the resultant new relaxation rate (Fig. 2.7). Blood has a T1 of 1200 ms, hence with a conventional extracellular fluid (ECF) gadolinium contrast agent with a T1-specific relaxivity rate of 4 mmol-1 s−1 a concentration of 1 mmol l-1 will reduce this to approximately 200 ms, and 2 mmol l-1 will reduce it to 100 ms – a useful contrast difference. Unfortunately, at concentrations above 5 mmol l−1 the T2* shortening effects of gadolinium come to predominate and signal on a T1w (T1-weighted) sequence rapidly declines. These T2 and T2* effects also act over a wider region than the T1 effects that are very local to the agent; this explains the artefacts seen from residual ‘neat’ venous contrast after arm injection, resulting in loss of signal in adjacent arteries on aortic arch and subclavian studies. To help combat this and ensure efficient contrast use, IV contrast medium should be immediately flushed with a suitable volume of saline (e.g. 30 ml) at the same injection rate to move the bolus into the central venous pool, where it is propelled on to the right heart, maximizing efficiency of the dose. To further minimize these effects, it is recommended that contrast agents are administered via a vein in the right arm since any artefact from the left brachiocephalic vein (where flow tends to be sluggish) across the origins of the aortic arch branches may seriously degrade their depiction. A right arm injection is also beneficial even when studying territories remote from the arch, since the route is shorter with less contrast travel time variability than with left arm intravenous injections (Fig. 2.8). Contrast agent injection for MRA studies is simplified by the use of an MRI compatible pump injector ensuring reproducibility while allowing medical and radiographic staff

Figure 2.8 First-pass arterial phase MRA aortic arch and subclavian arteries – normal study. Note the signal dropout in the right axillary artery where the axillary vein crosses and more medially in the right subclavian artery where the subclavian vein is directly adjacent due to T2* effect of residual concentrated gadolinium contrast in veins.

to remain in the MRI control room. The phenomenon of signal abolition at high agent concentrations also informs the use of contrast for ‘direct’ MR venography when injecting into the relevant veins, e.g. pedal injection for a lower limb study or for arteriovenous fistula studies. For this, an injection of diluted contrast is required and this also applies for direct MR arteriography studies.

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Peak arterial blood pool contrast concentration after IV injection is proportional to contrast agent flux (which is the product of agent concentration and rate of injection) and inversely proportional to cardiac output. The effect of differing cardiac output is lessened at injection rates over 1 ml s−1 and so, on modern systems, apart from trying to minimize patient anxiety, cardiac output is generally not directly manipulated (reduced) and varying the rate of contrast administration is used to maximize contrast efficacy. ‘Double dose’ (0.2 mmol kg−1) is usually employed for single station examinations (e.g. carotid, thoracic aortic, renal arteries etc.) and to maximize efficient use of contrast medium, the bolus should last for the duration of the CE-MRA sequence covering central and peripheral k-space data acquisition. For example, for a typical 20-second breath-hold sequence then 30 ml of a 0.5 mol l−1 ECF agent could be injected at a rate of 1.5 ml s−1. In practice somewhat higher rates are used as the bolus injected peripherally becomes ‘stretched’ as it passes through the pulmonary circulation and heart, hence rates of 2.0–2.5 ml s−1 are used. For a 1.0 mol l−1 agent (such as gadobutrol) half the volume is required for the same gadolinium flux at any given rate. If this reduced volume is injected at the same rate as a 0.5 mol l −1 medium, this will result in a higher intra-arterial concentration at the bolus peak. However, this may adversely affect image quality if the bolus no longer covers the whole of the acquisition in terms of peripheral k-space lines, hence a 30 per cent lower rate is recommended. Optimum timing of the data acquisition is required to ensure peak concentration in the vessel of interest, and therefore the most profound T1 shortening of blood, coincides with the contrast medium defining central lines of k-space data. Most MRI systems are now supplied with automated triggering and/ or fluoroscopic monitoring simplifying timing with centrically encoded pulse sequences acquiring low frequency contrast defining k-space profiles at the beginning of data acquisition34. These doses have generally been shown to be safe, particularly in terms of nephrotoxicity, even in renal impairment35,36 though the potential of some agents (Gadodiamide and Gadoversetamide) to interfere with assays for serum calcium may result in a spurious ‘pseudohypocalcaemia’ and this phenomenon is potentiated/prolonged in patients with chronic kidney disease. However, the recent reports of an association between Gd-CA (Contrast Agent) use in patients with severe renal impairment (particularly those on dialysis) and the subsequent development of nephrogenic systemic fibrosis (NSF) introduces a note of caution. Whether gadolinium is the specific and sole necessary factor triggering NSF is not yet resolved at the time of writing. Most patients with renal impairment do not develop the disorder, following Gd-CA,with an incidence of 3–5% in dialysis patients administered Gd-CA.37 While the exact relationship and any causality are not yet determined this is a particularly important issue for CE-MRA, which has hitherto been embraced as a safe means of investigation in these patients who are prone to vascular problems. This is especially so since there appears to be a dose-response relationship with

development of NSF more likely with high dose examinations. Despite no conclusive evidence of protection, the current ACR guidelines advise dialysis where practicable in such patients promptly after contrast administration.38,39 (Please see preceding section by Thomsen and Marcos).

PERIPHERAL RUN-OFF AND WHOLE-BODY CE-MRA CE-MRA is not restricted to a single station field of view with the use of stepping table techniques (analogous to those used in conventional angiography) and that allow scanning of the lower body arterial tree from abdominal aorta to feet40 with extended tables and whole body MRA. Mask images are used to allow subtraction of regions of high signal pre-contrast (such as subcutaneous and marrow fat) from the post-contrast images. Image subtraction is particularly crucial for the tibial station as the image scaling effect of high signal due to contrast agent is negligible within small crural arteries while both the subcutaneous and marrow fat are in close proximity to the vessels. Whereas a total dose of 0.2 mmol kg−1 standard (0.5 mol l−1) ECF gadolinium contrast is used for single station studies, this is increased to 0.3 mmol kg−1 for multiple-station CE-MRA of the aorta and lower limb arterial run-off. Biphasic injection protocols are also recommended, for example a rate of 1.5 ml s−1 for the first phase dropping to 0.8 ml s−1 for the second phase. This second phase lower rate helps to prolong the bolus and also reduce venous enhancement. The main problems with peripheral run-off CE-MRA revolve around the classic MRI trade-offs of spatial resolution versus signal-to-noise ratio (SNR) versus image acquisition time. There is also the drawback of so-called ‘venous contamination’ whereby the crural arteries may be obscured by early enhancement of calf veins. High spatial resolution is particularly critical for the small tibial arteries yet this is also the region where venous contamination is most likely, particularly in patients with Fontaine class III and IV critical ischaemia who demonstrate rapid arteriovenous shunting bypassing tissue diffusion/extraction. IV contrast agent travel time variability depends on passage from the venous compartment through the pulmonary circulation and heart. Once in the abdominal aorta, the leading edge of the contrast bolus travels at approximately 6 s per station down the lower limbs (groin, popliteal, ankle). However, the arterial-venous transit times at the calf vary depending on disease status with much more rapid transit in those with critical ischaemia and/or cellulitis compared to simple claudication.41 This has led to different strategies to optimize tibial arterial imaging. 1 Imaging the crural arteries first with a dedicated injection of contrast and a centric-encoded sequence designed for venous suppression, e.g. elliptic centric, CENTRA. Aortoiliac and femoral levels are subsequently imaged with standard stepping table technique and a second injection of contrast medium. 2 The imaging time window at the calf can be prolonged by venous compression techniques, whereby compression

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Table 2.7

•

INTRAVASCULAR CONTRAST MEDIA FOR RADIOLOGY, CT AND MRI

SUMMARY OF CONTRAST AGENTS AVAILABLE COMMERCIALLY SUITABLE FOR MR ANGIOGRAPHY43

Generic name

Trade name

Protein binding

Dose for MRA – mmol kg-1

T1 specific relaxivity in blood at 1.5 T – mmol−1 s−1

Gadopentate dimeglumine

Magnevist

Negligible

0.2–0.3

4.3

*

Gadobutrol*

Gadovist

Negligible

0.2–0.3

5.3

Gadoteridol

Prohance

Negligible

0.2–0.3

4.4

Gadoterate meglumine

Dotarem

Negligible

0.2–0.3

4.2

Gadodiamide

Omniscan

Negligible

0.2–0.3

4.6

Gadoversetamide

Optimark

Negligible

0.2–0.3

5.2

Gadobenate dimeglumine

Multihance

Weak

0.1–0.3

6.7

Gadofosveset trisodium

Vasovist

Strong

0.03–0.05

19

*Gadobutrol is formulated as a 1.0 mol l-1 preparation compared to 0.5 mol l-1 for the other ‘standard’ ECF gadolinium contrast agents with negligible or weak protein binding.

cuffs around the thighs inflated to subdiastolic pressures prolong arterial and venous transit times and also increase the venous volume of dilution (reducing contrast conspicuity in the venous compartment). 3 Another strategy is to image the calf vessels first with either fast thick slab 2D subtraction imaging or 3D time-resolved imaging, followed by a standard bolus chase 3D acquisition for the aortoiliac and femoropopliteal stations. 4 The use of parallel imaging and phased array coils at all stations with variable acquisition parameters can allow tailored and more rapid bolus chase such that venous contamination is minimized, e.g. following the aortoiliac station, the femoral segment is rapidly acquired with lower resolution and then highest resolution for the tibial station.42 These techniques listed are not mutually exclusive and the strategy used will depend upon MR machine capability, coil availability and operator experience/preference.

BLOOD POOL CONTRAST AGENTS Blood pool contrast agents (BPCAs – also known as persistent vascular agents) are specifically designed for vascular imaging through persistence in the intravascular space without the passage to extravascular interstitial extracellular fluid spaces that is the hallmark of conventional ECF agents (i.e. their vascular half-life is much greater than ECF agents). Blood pool persistence is achieved either by high molecular weight preventing diffusion through the capillaries or through protein binding to serum albumin, which has the same effect. BPCAs are either based on gadolinium compounds or ultrasmall paramagnetic iron oxides (USPIOs) though these latter iron based agents are as yet not commercially available for blood pool imaging. Gadobenate dimeglumine is predominantly an ECF fluid agent but its weak protein interaction provides higher relaxivity compared with standard 0.5 mol l−1 ECF agents and this weak (but still positive) protein binding also helps prolong blood pool residency. This results in a relatively higher signal intensity peak and longer duration of enhancement compared with standard agents for similar dose and injection rates. The safety profile of gadobenate dimeglumine is similar to the conventional ECF agents.

Gadofosveset has been specifically designed for CE-MRA as a blood pool agent with strong protein binding to albumin, up to 20 molecules of gadofosveset can bind to one molecule of albumin but in clinically used doses it is usually only one or two per albumin molecule. Binding to albumin is rapid with 75 per cent bound 1 min after IV injection, reaching a steady state with approximately 88 per cent bound. The high protein binding has two effects; firstly it markedly increases relaxivity (hence the lower doses/concentrations required) and secondly it results in the agent effectively being confined to the vascular space with little passage to extracellular fluid spaces and an increased half-life in the blood pool. This binding is reversible and the agent is ultimately excreted unchanged by the kidneys through glomerular filtration; it is also eliminated by routine renal dialysis, with 92 per cent cleared after three sessions. Gadofosveset has a safety profile similar to other gadolinium contrast compounds with a low rate of adverse reactions consisting of transient mild tingling sensations (particularly in the perineum), pruritus, flushing, headache and taste perversion. One possible serious adverse anaphylactoid event to gadofosveset was reported, resolving within 5 min. There is apparently no potential nephrotoxicity at clinically used doses nor interference with proper binding drugs such as warfarin. As a commercial BPCA for CE-MRA, gadofosveset provides currently unique properties, particularly through the very prolonged time window available for imaging as with a T ½ of 28 min (compared to 90s for Gad-DTPA) diagnostic imaging can be performed up to 1 h post injection. Arterial first-pass CE-MRA can be performed just as with the ECF agents, perhaps even with greater signal through the increased specific relaxivity of the agent. It is the possibilities of delayed imaging performed with high intravascular signal that are intriguing since this is not limited to the short time of a first-pass bolus and can therefore be performed at very high spatial resolution. Steady state imaging can start from 2 min after injection up to approximately 1 h. Imaging should be designed for high spatial resolution, ideally to create an isotropic dataset to allow for reformatting of images in any plane. In order to obtain best contrast, the sequence used should be optimised with a reduced f lip angle of 20°–25°, increased TR of 10–12 msec, TE of 2–4 msec and appropriately reduced bandwidth (increased water-fat shift). Since imaging is no longer confined

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Figure 2.9 CE-MRA detail of right calf in patient with critical lower limb ischaemia and tibial arterial disease using the blood pool agent Gadofosveset trisodium (Vasovist) in steady state phase with high spatial resolution - 1024 matrix reconstruction on 520 mm field of view with 0.6 mm slice thickness = voxel size of 0.5 × 0.5 × 0.6 mm = 0.15 mm3 voxel volume. (A) Overview MIP where overlying veins tend to obscure detail, (B) representative coronal slice from acquisition and (C) subvolume MIP of right calf showing clear separation of arteries and veins acheived through high spatial resolution. Note satisfactory peroneal artery but heavily diseased posterior tibial artery with irregular wall thickening seen between paired posterior tibial veins.

to a brief period of contrast enhancement, these high spatial resolution sequences can be used with phase oversampling for wrap suppression, increased signal averages for improved SNR etc. However, since these will now be longer sequences than usual MRA, they should be used with respiratory gating and/ or cardiac pulsatility compensation in appropriate body regions (Fig. 2.9). For imaging of systemic and portal venous disorders, as well as arteriovenous malformations, this prolonged time window and excellent contrast are a major advantage over the rapidly excreted conventional agents. Gadofosveset or similar products may well become the optimum agent in suspected pulmonary embolism for it will demonstrate both the pulmonary areterial tree in the early phase and the pelvic and leg veins (where the thrombus may arise) in the later phase of the same examination. The use of blood pool agents in coronary MRA may boost signal from the coronary vessels with respect to myocardium and might dramatically improve image quality. With its high relaxivity and long imaging window gadofosveset may well more reliably detect endoleak after aortic stent grafting than conventional imaging with ECF gadolinium agents and CT techniques. It is also thought that the use of a blood pool agent may allow the detection and localization of occult gastrointestinal haemorrhage. The downside to the use of a blood pool agent is that during later phase imaging, arteries and veins will now be equally conspicuous. Software programs are being developed to isolate arterial structures from the veins though these are currently not commercially available. An interesting further observation with gadofosveset is that late imaging reveals uptake in atherosclerotic plaques: no doubt this will provide fertile ground for further research as regards plaque activity and morphology.

REFERENCES 1. Grainger R G 1982 Intravascular contrast media. The past, the present and the future. Br J Radiol 55:1–18 2. Morris T W 1993 X-ray contrast media. Where are we now and where are we going? Radiology 188:11–16 3. Aspelin P, Aubrey S G, Fransson S G et al 2003 Nephrotoxic effects in high risk patients undergoing angiography. N Engl J Med 348:491–499 4. Heinrich M C, Kuhlmann M K, Grgic A et al 2005 Cytotoxic effects of ionic high-osmolar, nonionic monomeric and non-ionic iso-osmolar dimeric iodinated contrast media on renal tubular cells in vitro. Radiology 235:843–849 5. Nelson R C, Anderson F A Jr, Birnbawm B A et al 1988 Contrast medium extravasation detection accessory. Radiology 209:837–843 6. Dawson P, Cosgrove D O, Grainger R G 1999 Textbook of Contrast Media. Ch. 11 (Nephrotoxicity), Ch 15 (Endothelium and Blood) Isis Medical Media, Oxford 7. Katayama H, Yamaguchi K, Kozuka T, Takashima T, Seez P, Matsuura K 1990 Adverse reactions to ionic and non-ionic contrast media. A report from the Japanese committee on the safety of contrast media. Radiology 175:621–628 8. Yoshikawa H 1992 Late reactions to non-ionic contrast media. Radiology 183:737–740 9. Spring D B, Bettman M A, Barkan H E 1997 Deaths related to iodinated contrast media reported spontaneously to US Food and Drug Administration 1978–1994: Effect of availability of low osmolality contrast media. Radiology 204:333–337 10. American College of Radiology 1999 Manual on Contrast Media, 4th edition. American College of Radiology. Reston, Virginia, USA 11. Thomsen H S, Muller R N, Mattrey R F 1999 Trends in Contrast Media. Springer-Verlag, Germany 12. European Society of Urogenital Radiology 2003 Version 3.0 13. Thomsen H S 2006 Contrast media, safety issues and ESUR guidelines. Springer, Berlin Heidelberg New York 14. Prince M R, Grist, T M, Debatin J F 2003 3D Contrast MR Angiography, 3rd edn, Springer, Berlin Heidelberg New York 15. Schneider G, Prince M R, Meaney J F M, Ho V B 2005 Magnetic Resonance Angiography – Techniques, Indications and Practical Applications. Springer, Berlin Heidelberg New York

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16. Thomsen H S, Almén T, Morcos S K and the members of the Contrast Media Safety Committee of the European Society of Urogenital Radiology (ESUR) 2002 Gadolinium-containing contrast media for radiographic examinations: A position paper. Eur Radiol 12:2600–2605 17. Grobner T 2006 Gadolinium - a specific trigger for the development of nephrogenic fibrosing dermopathy and nephrogenic systemic fibrosis? Nephrol Dial Transplant 21:1104–1108 18. Marckmann P, Skov L, Rossen K et al 2006 Nephrogenic systemic fibrosis: Suspected etiological role of gadodiamide used for contrastenhanced magnetic resonance imaging. J Am Soc Nephrol 17: 2359–2362 19. Sadowski EA, Bennett LK, Chan MR, Wentland AL, Garrett AL, Garrett RW, Djamali A 2007 Nephrogenic systemic fibrosis: risk factors and incidence estimation. Radiology Jan 31; [Epub ahead of print] 20. Thomsen H S 2006 Nephrogenic Systemic Fibrosis: a serious late adverse reaction to gadodiamide. Eur Radiol 16:2619–2621 21. Morcos SK 2007 Nephrogenic systemic fibrosis following the administration of extracellular gadolinium based contrast agents: Is the stability of the contrast agent molecule an important factor in the pathogenesis of this condition? Br J Radiol 80:73–76 22. Morcos S K, Thomsen H S, Exley C M and members of the Contrast Media Safety Committee of the European Society of Urogenital Radiology (ESUR) 2005 Contrast Media: Interactions with Other Drugs and Clinical Tests. Eur Radiol 15:1463–1468 23. Quillin S P, Atilla S, Brown J J, Borrello J A, Yu C Y, Pilgram T K 1997 Characterization of focal hepatic masses by dynamic contrastenhanced MR imaging: findings in 311 lesions. Magn Reson Imaging 15:275–285 24. Ba-Ssalamah A, Schima W, Schmook M T et al 2002 Atypical focal nodular hyperplasia of the liver: imaging features of nonspecific and liver-specific MR contrast agents. Am J Roentgenol 179:1447–1456 25. Grazioli L, Morana G, Kirchin M A, Schneider G 2005 Accurate differentiation of focal nodular hyperplasia from hepatic adenoma at gadobenate dimeglumine-enhanced MR imaging: prospective study. Radiology 236:166–177 26. Wang C 1998 Mangafodipir trisodium (MnDPDP)-enhanced magnetic resonance imaging of the liver and pancreas. Acta Radiol Suppl 415:1–31 27. Koh D M, Brown G, Meer Z, Norman A R, Husband J E. Diagnostic accuracy of rim and segmental MRI enhancement of colorectal hepatic metastasis after administration of mangafodipir trisodium. Am J Roentgenol. 2007 Feb; 188(2):W154–61 28. Ward J, Naik K S, Guthrie J A, Wilson D, Robinson P J 1999 Hepatic lesion detection: comparison of MR imaging after the administration of superparamagnetic iron oxide with dual-phase CT by using alternativefree response receiver operating characteristic analysis. Radiology 210:459–466

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29. Birchard K R, Semelka R C, Hyslop W B et al 2005 Suspected pancreatic cancer: evaluation by dynamic gadolinium-enhanced 3D gradient-echo MRI. Am J Roentgenol 185:700–703 30. Obuz F, Dicle O, Coker A, Sagol O, Karademir S 2001 Pancreatic adenocarcinoma: detection and staging with dynamic MR imaging. Eur J Radiol 38:146–150 31. Smedby O, Riesenfeld V, Karlson B et al 1997 Magnetic resonance angiography in the resectability assessment of suspected pancreatic tumours. Eur Radiol 7:649–653 32. Prince M R, Yucel E K, Kaufman J A, Harrison D C, Geller S C 1993 Dynamic gadolinium-enhanced three-dimensional abdominal MR arteriography. J Magn Reson Imaging 3:877–881 33. Prince M R 1994 Gadolinium-enhanced MR aortography. Radiology 191:155–164 34. Riederer S J, Bernstein M A, Breen J F et al 2000 Three-dimensional contrast-enhanced MR angiography with real-time fluoroscopic triggering: design specifications and technical reliability in 330 patient studies. Radiology 215:584–593 35. Tombach B, Bremer C, Reimer P et al 2002 Using highly concentrated gadobutrol as an MR contrast agent in patients also requiring hemodialysis: safety and dialysability. Am J Roentgenol 178(1):105–9 36. Prince M R, Arnoldus C, Frisoli J K 1996 Nephrotoxicity of high-dose gadolinium compared with iodinated contrast. J Magn Reson Imaging 6:162–166 37. Collidge T A, Thomson P C, Mark P B et al 2007 Gadolinium-Enhanced Magnetic Resonance Imaging and Nephrogenic Systemic Fibrosis – A Retrospective Case-Control Study. Radiology, in press 38. Kuo P H, Kanal E, Abu-Alfa A K, et al 2007 Gadolinium-based MR contrast agents and nephrogenic systemic fibrosis. Radiology. 242(3):647–9 39. Kanal E, Barkovich J A, Bell C et al 2007 ACR Guidance Document for Safe MR Practices: 2007 AJR 188: 1–27 40. Meaney J F, Ridgway J P, Chakraverty S et al 1999 Stepping-table gadolinium-enhanced digital subtraction MR angiography of the aorta and lower extremity arteries: preliminary experience. Radiology 211:59–67 41. Wang Y, Chen C Z, Chabra S G et al 2002 Bolus arterial-venous transit in the lower extremity and venous contamination in bolus chase threedimensional MRA. Invest Radiol 37:458–463 42. Bezooijen R, van den Bosch H C, Tielbeek A V et al 2004 Peripheral arterial disease: sensitivity-encoded multiposition MR angiography compared with intraarterial angiography and conventional multiposition MR angiography. Radiology 231:263–271 43. Rohrer M, Bauer H, Mintorovitch J, Requardt M, Weinmann H J 2005 Comparison of magnetic properties of MRI contrast media solutions at different magnetic field strengths. Invest Radiol 40:715–724

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CHAPTER

Ultrasound: General Principles

3

David O. Cosgrove, Hylton B. Meire, Adrian Lim and Robert J. Eckersley

• • • • • • • • •

Nature of ultrasound Imaging methods Artefacts: clinical insights into ultrasound physics Interpretative principles Doppler ultrasound Volume flow measurement Doppler artefacts, errors and pitfalls Safety of ultrasound Developments in ultrasound

Ultrasound is an important technique for tomographic imaging of soft tissues. It provides images in real-time and so can also be used to interrogate the movement of structures such as cardiac valves, interactive guidance of biopsies and drainage procedures and, using the Doppler mode, the patterns of blood flow in both large and small vessels. Although ultrasound contrast agents are now available, images are generally obtained without them and thus are not dependent on organ function. As far as is known, ultrasound used at diagnostic intensities is harmless and it can be used without concern for injury to sensitive tissues such as the eye and the developing fetus. Ultrasound technology continues to improve; even low-end and portable systems are now digital and new transducer and beam-forming technologies allow higher frequencies to be used at depth. Although ultrasound has found major applications in the heart, abdomen and pelvis, as well as in the neck, breast, peripheries and neonatal brain, perhaps its most important clinical impact is in obstetrics, where the combination of safety and tomographic imaging of the fetus has rendered it indispensable. The parts of the body that can be imaged with ultrasound are, however, limited because it does not readily cross a tissue– gas or tissue–bone boundary, so they and deeper-lying structures cannot be studied.Thus, ultrasound is not generally useful for the lungs and is difficult to use in the head—except in the neonate, when the open fontanelles provides an excellent ‘window’. In other areas, overcoming the barrier caused by the

bony skeleton and gas-containing viscera requires technical expertise. Ultrasound is also subject to many artefactual signals which complicate interpretation and add to the operator skills required. Patients readily accept an ultrasound examination because the procedure requires only light pressure on the skin and preparation is minimal: bladder filling is required for pelvic ultrasound and fasting is necessary for the gallbladder, but otherwise the patient may be examined with this technique as and when convenient, a major advantage for emergency uses. In addition, mobile ultrasound systems that can be taken to the bedside, to intensive care and into the operating room are widely available.

NATURE OF ULTRASOUND Ultrasound is simply sound with a frequency above the limits of human hearing. In most diagnostic applications, frequencies in the 2–20 MHz (megahertz = million cycles per second) range are used, corresponding to wavelengths of 1–0.1 mm in tissue.

Ultrasonic transducers Ultrasound is generated by piezoelectric materials, which have the property of changing thickness when a voltage is applied across them. Lead zirconate titanate (PZT) is the most widely used. Its piezoelectric effect derives from movements of the heavy charged lead atom that is loosely bound within a complex crystal: when an electrical field is applied, the atom moves and distorts the crystal. PZT is a ceramic that is cast as a thin plate that may be disc shaped or, more usually, is formed into a strip that is then sliced into several hundred tiny elements as an array, with metal electrodes on the two surfaces. It is polarized by heating it above a critical temperature and then allowing it to cool in an electric field.When electrically pulsed, the crystal rings at a resonant frequency that is mainly determined by its thickness. Higher frequency crystals are thinner and thus more difficult to manufacture. The piezoelectric effect is reversible, so the same or a similar crystal is used as the receiver to

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produce small electrical signals when struck by an ultrasound wave. The crystal is mounted in a conveniently shaped holder, which contains the electrodes and any associated electronics, as well as the lenses and matching layers required to improve the beam shape. The whole assembly is known as the probe, transducer or head. An interesting new approach to designing transducers relies on photo etching, as used in semiconductor construction, to produce arrays of capacitance elements in which a pair of electrodes is separated by a small air gap; these capacitative microfabricated ultrasound transducers (CMUTs) are easily mass produced and integration of the associated electronics is simple (MEMS and Nanotechnology Clearinghouse website: www.memsnet.org/mems/what-is.html).

Propagation in tissue Ultrasound travels through tissue as a beam that is usually focused to improve lateral resolution. It propagates as a sequence of compression and rarefaction waves, transmitted by the elastic forces between adjacent tissue particles. The particles move in the same direction as the wave—thus ultrasound is a longitudinal wave in comparison to the transverse wave that occurs at the surface of water where the particles move up and down as the wave travels horizontally. The frequency of the oscillations is inversely proportional to the wavelength ( f = c/λ, where f is frequency, c is the velocity of ultrasound and λ is the wavelength). The way in which the ultrasound wave is transmitted varies with the strength of the elastic forces between adjacent particles (which relates to the elasticity of the tissue and thus to the velocity of ultrasound) and the masses of the particles (which determine density). These two factors determine the acoustic impedance (Z ) of the tissue (Z ≅ ρc, where ρ is density and c is the velocity of ultrasound). When the particles are heavy, a given amount of energy is transmitted with small movements of the particles: when they are light, larger excursions are involved; however, the actual distance a particle is moved at diagnostic ultrasound intensities is only a few Ångstroms. In clinical practice, since the velocity of ultrasound in tissue is almost constant (at 1540 m s−1), changes in impedance are mainly attributable to differences in density. The constant speed of ultrasound in soft tissues allows the depth of reflecting structures to be calculated by measuring the delay in the return of echoes after the ultrasound pulse has been transmitted.This is the essence of the pulse-echo method used in both ultrasound imaging and most forms of Doppler ultrasound. The position of reflecting structures across the image is determined by the direction in which the ultrasound beam is transmitted.

Attenuation Provided that the constituent particles of a tissue are small enough to move as a single entity, the acoustical vibrations are transmitted in an orderly and efficient manner. However, when very large molecules are involved, the vibrations become disorganized, one part of the molecule responding more or

less than another. While coherent vibration is what we know as sound, chaotic vibration is heat. This loss of coherence, the most important cause of dissipation of ultrasound energy, is known as absorption and is approximately proportional to the concentration of large molecules which correlates with viscosity. Absorption is also highly dependent on the ultrasound frequency, higher frequencies being more strongly absorbed. For average soft tissues, the loss amounts to approximately 1 dB per cm tissue depth for each MHz.Thus, when using a 3 MHz probe, for every 2 cm of tissue penetration there will be a loss of 6 dB, which is a halving of the signal’s strength. The noise floor (produced by random vibrations in the tissue and the transducer as well as by imperfections in the scanner electronics) typically lies some 60–90 dB below the peak signal, so the penetration of such a probe would be limited to about 20 cm depth. Ultrasound energy is also lost to the receiving transducer if it is reflected or refracted away from the returning line of sight or if the beam diverges. The total loss from all these mechanisms is called attenuation. High-frequency ultrasound gives good resolution because of its short wavelength, but the rapid attenuation of highfrequency ultrasound by tissue is the limiting factor to the maximum frequency that can be used in any given clinical application. Frequencies as high as 20 MHz can be used when only a few millimetres of tissue are to be traversed, such as for imaging the eye and skin, and for intravascular ultrasound (IVUS). For superficial tissues, 10–15 MHz is appropriate. For the heart, abdomen and second and third trimester obstetrics, 3–7 MHz is optimal; however, for some difficult applications, such as the abdomen in obese subjects, and for transcranial studies (most of which use Doppler), it is necessary to resort to 1 or 2 MHz transducers, with consequent reduction in spatial resolution. This frequency limitation can be mitigated by the use of coded transmit pulses, which impose a signature on the pulse (e.g. by making its frequency rise during the pulse, so called chirp encoding); long pulses are used and the scanner contracts them back to a short pulse using a matched decoding filter. Thus, the spatial resolution is regained but more acoustic energy can be transmitted without an increase in acoustic pressure so that the signal-to-noise ratio is improved. Compensation for the loss of signal with depth is achieved by applying progressively increasing amplification (gain) to later echoes using a time-varying amplifier that is triggered when each ultrasound pulse is sent. This is the time gain compensation (TGC), an important user control that must be set to equalize the image brightness for superficial and deep structures. Many modern ultrasound machines incorporate automatic TGC software that simplifies image optimization. This trade-off between depth and resolution is the chief reason for the development of intracavitary probes. Examples are transrectal probes for the prostate and transvaginal probes for gynaecology and early obstetrics, as well as transoesophageal probes for echocardiography. These can operate at 7–15 MHz to give good quality images and, as the barriers caused by

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impenetrable overlying structures are obviated, high-quality images are reliably achieved. Intracavitary probes can be combined with endoscopes for transoesophageal and transgastric data acquisition to evaluate the submucosal and deeper layers of the viscera and to provide high-resolution images of adjacent structures such as regional lymph nodes.

Reflection The other important ultrasonic interaction with tissue is reflection. Some of the transmitted energy is reflected whenever the beam crosses an interface where the transmission properties change, the proportion depending on the degree of impedance mismatch (change of Z ). The phenomenon is similar to that occurring in a rope that is fixed at one end and made to oscillate by shaking the free end up and down: provided the rope is of uniform construction, waves travel along it until they die out or reach the fixed end. However, if a part of the rope is thicker or more rigid, then the waves are partly reflected back when they reach this ‘obstruction’. The intensity of the reflection mainly depends on the degree of change in the tissue density. Ultrasound that is not reflected passes through and is available for imaging deeper tissues. Only a small fraction (2–10%) is reflected at each soft tissue interface but almost total reflection occurs at tissue–gas interfaces, and some two-thirds of the incident ultrasound is reflected at a tissue–bone or tissue–calculus interface so that an acoustic ‘shadow’ is produced deep to these surfaces. They are essentially opaque to ultrasound, which therefore cannot be used to interrogate aerated lung, and imaging through bone is difficult. Two main types of echo are encountered clinically, depending on the structure of the reflecting surface (Table 3.1). Where the interface is smooth compared with the ultrasonic wavelength (i.e. the surface is flat over an extent of several millimetres), the reflected wave obeys Snell’s law just as in optics and is reflected at an angle equal to the angle of incidence. By analogy, these echoes are known as specular or mirrorlike echoes. In the usual arrangement for ultrasound imaging, when the transducer is used both for generation and reception of the ultrasound, only those smooth surfaces that lie at right angles to the sound beam return specular echoes. The directed reflection means that the echo is of high intensity: strength and directionality are the cardinal features of these echoes from flat surfaces.They arise from the linings of hollow viscera and blood vessels, from the valves of the heart, from organ capsules and fascial planes, from the skin and from gas and bone surfaces. When the irregularities in the surface are similar in size as the ultrasound wavelength itself, i.e. 0.01–1 mm, a different mechanism, known as scattering, produces echoes; here each small

Table 3.1

ECHO-PRODUCING MECHANISMS

Specular: mirror-like reflections from flat surfaces

Strong, directional echoes

Scattered: interference patterns from small parenchymal discontinuities

Weak, nondirectional echoes

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interface (e.g. a lobule of an organ, arterioles, venules, etc.) is vibrated by the mechanical shock it has received from the incident ultrasound pulse. The energy is re-radiated more or less equally in all directions, so that each discontinuity behaves as an isolated point source of ultrasound.The echoes are isotropic and potentially detectable from any direction but, because the energy is distributed across the surface of an expanding sphere, the signal received by the transducer from any single vibrating particle is extremely weak and, in practice, is only detectable when several wavelets from adjacent particles happen to be superimposed and produce an additive effect. Low intensity and detectability from any angle are the cardinal features of these scattered echoes. They arise from soft tissue parenchyma and thus are important diagnostically.The texture in the resulting image is an interference pattern, known as speckle and, although deterministic, is not a one-to-one representation of the histological reality. This is why tissues of very different composition (e.g. liver and spleen) can produce indistinguishable echo textures. Many body tissues actually display properties intermediate between specular and scattered echoes. If the beam is sent into a tissue region from several angles and the echoes combined, the speckle pattern is reduced and specular echoes are more likely to be picked up.This approach, known as compounding, can improve image contrast and is widely used in high-end machines. Ultrasound is generally considered to be conducted in a linear manner whereby the waveform of the pulse is preserved with depth. In fact, this is not quite true and nonlinear propagation is an important phenomenon in which the wave of the original pulse from the transducer becomes distorted so that it comes to contain higher frequency components or harmonics. This occurs because the speed of sound increases with increasing density of the conducting medium so that the compression part of the cycle travels slightly faster than the rarefaction part; this distorts the wave much the same way that an ocean wave builds and breaks as it approaches the shore. The development of these harmonics depends on the sound intensity: higher acoustic powers produce stronger harmonics and therefore they are weak in the parts of the sound beam away from the central, desired portion. They are also weak for the first few millimetres of tissue depth because they take a few cycles to build up.These facts have been exploited in machines that use filters to select for the harmonics in the returning echoes by transmitting at, say, 1.5 MHz and tuning the receiver to 3 MHz to remove the fundamental echoes. In doing so, the beam aberrations from side lobes and from reverberations in the superficial tissue layers are suppressed (see below, Resolution). The tissue harmonic image is cleaner with higher contrast: this has been especially useful in echocardiography and in abdominal data acquisition (Fig. 3.1).

IMAGING METHODS Pulse-echo method In conventional pulse-echo ultrasound imaging, signals are displayed according to the depth calculated from the time

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Figure 3.1 Harmonic imaging. (A) Conventional grey scale image through the gallbladder hints at an echogenic mass within it. (B) In harmonic mode the image has much higher contrast and the gallstones (arrowhead), together with the shadowing they produce, are clearly visualized. Note also the thickened gallbladder wall that was not obvious on the fundamental image.

elapsed between transmission and receipt of the echoes, using the speed of sound in tissue to convert from time to depth; the same principle is used in radar and sonar. In the simplest system, only the depth of the interface is determined. It is displayed as a vertical deflection on a monitor and is known as A-mode (A for amplitude). It has one spatial dimension and the strength of the echoes is indicated by the height of the deflection (Fig. 3.2). If the transmitted pulse is repeated rapidly, then the position and intensity of any interfaces that arise from moving structures change with time. A simple way to display these changes is to modulate the intensity of the spot on the monitor in proportion to the intensity of the echoes and then to sweep the line of echoes across the screen. The resultant trace shows depth versus time and is known as an M-mode display (M for movement). It is especially useful in echocardiography for evaluating the rapid movements of valve leaflets (Fig. 3.3). A two-dimensional tomographic ultrasound image is formed by sweeping the beam through a slice of tissue and mapping the echogenicity of the reflectors as shades of grey to form a B-mode image (B for brightness), also known as a grey scale display. This is the commonest mode used for ultrasound imaging. While the depth of the reflectors is determined by the delay in the return of the echoes to the transducer (the

pulse-echo principle), their lateral position is determined by the direction in which the ultrasound beam was sent. Combining a series of B-mode images swept through a volume provides a three-dimensional display and, if the sweeping is sufficiently rapid, a real-time 3D display can be achieved, known as 4D. This has proved to be valuable in echocardiography and obstetrics.

Figure 3.2 A-mode trace. The A-mode is a trace indicating echo intensity tissue with depth. In this example, there is a fluid space (6–10 cm) from which no echoes arise. Tissues superficial and deep to this produce echoes of varying intensities and there is a particularly strong echo from the skin (0–5 cm). The time gain compensation (TGC) curve is also shown.

Figure 3.3 M-mode trace. The echo intensity is displayed as brightness and the trace is swept across the screen so that the x-axis represents time. This is an M-mode echocardiogram showing the rapid movement of the mitral valve apparatus within the left ventricle (LV), with thicker proximal and distal moving bands representing the myocardium. RV = right ventricle.

Beam steering The direction in which the beam is sent can be determined mechanically or electronically. Mechanical steering systems are used in intracavitary examinations where their small size is an advantage. They use a simple, single crystal transducer, which is mechanically swept through an arc (typically of 360 degrees) by spinning it on a wheel (Fig. 3.4). The resulting image has a circular shape, with the transducer itself at the centre and the tissues arranged around it, with the mucosa or intima closest to the transducer. Electronic sector systems (also known as phased arrays) have replaced mechanical systems for other applications. The beam direction is controlled by building up interference patterns between the waves transmitted from an array of a large number of small transducer elements.The usual arrangement consists of numerous PZT crystals (512 in stateof-the-art systems), each 1 mm or less in width and 5–10 mm in length, stacked up in a block. Each has separate electrical contacts and the pulses to each element are serially delayed from one end of the array to the other. These minute differences in timing occur within the short time needed to form

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Delay lines Pulser

128

Array elements

Figure 3.4 Mechanical transducer. (A) Diagram of a mechanical endoprobe transducer. A motor rotates the transducer assembly to sweep out a circular path. (B) A transrectal image covering a 2.2 cm axial section showing the layers of the rectal wall with defects of both the internal and the external sphincter (arrowheads) anteriorly. This is a classic sphincter tear that resulted from a forceps delivery. A = anterior, ES = external sphincter, IS = internal sphincter, M = mucosa, R = right, Tx = transducer. (Courtesy of Professor Clive Bartram, St Mark’s Hospital, London.)

an individual pulse and they produce interference patterns in which the troughs and peaks of the pressure waves add where they coincide and subtract where they are out of phase (hence the title ‘phased array’) (Fig. 3.5). This results in a beam that is angled away from the straight-ahead direction. For the next pulse, a slightly different set of delays is applied to adjust the steering. During receipt of the returning echoes, the electrical signals from the individual elements are delayed before they are summed in the beam former, exploiting the same summing and cancelling approach; the resulting signal is the electronic equivalent of angling the transducer face in the required direction. Although the concept is easy to understand, the underlying mathematics is complex and its implementation generally requires dedicated hardware, although software emulations have been developed so that beam forming can be performed on a PC, thus opening the way to cheaper and more flexible scanners. The resulting image has a triangular or sector shape, which has the important practical advantage of requiring only a small skin-contact area (or ‘footprint’), although it provides only a limited display of the superficial tissues. It is widely used in echocardiography. The operation of electronic linear array transducers is a little simpler.They are made longer than phased arrays and the beam is moved across the scanned plane by firing the elements in groups, starting from one end of the array and stepping along to the other (Fig. 3.6). This produces an image with a

Bea m

ep swe

A

Figure 3.5 Phased array. (A) Diagram illustrating the steering of an ultrasound beam with a phased array transducer. The delay lines introduce small timing differences in the pulses driving the elements so that those at one end are fired earlier than those further along the array. This has the effect of steering the beam away from the centre line, rather as if the transducer face had been tilted. (B) Right lobe of the liver in a patient with a pleural effusion imaged with a phased array ultrasound system. D = diaphragm, E = effusion, k = kidney.

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Figure 3.7 Breast carcinoma (arrows). Image obtained with a 14 MHz linear array probe. The rectangular format is particularly suitable for superficial structures. F = fatty tissue, G = glandular tissue, P = pectoralis muscle.

Array elements

Beam sweep

Figure 3.6 Linear array. Diagram of a linear array showing the formation of one ultrasound beam by triggering a set of elements at one end of the probe. The next beam would be formed by the adjacent or partially overlapping set of elements so that the beam is swept along the transducer face to give a rectangular image.

rectangular shape, best suited for the superficial structures that are of prime interest in small parts imaging (Fig. 3.7). In a simple variant, the linear array is curved so that the field of view is trapezoidal, as with a sector scanner, but with a longer skin line. This compromise is particularly useful when both superficial and deeper structures need to be imaged, for example in obstetrics (Fig. 3.8). An important limit to the rate at which ultrasound information can be acquired is set by its speed in tissue (average in soft tissue, 1540 m s−1). It is necessary to wait for each pulse

Figure 3.8 Convex array. (A) Image of a fetal skull obtained using a 3.75 Hz convex array. (B) Transvaginal image of a uterus with a fibroid (arrows) obtained using a tightly curved array at 7 Hz. The convex format is a compromise between the footprint requirements of sector and linear probes and has the advantage of giving a wide field of view for deeper structures. M = midline, P = placenta, S = skull, U = uterus.

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to have faded away before the next is transmitted because, if the second pulse is sent too early, the last echoes from the first overlap early echoes from the second and are falsely registered on the image as superficial structures. To avoid this range ambiguity artefact, the pulse repetition frequency (PRF) must allow a delay of at least the time taken for the most distant echoes to return. In practice, this limits the PRF to about 1000 pulses per second. The clinical application dictates the way these pulses are distributed in time and space. They are spread apart and the field of view restricted to maximize the frame rate for echocardiography but the sacrifice in spatial resolution and the small image size are not ideal for general applications in which image quality is paramount. Here, the low frame rate is generally less of a sacrifice, so a high line density and wider field of view provide a reasonable compromise.

Resolution Ultrasound resolution must be considered separately for the two dimensions, along and across the beam. Depth or range resolution is simply determined by the length of the ultrasound pulse, which is kept as short as possible, usually about two wavelengths in length; this equates to a shorter pulse if higher frequencies can be used and this is the main reason for the improved resolution of high-frequency systems. Lateral resolution, in contradistinction, depends on the width of the ultrasound beam. This is mainly controlled by mechanical or electronic lenses. The ultrasound emitted from a point source spreads in a hemisphere; however, as the source is enlarged, it forms a beam with near-parallel sides for several centimetres before the beam diverges again (Fig. 3.9). The changeover is the transition between the near (Fresnel) and far (Frauenhofer) fields. In fact, it is the point beyond which the ultrasound waves from the edge of the transducer arrive less than a wavelength later than those from the transducer’s centre, and this is why larger aperture transducers have a longer near field. The beam is actually formed by the same interference effects that are exploited to steer the beam in electronic transducers and which occur because the compression and rarefaction phases of the ultrasound pulse summate in the direction of the beam but cancel elsewhere. Adding a converging lens (usually a curved layer of plastic bonded to the front of the piezo material), or simply shaping the transducer face into a shallow dish, acts to narrow the beam. This can be thought of as working in the same way as an optical lens, i.e. by beam refraction, but a more useful concept is based on the fact that the lens conducts ultrasound more rapidly than tissue, so that waves emanating from the transducer edge are fractionally ahead of those from the centre as they enter the body. Thus, interference patterns are set up across the sound field that emphasize the centre region of the ultrasound beam and cancel ultrasound waves that would otherwise spread laterally. The same effect can be produced with a multi-element array by sending the transmit pulses to the outer elements fractionally ahead of those to more central elements; this is the

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Figure 3.9 Ultrasound beam shapes. The shapes of the ultrasound beam from four transducers are indicated. All are of the same frequency (3.5 MHz). The beam from the 10 mm probe in (a) has a complex region close to the transducer face, then a mid portion with near parallel sides, before the beam spreads out in the far field. The white lines indicate the half-power limits. Increasing the probe diameter (b) improves the overall beam width. In (c), weak focusing has been added by concave shaping of the crystal. This further improves the beam shape in the focal zone but causes it to spread more severely further out. This would be a useful compromise for general abdominal data acquisition. Stronger focusing (d) exaggerates these effects, producing a fine beam but only over a short distance. This would be useful for data acquisition in superficial structures.

electronic equivalent of shaping the transducer surface into a dish (Fig. 3.10). Focusing also occurs on receipt of the echoes, in a reciprocal fashion. While electronic focusing is complex, it confers two benefits. First, the focal position can be altered by changing the delays within the transmit pulse so that the probe can be optimized for each clinical situation. Secondly, on receipt of the echoes, the focus can be set close to the transducer initially (to optimize resolution of superficial tissues) and then be progressively refocused deeper into the body to track the train of echoes returning from deeper interfaces. This dynamic focusing optimizes resolution over the entire depth of the image. Tight focusing of the ultrasound beam has the undesired effect of accentuating its divergence in the far field, so that the beam is only optimal over a short depth (Fig. 3.9d).While such a transducer would be appropriate for structures where only a narrow strip of tissue is of interest (e.g. the eye), generally a better compromise is to use weak focusing together with as large an aperture as can reasonably be maintained in contact with the skin. In practice, the ultrasound beam achieved is far from perfect (Fig. 3.11). Not only is the optimum beam width achieved over only a relatively short focal depth, but its profile also leads to smearing of signals from strong reflectors because they are detected from some distance off-axis (Fig. 3.12). In addition, any real beam has side lobes that are emitted at steep angles from the main beam so that unwanted echoes may be received from interfaces that lie a long way off-axis. These signals further smear the image, particularly when strong reflectors such as gas bubbles lie in their direction.These limitations also apply across the imaging plane, so that out-of-plane reflectors may

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also interfere. Minimizing the loss of spatial and contrast resolution that results from these deficiencies is a major emphasis of the art and science of transducer design and is an area in which marked progress continues to be made (e.g. by using tissue harmonic imaging).

Delay lines Pulser

ARTEFACTS: CLINICAL INSIGHTS INTO ULTRASOUND PHYSICS

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Array elements

Focal zone

Figure 3.10 Electronic focusing. Diagram to illustrate the principle of electronic focusing. The delay lines are set to send the pulses from the outer elements fractionally ahead of those from more central elements. The resulting interference patterns accentuate the central part of the beam and cancel the off-axis portions. The effect is greatest at the focal zone.

The assumptions made in generating ultrasound images are that ultrasound travels at a constant speed in tissue and that the beam is ‘well behaved’. Artefacts arise whenever these principles are violated. The speed of ultrasound varies by only a few per cent in watery soft tissues but fat conducts approximately 20% more slowly. This means that the depth of echoes arising beyond large fatty regions is overestimated so that they are depicted as being further away from the skin than they actually are. It may be seen as a shelf in the diaphragm behind a fatty tumour in the liver (e.g. a lipoma). An even more dramatic illustration of this same artefact occurs with silicone implants, in which the ultrasound velocity is about half that of soft tissue (Fig. 3.13).The chest wall deep to a breast prosthesis, for example, can appear to be twice its true distance from the skin. The difference between the velocity in adipose and watery tissues can refract the ultrasound beam so that it deviates from its straight-line path. A gross version of this artefact occurs in the hypogastrium, where the fatty wedge lying between the lower ends of the two rectus abdominis muscles acts like an optical prism. The refraction of the ultrasound beam in transverse suprapubic images can produce double outlines of pelvic structures and even, in extreme cases, complete double images. This artefact may well be the cause of the ‘vanishing twin’ phenomenon (when one of a pair of twin embryos disappears early in pregnancy), described when ultrasound was first introduced in obstetrics. More generally, the different velocities at multiple small tissue–fat interfaces produce minor deviations of the ultrasound beam that defocus and disperse it. The clinical effect of this is loss of contrast, the image appearing noisy and blurred

Figure 3.11 Ultrasound beam plot. The complexity and marked deviation from the ideal of a practical diagnostic ultrasound beam is shown in this intensity plot. The beam is very complex for the first few millimetres from the transducer face (on the left of the figure) but improves towards the focal zone, where it reaches an effective diameter of a few millimetres, before spreading out again in the far field. Unfortunately, some ultrasound energy is also sent out as side lobes at angles to the main beam, further complicating the effective beam shape. These divergences from the ideal narrow shape limit both the spatial and contrast resolution of ultrasound images. (Beam plot kindly supplied by Dr Adam Shaw of the National Physical Laboratory, Teddington.)

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Figure 3.12 Beam width artefact. In this abdominal image in a patient with ascites, the intense echoes from gas in a loop of bowel have spread across into the fluid (arrowhead). This artefact results from the finite width of the ultrasound beam. A = ascites, B = bowel.

Figure 3.14 Beam dispersion by fat. Deep to this fatty renal hilum, the retroperitoneal tissue layers (arrowheads) are less clear than adjacent tissue planes because the hilar fat has defocused the beam. S = renal sinus.

(Fig. 3.14).The problem is worse when there is marked admixture of fatty and watery tissues, as may occur in fatty infiltration of muscle, a feature that may explain why poor images may be obtained from an out-of-training athlete. The degradation of ultrasound images and its variability from patient to patient are two of the most important limitations to the diagnostic use of ultrasound: strenuous efforts are being made by equipment manufacturers to investigate and minimize these problems. The use of tissue harmonics has greatly reduced this interpatient variability and thereby the number of unsuccessful examinations. A ‘well-behaved’ beam would be laser-thin and travel directly through the tissues to and from the transducer. The real-life situation is far from this ideal: the effects of beam spreading, side lobes, refraction and diffraction on spatial and contrast resolution have been mentioned. Multiple reflections are another common artefact that produces a series of

parallel false images of flat interfaces that happen to lie parallel to the skin. They are the result of repeated reflections of the incident ultrasound pulse between the transducer and the flat surface or between two such surfaces within the body. A common example is seen as parallel lines deep to the superficial fascia produced by the ultrasound being reflected from the transducer back into the tissue a second time. Similar linear artefacts are commonly seen in the superficial portions of the lumen of viscera such as the bladder and gallbladder and in blood vessels. The difficulty in measuring the thickness of the superficial intima–media layer of the carotid artery results from this artefact (Fig. 3.15). Rereflection from the transducer surface is reduced by applying matching layers which make it less reflective by reducing the impedance mismatch between the transducer and the skin, resulting in much less noisy images. A similar situation is produced by the closely packed bubbles in the foam that may occur in bowel gas. Here, the intense reflections repeat in near-random sequence, depending on the precise arrangement of the bubbles, to produce a train of echoes that fades over a few millimetres as the ultrasound energy dissipates. On the screen, a bright streak, the ‘comet tail’ artefact, is seen extending for a few millimetres deep to the gas. The same effect may occur deep to foci of calcification and behind metallic or plastic foreign bodies. Where a flat reflector lies at an angle to the ultrasound beam, ultrasound may be sent back into the tissue at an angle and then be reflected back from any interface it encounters. These echoes retrace the incident pathway and are picked up by the transducer, to be depicted as though lying deep to the angled reflector, exactly as objects are seen ‘through’ an optical mirror. This ‘mirror image’ artefact is commonly encountered at the air–pleura interface above the diaphragm: the echoes apparently above the diaphragm actually arise from the hepatic or splenic parenchyma (Fig. 3.16). They are quite different from the appearance of the lung examined intercostally when a very strong linear echo is seen, usually followed by strong reverberation echoes that fade over a few millimetres.

Figure 3.13 Velocity artefact from a silicone prosthesis. On this breast scan, the echofree ovoid is a prosthesis (S) inserted following removal of a tumour. The low speed of sound within it has distorted the pleural line (P) so that it appears to lie deeper behind the prosthesis (arrows) than to either side.

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Figure 3.15 Reverberation artefacts. (A) Multiple internal reflections within tissue layers give rise to false repeated signals which are most obvious where they fall over echo-free fluid spaces such as the gallbladder (arrowheads). (B) Although the intima–media layer is well delineated at the deeper surface of this normal common carotid artery (arrow), the superficial layer is mostly obscured by reverberation artefact. C = common carotid artery.

Figure 3.16 Mirror image artefact. Structures in the liver, such as cysts (arrowhead) in this patient with dominant polycystic kidney disease, are mirrored at the air–pleura interface and appear in the position of the lower lobe of the lung. When this surface is absent, for example when a pleural effusion is present, the effect does not occur (compare Fig 3.5).

INTERPRETATIVE PRINCIPLES Shadowing and increased through transmission Acoustic shadowing and increased through transmission of sound (often referred to as enhancement, a term better reserved for the signal-augmenting effects of microbubble contrast agents) are important components of the ultrasound image. Shadowing occurs when little or no ultrasound can penetrate an interface and results in a dark band over the deeper tissues, bounded by the ultrasound beam lines, which are parallel for a linear transducer and radiating for a curved or sector transducer (Fig. 3.17). Shadowing is reduced when compounding is used. Absorption and reflection are the two main causes of shadowing. If a portion of the tissue being imaged absorbs ultrasound faster than the background for which the TGC is adjusted, tissues deeper to the highly attenuating region are undercorrected and appear darker than adjacent tissue. Fibrous tissue and, to a lesser extent, fat attenuate more than average and are common causes of acoustic shadowing (e.g. in a fatty liver, behind scars, and behind scirrhous breast carcinomas). High attenuation also partly accounts for the shadows seen deep to calcific lesions such as biliary and renal stones; note that in this case the intense reflection is the main factor. Whatever proportion of the incident ultrasound is reflected is not available to continue through

for imaging. For stones, this amounts to about 60% of the incident energy, but for tissue–gas interfaces almost all the incident ultrasound is reflected and these produce dense shadows. In this case, the shadows are often partially filled in by reverberant echoes caused by the efficiency of these gas–tissue boundaries as reflectors, so that reverberation artefacts commonly occur. This noise in gas shadows has given rise to their designation as ‘dirty shadows’, as compared to the ‘clean shadows’ behind stones, and this is a useful differential diagnostic feature. Whether a stone or gas bubble actually casts a shadow depends on its size because about three-quarters of the beam must be obstructed to cause a shadow. If a stone is smaller than this, or lies away from the central axis of the beam, enough ultrasound passes beside it to insonate the deeper tissues. In practice, shadowing is usually apparent behind renal and biliary stones of 5 mm or more in diameter; much smaller calcifications may shadow if high-resolution transducers are used. Groups of fine calcifications can also shadow if their aggregate size and density is high enough (e.g. in nephrocalcinosis). A third important type of shadowing is ‘edge shadows’, sometimes known as ‘refractive shadows’, a description based on one explanation of their origin. They are seen as fine, dark lines extending deep to strongly curved surfaces. Cysts and the fetal skull are typical examples and fascia is often also responsible; for example, the fine shadows seen beyond Cooper’s ligaments in the breast and those caused by the neck of the

Figure 3.17 Acoustic shadowing. The dark band (arrows) deep to the gallstones (G) is an example of shadowing produced by a combination of high absorption and reflection. L = liver.

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gallbladder. These edge shadows must be recognized as being different from attenuating and reflective shadows to avoid errors in their interpretation. Increased through transmission is the opposite of attenuation shadowing: a region of tissue has a lower than average attenuation and so the TGC (which is adjusted to compensate for the average attenuation) is inappropriately high for that region so that echoes from deeper tissues are overamplified. The phenomenon is the hallmark of cystic spaces and the ‘bright up’ is often accentuated by the darker banding lines of the edge shadows typically formed from the cyst wall (Fig. 3.18). Even those fluid cavities that contain echogenic material, such as suspended crystals in the gallbladder, pus, blood or necrotic tissue, usually still produce increased transmission, depending largely on the proportion of fluid present. However, some solid tissues also show increased sound transmission, usually because they have a high proportion of fluid. Many tumours fall into this category, especially fibroadenomas in the breast, while lymphomatous deposits and inflammatory masses may behave in the same way.

Echogenicity The prime determinant of the strength of ultrasonic echoes is the impedance mismatch (Z ) between adjacent tissue components. At the risk of being somewhat simplistic, in practical clinical terms this may be understood as interfaces between tissues of different densities.The larger the mismatch, the stronger the echo, so that interfaces between soft tissues and bone, for example, give very strong echoes; however, within soft tissues, the most significant components are fibrous tissue (often in the form of the perivascular microskeleton) and fatty tissue. Thus, while uniform regions of fibre or fat are echo-poor (subcutaneous fat and retroperitoneal fibrosis are examples), admixtures between them and watery tissues give stronger echoes. A second important factor is the concentration of the scatterers: for a given impedance mismatch, a region that contains a large number of scatterers is more echogenic than one where they are spread out. Commonly, the ‘dilution’ of scatterers is caused by high water content. The low reflectivity of the

Figure 3.18 Increased sound transmission. The echoes (arrowheads) deep to this liver cyst (C) appear brighter than those from the rest of the liver; this is because the cyst fluid attenuates less than the solid liver and so signals from beyond it are relatively overamplified.

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Figure 3.19 Prostatic carcinoma. Small malignant tumours are usually relatively echo poor, as in this example (arrowheads). More advanced tumours have complex echo patterns as necrosis, fibrosis and calcification develop. This transverse image was obtained using a transrectal endoprobe.

congested liver in right heart failure is an example. Malignant tumours are a general case: until they grow large enough to undergo necrosis or calcification they tend to be echo-poor (Fig. 3.19). Similarly, the oedematous tissues in acute inflammation give low-level echoes (e.g. the echo-poor pancreas in acute pancreatitis). On the other hand, the high concentration of reflectors is the cause of the echogenic kidneys in recessive (infantile) polycystic renal disease—the interfaces between the innumerable cysts cause strong echoes (Fig. 3.20). Strong echoes are also obtained from the multiple interfaces of the vessel walls of haemangiomas, and even stronger ones from angiomyolipomas in which there is also admixed fatty tissue.

DOPPLER ULTRASOUND The Doppler principle has been utilized in diagnostic ultrasound since 1954. Colour Doppler is a refinement of the technique whereby the regions of blood flow are displayed

Figure 3.20 Echogenic kidney in recessive polycystic renal disease. The multiplicity of interfaces between the numerous small cysts results in high-intensity echoes from what would otherwise be expected to be an echo-poor process. k = kidney, L = liver.

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in colour as an overlay on the grey scale 2D image. ‘Duplex’ is used to imply that two types of information are being collected simultaneously: typically 2D imaging (possibly with colour Doppler) and spectral Doppler. The Doppler effect is a change in the frequency of sound being reflected from a moving target. In its simplest form in diagnostic ultrasound equipment, it is used to detect the beating fetal heart lying anywhere in the sound field of a large, unfocused and continually transmitting transducer.With these systems there is no way of determining either the direction of motion or depth of the fetal heart beneath the transducer, an advantage in this clinical application because the fetal heart beat can be detected successfully without accurate positioning of the transducer. In vascular applications, a wide, continuous ultrasound beam might result in confusing signals being received from several vessels simultaneously. This limitation is overcome by utilizing a narrow beam, similar to that used for imaging, and by transmitting the ultrasound in short pulses and processing only the information returning from a specific depth. Accurate positioning of the sensitive area of the beam (the sample volume) is achieved by positioning an electronic marker in the 2D image. Array transducers permit the collection of Doppler information while continuing with real-time imaging, albeit at a reduced frame rate since the machine has to share the pulses between the modes.

near the vessel wall so blood can be considered as flowing in a series of concentric layers (laminae), the blood in adjacent laminae having different velocities (Fig. 3.21). This type of flow is termed ‘laminar flow’ which is assumed in many of the calculations made from Doppler data. However, flow is nonlaminar in many vessel segments, including near bends, branches and junctions, and around plaques of atheroma. Values derived from Doppler signals in these circumstances may be inaccurate. If flow is not pulsatile and the velocity is not high, the ‘velocity profile’ across the vessel is approximately parabolic (Fig. 3.22). If the blood is flowing fast or is being accelerated rapidly, it tends to move with the same velocity, so-called ‘plug flow’ (Fig. 3.23). In a major artery the flow usually approximates to plug flow during systole and to parabolic flow during diastole.Venous flow is normally parabolic. The term turbulent flow is frequently misused to imply any form of flow disturbance: strictly speaking, it should only be applied to describe the complex flow disturbance that results from blood being forced at high velocity through a

Theoretical basis of Doppler studies

Figure 3.21 Laminar flow. Diagrammatic representation of the concentric layers of blood flowing at different velocities with the highest velocity in the centre of the vessel.

An ultrasound beam insonating a blood vessel is partially reflected by red blood cells. If these are moving, there is a change in the frequency of the reflected pulse: an increase in frequency if the flow is towards the probe and a decrease if it is away from it. This is the Doppler effect and the changes in frequency, referred to as the Doppler shift, may be calculated from the equation:

ƒd =

2ƒo ν cos θ c

This indicates that the Doppler frequency shift ( fd) depends on the transmitted frequency ( fo), the velocity of the blood (v) and the angle between the ultrasound beam and the direction of flow (θ).The speed of ultrasound (c) is constant. If the angle is known and the frequency shift measured, the blood flow velocity may be calculated by the scanner’s computer. The operator places a cursor appropriately in the vessel and marks the direction of the vessel axis. It is necessary to achieve an angle of considerably less than 90 degrees between the direction of blood flow and the direction of the ultrasound beam, as no Doppler signal results if the beam/vessel angle is 90 degrees (cos 90° = 0 in the above equation). The accuracy of the calculated velocities is inversely proportional to the beam/vessel angle. At an angle of zero degrees the error is up to 3%, at 70 degrees it is 30% and at 80 degrees it may approach 100% (see Fig. 3.38).

Figure 3.22 Parabolic velocity profile. The fastest flow is in the centre of the vessel, with a progressive reduction in velocity towards the vessel wall.

Characteristics of blood flow When blood is flowing in a straight vessel, the drag effects of the vessel wall and the viscosity of blood slow the flow

Figure 3.23 Plug flow. The flow velocities are almost equal across the whole vessel diameter.

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small orifice. The mildest form of flow disturbance (usually encountered near a moderate stenosis) results in an increase in the range of velocities present, giving rise to spectral broadening in the Doppler trace. More severe lesions may be accompanied by ‘disturbed flow’ in which eddies or vortices of flow occur beyond the lesion, giving rise to simultaneous forward and reverse flow components in the Doppler signal. The flow velocity waveform through the cardiac cycle depends on cardiac function, the type of blood vessel and the state of the distal vascular bed. Elastic arteries such as the femoral, supplying high-resistance beds, exhibit a pronounced triphasic flow pattern (Fig. 3.24). There is a rapid systolic upstroke but, because of the combination of elasticity and high peripheral resistance (if the supplied muscles are at rest), almost immediate flow reversal occurs, resulting in a transient negative Doppler component. There may be no forward flow during diastole. However, if the peripheral resistance is low, as in the internal carotid, hepatic and renal arteries, a uniphasic pattern is seen with forward flow throughout diastole (Fig. 3.25). A switch from the former to the latter occurs in the limb vessels with exercise and in the mesenterics with digestion. The waveform of the display may also be modified by many factors above, at, or below the sampling site. In a major artery the waveform is determined by cardiac contractility, aortic valve function, the proximal arterial tree, the state of the vessel at the examination site and disease in the distal arterial tree supplied by the vessel. When quantifying the severity of a stenosis, it is important to be clear as to whether a reduction in cross-sectional area (CSA) or diameter is being described; for example, a 70% diameter reduction gives rise to a 90% area reduction (Table 3.2). When the CSA of an artery is narrowed by up to 50% (diameter reduction 30%) there is no reduction in the volume of blood flowing and there must therefore be an increase in the peak flow velocity. This is usually accompanied by an increase in the range of velocities present (spectral broadening) (Fig. 3.26). If the stenosis is greater than 50%, there is a much greater increase in the velocity, and the flow beyond the steno-

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Figure 3.25 Low resistance flow. The flow is continuously forward throughout the cardiac cycle with moderately high flow throughout diastole. There is very little low-velocity flow throughout most of the cardiac cycle (compare with Figs 3.26 and 3.27).

sis forms eddies (disturbed flow) that give rise to 100% spectral broadening with small amounts of simultaneous reverse flow (Fig. 3.27). The disturbed flow propagates for a few centimetres beyond the lesion. Narrowing of the CSA by more than 75% (50% diameter reduction) reduces volume flow, which produces a fall in flow velocity in the proximal vessel segment. Stenoses of more than 90% CSA reduction (70% diameter reduction) give rise to a severe reduction in volume flow, very high jet velocities through and beyond the lesion and turbulence beyond the lesion (Fig. 3.28). In addition, there is an increase in the pulsatility of the flow proximal to the lesion and a decrease beyond. These indirect signs may be the only available indicators of disease in some clinical situations; for example, in renal artery stenosis, where the stenotic segment may not be accessible to ultrasound examination.

Information in the Doppler signal The spectral Doppler signal consists of a mixture of continually varying frequencies that, by chance, fall in the audible range so that the signal may be evaluated simply by amplifying it and feeding it to a loudspeaker. The frequency content and

Table 3.2 RELATIVE VALUES FOR AREA AND DIAMETER REDUCTION IN STENOSES

Figure 3.24 Normal common femoral artery flow. There is triphasic flow with early diastolic flow reversal.

Diameter reduction (%)

CSA reduction (%)

10 20 30 40 50 60 70 80 90

20 36 51 64 75 84 91 96 99

95

99.75

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Figure 3.26 Moderate spectral broadening. The spectrum throughout late systole and most of diastole has been filled in by an increased range of blood flow velocities.

waveform of the signal, however, contain a wealth of information about the nature of the blood flow in the vessel and it is easier to appreciate and interpret these if a visual display is created. This is derived from a ‘spectrum analyser’ which determines the strength of components in the Doppler signal that fall into each of numerous frequency bands during a time interval of typically 5–20 ms. The resultant display represents frequency on the vertical axis and time on the horizontal axis. The brightness of the tracing at any point indicates the amount of signal at each specific frequency over each time interval and is proportional to the number of red cells in the sample volume that are moving at each velocity (Fig. 3.29). The outline of the spectral display gives information about the direction of the flow, the maximum Doppler frequency at any time, and the nature and magnitude of any pulsations or other periodic changes in flow velocity. By convention, flow towards the transducer is displayed above the baseline and away from the transducer below. Clearly there is a great deal of information in the Doppler signal and currently we do not have techniques for extracting it all. There are, however, several simple observations that permit the diagnosis and grading of disease, and some of these can be supported by simple mathematical calculations.

Pulsatility measurements Many indices of waveform analysis have been devised but only two are in regular clinical use. These are the resistance index (RI) (also known as the Porcelot index) and the pulsatility index (PI) (also known as the Gosling index). Because they are ratios, they are independent of the beam/vessel angle (although obtaining a good quality Doppler trace from which to make the measurement does require a beam/vessel angle of < 60 degrees). Their derivations are:

Figure 3.27 Complete spectral broadening. There is now complete filling of the spectrum throughout the cardiac cycle, with brief periods of simultaneous reverse flow during systole.

Figure 3.28 Turbulence beyond a stenosis. There is high-velocity flow beyond this portal vein stenosis, with simultaneous forward and reverse velocities.

Figure 3.29 Close-up view of the spectral trace. The spectral trace is composed of increments on both the vertical and horizontal axes. The horizontal increments indicate the individual time intervals during which Doppler sampling occurs, 20 ms in this example. The vertical increments indicate increasing frequency. The brightness of the trace within each pixel indicates the number of blood cells moving with that velocity at that time.

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Resistance index =

Pulsatility index =

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maximum velocity − minimum velocity maximum velocity maximum velocity − minimum velocity time-averaged maximum velocity over a cardiac cycle

The PI was devised to characterize the triphasic flows typically seen in the femoral artery. Its main disadvantage is that it requires calculation of the mean peak frequency, whereas the RI only requires the measurement of two values. The RI is particularly sensitive to changes in downstream flow resistance and was devised to assess changes in diastolic flow in lowresistance vascular beds. The resistance may be increased by vascular stenosis or by disease in the organ supplied by the vessel. However, as it also depends on the ‘end diastolic’ velocity, the RI increases as the heart rate decreases, allowing a longer duration for the diastolic flow to fall (Fig. 3.30). In theory, it is possible to correct the RI for a normalized heart rate; in practice, however, this is seldom done.

Spectral content The distribution of the shades of grey in each time slot indicates the range of flow velocities present in the vessel at that time. If there is plug flow, this distribution is very narrow (Fig. 3.31A), whereas parabolic flow gives rise to a wider range of frequencies (Fig. 3.31B). The spectral broadening produced by a stenosing lesion is shown in Figure 3.28.

Colour Doppler With this form of display, areas of blood flow are represented as colour within the image. It has become common practice to represent flow towards the transducer as red and flow away as blue (Fig. 3.32). The operator can reverse these assignments and select alternative colours.The velocity information used is a cruder estimate of the mean velocity in each location compared to the spectral information described above. For some clinical applications the colour may be set to vary with the variance of the velocity rather than mean velocity; this may be helpful for the detection of flow disturbance or turbulence as an indication of the site of significant vascular disease. The Doppler

V

1

2

T

Figure 3.30 Dependence of RI on heart rate. The value of enddiastole is relatively high if the heart rate is rapid (1). A slower heart rate allows a greater time for diastolic deceleration, leading to a lower enddiastolic velocity (2) and an apparently higher resistance index. V = velocity, T = time.

Figure 3.31 (A) Spectral distribution in plug flow. A very narrow band of velocities is present throughout the cardiac cycle in this artery. (B) Spectral distribution in parabolic flow. The slower flow in this wide portal vein gives rise to a wide range of velocities during each time interval.

information is collected in rapid sequence from a large number of discrete picture elements. The derived velocity information displayed at each pixel is built up from a number of consecutive ultrasound echoes, and this necessarily reduces the number of complete frames of Doppler information that can be acquired in any time period.The frame rate for colour Doppler is therefore about one-quarter of that for grey scale imaging, although this problem can be alleviated by reducing the width of the Doppler gate and by increasing the size of each pixel. In practice, a compromise between frame rate, image area spatial resolution and accuracy of the colour velocity information has to be accepted. The main advantages of colour Doppler are the ease with which vessels can be detected and their patency confirmed. In many applications this may be all that is required for confirming that a structure is a vessel, or that a known vessel is patent. Where relevant, the direction of flow can also be easily confirmed (e.g. in the portal vein). The wealth of information contained in the Doppler spectrum is not available in the colour image, however, and it is often necessary to perform a conventional spectral Doppler study as well. The positioning

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Power Doppler is useful for depicting slow flow in smaller vessels—for example, in the kidney (Fig. 3.33)—and in malignancies.The reduction in signal strength when a volume element is only partially within a vessel gives an apparent improvement in vessel wall definition (Fig. 3.34). A further practical advantage is the reduced dependence on the beam/vessel angle. This is mainly because the direction information is averaged in colour Doppler and cancels out if the beam/vessel angle is close to 90 degrees (Fig. 3.35A). In power Doppler the direction information is ignored and the signal amplitudes are averaged. As these are all positive, a strong signal results (Fig. 3.35B).

VOLUME FLOW MEASUREMENT

Figure 3.32 Colour convention in colour Doppler. Fetal circle of Willis. The flow in the right middle cerebral artery, the left anterior communicating artery, the left posterior communicating artery and the right posterior cerebral artery are displayed in red as the flow is towards the transducer. The flow in the other vessels is passing away from the transducer and is therefore represented by blue.

of the Doppler gate is greatly facilitated by colour Doppler and made more accurate. In addition, colour Doppler readily shows up vessels that are too small to be resolved in the 2D image and that otherwise could not be studied. The major clinical applications of colour Doppler are the same as those for conventional spectral Doppler, although the ease of examination is improved and the operator’s confidence enhanced. In addition, colour Doppler permits the detection of very small vessels and the assessment of the number and distribution of vessels within a tissue volume. This is of relevance when attempting to record blood flow signals from vessels such as the renal arcuate arteries or the uterine arteries, and in the assessment of the vascularity in and around focal lesions. For example, the colour Doppler-detected vascular pattern around malignant breast lesions shows more numerous vessels, often with abnormal courses, than are seen in benign lesions, and primary hepatomas appear more obviously vascular than metastases in the liver.

It is theoretically possible to measure the volume of blood flowing within a vessel using Doppler ultrasound. This requires calculation of the mean blood velocity in the vessel

Figure 3.33 Power Doppler display of a right kidney. This extended field of view (Siemens ‘Siescape’) shows the flow in the IVC, main renal vessels and the intrarenal vessels right out to the capsule. The loss of directional information prevents differentiation of arteries from veins. (Courtesy of Siemens.)

Power Doppler The Doppler information used in power Doppler is the same as for colour Doppler but the velocity information is discarded and usually also the direction information.The colour hue is modulated by the strength of the Doppler signal, which, in turn, is proportional to the number of red cells moving within the sample volume.The display is thus a map of the distribution of moving red cells above the sensitivity threshold of the system. Large vessels give high signals, while the presence of both tissue and small vessels in the sensitive region gives a weaker signal. Fortunately, the random noise that impairs conventional velocity colour Doppler is self cancelling with this form of display, thus enabling higher gains (around 10 dB) to be used and increasing the sensitivity for vessel detection.

Figure 3.34 Vessel wall definition. Normal carotid artery. The power Doppler display colour intensity decreases near the vessel wall owing to the volume elements lying partly outside the vessel. This gives an apparent improvement in the definition of the intimal surface.

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Figure 3.35 (A) Colour Doppler with 90-degree beam-to-vessel angle. There is poor flow detection and direction ambiguity throughout the vessel. (B) Power Doppler display with 90-degree beam-to-vessel angle. As direction information is not used in the display, uniform flow detection is achieved throughout the vessel segment.

and the cross-sectional area of the vessel. The product of these two values should be the volume of blood flowing. In practice, there are significant uncertainties about both values, which may lead to errors in excess of 100%. Calculation of the mean velocity should take into consideration the distribution of velocities across the vessel in each time interval and the average of these over several cardiac cycles (Fig. 3.36). It is also assumed that the vessel is uniformly insonated and that signals are detected from all the blood in the cross-section of the vessel. In practice, slow-moving blood is missed and the range gate or beam width may encompass only a fraction of the cross-sectional area. These errors may not be obvious from visual inspection of the spectral display. If the crosssectional area of a vessel is calculated from a single diameter, any error is squared. Measuring the area from an image at right angles to the vessel axis may be more accurate, but this information cannot be obtained at the same time and site as

Figure 3.36 Calculation of time-averaged mean velocity. The weighted mean velocity during each time interval has been calculated. The average of these throughout the trace is a close estimate of the true mean velocity.

the Doppler signal, owing to the inappropriate beam/vessel angle. Volume blood-flow measurement in the carotid and femoral vessels and in the abdomen has not been shown to be of clinical value.

DOPPLER ARTEFACTS, ERRORS AND PITFALLS Doppler examinations are subject to a range of artefacts and potential errors that are not encountered in 2D imaging.

Sample volume size and position The shape of the sample volume is determined by the length of the range gate and the width of the ultrasound beam at the depth of the gate. The depth and length of the sample volume are determined by the operator and should be matched to the size of the vessel under investigation. If the sample volume is too small or does not encompass the whole cross-section of the vessel, the resulting spectrum underestimates the range of velocities present in the vessel. In scanners with array transducers, the beam width can be altered and many automatically focus the beam at the depth of the range gate. This is seldom advisable since, although it may improve sensitivity, it leads to undersampling in large vessels such as the portal vein. Similarly, if the sample volume is small and the vessel to be studied is mobile, a discontinuous Doppler signal results. If the movement of the vessel is due to respiration, the cause for the changes in the spectral display can usually be recognized. If, however, an artery moves as a result of intrinsic pulsations it may move in and out of the beam with each cardiac cycle. This can give rise to loss of the diastolic signal in each cycle, resulting in falsely high PI and RI values (Fig. 3.37). If the range gate is too large or the beam width too great, more than one vessel may lie in the beam at the same time and give rise to a confusing spectral display, especially if flow in both vessels is in the same direction. Sometimes such a feature is of positive advantage; for example, when searching for a

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Figure 3.37 (A) Normal superior mesenteric arterial trace. The vessel has remained within the range gate throughout the cardiac cycle, giving a satisfactory Doppler trace. (B) Artefactually abnormal SMA trace. Same case as Figure 3.37A. The range gate is too small and is misplaced. The vessel is moving out of the range gate during diastole, giving rise to the false appearance of increased vascular resistance.

renal vein signal alongside the renal artery.The actual length of the sample volume is usually greater than that displayed on the image. This has led to the false assertion that it is acceptable to have a sample volume smaller than the vessel diameter.

Velocity information An important source of error when attempting a velocity calculation is the effect of the beam/vessel angle. This shows the percentage error in velocity calculation that may be expected for different angles. At angles greater than 50 degrees there is a rapid increase in error that approaches 100% at 80 degrees (Fig. 3.38). Major errors may also arise from the way the mean velocity is calculated: calculating the mean value of the peaks is only accurate for pure plug flow (Fig. 3.39). Strictly, the mean velocity should be calculated for each time interval, taking into account the distribution of velocities in that time interval.The mean of a large number of these values over several cardiac cycles should then be calculated (Fig. 3.36). For a vessel with true laminar flow and a parabolic profile, this more accurate method produces a value half that of the mean peak method. The pulse repetition frequency (PRF) of the ultrasound system imposes limitations on the range of Doppler frequencies that can be measured. The maximum value of the PRF is limited by the depth from which echoes are to be recorded. In abdominal diagnosis this may be up to 15 cm. In order to allow time for the echoes to be received before the next pulse is transmitted, about 250 µs must elapse between pulses.Thus, the pulses can only be repeated at about 4000 s−1 (4 kHz).The Doppler detectors can only measure frequencies accurately if they are less than half of the PRF (the Nyquist limit), i.e. less than 2 kHz in this example. Doppler shift frequencies above this level are frequently encountered in the abdomen, are misinterpreted by the apparatus, and displayed on the wrong side of the baseline, giving rise to a characteristic discontinuous waveform display (Fig. 3.40A). The technical term for this phenomenon is ‘aliasing’ and it may sometimes be overcome by viewing the vessel from a steeper angle, using a lower ultrasound frequency,

a higher PRF or continuous Doppler beam, or by electronic correction (Fig. 3.40B).

Wall filters The pulsating walls of arteries give rise to high-amplitude, low-frequency Doppler signals that may overload the spectrum analyser and appear in the spectral display as high-intensity spikes. This ‘wall thump’ can be filtered out by rejecting the

100 90 80 70 Error in velocity

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60 50 40 30 20 10

10

20

30

40

50

60

70

80

Degrees

Figure 3.38 Dependence of velocity error on beam-to-vessel angle. The error in velocity calculation is less than 10% for angles of less than 50 degrees. There is a rapid and unacceptable rise in the error rate for angles greater than 70 degrees.

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of colour reversal are contiguous (Fig. 3.43). If the flow reversal is genuine, the forward and reverse flow colours are separated by a black margin.

SAFETY OF ULTRASOUND

Figure 3.39 Mean peak velocity calculation. Automatic software has calculated the peak instantaneous velocity for each time interval throughout the cardiac cycle. The mean of these values is the mean peak velocity.

very low frequencies in the Doppler signal. If the ‘wall filter’ is set too high, however, true flow information may be lost (Fig. 3.41). For abdominal venous studies, a filter value as low as 25–50 Hz is appropriate. In some makes of scanner, settings up to 800 Hz are available for arterial studies and would eliminate the Doppler information from many veins.

Colour artefacts The most common problem with colour Doppler is failure to appreciate that the colours represent flow direction with respect to the transducer. If a vessel curves or bifurcates in the image plane, the flow in different segments will be represented in different colours.The same artefact may result if a straight vessel is imaged with a curvilinear or sector transducer (Fig. 3.42). As with spectral Doppler, aliasing also occurs in colour flow imaging and results in reversal of colour and thus apparent flow reversal. Aliasing can be identified by noting that the areas

An important feature of diagnostic ultrasound, especially in obstetric applications, is its apparent safety. No study has shown any damaging effect of pulsed ultrasound at diagnostic intensities when applied to the intact fetus in utero. These studies include follow-up assessments of growth, the development of cataracts and hearing loss, and the induction of childhood malignancies. Occasional reports of effects (usually unimportant and found on searching the data collected for other purposes, a statistically unacceptable methodology) have failed to be substantiated on repeat studies. One curiosity that does seem to be statistically sound was the unexpected finding in a Norwegian study that male fetuses imaged with ultrasound were less likely to be right-handed (23% were left-handed or without a dominant hand, compared with 17% in controls). While unimportant in itself, this suggests that ultrasound can induce changes in cell migration at an early developmental stage. There have been several reports of biochemical alterations to cells in suspension, including changes to DNA and surface membrane behaviour, both of which are obvious at high intensities but are demonstrable also (at least in some studies) at diagnostic power levels. It is debatable whether these effects are significant for the intact animal and it is notable that many of them have proved to be elusive when repeat studies have been attempted.They do, however, inject a note of caution and reinforce the importance of using common sense in restricting the use of high-power modes and long imaging durations to situations in which the clinical benefit outweighs any possible ill effects: there is no place in medicine for a quick (i.e. inadequate) ultrasound or for ‘just to have a look’ (e.g. at a fetus).

Figure 3.40 Frequency aliasing. (A) The frequencies above the Nyquist limit have appeared on the wrong side of the baseline. (Note that no beam-tovessel angle has been set so the velocity values are uncorrected and therefore meaningless.) (B) The aliased peaks have been electronically transposed to their correct locations.

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Figure 3.41 Inappropriately high wall-thump filter. The wall-thump filter has been set at 200 Hz and has removed all the low velocity information in this venous study.

The entire subject of safety is under continuous review and up-to-date information is available from the European Federation of Societies for Ultrasound in Medicine and Biology (www.efsumb.org). Ultrasound intensity is the energy flowing across a surface in unit time; the usual units are watts per square centimetre. For pulsed ultrasound, a peak or average value may be taken. Because of focusing, the energy distribution along and across the beam is nonuniform, and the peak intensity at the maximum point is usually quoted. Mean spatial peak temporal average intensities for B-mode imaging are ~175 mWcm−2, and for spectral Doppler ~1600 mWcm−2. High intensities of ultrasound are capable of causing significant biological effects. These are used to therapeutic benefit

Figure 3.42 Artefactual flow reversal on colour flow imaging. Flow within the splenic vein (arrowheads) is from left to right. In this transverse epigastric scan, on the left side of the image the flow is towards the curvilinear array probe and is therefore displayed as red. On the right side, the flow is away from the probe and is therefore displayed as blue. There is a thin black area at the point where the colour changes, indicating a true flow direction reversal, rather than aliasing, as the cause of the colour change. A = aorta, I = inferior vena cava.

Figure 3.43 True and artefactual colour flow reversal. The highvelocity jet through this portal vein stenosis has given rise to aliasing. The aliased green signal passes through yellow to red in a continuous gradation (in the vertical limb of the vessel in this display). The coarse vortex within the poststenotic dilatation gives rise to a true area of flow reversal, colour-coded blue, which is separated from the forward flow red component by a black margin.

both in physiotherapy, when the changes sought (mainly local heating) are often reversible in nature, and in ultrasound surgery, when cell killing is the aim. The primary mechanism inducing biological change is thermal. The ultrasound beam is attenuated during its passage through tissue, some of the lost energy (20–40%) being scattered by tissue structures and the remainder absorbed, thus leading to heating. In the case of B-mode imaging and Doppler examinations of soft tissue, this phenomenon results in a biologically insignificant temperature rise. Significant temperature rises may, however, be induced when a Doppler beam at its highest power is incident on a bone surface because of the high absorption at this interface where the impedance changes markedly. The problem is worst in the case of spectral Doppler, in which the beam is held over one tissue region for long periods of time. It is therefore recommended that Doppler exposures should be performed at the lowest power level consistent with obtaining the desired information and that the time the beam is held stationary is kept to a minimum. Particular care is advised when Doppler is used during late pregnancy, when the skull has begun to calcify, and in examining the neonatal brain in the vicinity of the skull. The vibration of particles caused by a high-intensity ultrasound beam may produce mechanical disruption of intracellular membranes, and the pressures can cause fluids to flow in streaming movements. Extreme ultrasound powers can produce regions of such low pressure during the rarefaction phase of the cycle that dissolved gases (mainly nitrogen) can come out of solution, or water may vaporize to produce minute gas

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bubbles that pulsate in the sound field. This process, known as cavitation, can cause mechanical damage to the tissue and even lead to ionization. Cavitation is unlikely to occur at diagnostic imaging intensities in vivo but has been demonstrated with the higher powers and continuous-wave conditions used in ultrasound therapy. Safety-related information as presented to the operator is regulated by the FDA’s output display standards (ODS) which are calculated predictions of the possible effects in tissue. There are two indices, one of which, the mechanical index (MI), indicates the probability of mechanically induced damage: an MI of 1 warns of the possibility of mechanical effects in tissue. The second is the thermal index (TI), which is related to possible heating of tissues. Separate indices have been defined for soft tissue and for bone (TIS and TIB, respectively) because heating at a bone surface is likely to be higher. A TIS of 1 suggests that, in the worst case, tissue could be heated by 1°C; since damage is unlikely until temperature rises of 1.5°C or more are achieved (and maintained for some time), this is a conservative estimate, as is the MI indication. Although rather vague in their real-life meaning, the ODS does provide guidance to users, who can then decide whether any possible risk to, for example, the fetus, is outweighed by the potential benefit of the examination.

DEVELOPMENTS IN ULTRASOUND High frequencies, 3D, elastography Three-dimensional ultrasound (3D US) has become available on many machines and the data can be acquired rapidly enough to allow display in real-time at low frame rates (called 4D). It has found niche applications, mainly for the display of complex anatomy such as the fetal face and the heart. In most parts of the body the clean surfaces that are required for true 3D rendering are not present; however, reslicing in otherwise unobtainable planes has proved useful in some situations; for example, to display the uterus in the coronal plane for depicting developmental anomalies. Elasticity imaging (or elastography) is a new imaging method that is promising because of the very high contrast it offers between masses, especially tumours, and the host tissue. The principle is simple: obtain signals before and after applying a distorting force (stress) that moves the tissues by a few millimetres, and create an image of the tissue’s response (strain) by comparing the two. In principle, any imaging method can be used: ultrasound has the advantage that the transducer can be used to apply the stress and of working in real-time; magnetic resonance imaging has also been used successfully. The information used to create the elastograms is similar to clinical palpation except that it is much more sensitive, especially to deeper structures. It has been applied to the breast and the prostate and is an active research area. A commercial device that estimates liver elasticity as a measure of fibrosis to assess the need for antiviral therapy in hepatitis has been developed.

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Contrast agents An important development has been the introduction of microbubble agents to enhance ultrasound signals. They exploit the very high reflectivity of microbubbles small enough to cross the lung capillary bed, so that systemic (including myocardial) ultrasound enhancement can be achieved following an intravenous injection. Their gas content makes them very reflective, even at high dilution, and initially this was exploited using Doppler to enable signals to be obtained from small vessels and from those where the signals are attenuated by overlying tissue such as the intracranial arteries under the skull. However, by virtue of their compressibility, microbubbles display unique properties in an ultrasound beam, which sets them into resonance when there is a match between their diameter and the ultrasonic wavelength. Fortuitously, for microbubbles in the 2–7 µm range (needed to allow transcapillary passage) this occurs at ultrasound frequencies of 2–10 MHz. As the acoustic pressure increases and still well within diagnostic levels, the compression and expansion phases become asymmetrical because it is harder to compress a bubble than to expand it.This nonlinear behaviour produces echoes that contain frequencies not present in the transmitted pulses. Many of these are higher frequency harmonics at twice the transmitted frequency. An elegant way to extract these nonlinear signals is to send a series of pulses down each line, varying their phase and amplitude; the returning signals are combined to cancel the linear signals from tissue and the remaining bubble-specific signals are used to form an image that can be presented as a colour overlay on the B-mode image or shown on a side-by-side display, all in real-time. As these contrast-specific methods are exquisitely sensitive, low transmit powers can be used (MIs of 0.1 or 0.2) such that the fragile microbubbles are not destroyed, allowing continuous imaging for their life of around 5 min after injection. Since they do not depend on microbubble motion, they can be detected even in the microcirculation, a major advantage over Doppler techniques. A major technical difficulty with making small bubbles is that their surface tension is so high that they collapse and dissolve within a few seconds of formation.To produce enhancement of useful duration, the microbubbles must be stabilized by encapsulating them in a membrane such as denatured albumen or a phospholipid. The gas is also critical to achieving good resonance (and therefore echogenicity) as well as providing a useful duration after injection. Air or inert gases such as perfluorocarbon compounds are the most frequently used (Fig. 3.44). Microbubbles extend to ultrasound many of the features of angiography, such as demonstrating tissue perfusion where infarction is a clinical question and demonstrating bleeding points (the microbubbles are confined to blood vessels unless there is active haemorrhage). The temporal filling patterns of focal lesions such as liver masses can be studied in similar ways to dynamic CT and MRI methods and this has become an important clinical application. Contrast-enhanced ultrasound is now incorporated into European guidelines for hepatoma detection in cirrhosis. They have proved invaluable in monitoring the extent of liver ablation using ultrasound-guided

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the arterialization of the hepatic blood flow in cirrhosis and in malignancy by timing the delay from injection to the appearance of microbubbles in the hepatic veins: a long delay of around 40 s is normal and reflects the fact that three sets of microcirculation have to be crossed by portal blood. Arterialization, however, leads to a much earlier arrival (< 25 s) and this seems to be a sensitive marker for these diseases. An approach that may be especially important for the myocardium depends on the fragility of many types of microbubble such that they are destroyed by the ultrasound beam at the higher ranges of permitted intensity. If, after such

Figure 3.44 Fibrous nodular hyperplasia: enhancement after intravenous administration of a microbubble contrast agent. The baseline longitudinal section through the right lobe of the liver (A) shows a subtle subcapsular lesion (arrowheads). The system was then reset to display the contrast image on the left (using Contrast Pulse Sequences) and the B-mode image on the right, both with low MIs. SonoVue (2.4 ml) was given IV and the haemodynamics of the flow through the lesion observed in real time. At 19 s (B), the lesion had enhanced markedly—the ‘light bulb sign’ of a hypervascular lesion—and its feeding vessel was apparent (arrowhead). By 35 s (C), the liver had enhanced to the same extent as the lesion, which therefore disappeared, indicating that it contains normal liver tissue. This sequence is pathognomonic of focal nodular hyperplasia. (With thanks to Siemens Medical, Mountain View, CA for provision of the CPS software and Bracco, Milan for supply of SonoVue.)

RF or laser techniques because remaining perfused tumour can be detected immediately and treatment extended to complete ablation in the same session; previously, contrast CT was required and the patient would have to return for a second ablation session. In addition, using these agents as tracers to measure time– intensity curves analogous to those used in dynamic radioisotope studies is now possible, launching ultrasound into a new functional era of great potential. One promising approach reveals

Figure 3.45 Reperfusion study. (A) With an infusion of a microbubble (in this case SonoVue) to produce a steady blood level, a high-intensity power Doppler pulse has been used to inactivate the agent in this renal transplant; this is the first frame after the ‘flash’ whose position is indicated by the green box. (B) After a few seconds interrogation with a low intensity beam (an MI of < 0.2 was used here) in the pulse inversion mode, the hilum of the transplant has reperfused and is again returning a signal. (C) Several seconds later the entire cortex has filled with contrast agent. The rate of recovery of these signals relates to the speed of blood flow, an important physiological measurement. (With thanks to Siemens Medical, Mountain View, CA for provision of the PI software and Bracco, Milan for supply of SonoVue.)

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a destructive pulse (or ‘flash’), the tissue is interrogated with a low-power, nondestructive beam, the speed with which microbubbles (and therefore red blood cells) reperfuse the slice can be measured (Fig. 3.45). The slope of this refill curve relates to flow rate, while its peak level relates to the fractional vascular volume, the two measurements needed to estimate true perfusion. Although its potential importance in ischaemic heart disease is obvious, it also illustrates the innovative ways that the special properties of microbubbles can be used—in this example to create a ‘negative bolus’ for haemodynamic measurements. Microbubbles can be modified as vehicles to carry useful compounds such as therapeutic agents, including DNA and oligonucleotides, as well as ligands for biochemical processes such as activated endothelium. This opens up possibilities for molecular imaging and for directed drug/gene delivery. Physiotherapy-level ultrasound can be used not only to release the payload but also to create transient pores in cell membranes that facilitate cell uptake. The therapeutic applications of microbubbles could overtake their diagnostic uses.

Ultrasound therapy Ultrasound also seems likely to find a role in tissue ablative therapy, particularly in oncology. Therapeutic ultrasound uses the thermal effects of high-power ultrasound beams that are tightly focused so that very small target tissue regions can be heated and coagulated, and dedicated transrectal therapy systems have been marketed. There is also considerable research interest for the treatment of hepatic and renal lesions. Although the general approach is similar to radiotherapy, ultrasound therapy has the advantages of being much more precise, rapid (the entire process could be completed in one session), easy to handle and of not damaging adjacent tissues. It is also possible to monitor its effects in real-time with diagnostic ultrasound. The physical limitations to accessible anatomical regions are the same as for diagnostic ultrasound.

SUGGESTED FURTHER READING Physics Evans D H, McDicken W N 1999 Doppler ultrasound: physics, instrumentation and signal processing. Wiley, Chichester Hill C R (ed.) 2001 Physical principles of medical ultrasound, 2nd edn. Wiley, Chichester

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Hoskins P R, Thrush A, Whittingham T 2000 Diagnostic ultrasound: physics and equipment. Greenwich Medical Media, London Hykes D L, Starchman D E, Hedrick W R 2004 Ultrasound physics and instrumentation. Mosby, St Louis Kremkau F W 1997 Diagnostic ultrasound: principles and instruments. W B Saunders, Philadelphia

Clinical Baxter G M, Allan P L, Morley P (eds) 1999 Clinical diagnostic ultrasound. Blackwell, Oxford Bluth E I, Peter H, Arger P, Carol B, Benson C B 2000 Ultrasound: a practical approach to clinical problems. Thieme, Stuttgart Callen P W 2000 Ultrasonography in obstetrics and gynecology. Elsevier, Philadelphia Rumak C M, Wilson S R, Charbonneu J W (eds) 2005 Diagnostic ultrasound, 3rd edn. Mosby-Year Book, St Louis Zwiebel W J, Pellerito J S 2004 Introduction to vascular ultrasonography. Elsevier-Saunders, Philadelphia

Safety Abbott J G 1999 Rationale and derivation of MI and TI: a review. Ultrasound Med Biol 25: 431–442 Barnett S B, ter Haar G R, Ziskin M C, Duck F A, Maeda K 2000 International recommendations and guidelines for the safe use of diagnostic ultrasound in medicine. Ultrasound Med Biol 26: 355–366 ter Haar G, Duck F A 2000 The safe use of ultrasound in medical diagnosis. British Institute of Radiology, London

New developments Albrecht T, Cova L, Frauscher F, Solbiati L, Thorelius L 2005 Contrastenhanced ultrasound in clinical practice. Springer, Berlin Goldberg B B (ed.) 2001 Ultrasound contrast agents, 2nd edn. Martin Dunitz, London Nelson T R, Downey D B, Pretorius D 1999 Three-dimensional ultrasound. Lippincott Williams & Wilkins, Philadelphia Quaia E (ed.) 2005 Contrast media in ultrasonography. Springer, Berlin

Acknowledgements Dr Gail ter Haar (Royal Marsden Hospital, London) kindly reviewed the section on Safety.

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Ashley S. Shaw and Adrian K. Dixon

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Computed tomography: a brief history Technical overview of MDCT Image reconstruction and analysis Clinical applications of MDCT Whole body MDCT in asymptomatic adults Radiation dose considerations Radiotherapy Future directions

COMPUTED TOMOGRAPHY: A BRIEF HISTORY The first CT machine used for clinical purposes, developed by the late Sir Godfrey Hounsfield, was installed at the Atkinson Morley Hospital, London, in the early 1970s1. Each axial image of the head took several minutes to acquire and days to reconstruct at the EMI laboratories. Over the next decade, CT machines became faster as the processing power of computers improved. In the mid-1980s, the development of slip-ring technology enabled continuous revolution of the X-ray unit, which reduced the acquisition time and allowed helical data to be acquired2. Since the late 1990s, the focus of development has been on reducing the size of the detectors in order that they could be arranged in multiple rows along the z-axis.This permits the acquisition of multiple ‘slices’ simultaneously, thus reducing acquisition time yet further. At the time of writing, ‘state of the art’ CT systems have up to 64 rows of detectors, allowing up to 4 cm to be imaged per revolution, a revolution time around 0.4 s, and with a resolution of approximately 0.4 mm.

TECHNICAL OVERVIEW OF MDCT Before multidetector CT, imaging protocols and patient throughput were frequently limited by the time taken for the tube to cool between exposures. Refinements in tube design now ensure that multiple phases of imaging can be performed without overheating, as the tube cools faster than the throughput of patients. These design improvements also allow increas-

ingly complex studies of small structures, which require the focal spot to be kept to an absolute minimum even when the mA is increased. For example, one manufacturer offers a focal spot measuring 0.6–0.9 mm, depending on the mA used. The X-rays are detected by various forms of detector. A ceramic scintillator, for example, absorbs the X-rays and produces visible light. This stimulates a photodiode to produce an electrical signal from which the image can be constructed. In order to reach the spatial resolution, detectors have become smaller (typically around 0.6 mm), more sensitive (allowing some dose reduction) and react faster (allowing increased gantry rotation speeds). In short, the smaller each detector is, the smaller the possible collimation, whereas the wider the bank of detectors, the greater the potential coverage per rotation. Some manufacturers arrange detectors of varying width such that the smallest are positioned centrally and the largest positioned peripherally, whereas others use identically sized detectors across the array.The majority of machines in use are characterized by the number of detectors (e.g. 16-, 40- or 64-slice) and use a single focal spot on the tube. By contrast, one 64-slice machine has a 32-detector array with two rapidly alternating focal spots in the tube. In effect, this produces overlapping projections that are acquired and processed simultaneously. The end result of these advances is that the latest technology from all manufacturers is able to produce isotropic images with sub-mm resolution in all planes with ever-increasing rapidity.

IMAGE RECONSTRUCTION AND ANALYSIS The rapid acquisition of such large datasets has necessitated a change in radiologists’ approach to CT reporting. As a bare minimum, both thick and thin section axial images should be produced and stored. Images need to be viewedwon a workstation capable of producing images in all planes, with the ability to change window level and width. When viewing images, our experience has shown that, in addition to standard lung or soft tissue windows, review of the data on bone windows (preferably in the sagittal plane) often demonstrates unsuspected abnormalities (Fig. 4.1).The optimal

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Figure 4.1 Sagittal MPR imaged on bone windows, demonstrating multiple sclerotic lesions throughout the vertebrae in a patient presenting with lymphoma.

slice thickness will depend on the clinical question posed. Cardiac imaging, for example, will require sub-mm images, whereas renal calculi will be reliably visualized on images up to 3 mm thick, and the detection of lung nodules can be enhanced by scrolling through overlapping 10 mm images (allowing differentiation from branching blood vessels). For 3D viewing of datasets, there are four principal postprocessing techniques: multiplanar reconstructions (MPR), maximum intensity projection (MIP), surface shaded display (SSD) and volume rendered (VR) images (Fig. 4.2). MPR— and associated curved planar reconstruction (CPR)—simply allows the image to be viewed in any plane chosen, including a curved plane with CPR (Fig. 4.3). MIP displays the highest attenuation voxels in an image only, thus allowing visualization of structures (usually vessels) not in a single plane. However, visualization of arterial wall calcification can be difficult and vessels near to bony structures may be obscured. SSD gives 3D representations of the anatomy in grey scale, but is prone to noise and artefacts. VR allows visualization of the vessels distinct from the surrounding anatomy and is of particular value in cerebral angiography, where the vessels pass through and lie in close proximity to the skull base. Computer aided detection (CAD) of lung nodules is currently under development by a number of groups. Preliminary results indicate that, while more nodules can be detected,

Figure 4.2 (A) Coronal MPR demonstrating renal cell carcinoma at the lower pole of the left kidney extending to the hilum. (B) Coronal MIP demonstrating renal cell carcinoma at the lower pole of the left kidney extending to the hilum. (C) Coronal SSD demonstrating renal cell carcinoma at the lower pole of the left kidney extending to the hilum. (D) Coronal VR demonstrating renal cell carcinoma at the lower pole of the left kidney extending to the hilum.

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Dynamic perfusion CT (PCT)5 evaluates regional blood flow within the brain by serially imaging a limited area following the infusion of intravenous contrast medium.The resulting images give a computerized map of relative cerebral blood flow, mean transit time and cerebral blood volumes, albeit with limited coverage of the brain (Fig. 4.4)6. In the context of patients presenting with acute stroke, PCT is significantly more accurate than non-enhanced CT in detecting stroke

Figure 4.3 artery.

CPR demonstrating a normal calibre right coronary

there is a significant rise in the number of false-positive nodules. Although not current routine clinical practice, CAD may become a useful prompt in the evaluation of resectable metastases and on lung cancer screening programmes (should they prove to be of clinical value).

CLINICAL APPLICATIONS OF MDCT Neuroradiology In many areas, CT has been superseded by magnetic resonance imaging (MRI) for the evaluation of neurological disease as it gives superior differentiation of the soft tissues. However, CT remains the initial imaging investigation for many patients, particularly those presenting following trauma or with acute neurological syndromes. Digital subtraction angiography (DSA) has long been regarded as the gold standard for detection of an intracranial aneurysm. Single row detector helical CT could not acquire data sufficiently quickly to obtain reliable diagnostic angiographic studies. MDCT has been shown to be highly accurate in the detection of aneurysms, with a sensitivity of over 96% and specificity of 100% on a 16-detector row machine, identical to the DSA results3. Even for small aneurysms (< 3 mm), sensitivity remained over 90%. This has led to a significant change in neuroradiological and neurosurgical practice, enabling treatment to be planned from 3D images at a workstation, with a consequent reduction in the number of diagnostic angiograms performed. Furthermore, it is now possible to perform multiple phases of vascular imaging within the brain, obtaining both an arterial and a venous phase4. This may provide additional information in patients with an arteriovenous malformation, but needs further evaluation.

Figure 4.4 (A) Unenhanced CT of the head in a patient presenting with acute aphasia. (B) Cerebral perfusion map demonstrating delayed perfusion in the left middle cerebral artery territory. Images courtesy of Rosie Farmer, Addenbrooke’s Hospital, Cambridge, UK.

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(75–86% versus 66%), the extent of stroke (94% versus 43%), and has a high interobserver agreement7. This may prove to be of clinical use in those patients in whom thrombolysis or other novel therapies are being considered. Potentially, the technique can differentiate the ischaemic penumbra from infarcted tissue, identifying patients who would benefit from revascularization and those in whom unnecessary treatment could be avoided. The data obtained from this technique are very similar to that obtained with perfusion and diffusion-weighted MRI but CT is considerably easier in the acute situation. It is likely that routine clinical use will be guided by developments in therapy for acute stroke. Similarly, the use of perfusion CT following severe head trauma has been reported to yield independent prognostic information regarding functional outcome8, although again this is not part of current routine clinical practice.

Cardiac The rapid acquisition times and increased spatial resolution of MDCT probably holds the greatest potential for cardiac imaging. Imaging of the coronary arteries with CT has two major components: unenhanced coronary artery calcium scoring and contrast-enhanced coronary artery angiography. Electron-beam CT (EBCT) has been used for coronary calcium scoring for approximately two decades and is a highly sensitive method of detecting coronary artery atherosclerosis. A number of authors have shown there to be a high correlation between EBCT and MDCT. However, there is a well-recognized high interscan, interobserver and intraobserver variability and the consensus of the American Heart Association/American College of Cardiology is that the incremental value of calcium scores over traditional risk assessments has not been proven and that its use could only be justified in selected groups at intermediate risk9. CT coronary angiography is typically performed in a single breath-hold, with ECG gating, during a rapid (≤ 5 ml/s) infusion of iodinated contrast material (Fig. 4.5). The sensitivity and specificity of 16-detector row CT for detection of ≥ 50% luminal stenosis is around 90%10,11, although a significant proportion of the arteries (12%)11 were not demonstrated sufficiently well to be evaluated. However, the negative predictive value (NPV) of a normal study is reported to be 95%12 and thus could be potentially used to exclude cardiac disease in patients with chest pain.There are, of course, limitations: problems occur in patients with arrhythmias (e.g. atrial fibrillation), the heart rate should ideally be 60 beats per minute or below13, and beta-blockers may be necessary for optimal imaging. In order to detect a 10–20% change in stenosis, imaging will require an isotropic spatial resolution of 0.3 mm and an acquisition time of approximately 0.1 s—technical specifications that, as detailed above, are becoming possible with double tube 64-slice systems CT. It is anticipated that this development will further enhance the accuracy and reproducibility of coronary artery angiography and, importantly, give an overview of cardiac morphology and function. By using all the data during the various phases of the cardiac cycle, cine loops can provide ejection fractions and other objective measurements.

Figure 4.5 VR image of the heart depicting the coronary arteries.

Assessment of perfusion can be made on delayed imaging.The potential impact of noninvasive coronary imaging on clinical practice is clearly enormous, and will depend on the results achieved with newer MDCT technology.

Thoracic The imaging of patients with suspected pulmonary venous thromboembolism (PE) had, for many years, relied on either ventilation–perfusion (VQ) studies or pulmonary angiography. However, once CT became sufficiently fast to image the thorax with thin collimation (≤ 2 mm) in a single breath-hold, CT pulmonary angiography (CTPA) rapidly became the test of choice for many clinicians. Indeed, the number of patients referred for CT with suspected PE has increased dramatically since its introduction. This is likely to be due to increased awareness of the diagnosis in combination with easy access to an accurate noninvasive study. VQ studies expose the patient to a significantly lower radiation dose than CTPA, but have a number of drawbacks. There is a significant proportion in which the result will be indeterminate due to cardiorespiratory disease or an abnormal radiograph14. Once these are taken into account, the overall radiation burden to the population is only slightly (approximately 4%) higher using CTPA for all15. Moreover, the greater specificity of a positive CTPA, compared with a ‘high probability’ of PE, gives the clinician greater confidence in the diagnosis. The negative predictive value of a CTPA for PE is 99.4% in patients at 3-month follow-up, with a mortality attributable to PE within this time period of 0.01%16. The current generation of MDCT machines produces consistently high quality images of the pulmonary arteries down to the subsegmental level. However, the clinical significance of tiny peripheral emboli is unclear

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as they may not affect outcome17. Consequently, CTPA is now widely considered to be the equal of pulmonary angiography in terms of accuracy without the associated risks. Imaging of the venous system 2 min postinjection from the iliac crests to the popliteal fossa has been advocated by several groups for all patients undergoing CTPA. The evidence suggests that 20% of patients with a negative CTPA will have venous thrombosis in the legs18. However, it is difficult to justify replacing a highly accurate, noninvasive and radiationfree test (ultrasound) with CT, particularly as the gonads are unavoidably included in the imaging.

Vascular The high resolution and speed afforded by MDCT has a number of potential applications in the assessment of the vascular system elsewhere. The increased speed allows a ‘pure’ arterial phase study to be obtained; however, the faster the machine, the less forgiving it is. It is imperative that all such studies are either triggered by the arrival of contrast medium at a defined site or are based on a test bolus. In our practice, unenhanced and arterial-phase images are acquired. This aids detection of intramural haematoma or haemorrhage in acute patients and differentiates calcification from endoleak in treated patients. The commonest vascular application clinically is the assessment of aortic aneurysms, both before and after treatment.The enhanced resolution, particularly in the z-axis, together with multiplanar reconstruction (Fig. 4.6), allows accurate planning of endovascular or surgical repair. Following endovascular aneurysm repair (EVAR), MDCT is regarded by the Society of Interventional Radiology (SIR) as the gold standard for monitoring the aorta. Particular attention should be paid to the presence of any fractures or kinking of the stent on the unenhanced study. Following intravenous contrast medium,

Figure 4.6 Colour volume rendered image of the aorta demonstrating an infrarenal aneurysm extending down to the bifurcation.

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the presence of an endoleak will be seen as contrast medium within the aorta tracking into the thrombus and implies failure of the stent graft. Enlargement of the aneurysmal sac without an apparent endoleak is termed endotension. Visceral infarction, particularly involving the kidneys, should also be apparent on the arterial study. When using MDCT to investigate patients with peripheral vascular disease, it is important to appreciate that the flow rate down both normal and atherosclerotic peripheral vessels can be highly variable. Consequently, it is our practice to perform two test bolus studies: one in the lower aorta and a second in the popliteal fossa. We then subtract the time delay between the two and set the table speed accordingly. Post-processing of the data can be time consuming and the images in the distal vessels can be variable, as with all forms of peripheral vascular imaging. In our experience, the results from 16-detector row and above are comparable with magnetic resonance angiography (MRA) and have the advantage of providing bony landmarks to help treatment planning. Both of these techniques yield a 3D dataset, unlike conventional arteriography, and have the advantage of being noninvasive. It would seem likely that the role of MDCT in this field will grow in coming years, with arteriography increasingly reserved for therapeutic cases.

Trauma Accurate assessment of the cervical spine is of critical importance in trauma cases: the consequences of missing a significant injury are potentially devastating. The reported accuracy of plain radiographic studies is highly variable, but approximately 10–20% of significant cervical spine injuries will be missed through a combination of suboptimal radiographs, interpretation error and absence of radiographic signs19. In the obtunded patient, these problems are compounded by the absence of any clinical information from physical examination. Single detector row CT is superior to plain radiography, but the relatively poor resolution in the z-axis could prove problematic in identifying fractures in the axial plane. MDCT has overcome this with increased speed and resolution, such that the entire cervical spine can now be imaged and viewed in any plane within a matter of seconds. The sensitivity, specificity and NPV of MDCT are reported to be 98.1%, 98.8% and 99.7%, respectively, with many major trauma centres now using MDCT as the primary imaging method for evaluating the cervical spine in the obtunded trauma patient20,21. Imaging of the thorax and abdomen following trauma is usually performed with MDCT; ultrasound may be of use in confirming the presence of a haemoperitoneum in the unstable patient before laparotomy, but cannot reliably assess the full extent of solid organ or GI tract injuries. In addition to the non-enhanced imaging, the speed of image acquisition with MDCT allows the radiologist to perform each phase of the study at an appropriate delay following intravenous enhancement. The thorax should be imaged in the arterial phase, as deceleration injuries may lead to shearing injuries of the aorta (most commonly occurring at the isthmus, just distal to the origin of the left subclavian artery). The abdomen should ideally be imaged in the arterial and portal venous phases, in order

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that both vascular injuries (dissection, pseudoaneurysm, avulsion) and parenchymal injuries may be assessed. The pelvis is usually imaged in the portal venous phase. The sensitivity and specificity of CT for detecting bowel and mesenteric injury are approximately 80%22, while for solid organ injuries these approach 100%.

Gastrointestinal CT colonography (CTC) has been extensively investigated over the past decade, with the ultimate goal of providing an accurate, minimally invasive screening test for detecting colonic polyps before they become frankly malignant. The evolution of MDCT has enabled the development of CTC. Current best practice involves thin collimation (≤ 3 mm) and low dose imaging of the insufflated colon in both prone and supine positions following either purgative or faecal tagging with multiple doses of oral contrast media for 48–72 h23. The use of intravenous hyoscine is recommended by some authors. Reporting is usually, though not always, performed from the axial and MPR images with 3D endoluminal fly-through views used for problem solving. However, the precise techniques employed are as variable as the results achieved, which range from as low as 59% sensitivity for polyps > 10 mm24 to 93% (outperforming colonoscopy) in the detection of polyps > 8 mm25. CTC outperforms double contrast barium enema (DCBE) in all studies. The clinical significance of ‘missed’ and diminutive polyps has also been questioned, given the slow progression of polyps over time26. Moreover, there is wide variation in the reporting of polyps even among ‘expert’ readers27, illustrating that this truly is a technique in evolution from a technical perspective, with a steep learning curve for the radiologist. Multicentre trials are currently underway in the USA (ACRIN) and the UK (SIGGAR 1) to evaluate the use of CTC in the context of a screening programme, but at present the majority view is that colonoscopy should be viewed as the gold standard, with CTC likely to replace the DCBE in the future. Again CAD techniques are under evaluation.

identifying the arterial and venous anatomy of the donor for surgical planning. Following transplantation, we have found MDCT to be particularly useful in evaluating the vascular supply to the graft in cases of suspected ischaemia where US is nondiagnostic (Fig. 4.7).

Genitourinary The evaluation of a patient presenting with haematuria may involve intravenous urography, ultrasound, CT or MRI in combination with cystoscopy. Increasingly, MDCT is being used as the primary imaging investigation as it has a high sensitivity for detecting malignant lesions, calculi and traumatic injuries. The other major causes (infection, coagulopathy, instrumentation) may be diagnosed on clinical history and blood and urine testing.The technique employed should encompass an unenhanced study (to detect calculi), a nephrographic phase (approximately 100 s, to assess the renal parenchyma) and an excretory phase (8–10 min, to assess the ureters). The latter two stages may be combined by giving the intravenous contrast medium in two parts, 8–10 min apart, and then imaging 100 s after the second dose. This will significantly reduce the radiation dose to the patient. As MDCT may not depict flat tumours of the bladder wall, cystoscopy must not be omitted. The increased anatomical resolution and multiplanar reconstructions possible with MDCT have enabled its use for surgical planning. In particular, urologists are able to accurately assess whether resection of a malignant lesion requires partial or complete nephrectomy, whilst potential organ donors may be noninvasively evaluated with an accurate depiction of the renal vasculature.

Paediatric Imaging children with CT poses a number of problems. It requires cooperation with instructions and the need to remain still, requires a high degree of spatial resolution to depict smaller organs and their vessels and exposes patients to a relatively high dose of radiation. Some hold that around 1 per

Hepatobiliary The use of multiple phases of contrast-enhanced imaging has probably had the greatest impact in liver imaging, which is reflected in the extensive US, CT and MR literature on the subject. The detection and characterization of focal liver lesions is largely based on patterns of vascular enhancement, with the hypervascular lesions—hepatocellular carcinoma (HCC), regenerative nodules, focal nodular hyperplasia and adenoma—in particular providing diagnostic dilemmas. In order to optimize contrast enhancement, our practice is to trigger the CT study from the arrival of contrast in the abdominal aorta. CT-based surveillance programmes for early detection of HCC with cirrhosis consist of an arterial and portal venous phase study at the very least, with many centres also obtaining unenhanced images. Both early and late arterial phase imaging can be performed, but no additional benefit for this has yet been shown. In living related liver transplantation, MDCT may be used to determine liver volumes and is increasingly utilized for

Figure 4.7 Coronal MIP image demonstrating hepatic artery thrombosis with collateral arteries in a patient following liver transplantation.

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1200–2000 patients undergoing CT might develop cancer because of the effects of the CT radiation, and these risks are greater in children28. Advances in MDCT have led to a significant reduction in acquisition time, which has been shown to reduce the need for sedation of paediatric patients29. Simultaneously, it has become possible to reduce the collimation below 1 mm, dramatically improving the resolution, and thereby not only aiding the diagnosis but also enabling accurate assessment of congenital abnormalities and surgical planning, particularly in patients with malignancy. Unlike in adult patients, imaging in different vascular phases is discouraged because of radiation issues. Therefore it is important to select a single optimal imaging sequence whenever possible (e.g. in a Wilms’ tumour, the late arterial phase will also opacify the renal veins and upper IVC). Optimizing an MDCT study for a paediatric patient should also include reducing the radiation dose as much as is possible while maintaining diagnostic quality. Reducing the kV and/or mAs will significantly reduce the dose to the patient30, and this can now be modulated automatically during imaging on newer machines, giving a constant image signal-to-noise ratio throughout the study. Additionally, increasing the pitch reduces the radiation dose significantly without loss of diagnostic quality30.

WHOLE BODY MDCT IN ASYMPTOMATIC ADULTS Whole body CT imaging of asymptomatic patients is controversial, yet is becoming commonplace in several countries, particularly the United States. Clearly, the potential benefits of identifying early stage malignant lesions or coronary heart disease with one noninvasive test could be enormous, but as yet the case for whole body CT screening is far from proven. When considering these studies, it should be borne in mind that a single CT examination cannot be optimized for looking at all organs at once. Consequently, the sensitivity and specificity of a whole-body study is likely to be significantly lower than with the dedicated organ-specific studies currently being evaluated. Inevitably, further follow-up investigations will be required for incidental findings, resulting in anxiety and distress for the patient and occasionally lead to morbidity and mortality. Moreover, the radiation dose is nontrivial. A recent publication exploring the hypothetical situation of a 45-year-old man undergoing annual screening CT until the age of 75 calculated the overall estimated lifetime attributable risk of cancer mortality to be approximately 1.9% (1 in 50)31. This combination of high cost, low specificity and high risk indicates that it is unlikely that whole body CT screening would be of benefit to the population. In the UK, such CT examinations are not recommended by the National Screening Committee32.

RADIATION DOSE CONSIDERATIONS In the late 1980s, CT represented approximately 2% of radiological investigations and 20% of the collective dose to the

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population. In 2003–2004, it is estimated that this has increased to 9%, with the dose from CT contributing around half of all radiation from medical exposures33. Since 1997, CT has been designated a high-dose procedure by the European Union, along with interventional radiology and radiotherapy. The reasons for this expansion in practice have been outlined above, with MDCT able to perform increasingly complex diagnostic procedures noninvasively and, in some instances, requiring multiple phases of imaging. Furthermore, unlike conventional plain radiography, where the radiation dose from each exposure is to some extent regulated by automatic detectors (increasing the dose will result in an overexposed, dark radiograph of little diagnostic value), increasing the exposure factors (and patient dose) for a CT study will provide the user with higher quality images. More difficult to assess is the effect that an increasingly litigious society has had on the practice of medicine and in particular the number of requests for studies ‘to rule out’ underlying tumour, pulmonary embolism, etc., even when these are clinically unlikely. The responsibility for reducing patient dose should be shouldered by all parties. The referring clinician should ensure that the radiologist is given full clinical information to ensure that CT is indeed the most appropriate test. The radiologist should ensure that each study is justified, that the imaging protocols are optimized to answer the clinical question and that the dose to the patient follows the ALARA (as low as reasonably achievable) principle. CT manufacturers also play a key role in this area through the continued development of dose modulation and the installation of low-dose preset protocols. In order to maximize patient safety, it is essential that all these issues are addressed in each case.

RADIOTHERAPY Over the past decade, significant advances have been made in the planning and delivery of radiotherapy. The introduction of MDCT enabled the oncologist to map the extent of tumour in three planes, allowing accurate dose planning of irregular shapes while minimizing the dose to the surrounding normal tissues. This 3D conformal radiotherapy (CRT) results in a significant reduction in side-effects. However, there were limitations in corrections that could be made to the delivered dose. More recently, intensity modulated radiotherapy (IMRT) has been developed, whereby each radiation beam is divided into 1-cm2 ‘beamlets’, each delivering a different prescribed dose. This has reduced even further the dose delivered to the surrounding normal tissues through increasing accuracy of delivery. Current research in radiotherapy includes the development of techniques to account for movement during radiotherapy, particularly respiration, with 4D CRT. Moreover, techniques are being developed which take account of the changes to the patient and tumour (weight loss and tumour shrinkage) that occur during a course of radiotherapy, so-called adaptive radiotherapy. It is anticipated that both of these will again further the accuracy of radiotherapy with more precise dose delivery and reduced side-effects. The increasing use of functional imaging of tumours with 18FDG-PET is also beginning to impact on

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radiotherapy planning. The co-registration of anatomical and physiological images using CT-PET has resulted in significant advances in lung tumour radiotherapy in particular, as it is now possible to differentiate between the central obstructing tumour and distal collapse.

10.

11.

FUTURE DIRECTIONS The gantry speeds and coverage currently achieved in MDCT are more than adequate for almost all clinical applications. The most notable exception, cardiac imaging, will undoubtedly drive manufacturers to produce yet faster machines with greater coverage until high-quality,‘real-time’ images are achieved, which could potentially directly replace all diagnostic coronary angiography. A significant improvement in resolution will be more difficult to achieve using the current technology with up to 64 banks of detectors.This would require significant improvements in detector efficiency and size reduction. However, the next generation of CT machines may instead employ flat panel detectors. Early reports of prototype machines indicate that spatial resolution is greatly improved, with isotropic voxels of the order of 0.25 mm, with coverage of approximately 20 cm in the z-axis per revolution. However, contrast resolution is relatively poor and data acquisition is still slow on these machines at present. One of the most important potential applications for such high-resolution studies, in addition to those described above, would be imaging of the breast. Early reports indicate that the technique is feasible at a dose comparable to that from 2-view mammography, with the added advantages of better tumour localization and potentially better detection of tumours. The tumour detection and dose characteristics will clearly be of greater importance than the speed of data acquisition for breast work, but increased speed will be crucial for cardiac applications.

REFERENCES 1. Ambrose J, Hounsfield G N 1973 Computerized transverse axial tomography. Br J Radiol 46: 148–149 2. Kalender W A, Seissler W, Klotz E, Vock P 1990 Spiral volumetric CT with single-breath-hold technique, continuous transport, and continuous scanner rotation. Radiology 176: 181–183 3. Tipper G, U-King-Im J M, Price S J et al 2005 Detection and evaluation of intracranial aneurysms with 16-row multislice CT angiography. Clin Radiol 60: 565–572 4. Matsumoto M, Kodama N, Sakuma J et al 2005 3D-CT arteriography and 3D-CT venography: the separate demonstration of arterial-phase and venous-phase on 3D-CT angiography in a single procedure. Am J Neuroradiol 26: 635–641 5. Miles K A, Hayball M, Dixon A K 1991 Colour perfusion imaging: a new application of computed tomography. Lancet 337: 643–645 6. Gillard J H, Minhas P, Hayball M P et al 2000 Evaluation of quantitative computed tomographic cerebral perfusion imaging with H215O positron emission tomography. Neurol Res 22: 457–464 7. Wintermark M, Fischbein N J, Smith W S, Ko N U, Quist M, Dillon W P 2005 Accuracy of dynamic perfusion CT with deconvolution in detecting acute hemispheric stroke. Am J Neuroradiol 26: 104–112 8. Wintermark M, van Melle G, Schnyder P et al 2004 Admission perfusion CT: prognostic value in patients with severe head trauma. Radiology 232: 211–220 9. O’Rourke R A, Brundage B H, Froelicher V F et al 2000 American College of Cardiology/American Heart Association Expert Consensus document

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on electron-beam computed tomography for the diagnosis and prognosis of coronary artery disease. Circulation 102: 126–140 Nieman K, Cademartiri F, Lemos P A, Raaijmakers R, Pattynama P M, de Feyter P J 2002 Reliable noninvasive coronary angiography with fast submillimeter multislice spiral computed tomography. Circulation 106: 2051–2054 Ropers D, Baum U, Pohle K et al 2003 Detection of coronary artery stenoses with thin-slice multi-detector row spiral computed tomography and multiplanar reconstruction. Circulation 107: 664–666 Schuijf J D, Bax J J, Salm L P et al 2005 Noninvasive coronary imaging and assessment of left ventricular function using 16-slice computed tomography. Am J Cardiol 95: 571–574 Cademartiri F, Mollet N R, Runza G et al 2006 Diagnostic accuracy of multislice computed tomography coronary angiography is improved at low heart rates. Int J Cardiovasc Imaging 22: 101–105; discussion 107–109 Cross J J L, Kemp P M, Walsh C G, Flower C D R, Dixon A K 1998 A randomized trial of spiral CT and ventilation perfusion scintigraphy for the diagnosis of pulmonary embolism. Clin Radiol 53: 177–182 O’Neill J, Murchison J T, Wright L, Williams J 2005 Effect of the introduction of helical CT on radiation dose in the investigation of pulmonary embolism. Br J Radiol 78: 46–50 Quiroz R, Kucher N, Zou K H et al 2005 Clinical validity of a negative computed tomography scan in patients with suspected pulmonary embolism: a systematic review. JAMA 293: 2012–2017 Prologo J D, Gilkeson R C, Diaz M, Cummings M 2005 The effect of single-detector CT versus MDCT on clinical outcomes in patients with suspected acute pulmonary embolism and negative results on CT pulmonary angiography. Am J Roentgenol 184: 1231–1235 Cham M D, Yankelevitz D F, Henschke C I 2005 Thromboembolic disease detection at indirect CT venography versus CT pulmonary angiography. Radiology 234: 591–594 Tins B J, Cassar-Pullicino V N 2004 Imaging of acute cervical spine injuries: review and outlook. Clin Radiol 59: 865–880 Brohi K, Healy M, Fotheringham T et al 2005 Helical computed tomographic scanning for the evaluation of the cervical spine in the unconscious, intubated trauma patient. J Trauma 58: 897–901 Griffen M M, Frykberg E R, Kerwin A J et al 2003 Radiographic clearance of blunt cervical spine injury: plain radiograph or computed tomography scan? J Trauma 55: 222–226 Elton C, Riaz A A, Young N, Schamschula R, Papadopoulos B, Malka V 2005 Accuracy of computed tomography in the detection of blunt bowel and mesenteric injuries. Br J Surg 92: 1024–1028 Taylor S A, Halligan S, Bartram C I et al 2003 Multi-detector row CT colonography: effect of collimation, pitch, and orientation on polyp detection in a human colectomy specimen. Radiology 229: 1–2 Rockey D C, Paulson E, Niedzwiecki D et al 2005 Analysis of air contrast barium enema, computed tomographic colonography, and colonoscopy: prospective comparison. Lancet 365: 305–311 Pickhardt P J, Choi J R, Hwang I et al 2003 Computed tomographic virtual colonoscopy to screen for colorectal neoplasia in asymptomatic adults. N Engl J Med 349: 2191–2200 Macari M, Bini E J, Jacobs S L et al 2004 Significance of missed polyps at CT colonography. Am J Roentgenol 183: 127–134 Johnson C D, Harmsen W S, Wilson L A et al 2003 Prospective blinded evaluation of computed tomographic colonography for screen detection of colorectal polyps. Gastroenterology 125: 311–319 Brenner D J, Elliston C D, Hall E J, Berdon W E 2001 Estimated risks of radiation-induced fatal cancer from pediatric CT. Am J Roentgenol 176: 289–296 Pappas J N, Donnelly L F, Frush D P 2000 Reduced frequency of sedation of young children with multisection helical CT. Radiology 215: 897–899 Donnelly L F, Emery K H, Brody A S et al 2001 Minimizing radiation dose for pediatric body applications of single-detector helical CT: strategies at a large children’s hospital. Am J Roentgenol 176: 303–306 Brenner D J, Elliston C D 2004 Estimated radiation risks potentially associated with full-body CT screening. Radiology 232: 735–738 Dixon A K 2004 Whole-body CT health screening. Br J Radiol 77: 370–371 Shrimpton P C, Hillier M C, Lewis M A, Dunn M 2005 Doses from computed tomography (CT) examinations in the UK—2003 review. National Radiation Protection Board, Oxon, UK

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Magnetic Resonance Imaging: Basic Principles

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Iain D. Wilkinson and Martyn N. J. Paley

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Historical perspective Spin physics Excitation and relaxation: free induction decay and echoes Signal localization: techniques for building images Instrumentation: magnets, coils and computers Physical parameters that provide contrast Pulse sequences Safety considerations Conclusion

This chapter considers the basic principles of the MRI technique. Magnetic resonance imaging (MRI) is a noninvasive method of mapping the internal structure and certain aspects of function within the body. It uses nonionizing electromagnetic radiation and appears to be without exposure-related hazard. It employs radiofrequency (rf) radiation in the presence of carefully controlled magnetic fields in order to produce high quality cross-sectional images of the body in any plane. It portrays the distribution of hydrogen nuclei and parameters relating to their physical surroundings in water and lipids. MRI has progressed over 30 years from being a technique with great potential to one that has become the primary diagnostic investigation for many clinical problems. Its application, initially limited to the neuro-axis, now covers all regions of the body and an increased knowledge base has provided a better understanding of how it can best be utilized, either alone or in conjunction with other techniques, in order to maximize diagnostic certainty. Technical advances have included improvements in spatial resolution, types of contrast and, in particular, speed of imaging. The range of information that can be obtained from a multitude of different types of image contrast is one of its biggest selling points: modern MR is truly multimodal, capable of depicting function as well as anatomy with a high sensitivity to the presence of disease.

HISTORICAL PERSPECTIVE The phenomenon of ‘nuclear induction’, later to be termed nuclear magnetic resonance (NMR), was described independently but almost simultaneously by Bloch1 and Purcell2 and their colleagues in 1946 and for this they were jointly awarded the Nobel Prize for Physics in 1952. Much later, for some reason the term ‘nuclear’ was dropped. In applications to medicine, it is now commonly referred to as magnetic resonance (MR). Since its discovery, NMR has been used extensively as a laboratory method for studying the properties of matter at the molecular level (NMR spectroscopy). In 1971, Damadian noted that in vitro, animal tumours had elevated MR relaxation times when compared to normal control tissue, and in the following year he filed a US patent: ‘apparatus and method for detecting cancer in tissue’. The application of MR to imaging required a method of spatial localization. Published in 19733, Lauterbur’s paper showed how this could be done by applying a linearly varying magnetic field across a liquid. In the same year, Mansfield and Grannell reported revealing structure within a solid by the use of a linear magnetic field gradient4. Human in vivo images were first published in 1977 by Mansfield and Maudsley5, Damadian et al6 and Hinshaw et al7. Multiplanar imaging ability was first demonstrated by Hawkes et al in 1980, who also reported the first demonstration of intracranial lesions8, and 1980 also saw the introduction of the basis of the most common spatial encoding method – spin-warp imaging by Edelstein et al9.The introduction and subsequent improvements in diffusion, perfusion, flow and spectroscopy from a large number of groups have been providing fascinating new insights into cerebral and cardiac pathology10–16. Functional MRI (fMRI) of the brain, which uses endogenous blood oxygen level dependent (BOLD) contrast, was introduced by Ogawa et al in 199217 and developed by Rosen et al18. Methods have been developed that have enabled areas that have previously been difficult to assess, such as pulmonary ventilation, to be imaged with the aid

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of hyperpolarized helium gas19. New parallel encoding methods used with phased-array rf coils, such as SMASH introduced by Sodickson et al in 199720 and SENSE introduced by Pruessmann et al in 199921, are changing the rules of imaging speed, particularly when coupled with the introduction of 3T whole-body systems.

SPIN PHYSICS MR describes the phenomenon whereby the nuclei of certain atoms, when placed in a magnetic field, absorb and emit energy of a specific or resonant frequency. Nuclei suitable for MR are those which have an odd number of protons or neutrons and therefore possess a net charge distribution. They also exhibit the property of nuclear spin, which gives them angular momentum. The combination of charge and angular momentum causes these nuclei to behave as magnetic dipoles. Almost all medical images produced to date have used the simplest of all nuclei, that of hydrogen (a single proton), which is present in virtually all biological material and exhibits relatively high MR sensitivity. Other relevant naturally occurring nuclei include phosphorus (31P), sodium (23Na), carbon (13C) and potassium (39K). Noble gases such as helium (3He) and xenon (129Xe) can also be made sufficiently sensitive by laser pre-polarization techniques. The proton can be regarded as a small, freely suspended bar magnet spinning rapidly about its magnetic axis. Place a group of protons in a uniform magnetic field and their magnetic moments experience a torque, tending to line them up with the applied field. Due to thermal energy not all of the spins line up and at body temperatures the difference in numbers between those that do and those that do not (the net magnetization) is small. The stronger the applied magnetic field, the larger the net magnetization and the higher the available MR signal.The direction of the applied magnetic field conventionally defines the z-axis, which is generally in the craniocaudal direction in the common cylindrical imaging magnet configuration. Because the nuclei are spinning, they respond to the magnetic torque like a gyroscope and their axes are tilted so they rotate about the magnetic field’s axis in a movement termed ‘precession’. The frequency of precession is directly proportional to the applied magnetic field. For protons in a field of 3T, it is almost 128 MHz. This relationship is expressed by the Larmor equation: f0 = γB0 where f0 is the resonant frequency, γ is the gyromagnetic ratio (constant for a nuclear species) and B0 is the applied magnetic field.

EXCITATION AND RELAXATION: FREE INDUCTION DECAY AND ECHOES If rf energy is then applied, there is a strong interaction or resonant effect, providing that the frequency of the rf is equal

to the proton precession frequency. This is called magnetic resonance (MR) and manifests itself in the following way: rf energy is absorbed by the nuclei causing the motion of the nuclear dipoles to be disturbed (they become ‘excited’). Their net magnetization along the z-axis deviates through an angle that depends upon the amount of rf energy absorbed. Once the applied rf is turned off, the magnetization gradually returns to its equilibrium position. As it does so, the changing magnetization induces a small voltage in a receiver coil. In its simplest form, the electrical signal detected following an rf pulse is known as the free induction decay (FID), the size and length of which is determined by the sample’s proton density and relaxation times. The time it takes for the net magnetization to revert back to its equilibrium position along the longitudinal z-axis is governed by the longitudinal or T1 relaxation time. In this way, energy is given up by the spins to the overall structure or lattice (T1 can be called the spin–lattice relaxation time) and once the energy associated with a given excitation pulse has decayed away by this mechanism, it cannot be retrieved: you have to start again by applying another rf excitation pulse, imparting more energy to the system. In addition to the transfer of energy to the overall structure, the spins interact with each other. As this occurs, their net magnetization in the x–y plane (perpendicular to the z-axis) becomes less coherent or dephases. The rate at which it does so is governed by T2, the transverse or spin–spin relaxation time. As will be described below, for a given rf excitation pulse, part of the T2 decay or dephasing can be reversed, leading to the production of echoes. Due to the different nature of T1 and T2 decay, T2 can never be longer than T1. In liquids, T2 starts to approach the value of T1 and in human tissue, T2 is approximately a tenth of T1.The spin–spin interaction (T2 decay) can be hastened by the presence of local changes in magnetic field homogeneity. When this occurs, the transverse magnetization decays with time constant T2*, which is shorter than its intrinsic T2. So far, the production of a FID has been described, occurring after termination of an rf excitation pulse.To facilitate the spatial encoding process and to enable ‘weighting’ of the signal by various contrast mechanisms in differing proportions, it is necessary to delay the production of the signal. Within the overall decay envelope given by T2 relaxation, a signal echo (or echoes) can be formed in different ways. The two most common types of echo are the spin echo and the gradient echo (or field echo) (Fig. 5.1). Both cause rephasing or refocusing of the transverse magnetization (in the x–y plane) at an echo time (TE) following the excitation pulse. Spin echoes are produced by an 180-degree, second rf pulse which flips the spins over in the x–y plane: instead of spreading further apart, they refocus, moving into phase with each other again, pass through a maximum and then dephase once more. Gradient echoes are, as the name suggests, produced by the introduction of a magnetic field gradient: the gradient causes the spins to realign in the x–y plane while they are ‘pointing’ in the same direction (i.e. without flipping them over). At sufficiently long TE, spin echoes and gradient echoes produce signals that are

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90
rf

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Echoes

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Figure 5.1 Schematic depiction of signal formation: (A) application of a 90-degree rf pulse leads to the production of the free induction decay (FID); (B) a spin echo (SE) is produced at the echo time (TE) by the application of a 180-degree rf pulse at TE/2; (C) a gradient echo (GE) is produced by the application of a magnetic field gradient (G).

T2-weighted and T2*-weighted, respectively; that is, different tissues having different T2 and T2* relaxation times yield different signals that depend on the values of those relaxation times. There are other echo types: Hahn (or partial) echoes and stimulated echoes. The former will be briefly mentioned in the section outlining gradient-echo sequences.

SIGNAL LOCALIZATION: TECHNIQUES FOR BUILDING IMAGES The physical relationship that makes building up an image possible is the proportional relationship of the resonant frequency to the strength of the magnetic field (Larmor equation again). In understanding the MR spatial-encoding process, two concepts need to be separated: that of resonant frequency (as given by the Larmor equation: spins precessing at a high resonant frequency when in a high magnetic field) and spatial frequency (the high spatial frequency components of an image corresponding to the fine detail in that image). Imaging is performed using the properties of the Larmor relationship but, in addition to this, the data are encoded in the spatial frequency domain! The latter refers to the sampled signal being in the wonderful world of k-space. Different gradient schemes have been devised to traject through k-space in different ways but most images are acquired using a two-dimentional (2D) encoding technique which samples kspace as a set of rectangular Cartesian coordinates (like the regular spacing in a tight climbing net). The beauty of this method is that the image is then just the 2D Fourier transform of the sampled signal (Fig. 5.2). The basic components of this strategy can be split into three parts: 1 Slice selection is performed during the rf excitation process. The excitation pulse consists of a narrow bandwidth of frequencies. A linearly increasing field (field gradient) is applied across the object in the slice-select direction; the combination of the limited excitation bandwidth in the

Figure 5.2 (A) Two-dimensional (2D) ‘k-space’ magnitude data array and (B) corresponding sagittal T1-weighted proton density image formed after 2D Fourier transformation.

presence of the gradient means that only spins within a finite distance in the slice-encode direction (i.e. a 2D slice) will resonate and be excited. The slice position and thickness can be altered by adjusting the rf centre frequency, pulse bandwidth and/or gradient strength (rate of change of magnetic field across the object). A negative gradient lobe is added after excitation so that once slice selection has been accomplished, the spins across the chosen slice precess in-phase with one another. 2 Phase encoding is performed following excitation and prior to the formation of the signal echo. A second gradient is applied along one direction within the plane of the excited slice. While it is on, the presence of the gradient causes a linear change in spin precession frequency along its direction. As with slice selection, once it has been switched off, all spins within the slice revert to precessing at the static field’s resonance frequency. However, the phase-encoding gradient will have introduced an incremental phase shift along its direction, the size of which depends upon its magnitude. This corresponds to one sampled spatial frequency (one value of ky) along the phase-encode direction. In order

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to sample other spatial frequencies in that direction, the excitation process is repeated to fill the desired matrix size. During each repeat, which occurs every repeat time (TR), the strength of the phase-encode gradient is altered, so that a different phase shift is introduced, thereby encoding a different spatial frequency in the phase-encode direction. 3 Frequency encoding is performed during data sampling of the signal echo. A third linear field gradient is applied in the imaging plane, perpendicular to the phase-encode and slice-selection directions. At different positions along the frequency-encoding direction, spins will be resonating at different resonant frequencies as the signal is sampled. As the echo evolves, successive spatial frequencies of the object are encoded along this third dimension. Thus, in a basic spin-echo or gradient-echo technique (Fig. 5.3), during any one TR, the slice is selected, the total number of spatial frequencies (say 256 or 512) is sampled along the frequency-encode direction at one spatial frequency along the phase-encode direction. This process is repeated for the total number of spatial frequencies required in the phase-encode direction (say 256 or 512). This can lead to inappropriate, excessive imaging times if high spatial resolution is required (> 25 min for 512 phase encodes with a TR of 3 s). However, developments in sequence design have enabled multiple phase-encoding steps to be performed within each TR, speeding up the process using, for example, fast spin-echo (FSE) techniques (Fig. 5.4). It is important to realize that individual elements in k-space do not correspond on a one-to-one basis with individual pixels in the final image, but rather that each k-space element contains information about that spatial frequency component throughout the image. An alternative method to slice selection is to use an additional (secondary) phase-encoding gradient along the slice selection axis, yielding true three-dimensional (3D)-encoded datasets that typically have high spatial resolution in all three directions (often submillimeter). Putting together phase and frequency encoding plus slice selection has introduced the concept of a pulse sequence that consists of a series of applied rf and gradient pulses that lead to

TR 90


90


RF Slice Phase Frequency TE Signal

Figure 5.3 Basis of the gradient-echo (spin-warp) pulse sequence. Three orthogonal gradients are used for: slice selection (while the rf excitation pulse is applied); phase encoding (the amplitude of which is iterated every repeat time [TR]) and frequency encoding (leads to echo formation at the echo time [TE]). The process is repeated after time TR for the total number of required phase-encoding steps.

RF

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Figure 5.4 Train of 180-degree refocusing rf pulses that form the basis of the fast spin-echo (FSE) sequence. Each 180-degree rf pulse has a different magnitude phase-encoding gradient. A whole train of 2, 4, 8, etc. echoes are produced per TR, reducing the number of TRs required for a given total number of phase-encoding steps.

the production of the signal in the form of an echo which has been encoded in k-space.Variations in the timing of the component parts of these pulse sequences can be used dramatically to alter image contrast; this being one of the major advantages of MR, helping to identify anatomy and characterize normal and pathological tissue types. Adjustments to the building blocks and parameters that define a pulse sequence can be used to manipulate image contrast. Most sequences in use today are based on variants of the above. There are several other methods of moving or trajecting through k-space that have potential advantages. Of these, the most commonly used is echo-planar imaging (EPI). EPI was one of the first imaging schemes to be proposed. It requires very rapid gradient switching and is thus technically demanding and, in its early forms, prone to various artefactual complications (such as Eddy current effects). Technical developments have enabled its introduction on commercial clinical systems over the last 10 years. Its speed is a great asset, being capable of yielding an image in under 100 ms, and has led to its use in areas such as imaging of water diffusion and brain function. It is important to note that MRI is a true digital imaging technique as discrete data points are sampled which, when transformed, form an image made up of 2D picture elements (pixels) or 3D volume elements (voxels).

INSTRUMENTATION: MAGNETS, COILS AND COMPUTERS The major parts of an imaging system comprise (A) the magnet subsystem to produce a spatially and temporally constant B0; (B) the gradient subsystem to produce time-varying magnetic field gradients primarily for spatial encoding; (C) the rf subsystem to transmit and receive the rf energy; and (D) the associated microprocessors/computers to specify and control the pulse sequence, calculate, process, display, store and transfer the resultant images. The magnet, gradient and rf coils are all situated within an rf-shielded examination or magnet room. This shielding (a Faraday cage) consists of a conductive metal lining (copper or aluminium) through which rf electromagnetic radiation will not pass. It serves to keep external electrical noise out and the generated rf energy within the examination room. All electrical connections into the room must be filtered appropriately as they pass through this shield. Most screened rooms possess wave guides or tubes (with a high length-to-diameter ratio; > 5:1) fixed into the wall of the

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screened room, through which nonconductive tubing may be passed (plastic gas tubing or fibre-optic cabling).The examination room door should be shut while acquiring data to enable the shield to work effectively.

Static magnetic field (Fig. 5.5) Although it is possible to perform MR experiments within the earth’s magnetic field (being approximately 0.00004T!), most clinical MRI systems are constructed around magnets which produce fields in the range 0.2–3T. Research systems capable of human in vivo imaging at 8T and above are available. There has been much debate over the optimum B0 for clinical imaging and this debate continues.The vast majority of installed systems operate at 1.5T, although there is presently a surge in systems operating at the new high-field strength of 3T. Low-field, reduced cost systems (0.2–0.5T) are often marketed for various ‘niche’ applications (e.g. musculoskeletal extremity imaging). At this low end of the range, permanent ferromagnetic composites or room temperature electromagnets tend to be utilized. Above 0.5T, superconducting windings (a niobium–tin alloy embedded in a copper matrix) are supercooled in a bath of liquid helium (at 4.2 K or −269°C) such that they offer zero resistance to the high currents required to produce high magnetic flux densities. Superconducting magnets, which are constantly ‘on’ or ‘at field’, are mostly cylindrical in geometry, especially at 1.5T and above, although more ‘open’ designs are becoming available at 1T and below. Developments in magnet technology have meant that, provided the helium refrigeration pump remains in operation (usually a rhythmical thud can be heard while the system is not acquiring data), one fill of liquid helium should last several years. The overall size of these superconducting magnet systems has decreased substantially in recent years, yielding friendlier designs with reduced levels of patient anxiety/claustrophobia. Why buy a high B0 system? The protons’ net magnetization increases with B0, leading to an overall increase in imaging signalto-noise ratio (approximately a linear relationship). Various other parameters also change with increasing B0, leading to

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potential benefits such as increased magnetic susceptibility contrast (useful for depicting nuclei within the brainstem or fMRI responses), greater resolution between resonance peaks in spectroscopy (potentially helpful for water/fat imaging) and an increase in tissue T1s (prolonged blood signal and better background suppression for time-of-flight angiography). However, higher is not automatically better as the aforementioned can also produce increased metallic foreign-body artefact, increased chemical shift artifact and changed T1 contrast. At high resonant frequencies, rf homogeneity also provides greater technical challenges. A high degree of homogeneity of the main magnetic field together with gradient linearity are essential for correct geometrical representation, this being especially pertinent when MR is used directly for therapy (e.g. in stereotactic radiosurgery to the trigeminal nerve or the placement of deep-brain stimulators). To maximize B0 homogeneity, a combination of passive and active shimming is employed. The former can take the form of a series of metallic coin-like discs or nougats, the configuration of which is usually site specific, determined and placed during installation. In high-end systems, active shimming can be computer optimized on each imaging volume for each patient episode.

Magnetic field gradients In cylindrical systems, the gradient coil sets required to produce the three orthogonal, linear gradient fields (gradients), are located adjacent to the inner surface of the magnet cryostat. The maximum gradient amplitude that can be generated (given in milliTesla per metre) has a significant effect on the minimum slice thickness and field of view and, thereby, the highest spatial resolution that can be achieved. The rate at which the gradients can achieve the desired amplitude is known as the gradient slew rate (given in milliTesla per metre per second). This sets a limit to the minimum TR and TE values that can be achieved, and thereby the minimum scan acquisition time. Applications such as fast cardiac, MR-digital subtraction angiography (MR-DSA), or high sensitivity diffusion tensor imaging require high amplitude gradients with a

Figure 5.5 MRI systems: (A) Modern system based on a cylindrical cryogenic magnet; (B) open magnet for interventional MRI; (C) low-field system for niche applications.

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fast slew rate. Fast gradient switching can lead to the induction of artefact-producing eddy currents within the conductive structures of the system. Improvements in gradient coil design and compensation methodology have significantly reduced this problem, although correct set-up during installation is essential for artefact-free use of gradient-intensive sequences such as echo-planar diffusion or sensitive techniques such as gradient localized spectroscopy. Powerful gradient amplifiers are an essential component of the gradient subsystem and are sited in a room with restricted access, adjacent to the magnet room.

Radiofrequency field Many of the rf transmit and receive subsystem components (digital-to-analogue converters, frequency mixers, power amplifiers, analogue-to-digital converters, etc.) are also sited within the electronics/cabinet room. These comprise all but the rf coils and associated receiver pre-amplifiers, which are placed close to the anatomical area under investigation in order to maximize their effectiveness. A typical system will have been purchased with a number of signal-detecting rf ‘receive’ coils, each dedicated to investigating a particular anatomical area (Fig. 5.6); often used in conjunction with the main rf transmitter body coil located adjacent to the gradient coil sets. The body coil is designed to provide a homogeneous rf excitation field throughout the required imaging volume. Many of the receiver coils will comprise several coil ‘elements’. Coil sensitivity needs to be maximized and the best signal-to-noise ratio can be obtained from a small volume adjacent to a close

fitting coil, but small coils only detect from a limited volume. To achieve high sensitivity over a large volume or area, several coil elements can be used in ‘phased-array’ configurations (e.g. for imaging the entire length of the spinal cord) or signals from multiple coils themselves (e.g. a coil ‘matrix’ covering the whole body) can be used in tandem.To facilitate this multicoil technology, many rf channels are used to pipe the detected signals away from the examination room. Thirty-two channel systems are available as standard product while those under research development can contain over 100. The number of coils/elements used is likely to grow substantially over the next few years. In addition to signal-to-noise ratio considerations, the inherent spatial information provided by arrays of coils/elements can be put to good use in parallel imaging techniques, where SENSE and SMASH technology can reduce the number of required phase-encoding steps and hence imaging time to yield a prescribed resolution. This is particularly pertinent to high-field strength systems where there is sufficient signal-to-noise ratio.

Pulse sequence control, data manipulation and image handling MR depends on highly sophisticated computing. The relentless increase in computing power has had a direct beneficial effect upon performance. Real-time image reconstruction is common without noticeable delay following data acquisition, which is important when hundreds (e.g. tissue perfusion studies) or thousands of images (e.g. BOLD fMRI studies) are acquired during a single examination. It is also now possible to

Figure 5.6 Examples of dedicated rf receive coils: (A) ‘CTL’ phased-array coil for imaging the whole spine plus typical anatomical coverage (B).

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Figure 5.6 cont’d (C) a neurovascular coil that allows MR angiography from the circle of Willis to the aortic arch (D); (E) a dedicated wrist coil designed to produce high-resolution imaging of the wrist joint (F).

register two or more datasets accurately with each other using complex computationally intensive algorithms. Slice positions can be prescribed automatically using such image registration techniques. Highly accurate comparisons between images from the same examination can be obtained ‘on the fly’ (e.g. statistical analysis of BOLD fMRI data) or from different examinations (e.g. subtraction of images to determine parenchymal volume changes in dementia or slowly evolving lesions such as meningioma). The real-time manipulation of 3D datasets (maximum intensity projections, surface rendering or general multiplanar reformatting) enables complex interactive postacquisition viewing strategies to be employed by the radiologist, which is particularly important when assessing complex vascular structures such as arteriovenous malformations or visualizing intracranial cortical developmental malformations. Digital data processing speeds will become even more critical if

ultra-high resolution imaging at very high field strengths (≥ 3T) with acquisition matrices of at least 2048 × 2048 becomes widespread. The common PC–user interface can be found on the majority of new MR systems. The introduction of picture archiving and communications systems (PACS) and the data standard, DICOM, has been evident. Local archive is still possible (commonly to DVD in new systems), but the provision of a central PACS archive is often assumed. The introduction of soft copy reviewing and reporting has meant that the radiologist is no longer confronted with fixed viewing parameters on sheets of film, but has the advantage of being able to interact with window widths and levels during the review process, enabling the interrogation of features that may be outside the acuity of the human visual system when observed on hard copy format. The incorporation of a plethora of functional investigative techniques into clinical practice (perfusion,

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exogenous contrast medium uptake, water diffusion, etc., and not just BOLD fMRI) often benefits from representation in colour overplayed on high resolution anatomical images. The rapid cinematic presentation of multiple contiguous sections can facilitate integration of information across several sections. Ciné presentations are frequently used in the analysis of dynamic studies, particularly of the vasculature.

PHYSICAL PARAMETERS THAT PROVIDE CONTRAST MR offers a vast range of possible image contrasts (Fig. 5.7), which, in turn, provides high sensitivity about the presence of many abnormalities. However, pixel values do not represent an absolute measure (e.g. absolute T2, in milliseconds). They are only absolute on the rare occasions when specialist quantitative techniques are invoked (e.g. calculation of absolute T2 of the hippocampal structures for the definition and lateralization of sclerosis in temporal lobe epilepsy). Rather, they reflect various physical parameters, including proton density, T1, and T2, and their scaling is influenced by factors such as body size (which affects rf coil loading and hence the induced voltage in the receiver coil). This is unlike CT, where pixel values in Hounsfield units correspond directly to X-ray attenuation (electron cloud density). When a particular physical parameter, e.g. the spin–spin relaxation time T2, dominates the relative pixel intensities in the image (i.e. the image contrast mostly reflects differences in T2), that image is said to be T2 weighted. The diversity of available weightings can lead to difficulties assessing what is the optimum amount of information to gather. As well as in-depth anatomical knowledge, interpretation requires knowledge regarding the type of weighting present in an image, how that relates to a particular physical contrast mechanism and how that contrast varies in the normal population and with disease. This section outlines the most common contrast mechanisms in common clinical use and emphasizes how they can manifest as artefacts or false image information if not properly analysed. Some examples are given.

Proton density It seems intuitive that the proton density (PD or ρ) of an object will fundamentally be linked to the size of the signal returned by that object. Indeed, this is the case; the higher the number of ‘MR visible’ protons, the greater the potential signal becomes. To provide image weighting so that the PD dominates pixel intensity, the effects of other potential influences should be minimized. This approach to providing the majority weighting by one factor is common practice in sequence design and protocol implementation (see ‘sequence’ section below). It is useful to rank body components according to their relative PDs. Highest are cerebrospinal fluid (CSF) and other fluids. These are followed in order by various soft tissues such as liver, kidney, spleen, grey matter and white matter, then articular cartilage, fibrocartilage, membranes, cortical bone and, lowest of all, air. It is important to recognize that the ‘MR visible’ proton pool consists of mobile protons. Hydrogen nuclei that form part of large molecular structures or are immobilized in solids have tight chemical bonds that ensure that they interact with each other so rapidly (i.e. have very short T2 values) that they provide no detectable signal over the timescales involved in standard clinical MR technique, and appear to have a PD of zero. PD is increased in some lesions.These include oedema, infection, inflammation, acute demyelination, acute haemorrhage, various tumours and cysts. Since the formation of oedema is a common response to a wide variety of insults, an increase in PD is a frequent occurrence. A decrease in PD may be seen in scar formation, fibrosis, some tumours, capsule and membrane formation, as well as with calcification. The changes in PD are often relatively small in comparison to the changes seen on T2-weighted images.

T1 (longitudinal relaxation time or spin–lattice relaxation time) This time constant characterizes the rate of the protons’ recovery from their excited state to their resting or equilibrium state, in the longitudinal or z-direction (parallel to the main field), for each component of the object. It relates to the transfer of energy from the protons to the tissue’s overall structure. Body

Figure 5.7 Patient with multiple sclerosis with plaques of demyelination shown on (A) fast spin-echo (FSE) proton density; (B) FSE T2; and (C) FSE FLAIR. There is no discernible abnormality on T1-weighted images without contrast medium (D).

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fluid T1 is long (typically 1500–2500 ms) compared to soft tissues (typically 300-1200 ms), while those of fatty structures are shorter still (100–150 ms). T1 increases with B0, requiring an alteration in sequence parameters to provide equivalent T1 image contrast at different field strengths. For any tissue, T1 is always longer than T2. Although the relaxation processes characterized by T1 and T2 are physically different, both tend to increase with most disease processes, primarily due to the accumulation of fluid associated with oedema, inflammation and infection. Despite an overall increase in T1 with pathology, T1-weighted images are often favoured for the depiction of anatomical structure. Standard T1-weighted spin-echo images can be acquired in less time than their corresponding T2-weighted counterparts, due to the reduction in TR used to introduce the degree of T1 weighting. Images with T1 weighting are heavily influenced by the presence of the most common exogenous contrast agent, chelated gadolinium (Gd) (see section below on ‘exogenous contrast media’), which can cause a dramatic increase in signal on T1-weighted imaging (Fig. 5.8).

T2 (transverse relaxation time or spin–spin relaxation time) This time constant characterizes the rate of phase dispersion of the spins’ magnetization in the transverse plane (perpendicular to the main field). This fanning out of the x–y magnetization occurs as the protons exchange energy (or quantum, mechanical chatter) with each other. For a given tissue, the dependence of T2 on B0 is small when compared to that of T1. Fluids have long T2 values (typically 600–1200 ms) those of soft tissues being shorter (typically 50–200 ms) and fat shorter still (10–100 ms). All of these are shorter than their corresponding T1 values: for a given tissue, absolute T2 can never exceed T1. The elevation of T2 in disease processes leads to a very high

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sensitivity using a standard T2-weighted spin-echo imaging technique, which is often the ‘bread and butter’ sequence for detecting lesions. When exogenous contrast media are administered, T2 is shortened. This mechanism is commonly used in abdominal imaging. For the use of Gd-chelates in the central nervous system (CNS), this effect is generally less and rarely used by comparison to its corresponding T1 effect. However, monitoring the passage of rare earth metals such as Gd is routinely performed using its modulatory effect on a closely related parameter, T2*.

T2* and endogenous susceptibility The concept of T2* was introduced above. When an image is weighted by T2*, not only is it weighted by tissue T2, but also by another component that can be ascribed to the hastening of the intrinsic spin–spin interaction by perturbations in the local magnetic field. Such perturbations are caused by the presence or introduction of objects that have different magnetic susceptibilities (χ). Magnetic susceptibility defines how magnetized an object becomes when placed in a magnetic field. When adjacent structures inside the homogeneous MR magnet have different χ, a small magnetic field gradient will be present. This gradient is the perturbing agent. Most of the constituents of the body are diamagnetic, having weak susceptibility, and they lead to a subtle reduction in the field that passes through them. Exogenous contrast agents (e.g. Gd-chelates) and some endogenous agents (such as deoxyhaemoglobin) exhibit paramagnetic properties which enhance local magnetic flux. Ferromagnetic objects exhibit an extreme form of paramagnetism, experiencing strong forces when placed in a magnetic field. Increasing B0 leads to shortening in T2* and an increase in T2* contrast. Artefacts can result from local decreases in T2*. Signal dropout adjacent to nonferrous metallic implants can be dramatic, even on low- and mid-field strength systems

Figure 5.8 Typical appearance of a brain tumour, specifically a cerebellar medulloblastoma. (A) The tumour has a low signal on spinecho T1-weighted images owing to a prolonged T1. (B) After the injection of Gd-DTPA the tumour enhances avidly, depicting breakdown of the blood–brain barrier.

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(0.2–1T). At high field, signal dropout occurs at the edges of various intracranial structures, such as the base of the frontal lobes, due to differences between the χ of cortical parenchyma and air within the sinuses. Dropout can be minimized by the use of spatially localized high-order shimming and increasing sequence bandwidth (number of Hertz per pixel). The presence of endogenous susceptibility gradients can be put to good use, providing a valuable contrast mechanism. One such example is BOLD f MRI, used not just within clinical neuroradiology, but also within associated neuroscientific specialities such as psychiatry and psychology. It is able to detect regional haemodynamic responses to simple or complex stimulation tasks. Although still not totally understood, neuronal/synaptic activity results in a change in local energy consumption which gives rise to an alteration in the ratio of intravascular oxy- and deoxy-haemoglobin and blood flow: the result being a localized net increase in the amount of fresh oxyhaemoglobin and hence T2*-weighted signal.This contrast is very subtle: activity within the primary visual or motor areas leads to signal changes of the order of 2–3% at 1.5T which approximately doubles to 4–6% at 3T. Much data have to be collected, processed and compared using computer-intensive algorithms to obtain statistically significant contrast. Clinical applications include aiding pre-operative neurosurgical planning and intra-operative neurosurgical guidance (Fig. 5.9). There are numerous other important uses of endogenous susceptibility-related, T2*weighted contrast, such as the depiction and timing of blood breakdown products following haemorrhage (Fig. 5.10).

Exogenous contrast agents Several agents have been tested, approved and are commercially available. Two groups will be considered: Gd compounds (the most common) and super-paramagentic iron oxide particles (SPIOs). Being a rare earth metal, Gd is toxic in isolation. Its toxicity is negated by appropriate chelation: the initial product to be introduced uses diethylenetriamene penta-acetic acid for this purpose (Gd-DTPA, Magnevist, Schering AG). Predominantly administered by intravenous injection, Gd-DTPA is rapidly distributed

Figure 5.9 Blood oxygen level dependent functional MRI is being investigated for its use in aiding pre-operative neurosurgical planning and intra-operative neurosurgical guidance. These data were obtained from a patient with a glioma in the left parietal lobe. Areas of activation due to finger tapping are overlaid in orange/yellow. Two anatomical slices are shown for right-hand motion and left-hand motion. Activation that correlates with right-hand movement can be seen close to the top of the lesion (bottom right).

throughout the vasculature. It is excreted by the kidneys, having a biological half-life of approximately 90 min. The agent lowers localized T1,T2 and T2*, as outlined above. Decreased T1 leads to bright pixels on T1-weighted spin-echo images (the most commonly used contrast mechanism [see Fig 5.8]), and decreased T2 can lead to less intense pixels on heavily T2-weighted images, particularly at high Gd concentrations. The effect on T2* is dramatic: the decrease causes low signal on T2*-weighted images which, on a suitable dynamic dataset, depicts parenchymal perfusion in real-time (Fig. 5.11; see section on ‘perfusion’ below).

Figure 5.10 Haematoma. A patient presents with a history suggestive of subarachnoid haemorrhage but MRI shows an extensive left frontal haematoma. The haemoglobin in the haematoma is in different stages of breakdown as shown on (A) the spin-echo T1- and (B) fast spin-echo T2-weighted images. (C) Note the high signal rim on FLAIR imaging indicative of oedema. (D) Gradient-echo sequences are very sensitive for acute haemorrhage and show prominent ‘blooming’ of reduced signal due to susceptibility effects.

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A

INTENSITY v, TIME

B

INV. INTENSITY v, TIME

0,00 13,2 28,6 39,7 52,9 66,2 79,4 92 1

2

C

ROI AREA 1. 426 4 2. 363 9

PEAK 96.2 87.6

D

0,00 13,2 28,6 39,7 52,9 66,2 79,4 92,6sec 1

2

Figure 5.11 Exogenous perfusion data obtained from a time series of T2*-weighted echo planar images in a 70-year-old woman who presented with amaurosis fugax and was found to have a 95% stenosis of the right internal carotid artery. (A) A base image shows two regions of interest within the middle cerebral artery territories. (B) A drop in signal intensity can be seen due to the first pass of the Gd-chelate from which (C) a concentration–time curve is calculated and a gamma-variate fit is performed (solid lines). (D) The resultant time-to-peak (TTP) map shows prolonged TTP in the affected hemisphere shown as high signal.

Gd-chelates do not cross the intact blood–brain barrier, but accumulate where it has been compromised (e.g. intracranial tumours, abscesses and acute demyelination).This has been one of their primary clinical uses. They are particularly effective in the delineation of small tumours, such as acoustic neuromas. High-grade malignancies generally show more enhancement than tumours of lower grade. Areas of cystic necrosis may also show enhancement but the greatest clinical value of chelated Gd has been in distinguishing tumour from oedema. Care should be exercised as the integrity of the blood–brain barrier may be altered by the modulatory influences of corticosteroids. Applications outside the CNS benefit, for example, dynamic contrast uptake characteristics may aid breast lesion differentiation. Having Gd-chelates ‘on-board’ can increase overall vessel conspicuity when using standard TOF or phase-contrast angiography techniques, although the most common angiographic application is in contrast-enhanced MRA (CE-MRA)

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and MR digital subtraction angiography (MR-DSA), where blood T1-shortening effects are used to visualize the passage of a bolus of contrast medium. If a particularly ‘tight’ bolus is needed, higher concentrations are available (e.g. 1 M Gadovist, Schering AG). Oral preparations can be used to visualize the gastrointestinal tract. The other class of agents considered here, SPIOs, is primarily for use in liver (Fig. 5.12) or spleen imaging (Endorem, Guerbet; Resovist, Schering AG).These particles have also been used in oral form for labelling the gut. In clinical applications to date, relatively large particles have been used which, after intravenous injection, lodge in the reticulo-endothelial system (Kupfer cells). Once lodged, they cause strong perturbations to the main magnetic field, producing a loss of transverse coherence, shortening T2. Tumours or other lesions which do not have reticulo-endothelial cells do not take up the SPIO and retain their high MR signal on T2- or T2*-weighted imaging. Lesions down to a few millimetres in diameter can be delineated in the liver.These iron-based agents are metabolized into the iron pools of the body and are excreted over a period of weeks. A selection of other exogenous contrast agents is available or under development, including coated iron-oxide and Gdcomplex blood-pool agents/strong albumin binding agents, hepatocyte-specific Gd-chelates (MultiHance, Guerbet), manganese-containing agents (Teslascan, GE-Healthcare) for the liver, pancreas and cortex of the kidneys; Mn-DPDP may be useful for assessment of acute myocardial ischaemia. Natural compounds, such as blueberry juice, may act as negative contrast agent in upper abdominal MR investigations, such as MR cholangiopancreatography. Other targeted agents, which may be necrosis specific (bis-Gd-mesoporphyrin), provide lymphographic contrast, or are specific for inflammation detection. The exact role of these agents is currently unclear, but they could considerably alter the practice of radiology.

Figure 5.12 Liver contrast agents—super-paramagnetic iron oxide particles (SPIOs). (A) Pre and (B) post iron oxide single-shot fast spinecho images. Both areas of FNH take up the iron oxide. Notice the central scar which has high signal on T2 and low signal on T1. (Courtesy of Dr A Blakeborough, Royal Hallamshire Hospital, Sheffield, UK.)

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Chemical shift, spectroscopy and water/fat imaging The phenomenon of chemical shift (δ) has important consequences for imaging as well as MR spectroscopy (MRS). In addition to the provision of clinical information, an overview of MRS can aid the understanding of many aspects and underlying principles of MR. Analytical chemists have used bench-top spectrometers since the 1950s in their quest to identify, understand and model molecular structure. The technique relies on information provided by chemical shift. In proton MRS (H-MRS), a hydrogen nucleus attached to a particular molecule does not ‘see’ the exact field produced by the magnet. This is because it is not a free proton; rather it is a proton in a chemical bond attached to a molecule. Neighbouring atoms that form that molecule have associated electron clouds which shield the hydrogen nucleus from the main magnetic field. The shielding effects within different molecules will differ. Back to the Larmor equation, since the resonant frequency is proportional to the field, the hydrogen nucleus will resonate at a particular frequency that is characteristic of that molecular environment. A measurement of the frequency components that emanate from a sample (in the absence of a frequencyencoding imaging gradient) will yield a frequency spectrum. This usually contains a number of peaks, the position of which (or shift along the frequency axis) is given in units of parts per million (ppm) of B0. Each peak can be attributed to a chemical group. Thus, MRS provides direct biochemical information. This can be spatially localized using a number of methods: by the spatial sensitivity of a surface coil, by the intersection of three slice-select gradients (single-voxel selection [SVS]), or a 2D or 3D metabolite map can be created by phase encoding the spectroscopic signal (chemical shift imaging [CSI]). It has become feasible to assess clinically important chemi-

cals and metabolities which are present in tissues in fairly low concentrations (millimoles in proton spectroscopy of the CNS). Spectroscopy is demanding in terms of the signal-to-noise ratio of the system and B0 homogeneity. The homogeneity within the spectroscopic region of interest must be such that any changes in resonant frequency are dominated by the chemical shift and not any heterogeneity in the applied magnetic field itself. Spectroscopy packages (particularly proton) have developed over the last 25 years to a stage where complex high-order shimming (B0 homogeneity optimization), signal localization, and data sampling and post-processing can be performed by ‘one-press’ button automation. Since its implementation in whole-body medical imaging systems, MRS has sometimes struggled to find a true clinical role. In part, that may follow from being seen as a disparate, complicated technique compared to MRI. On the contrary, it is often simpler. The technique has become more robust, forming part of a clinical investigative brain examination in many centres. It is often sensitive to the presence of disease, but lacks specificity; its clinical use (as for many imaging techniques) relies on other contextual information. Just as with standard imaging, spectra can be metabolite concentration-, T1- or T2-weighted, or a mixture of the three, depending upon acquisition parameters such as TE and TR. The brain has been the most extensively studied organ to date by proton MRS (Figs 5.13 and 5.14). If left alone, the hydrogen nuclei from water will dominate the MR ‘visible’ cerebral metabolites. This large peak from water (at 4.77 ppm) needs to be suppressed, so that an appropriate dynamic range can be obtained for metabolite detection. The three main cerebral metabolites are attributed to choline-containing

Figure 5.13 Two examples of the applications of proton spectroscopy in neuroimaging. (A) The patient is a child with known mitochondrial abnormality who presents with seizures. MRI shows extensive areas of hypoxic damage and proton spectroscopy shows the doublet characteristic of lactate. Lactate is not normally seen. The other three major peaks (choline, creatine and N-acetyl aspartate [NAA]) are normal except for a mild reduction in the Na/Cr ratio. (B) In comparison, a child with a cerebellar tumour shows a massive increase of choline and complete loss of NAA relative to creatine. This is highly suggestive of tumour and characteristic of medulloblastoma, which was confirmed on histology.

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Figure 5.14 A ‘metabolite map’ or chemical shift image (CSI) where phase encoding is used in two dimensions to map the spectral metabolites. This example shows the distribution of N-acetyl aspartate in a patient with a periventricular space-occupying lesion. A colour overlay is used. (Courtesy of Dr P. Pattany, University of Miami, Florida, USA.)

compounds (Cho) centred at 3.2 ppm, the total creatine pool (creatine plus phosphocreatine) at 3.0 ppm, and N-acetyl (NA) groups at 2.0 ppm (such as N-acetylaspartate [NAA] and N-acetyl-aspartyl-glutamate [NAG]). The biochemical role of NA is not completely understood, but the peak serves as a putative marker of neuronal integrity/function, being predominantly confined within neuronal cell bodies and axons. In a variety of disorders the NA signal may be decreased when there is little or no change on MRI. These include diseases in

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which there is acute or chronic neuronal loss, such as some of the dementias, chronic head injury, temporal lobe epilepsy and infective encephalopathies. In the latter, in particular HIV-associated encephalopathy, spectroscopy has been shown to follow clinical neurological status more closely than imaging following the administration of antiretroviral medication. Proton MRS appears complementary to MRI, which is generally of most value in acute and subacute disease. At short TEs (where T2 and modulation effects are reduced) more detail on other biochemicals within the brain, such as glutamate, glutamine and myo-inositol can be obtained. Phase-encoded CSI yields metabolite images whose spatial resolution is sufficient to allow differentiation between normal tissues and pathology. Potential applications of proton MRS to other organ systems have been investigated. Most notably it has been investigated for the diagnosis and staging of prostate cancer where the level of citrate decreases, while the choline peak is high. This is the reverse of findings in the normal peripheral zone of the prostate, where citrate is high and choline low, and benign prostatic hypertrophy with intermediate citrate and choline levels. From an imaging point of view, the slightly different resonant frequencies of fat and water can lead to band artefacts at tissue boundaries (the component water and fat images are shifted with respect to each other).These can be minimized by increasing the range of resonance frequencies covered along the frequency-encode direction (acquisition bandwidth); as ever, there is a trade-off as more noise is sampled. Chemical shift can be used to provide images of just water or just fat, or their relative contributions can be altered. Fat- or water-only images can be produced by selective excitation, where the centre frequency of an rf excitation pulse (ChemSat) of narrow bandwidth is placed over either the water or fat peak (Fig. 5.15). Another method utilizes interference effects between fat and water signals. It is useful to consider the phase of the signal. Because of the difference in resonant frequency, their signals will alternatively be in and out of phase. The detected signals cancel out when they are 180 degrees out of phase and add together when they come

Figure 5.15 Saturation of orbital fat using a chemical shift selective pulse. (A) Axial fast spin-echo T2 images in a 2-year-old child showing proptosis of the right eye. (B) Axial images through the orbits show a mass intermingled with the orbital fat and conal musculation. (C) The anatomy is clarified using a post Gd-DTPA fat-saturation sequence.

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back into phase. The phenomenon is like the beating effect from two musical notes which are slightly out of tune. At 1.5T, the first water–fat cancellation point occurs at TE = 2.2 ms (Fig. 5.16). Such imaging can provide invaluable information in certain areas (e.g. adrenal imaging).

Flowing spins The given description of the spatial encoding process in the earlier section makes a bold assumption for the production of artefact-free images – that protons stay still while they are being excited and their position is encoded. This is often a tall order for the patient in a magnet with an itchy nose and, of course, impossible for humans to have totally still spins while they have a functioning circulatory system! The different MRI behaviour of fluids can be harnessed to yield information regarding fluid flow. There are various forms of in vivo flow or flow-like movement that that can be imaged to provide spin-movement related contrast.

Macrovascular flow (angiography) Flow within large vessels (arterial and venous) can lead to various appearances using standard imaging techniques: vessels can be void of signal (Fig. 5.17) or often they return hyperintense signal. An understanding of the principles underlying the appearance of flowing blood on MRI has led to the development of magnetic resonance angiography (MRA), where the protons in flowing blood produce a high signal against a background of little or no signal from stationary tissues. Such high contrast datasets can be manipulated and viewed, often using maximum intensity projection (MIP) algorithms, to reveal projected vascular anatomy. The common contrast mechanisms can be split into three groups: (A) TOF; (B) phase contrast; and (C) contrast enhanced. Time of flight effects utilize differences in magnetization when very short TR is used. Contrast is provided between stationary background spins in the imaging volume and inflowing, fresh

Figure 5.16 Colorectal metastasis in a fatty liver. (A) In-phase and (B) out-of-phase T1 gradient-echo axial images at same level. There is marked loss of signal from the liver parenchyma on the out-of-phase image indicating fatty infiltration. Notice the metastasis in the right lobe, and an artefact from aortic pulsation in the midline. (Courtesy of Dr A Blakeborough, Royal Hallamshire Hospital, Sheffield. UK.)

Figure 5.17 Flow effects. This patient presented with right-sided seizures. (A) Fast spin-echo T2-weighted images show a large area of signal void within the area of the left sensorimotor cortex, which was present on all other sequences. This is characteristic of a high-flow vascular malformation. (B,C) Time of flight angiograms before (B) and after Gd-DTPA (C) confirm the presence of a large pial arteriovenous malformation.

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blood.The stationary spins are repeatedly exposed to the excitation pulse and, due to the short TR, there is not enough time for their longitudinal magnetization to return to equilibrium: it becomes saturated and there is a lack of magnetization with which to form a signal. The inflowing blood, on the other hand, experiences its first excitation pulse and it returns a high signal. This phenomenon provides endogenous contrast. The imaging volume can be encoded as a set of 2D slices or a 3D dataset. High field strength systems (3T in particular) can provide excellent vascular depiction due to high signal-to-noise ratio and prolongation of both the T1 of blood and stationary background protons (Fig. 5.18). Gradient-echo sequences with high T1-weighting are often used as they can be acquired at very short TR, maximizing the TOF contrast. Phase contrast techniques utilize the basic phenomenon that phase changes are introduced to the transverse magnetization when spins are exposed to a magnetic field gradient.This is put to use in phase encoding where the phase difference in a single TR is used to encode the signal at one spatial frequency in the phase-encode direction. If, after a time delay, the phase-encode gradient were to be reversed, then the phase differences in the spins across the sample introduced by the initial gradient will be completely reversed or rewound. This is only true for stationary spins. If a spin has moved in the direction of the gradient between its initial application and its reversal, then the spin will either not have been rewound enough or it will have been rewound too much. It will have developed an overall phase change. Thus, phase-contrast encoding is performed by a pair of additional ‘bipolar gradients’, the magnitude, duration and time interval between which will determine the phase change experienced by a spin moving at a particular velocity. This information can be used to quantitate the velocity of the moving spins. In qualitative PC angiography, velocity encoding has to be applied in more than one direction. For a 3D implementation, separate flow encoding is applied in each direction. As with the TOF technique, PC is also an endogenous contrast mechanism.

Figure 5.18 3D Time of flight MR angiography projection of an intracranial, middle-cerebral artery berry aneurysm.

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Contrast-enhanced depiction of flow uses fast T1-weighted acquisitions to capture signal from a bolus injection of exogenous Gd-based contrast media. The addition of Gd into the vasculature causes a decrease in the T1 of blood which will appear bright on T1-weighted images. In order to delineate the arterial phase, the acquisition is required to be performed within one cardiac cycle. In practice, the low spatial frequency components are collected during the arterial phase as these form the major components of the image. Care is necessary in recognizing flow effects, for it is relatively easy to diagnose a solid lesion when all that is being seen is an unusual flow effect within a fluid-filled space.

Microvascular flow (perfusion) One of the key indicators of normal tissue function is that of normal arterial flow and tissue perfusion. Perfusion-weighted MRI has also been developed and can be performed with the aid of endogenous or exogenous contrast agents.The latter are by far the most common implementation. Techniques using endogenous contrast media rely on the rf labelling of proximal arterial blood, which can then be detected as it reaches an image plane selected through the region of interest. The noninvasive nature of this measurement makes it ideal in certain situations, e.g. in the assessment of placental perfusion in utero or when several repeat studies are needed. However, the contrast-to-noise ratios are less than those that can be obtained following the administration of exogenous contrast media. Gd-chelates are suitable exogenous agents for use in MR perfusion studies (see Fig. 5.11). The presence of the Gd–ion complex shortens both T1 and T2*. If a time series of T2*weighted images is acquired with high temporal resolution (approximately 1 frame s−1), the presence of Gd can be detected as a drop in signal as blood transporting the agent arrives at the image plane. A concentration–time curve can be obtained from which parameters such as bolus arrival time (BAT), mean transit time (MTT), time-to-peak (TTP) and relative blood volume (rBV) can be derived, and the application of the central volume theorem can yield information regarding relative cerebral blood flow (rCBF). Early MR perfusion studies performed using clinical systems were limited by image acquisition times. Fast gradientecho techniques (such as FLASH) could be used but usually on a single-slice basis. The introduction of EPI into the clinical setting has revolutionized MR perfusion imaging: enabling subsecond, multislice datasets to be acquired. As with diffusion-weighted imaging (DWI), most of the experience of perfusion-weighted imaging is in the brain, and particularly in stroke and other vascular abnormalities. Ischaemic stroke is usually due to occlusion of the cervicocranial vessels. The resulting reduced blood flow will have effects on the parameters measured by perfusion-weighted MR. In the acute setting the most commonly encountered changes are: increased time-to-peak; increased mean transit time; reduced cerebral blood volume and flow. It is hoped that a combination of DWI and perfusion-weighted imaging will permit delineation of the ischaemic penumbra in stroke, defining potentially

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salvageable tissue. Another area of interest is myocardial perfusion—used to provide information regarding the viability of the myocardium.

Very slow flow (molecular diffusion) The diffusion of water molecules is very slow random movement which can be thought of as slow directional flow when its randomness is restricted by a surrounding barrier (e.g. an axonal tract in brain parenchyma). Just as with arterial and venous intravascular blood flow and CSF flow, molecular DWI contrast can be introduced into MRI. The technique utilizes the same principles as outlined in the phase-contrast MRA section above: depending on the alteration in the phase of protons attached to water molecules as they travel along a magnetic field gradient. If a bi-lobed diffusion-encoding gradient pulse scheme is applied along one physical direction between the initial rf excitation pulse and data collection, any protons which are displaced along the direction of diffusion encoding will accumulate a relative phase change.Those that are stationary will not. The bipolar diffusion gradients are most commonly added to a single-shot EPI sequence, which enables a snap shot of the molecular diffusion to be obtained in one direction. This can be repeated along the three mutually orthogonal directions within the image plane. The degree of diffusion weighting can be changed by altering what is referred to as the b-value. The amplitude and duration of the bipolar gradients determine this b-value and images that have strong diffusion weighting require the application of very strong field gradients. Due to the inclusion of the time-consuming bipolar gradient scheme, data collection occurs at a relatively long effective TE and thus the diffusion-weighted images are also T2 weighted. This can lead to ‘T2 shinethrough’: high pixel values in areas of increased T2 as well as in areas of abnormal diffusion, which can lead to confusion! It is common practice to acquire a set of images with the bipolar gradients turned off, yielding a set of images with identical T2 but no diffusion weighting. Post-processing can then be performed to eliminate

the T2-dependent image contrast, yielding maps of apparent diffusion coefficients (ADCs). DWI is being assessed in many types of brain disease. Its major contribution appears to be in the distinction between cytotoxic oedema (which produces restricted diffusion) and vasogenic oedema (which does not restrict diffusion of water). The pathology that primarily produces cytotoxic oedema is ischaemia/infarction and DWI appears to be a sensitive and reasonably specific means of detecting ischaemic stroke. DWI is often used in conjunction with dynamic, Gd-enhanced perfusion imaging in the investigation of acute ischaemic stroke (Fig. 5.19).

PULSE SEQUENCES The most commonly used groups of pulse sequences are outlined below. Two main groupings are given: sequences where the echo results from the application of (A) an rf pulse and (B) a gradient.

Variations on a spin echo Basic two-dimensional spin-echo sequence In this pulse sequence; the digitized signal echo is produced by flipping the spins over in the x–y plane by a 180-degree rf pulse. This is applied at a time TE/2 after an initial 90-degree excitation pulse, producing a spin echo at time TE. Multiple slices are encoded by staggering each slice-selecting excitation pulse within each TR period and the Larmor frequency of each slice. Each of these slices experiences the application of one phase-encode step. Slice timing within each TR can be sequential (adjacent slices excited one after another: 1,2,3,4, 5…) or interleaved (1,3,5….2,4,6…): the latter is performed to reduce the effects of any overlap between slice edges that can occur at low TR (cross-talk). Parameters that alter image contrast are TE and TR. This is a very versatile sequence that can be PD, T2 or T1 weighted. PD-weighted images result if

Figure 5.19 Imaging of ischaemic stroke by multisequence MR. (A) An axial fast spin-echo T2-weighted image from a patient with an acute onset of left hemiparesis which shows a large area of T2 prolongation in the confines of the middle cerebral artery. The image is somewhat degraded by movement artefact as is often the case when using standard sequences in acutely unwell patients. (B) The single-shot fast spin-echo (imaging time < 1 s for one slice), which gives the same information. (C) A diffusion-weighted image acquired in 15 s shows high signal and (D) the apparent diffusion coefficient map shows a low signal, indicating an acute stroke.

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(A) TE is short (≤ 20 ms), thereby not allowing any differences in signal to evolve between tissues that may have different T2s, at the same time as (B) TR is long (≈ 3000 ms), enabling the spins to return to equilibrium between each rf excitation pulse. Differences in signal will be due to differences in the number of spins: a high PD will be represented by a bright pixel while a low PD will be represented by a dark pixel. If the TR is kept long, then as TE is increased (80–100 ms), different tissues with different T2 values will return different signals, and the T2 weighting of the resultant images will increase. In this case, tissues with long T2 values (e.g. CSF) will still have high signal at long TE and be represented by a bright pixel and vice versa. If on the other hand TE is short (to minimize T2 influences as in the PD-weighted case) but TR is decreased (≈500 ms), the T1 weighting of the sequence increases. The latter occurs due to different amounts of saturation experienced by tissues that have different T1 values. A tissue with a short T1 (e.g. fat) with respect to the specified TR, will relax back to equilibrium before the next excitation pulse: maximum magnetization will be available and so a bright pixel will ensue. A tissue with a long T1 (e.g. CSF) will not have time to relax back to its equilibrium state before the next excitation: only a small component of its longitudinal magnetization will and so the signal returned will be low and hence the corresponding pixels will be dark. Thus, pixel intensity resulting from a T1weighted spin-echo sequence (a partial saturation sequence) is inversely related to T1. A dataset can be obtained from the spin-echo technique in the order of several minutes.

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slices can produce motion artefact-free images, but care should be exercised as unwanted movement can occur between slices, altering the relative slice positions. Blurring can occur in the phase-encode direction. Every phase-encode step is acquired at a different TE, leading to a variation of T2-weighting across spatial frequency components. All resultant images appear heavily T2 weighted. Use of the single-shot FSE (SS-FSE) technique includes MR cholangiopancreatography, imaging of the diaphragm, bowel imaging and imaging of the fetus in utero (Fig. 5.20). It is also useful as a fast T2-weighted screening sequence in a moving patient!

Inversion recovery In this sequence type, the magnetization is ‘prepared’ before the initial 90-degree excitation pulse by the addition of an 180-degree ‘inversion’ pulse.The signal obtained will be influenced by the relative degree of recovery experienced by the spins along the z-axis.The inversion time (TI) is the time allotted for this recovery process to evolve between the 180-degree inversion and 90-degree excitation pulses. The inversion pulse introduces heavy T1 weighting into this sequence. Specific forms of the technique are fluid attenuated inversion recovery (FLAIR) where the TI is chosen so that the magnetization from fluids is nulled (≈ 2200 ms) in the detection (x–y) plane and TE is long (80–100 ms), introducing a lot of T2 as well as T1 weighting. The combination of fluid nulling with T2 weighting is particularly useful in assessing lesions with prolonged T2 adjacent to fluid structures (e.g. periventricular demyelination; see Fig. 5.7). In this same acquisition, T1 shortening brought

Multi-spin echo More than one echo can be recalled per 90-degree excitation using multiple 180-degree rf refocusing pulses which yield image datasets with different contrasts (as they are obtained at different TEs).The maximum number of slices per TR will reduce as the number of recalled echoes increases. The most common type used in clinical practice has been the dual spin-echo with short and long TE datasets (≈20 and 80 ms, respectively) at long TR, yielding PD- and T2-weighted datasets, respectively.

Fast-spin echo or turbo-spin echo or rapid acquisition with relaxation enhancement (RARE) As mentioned previously, the introduction of more than one phase-encode, refocusing pulse and frequency-encoded echo per excitation pulse (increasing the echo train length [ETL]), can speed imaging up by factors of between 2 and 32. During any one TR period, different spatial frequency components in the phase-encode direction will be acquired at different TEs, leading to a complex mixture of image contrasts. It is impossible to produce images that are as heavily PD weighted as can be achieved with the standard spin-echo technique above. Both T2- and T1-weighted images are commonly produced.

Single-shot fast-spin echo or half-Fourier acquired single-shot turbo-spin echo (HASTE) This is an extension of the FSE technique that enables all phase-encoding steps to be obtained following one initial excitation. One slice is acquired at a time, in approximately 1 s. Covering a volume prone to movement with multiple single

Figure 5.20 Imaging of the fetus in utero using single-shot fast spin-echo. Each set of 20 slices takes 20 s to acquire. This fetus has agenesis of the corpus callosum.

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about by the presence of exogenous Gd-based contrast agents will return high signal due to the sequence’s T1 weighting. In short-tau inversion recovery (STIR), T1 is kept short, and is typically used to null signal from fat. An example of its use in this context is in the delineation of breast lesions. There are several other methods of fat suppression; such sequences often demonstrate subtle lesions to best effect.

Spin-echo echo planar imaging In this variant of EPI, several or all lines of k-space are swept through at once during the evolution and decay of the main echo. The main echo is produced by an 180-degree rf pulse, i.e. it is a spin echo which is less influenced by field heterogeneities than the gradient-echo variety (below). This spin-echo EPI variant tends to be used for DWI as it is less prone to susceptibility-related artefacts than the gradient-echo EPI variant (see below).The sequence is actually a mix of gradient echoes, refocused by the EPI technique, that occur during the evolution of the spin echo that results from the refocusing rf pulse. Readers are referred elsewhere for further explanation and discussion of mixed gradient and spin echo (GRASE) sequences which have mixed spin-echo/gradient-echo contrast features.

Variations on a gradient echo/field echo Basic two-dimensional gradient-echo sequence This is as above for the 2D spin-echo sequences, except that the echo is produced by a gradient (rather than an rf pulse), causing the components of the spin magnetization to refocus in the x–y plane, producing an echo, without ‘flipping’ them over (i.e. without using an 180-degree pulse) (see Fig. 5.3). Unlike 2D spin-echo sequences, reduction in signal amplitude resulting from faster dephasing due to the presence of magnetic susceptibility differences/localized heterogeneity is not cancelled in this technique. Images can have mixed contrast consisting of PD, T1 and T2* weighting. When a different type of echo formation is present (the Hahn echo), gradient-echo sequences can also be T2 weighted. An understanding of the Hahn echo is needed to understand how different combinations of these four parameters can be made to dominate the image contrast. The TR used in a gradient-echo sequence is often shorter than the component tissue T2 values. In other words, the magnetization in the x–y plane will not have dephased completely before the next excitation pulse is applied (unlike the transverse magnetization in the spin-echo sequence which completely dephases, even when short TR is used to introduce saturation of the longitudinal component, i.e. to introduce T1 weighting). This net transverse coherence plus the addition of a further rf pulse produces the Hahn echo. So, two echoes can be present: the gradient echo and the Hahn echo. Each or both of these echoes can be used as follows. 1 Spoiled gradient-echo or spoiled gradient-recalled echo (SPGR) or RF spoiled Fourier acquired in the steady state (RF-FAST) or T1 fast field echo (T1 FFE). Any magnetization coherence in the x–y plane can be destroyed by the addition of a gradient pulse or rf ‘spoiler’, applied towards the end of the TR period. This sequence is

often termed a ‘spoiled gradient echo’. There is no Hahn echo, just the gradient echo. At short TE (< 8 ms) and short TR (≈ 50 ms), T1-weighting predominates (as with spin echo at short TR, due to differential saturation effects), but the weighting is not as heavy as the T1-weighted spin-echo variants. FLASH is a variant of the spoiled gradient echo. Instead of an initial 90-degree rf pulse, a low flip-angled pulse (α) is used. Very short repetition times can be used as saturation along the z-axis is minimized; enabling fast acquisition times and refresh rates for dynamic and angiographic studies. A relatively long TE (≈ 18 ms) increases the T2* weighting, with low α (≈ 25 degrees), decreasing the T1 weighting, allowing the T2* weighting to dominate. 2 Contrast-enhanced Fourier acquired in the steady state (CE-FAST) or PSIF (mirrored FISP) or T2 fast field echo (T2 FFE). T2-weighted images can be produced by imaging the steady-state coherence or Hahn echoes (that build up over a few repetitions) and not the gradient echo. The Hahn echoes are sampled towards the end of the TR period and the effective TE is approximately twice the TR. 3 Gradient recalled acquisition in the steady state (GRASS) or fast imaging with steady-state precession (FISP) or Fourier acquired in the steady state (FAST) or fast field echo (FFE). Residual phase coherence at the end of each TR is not destroyed and both the gradient and Hahn echoes contribute to the image. Complex, combined T1/T2 weighting results.

Three-dimensional gradient echo Secondary phase encoding is performed instead of slice selection. The penalty associated with this is an increase in acquisition time. However, as gradient-echo sequences can run with very short TR, this is manageable with typical standard 3D gradient-echo acquisition times in the order of 5–15 min. Acquisition times can be reduced substantially (with lower coverage and resolution), enabling breath-hold techniques (10–20 s). The spatial resolution in the second phase-encode direction is generally less than 2 mm (above this, 2D techniques tend to be used) and the signal-to-noise ratio at high resolution is superior to 2D techniques. Isotropic, submillimeter voxels facilitate true 3D visualization strategies.

Magnetization prepared rapid acquisition gradient echo (MP-RAGE) or turbo-FLASH or inversion recoveryprepped fast spoiled gradient-recalled echo As with the inversion recovery spin-echo sequence, an inversion pulse is applied before the excitation pulse, producing a T1-weighted sequence that is often used with rapid 3D encoding for high-resolution imaging of, for example, brain malformations. Coupled with its ability to depict exogenous contrast medium which has traversed the damaged blood–brain barrier, it is also useful for intra-operative neurosurgical guidance.

Gradient-echo echo planar imaging This is similar to spin-echo EPI but the main echo is produced by a gradient lobe. This sequence is heavily weighted to reflect differences in magnetic susceptibility. It is often used

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for BOLD fMRI studies and EPI-based exogenous perfusion assessment.

SAFETY CONSIDERATIONS Undergoing an MR investigation or working in a MR unit does not involve exposure to ionizing radiation and no clear long-term biological effects have been reported. However, the environment is quite unusual, presenting several potential risks to human safety that need to be fully addressed by those who work in it and for those who have occasion to visit it. Each unit should have a set of local rules based on advice outlined in documents from the relevant regulatory bodies (UK’s Medicines and Healthcare products Regulatory Agency (MHRA), European Union, etc.) and all MR unit staff should be familiar with these. Inner and outer ‘controlled areas’ that have restricted access should be detailed. Accountability to the hospital’s radiation safety committee is to be commended, ensuring that incidents are reported appropriately and that necessary safeguards/recommendations are acted upon. The invisible MR environment includes potential effects resulting from the following.

Magnet Ferromagnetic objects (mainly containing iron, nickel or cobalt), when near to a magnet, experience a force of attraction towards the magnet bore (magnetic isocentre). The force depends upon factors including magnetic field magnitude, proximity, object mass and composition.The danger associated with this hazard is real: fatalities have resulted. All patients, visitors, non-MR personnel and pieces of equipment should be screened appropriately before entry into the controlled area. Cardiac pacemakers contraindicate entry as do intracranial aneurysm clips of ferromagnetic/unknown composition. In very high fields (>> 1.5T) mild sensory activations (such as visual magnetophosphenes and balance alteration) have been reported. There is no evidence that these are harmful. If the magnet’s liquid helium/nitrogen Dewar is being replenished, the magnet room should be kept clear due to risk of cryogenic burns and suffocation. There may be occasion when the magnet needs to be shut down in an emergency, e.g. if a person becomes trapped between the magnet and a ferromagnetic object. Once the emergency run down unit (ERDU) button is depressed, liquid helium will vaporize and there is a valve/ pipe system to vent this gas to the outside. However, as a matter of caution, all personnel should exit the magnet room and the door should be kept closed.

Time-varying magnetic field gradients At high amplitudes and slew rates, these can cause unpleasant peripheral nerve stimulation. This is an acute effect and not harmful. Modern system software will alert the operator when there is risk of attaining such levels. The acoustic noise generated due to the fast switching of currents within the gradient coils while restrained in a high static magnetic field (an expensive loudspeaker!) may be considerable20, levels being up to

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108 dBA. Protective ear plugs/headphones should be worn by patients (and accompanying personnel), particularly in systems operating at 1.5T and above. Recent advances in gradient coil design and construction have led to a decrease in gradientassociated noise on some systems for given performance levels.

Radiofrequency field The transmitted rf field deposits energy within the body. Energy deposition is expressed as the specific absorption rate (SAR) in watts per kilogram. The levels given in MHRA guidelines (for SARs applied over a 15-min period) have been set to limit tissue heating to less than 1°C. Modern system software indicates the SAR to the operator for each acquisition. Care should be exercised when examining anaesthetized patients who require ECG monitoring. Burns can result from rf Eddy currents being induced under the skin electrodes (e.g. where clinical indications suggest SAR-intensive sequences such as FSE imaging of the lumbar spine). In these circumstances, pain sensation cannot be communicated while a burn is developing. Caution should also be made when the patient’s circulatory system is compromised. Clinical need should be noted in cases such as these. MR is not advised during the first trimester of pregnancy. This is precautionary as no deleterious effects to mother or child are known. Patients with various implants (from deep brain stimulators to stents) should be assessed for possible interactions with static field and rf heating: evidence from an up-to-date reference should be sought. At the time of writing, an EU Directive covering work exposure to electromagnetic radiation is due to become part of UK law in 2008. It will affect working practices in clinical MR units. The relevant exposure is likely to be related to the time-varying fields produced by the gradient subsystem. Staff will not be able to stand at the entrance to the magnet bore to comfort the patient, initiate examinations from the magnet fascia, perform close monitoring of anaesthetized patients or take part in MR-guided interventional procedures. Various professional bodies are presently attempting to address these issues and it is hoped that these restrictions will not be necessary.

CONCLUSION In the first 25 years of clinical practice, MRI has moved from being a curiosity to being the technique of first choice in a variety of diseases. The physics necessary for image interpretation is more complex than that for any other technique in radiology, yet the great diversity of the technique provides a wide range of possibilities, many of which are being developed or are yet to be explored. Clinical MR is truly 3D, and with the advent of modern workstations there is an opportunity to obtain radiographic views that have not been available previously.The recent development of fast imaging ciné studies, perfusion and diffusion techniques, hyperpolarized gas imaging (Fig. 5.21) and solid

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Figure 5.21 Image of a healthy lung obtained with a two-dimensional, rapid, steady-state free precession sequence with 300 ml. Hyperpolarized 3 He at 30% polarization. (Courtesy of Dr J.M. Wild, University of Sheffield, UK.)

state methods provide new possibilities. The development of new contrast agents, targeted at specific organs, diseases, cell or gene types is likely to result in a considerable expansion of the options now available. Parallel acquisition methods promise radically to increase the throughput of MR techniques and allow contrast mechanisms, which previously could only be acquired statically, to be acquired in dynamic mode. Specialized MR systems targeted at specific body regions or diseases may increase access to MR and provide options for monitoring and even performing minimally invasive therapy. The full potential of MR is yet to be realized.

REFERENCES 1. Bloch F, Hansen W W, Packard M E 1946 Nuclear induction. Phys Rev 69: 127. 2. Purcell E M, Torrey H C, Pound C V 1946 Resonance absorption by nuclear magnetic movements in a solid. Phys Rev 64: 37–38. 3. Lauterbur P C 1973 Image formation by induced local interactions: examples employing NMR. Nature 242: 190–191. 4. Mansfield P, Grannell P K 1973 NMR ‘diffraction’ in solids? J Phys C 6: 422–426. 5. Mansfield P, Maudsley A A 1977 Medical imaging by NMR. Br J Radiol 50: 188–194. 6. Damadian R, Goldsmith M, Minkoff L 1977 NMR in cancer: Fonar image of the live human body. Physiol Chem Phys 8: 97–108. 7. Hinshaw W S, Bottomley P A, Holland G N 1977 Radiographic thin section of the human wrist by nuclear magnetic resonance. Nature 270: 722–723. 8. Hawkes R C, Holland G N, Moore W S, Worthington B S 1980 NMR tomography of the brain: a preliminary clinical assessment with demonstration of pathology. J Comput Assist Tomogr 4: 577–586. 9. Edelstein W A, Hutchinson J M S, Johnson G, Redpath T 1980 Spin warp NMR imaging and application to human whole body imaging. Phys Med Biol 25: 751–756. 10. Thomas D L, Lythgoe M F, Pell G S, Calamante F, Ordidge R J 2000 The measurement of diffusion and perfusion in biological systems using magnetic resonance imaging. Phys Med Biol 45: R97–138.

11. Wilkinson I D, Griffiths P D, Hoggard N, et al 2003 Short-term changes in cerebral micro-hemodynamics following carotid stenting assessed by MR perfusion imaging. Am J Neuroradiol 24:1501–1507. 12. Neumann-Haefelin T, Moseley M E, Albers G W 2000 New magnetic resonance imaging methods for cerebrovascular disease: emerging clinical applications. Ann Neurol 47: 559–570. 13. Taylor A M, Keegan J, Jhooti P, Gatehouse P D, Firmin D N, Pennell D J 2000 A comparison between segmented k-space FLASH and interleaved spiral MR coronary angiography sequences. J Magn Reson Imaging 11: 394–400. 14. Wilkinson I D 2005 Perfusion and diffusion imaging in chronic carotid disease. In: Gillard J, Waldman A, Barker P (eds) Clinical MR neuroimaging: diffusion, perfusion and spectroscopy. Cambridge: Cambridge University Press. 15. Paley M 1996 Human brain proton spectroscopy. In: Bydder G M, Bradley W (eds) Advanced MR imaging techniques. London: Martin Dunitz. 16. Schaefer S, Balaban R S (eds). 1992. Cardiovascular magnetic resonance spectroscopy. New York: Springer. 17. Ogawa S, Tank D W, Menon R, et al 1992 Intrinsic signal changes accompanying sensory stimulation: functional brain mapping with magnetic resonance imaging. Proc Natl Acad Sci USA 89: 5951–5955. 18. Rosen B R, Aronen H J, Kwong K K, Belliveau J W, Hamberg L M, Fordham J A 1993 Related articles advances in clinical neuroimaging: functional MR imaging techniques. RadioGraphics 13: 889–896. 19. Middleton H, Black R D, Saam B, et al 1995 MR imaging with hyperpolarized 3He gas. Magn Reson Med 33: 271–275. 20. Sodickson D K, Manning W J 1997 Simultaneous acquisition of spatial harmonics (SMASH): fast imaging with radiofrequency coil arrays. Magn Reson Med 38: 591–603. 21. Pruessmann K P, Weiger M, Scheidegger M B, Boesiger P 1999 SENSE: sensitivity encoding for fast MRI. Magn Reson Med 42: 952–962.

FURTHER READING Overviews Edelman R R, Hesselink R R 2005 Clinical magnetic resonance imaging. Philadelphia: WB Saunders. McRobbie D W, Moore E A, Graves M J, Prince M R 2003 MRI from picture to proton. Cambridge: Cambridge University Press. Reimer P, Parizel PM (rds) 2006 Clinical MR imaging: A practical approach. Berlin: Springer-Verlag.

Sequences design/spatial encoding/hardware Haacke E M, Brown R W, Thompson M R, Venkatesan R 1999 Magnetic resonance imaging: Physical principles and sequence design. Chichester: J Wiley & Sons. Holland G N, MacFall J R 1992 An overview of digital spectrometers for MR imaging. J Magn Reson Imaging 2: 241–246. Landini L, Positano V 2005 Advanced image processing in magnetic resonance imaging. CRC Press. Nassaivier M 1996 All you really need to know about MRI physics. Simply physics. www.simplyphysics.com. Robitaille P-M, Berliner L (eds) 2006 Ultra high field magnetic resonance imaging. Berlin: Springer-Verlag. Schmitt F, Stehling M K, Turner R 1998 Echo-planar imaging: theory, techniques and application. Berlin: Springer-Verlag. Schoenberg S O 2007 Parallel imaging in clinical MR applications. Berlin: Springer-Verlag. Tweig D B 1983 The k-trajectory formulation of the NMR imaging process with applications in analysis and synthesis of imaging methods. Med Phys 10: 610–623.

Diffusion, perfusion and spectroscopy Danielson E B, Ross B 1998 Magnetic resonance spectroscopy diagnosis of neurological diseases. New York: Marcel Dekker.

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Gillard J, Waldman A, Barker P (eds) 2005 Clinical MR neuroimaging: diffusion, perfusion and spectroscopy. Cambridge: Cambridge University Press. Warach S, Davis S 2003 Magnetic resonance imaging in stroke. Cambridge: Cambridge University Press. Young I R (ed) 2000 Methods in biomedical magnetic resonance imaging and spectroscopy. Chichester: J Wiley & Sons.

BOLD fMRI Buxton R B 2001 Introduction to functional magnetic resonance imaging: Principles and techniques. Cambridge: Cambridge University Press. Frackowiak R S J, Friston K J, Frith C D, Dolan R J, Mazziotta J C 1997 Human brain function. New York: Academic Press.

Exogenous contrast Jackson A 2005 Dynamic contrast-enhanced magnetic resonance imaging in oncology. Berlin: Springer-Verlag. Kauczor H-U (ed) 2000 Special issue: hyperpolarized gases in MRI. NMR Biomed 13: 173–264. Stark D D, Weissleder R, Elizando G, et al 1989 Superparamagnetic iron oxide: clinical application as a contrast agent for magnetic resonance imaging. Radiology 168: 297–301. Weinmann H-J, Brasch R C, Press W R, Wesbey G E 1984 Characteristics

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of gadolinium-DTPA complex: a potential NMR contrast agent. Am J Roentgenol 142: 619–624.

Angiography/cardiology/interventional Bogaert J, Dymarkowski S, Taylor A M (eds) 2005 Clinical cardiac MRI. Berlin: Springer-Verlag. Graves M J 1997 Magnetic resonance angiography. BJR 70: 6–28. Lee V 2006 Cardiovascular MRI: Physical principles to practical protocols. Philadelphia: Lippincott, Williams & Wilkins. Lufkin R B 1999 Interventional MRI. St Louis: Mosby. Morris E, Liberman L 2005 Breast MRI: Diagnosis and intervention. Berlin: Springer-Verlag.

Safety Medical Devices Agency 2002 Guidelines for magnetic equipment in clinical use with particular reference to safety. London: Medical Devices Agency. Moseley I F 1994 Safety of magnetic resonance imaging. Br Med J 308: 1181–1182. National Radiation Protection Board 1991 Principles for the protection of patients and volunteers during clinical MRI. NRPB 2 (no 1). Shellock S G 2006 Reference manual for magnetic resonance safety, implants, and devices: 2007 edition. BRPG. Shellock R, D Services Inc. www.mrisafety.com.

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Angiography: Principles, Techniques (Including CTA and MRA) and Complications

6

James E. Jackson, David J. Allison and James Meaney

Multidetector CT angiography • Clinical applications Magnetic resonance angiography • Background • Contrast mechanisms • Post-processing • MRA in clinical practice • ‘Supplemental’ imaging: when is imaging of the lumen not enough? • Future directions Catheter arteriography • Technique • Preparation of the patient

• Contraindications • Anaesthesia • Arterial puncture • Digital subtraction angiography • Intravenous digital subtraction angiography (IVDSA) • Aftercare • Complications Catheter venography • Techniques • Complications Embolization techniques • Embolic materials and techniques • Indications for therapeutic arterial embolization

The imaging of blood vessels has changed considerably since the first edition of this textbook and, indeed, there have been significant new developments in cross-sectional imaging techniques even since the 4th edition. These have made many of the diagnostic catheter angiographic techniques described in previous editions almost obsolete. On the whole this is clearly a welcome advance; the newer multidetector CT angiographic techniques are obviously less invasive and, therefore, safer. Furthermore, in many instances these techniques will give more diagnostic information than could be obtained by conventional catheter arteriography because of the concurrent visualization of surrounding tissues and the ability to reconstruct the data in any plane. One of the disadvantages of the decline in the number of diagnostic catheter

angiograms performed, however, is that it has become more difficult for radiologists to acquire suitable expertise in the catheter techniques that are still required for more complex therapeutic interventional procedures. As a good understanding of the basic principles and techniques of catheter angiography remains essential for those intending to become interventional radiologists (and it becomes less likely that they will be able to obtain sufficient practical experience during their training for the reasons given earlier) it perhaps becomes more important that this information is available in this textbook. The newer cross-sectional techniques for imaging blood vessels will, however, be discussed first as these will, quite rightly, be requested before (and often instead of ) conventional catheter angiography.

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MULTIDETECTOR CT ANGIOGRAPHY (MDCTA) The development of CT machines combining a fan-shaped x-ray source and multiple detector rows has led to the ability to acquire image data from a large tissue volume in a single breath hold. With IV contrast medium and appropriate timing, exquisite images can be obtained of blood vessels during any particular vascular phase. Optimal imaging of the vessels requires the relatively rapid IV injection of iodinated contrast medium (usually 3–5 mls-1) and the acquisition of data at the appropriate time of vascular enhancement. The latter can be estimated based upon the ‘normal’ time of arrival of the contrast medium within the organ being imaged or, more commonly nowadays, by the more accurate use of contrast bolus detection technology. ‘Tight’ boluses of contrast medium using a chaser of normal saline may be useful not only to improve vascular opacification but also to reduce the total volume of contrast medium required. Depending upon the region and volume of the body being imaged hundreds, if not thousands, of axial images will be acquired; whilst all the diagnostic information is available in this data-set, evaluation of the axial images alone can be extremely time-consuming and is helped considerably by reconstruction of the data in axial, coronal, and oblique planes without loss of resolution, so-called multiplanar reconstruction (MPR). Tortuous vessels can be ‘straightened’ by curved MPR to aid in the assessment of luminal narrowing due to, for example, atheromatous disease or encasement by tumour. Maximum intensity projection (MIP) and volume rendering (VR) techniques are additional tools that help greatly in the assessment of blood vessels. Each of these reconstruction techniques has its advantages and disadvantages: 1 MPR is very useful for the rapid review of blood vessels in any plane including the surrounding bone and soft tissues, and will allow the assessment of vessel walls that might be obscured in MIP and VR techniques by the presence of, for example, calcification or an endoluminal stent. Each image, however, gives

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only one ‘slice’ of information and multiple separate images are required to see the vessel in its entirety (Fig. 6.1). 2 MIP techniques produce a planar image from a volume of data within which the pixel values are determined by the highest voxel value in a ray projected along the data set in a specified direction.The images obtained are those most similar to a conventional arteriogram but one of the clear disadvantages of this technique is that any tissue of high density (such as bone or vascular calcification) lying within the ray through a vessel will determine the pixel intensity instead of the contrast medium itself within the blood vessel (Fig. 6.2).This is a common cause of overestimation of vascular stenoses. 3 Volume-rendering techniques assess the entire volume of data with an attenuation threshold for display and produce a three-dimensional image. Typically, tissues are assigned a

Figure 6.2 An MIP image from an MDCT of a renal transplant artery clearly demonstrating the vascular anatomy.

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Figure 6.1 MDCT images of a gastroduodenal artery pseudoaneurysm. (A) An axial image demonstrates an enhancing pseudoaneurysm cavity surrounded by a large haematoma in the region of the pancreatic head. (B) A sagittal MPR image demonstrates a pseudoaneurysm arising from the gastroduodenal artery. (C) A further sagittal MPR image demonstrates that the right hepatic artery, from which the gastroduodenal artery arises, originates from the superior mesenteric artery.

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• ANGIOGRAPHY: PRINCIPLES, TECHNIQUES (INCLUDING CTA AND MRA) AND COMPLICATIONS

colour that is dependent upon their attenuation values, facilitating the differentiation of structures of differing density (Figs 6.3, 6.5). The final images can be rotated in real time to find the best projection to display anatomy and pathology and this is the most important feature of this technique. Vascular stenoses can be overestimated, however, and small vessels may not be clearly visualized. It should be remembered that, whilst these post-processing techniques are very helpful for diagnostic assessment and for display in multidisciplinary team meetings, the axial source images are essential and often allow the operator to distinguish between artefact and disease when an abnormality is suggested on reformatted views.

CLINICAL APPLICATIONS MDCT angiography (MDCTA) is replacing conventional angiography in many, if not all, body areas and is indicated, therefore, in any disease process that requires the visualization

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of blood vessels to improve diagnosis and outcome. Within the thorax, for example, this would include the assessment of: pulmonary embolic disease1,2 (Fig. 6.4); thoracic aortic disease3–5; coronary artery graft patency6–10; and bronchial artery anatomy and pathology in massive haemoptysis11. In the abdomen common indications include the pre-operative planning and posttreatment assessment of abdominal aortic aneurysmal disease12,13 (Fig. 6.5); the assessment of native and transplant renal arteries14–16; the staging of hepato-pancreaticobiliary neoplasms17–20; and the assessment of vascular complications in patients suffering severe trauma21. It is also used increasingly in the assessment of peripheral arterial disease in the lower (and upper) limbs where it is less invasive, less expensive and exposes the individual to less radiation than conventional catheter angiography22–26. The scanning technique (positioning of the patient, rate of contrast medium administration, time of image acquisition), and the post-processing techniques most suited to the different indications listed earlier will clearly vary and lie outside the scope of this chapter; interested readers are referred to other chapters within this book and to other texts cited in the reference list.

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Figure 6.3 The value of MDCT MPR and VR images in the assessment of pulmonary sequestration. (A) The CXR demonstrates a mass projected through the left side of the cardiac silhouette. (B) An axial image from an MDCT examination demonstrates a left paraspinal soft-tissue mass. A feeding vessel arising from the thoracic aorta is visible. (C) MPR and (D) VR images show the full length of the oblique course of feeding artery.

Figure 6.4 MDCT demonstration of bilateral pulmonary emboli. (A) Axial and (B) coronal MPR images demonstrating extensive bilateral pulmonary emboli in central pulmonary arteries.

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Figure 6.5 MDCT images of abdominal aortic aneurysmal disease before and after stent insertion. A large infrarenal abdominal aortic aneurysm is seen on axial (A and D), coronal MPR (B and E) and VR (C and F) images before and after the insertion of an endoluminal bifurcated stent

MAGNETIC RESONANCE ANGIOGRAPHY BACKGROUND Magnetic resonance angiography (MRA) is a method for generating images of blood vessels with magnetic resonance imaging (MRI)27,28. With improved understanding of the nature of the signals emanating from blood vessels on MR images, it became evident that rather than simply contribut-

ing a source of artefacts on MR images, flow phenomena could be harnessed to generate diagnostic ‘angiograms’27–29. MRA has undergone a revolution over the last decade, replacing catheter angiography as the primary diagnostic tool for the evaluation of most vascular territories (apart from the coronary arteries); this change is due mainly to the success of contrast-enhanced techniques30.

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CONTRAST MECHANISMS Unenhanced (time-of-flight [TOF] and phase contrast [PC]) MRA With these techniques, the intravascular signal depends on inherent properties of flow, and the MR parameters must be carefully tailored to ensure a high intravascular signal27,28. In the case of TOF MRA, for example, data must be acquired perpendicular (and ideally orthogonal) to the direction of flow, and the time between successive radio frequency (RF) pulses (the repetition time [TR]), must be sufficiently long to allow an adequate ‘inflow’ of fully relaxed protons into the imaging slice27.The TR, therefore, is dictated by expected flow rates within the regionof-interest and, typically, should be 35 ms or greater. As the data acquisition time is directly proportional to the TR (acquisition time = TR × number of phase-encoding steps × number of slices × number of excitations), a TR substantially greater than the shortest possible for gradient-echo imaging (3–5 ms) must be used, with resultant prolongation of scan time. Additionally, owing the predominant head–foot orientation of the major arteries (e.g. the aorta, ilio-femoral and infra-popliteal vessels), the axial plane must be used, which also prolongs the data acquisition owing to the large number of slices required. Despite ‘faster’ techniques, therefore, acquisition times for TOF MRA are artificially prolonged, firstly as a result of physiological bloodflow rates that mandate the use of a relatively long TR and secondly by the requirement for axial imaging, which affords poor spatial coverage in comparison with sagittal or coronal imaging30. Selective arteriograms or venograms can be acquired by employing a (travelling) saturation pulse placed downstream of the imaging slice for MRA (to eliminate venous return from the opposite direction) or upstream of the imaging slice to generate MRV’s (venograms). If no saturation pulses are employed, both arteries and veins are identified on the same image.

Time-of-flight MRA methodology and limitations TOF angiography relies on the fact that the blood enters the volume under consideration with relatively high velocity and traverses it quickly, so that it receives very few RF pulses27,28. In order to maintain a highest possible inflow effect, all protons within the imaging volume must be replenished between successive TRs, though maximal inflow may not be necessary in clinical practice and some trade-offs can be accepted.An oblique course of the blood vessel being imaged in relation to the slice orientation and short TRs both adversely affect signal-to-noise ratios (SNR) as a result of protons under these circumstances experiencing more RF pulses whilst in the imaging slice. The severity and length of stenoses also tend to be overestimated on TOF MRA images because of intra-voxel dephasing secondary to turbulent, slow or pulsatile flow. As a result of these limitations,TOF MRA has failed to offer a viable non-invasive screening alternative to conventional arteriography, and has not had a major impact on clinical practice.

Phase-contrast MRA Phase-contrast angiography (PCA) is now seldom used in clinical MRA. The methodology underpinning the technique

is somewhat complex and, like TOF MRA, images are prone to artefacts and data acquisition is lengthy29. The term ‘phase’ refers to the angle that the transverse magnetization makes relative to a reference axis. When protons are moving, for example, in flowing blood, a change in phase proportional to the distance the proton moves (and, therefore to blood velocity) is induced by the imaging gradients, in particular the slice-selection gradient and the frequencyencoding gradient. As these gradients actually consist of a pair of gradient pulses applied in opposite directions (so-called bipolar gradients), the effect of each gradient on the phase of protons for stationary (i.e. background, motionless) tissues is equal and opposite, and the effect cancels out. For moving (flowing blood) protons, the position of each proton will change between consecutive pulses resulting in a phase change relative to stationary protons that is proportional to velocity. This observation of gradient-induced phase change provides the basis for phase-contrast angiography in which the gradient pulses are designed to produce phase changes for a given velocity range. In this way, the signals do not cancel, phase information is preserved during the image reconstruction. Pixel brightness is directly proportional to the phase-shift acquired by a moving proton in the magnetic field and, therefore, to velocity. In practice, as the method is only sensitive for the velocity component applied along the flow-encoding gradient, the acquisition must be repeated a total of four times in order to generate an ‘angiogram’: an initial flow-compensated sequence is followed by three flow-encoded acquisitions one for each direction of flow (head–foot, left–right, anteroposterior), followed by a complex subtraction to generate the final angiographic image. As the phase is unchanged for static protons, the subtraction completely suppresses the signal from the background tissue thus facilitating the generation of high-quality images29. One of the major challenges of PCA relates to the requirement of the operator to appropriately select the velocity-encoding gradient.As the signal intensity is proportional to velocity, the range of velocities present within the vessels of interest must be inferred or measured to allow the operator to set the velocity-encoding gradient (Venc) correctly.This assumes an a priori knowledge of the blood flow velocities within the relevant artery and, though this can be rapidly measured directly by acquiring a series of 2D PC images with different phase-encoding values, it is time-consuming and the presence of different flow velocities within the arteries enclosed within a single field of view may introduce artefacts29.

Limitations of unenhanced MRA and requirement for contrast agents At each stage of the evolution of MRA, high accuracy was reported for almost all techniques when compared with catheter angiography27–29 but TOF and PC MRA did not gain widespread acceptance in clinical practice because of long examination times, suboptimal resolution and frequent artefacts. TOF MRA was used to evaluate disease involving the carotid bulb and the femoro-popliteal and pedal arteries as these vascular territories were ideally suited to this technique due to the relatively straight course of their vessels, which

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meant that adequate ‘inflow’ could be ensured. MRA was also established as an accurate modality for portrayal of the proximal intra-cranial arteries (despite the fact that the arteries demonstrate marked tortuosity) because of the relatively small volume of tissue that needed to be covered, coupled with the fact that constant blood flow within the cerebral arteries over the cardiac cycle optimizes intravascular signal29,30.

Contrast-enhanced MRA (CEMRA) Because of their unmatched high contrast-to-noise ratios, high spatial resolution, rapid speed of acquisition and relative freedom from artefacts, contrast-enhanced techniques have almost universally replaced non-contrast techniques in clinical practice30. Unlike TOF and PCA techniques, where the intravascular signal is dependent on inherent properties of flow and is, therefore, at the mercy of alterations in flow rate secondary to vascular disease, intravascular signal for contrast-enhanced MRA (CEMRA) depends on a T1 shortening effect induced by the injection of a paramagnetic contrast agent (usually gadolinium based). Images can, therefore, be acquired in any plane including coronal, which affords the best anatomical coverage for virtually all vascular territories outside the brain (Fig. 6.6). In addition, the ability to exploit ultrafast 3D acquisitions (by using the shortest TRs possible), allows rapid image acquisition that can easily be accommodated within a single breath-hold, an important factor when imaging in the chest and abdomen. In order to generate ‘selective’ arteriograms, images are acquired during the first arterial passage of the contrast agent before its arrival within the veins.The synchronization of data acquisition with the peak arterial bolus is one of the major challenges of CEMRA as the rate of transit of contrast medium from the peripheral vein injection site to the vessel of interest is affected by a number of factors including heart rate, stroke volume and the presence or

Figure 6.6 A surgically created dialysis (arteriovenous) fistula in the left arm of patient with chronic renal failure.

absence of proximal steno-occlusive lesions. Although the circulation time can be measured using a test bolus, or can be inferred by making some assumptions about the patient’s cardiovascular status, the process is now automated by employing an MR fluoroscopic approach – a technique that demonstrates contrast medium arrival in real time on the display monitor, thus signalling the appropriate time for data acquisition31. The unique nature of k-space (the array of data from which the final image is generated) whereby the central lines determine image contrast and the peripheral lines determine image resolution, can be uniquely exploited to generate CEMRA images with unrivalled signal-to-noise ratios32. In situations where breath-holding is not required (e.g. peripheral MRA and carotid artery imaging) as long as collection of the contrast-defining central lines of k-space is completed during the arterial peak before contrast medium reaches the veins, the continued collection of resolution-defining peripheral lines of k-space during venous enhancement does not result in venous contamination of the images32. CEMRA is now the standard of reference for MRA against which all new techniques must be measured.

POST-PROCESSING Regardless of which method is employed to generate MR angiograms, the aim of all techniques is to make the arteries the brightest structures on the images, and to extract the vascular data by means of a maximum intensity projection (MIP) computer algorithm28 (Figs 6.7, 6.8). Other methods of post-processing include multi-planar reformatting, volume rendering and surface-shaded displays. For phase-contrast MRA, there is inherent complete background suppression because of the absence of bulk motion in background tissues.

Figure 6.7 Normal MRA. Note clear depiction of the abdominal aorta, iliac arteries and renal arteries on the frontal MIP.

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monary angiography; and the fact that catheter pulmonary angiography, the reference standard against which new and improved MRA techniques should be compared, has largely disappeared from clinical practice due to the success of CTA, thus depriving MRA of a valid arbiter for comparative studies. Improvements in spatial resolution bring the subsegmental arteries within the realm of MRA (Fig. 6.9) and further refinements including MR perfusion and ventilation (mirroring the ventilation and perfusion components of nuclear medicine studies albeit at much higher resolution) offer additional functionality to determine the location and distribution of small emboli36.

Abdominal aorta, renal and mesenteric arteries

Figure 6.8 Severe left renal artery stenosis and right common iliac artery occlusion.

For TOF MRA, the background is suppressed by virtue of the short TR in relation to the longish T1s of background tissues. Although fat remains bright, it can be eliminated by use of fat-saturation techniques, albeit with an additional time penalty. For CEMRA, there is the additional benefit that the background tissues can be completely eliminated by the subtraction of a mask acquired before the injection of contrast material.

CEMRA is the MR technique of choice for imaging the aorta in its entirety and for evaluating its large and medium-sized branches including the renal (Fig. 6.8) and proximal mesenteric arteries30,37–40. Because of the need for breath-holding the technique’s spatial resolution remains inferior to that of catheter angiography because of the inability to collect a highresolution data-set that matches that of DSA during a breathhold. Nonetheless, numerous studies and meta-analyses attest to the accuracy of MRA in clinical practice38. An additional benefit of MRA lies in its ability directly to measure the flow rate to each kidney using a two-dimensional (2D) cardiactriggered, phase-contrast approach, which facilitates both the assessment of both end-organ damage and the likelihood of success of transluminal angioplasty37. MRA is also highly accurate for depicting the mesenteric arteries in patients with suspected chronic mesenteric ischaemia39. In patients with abdominal aortic aneurysms, the external dimensions of the aneurysm can easily be delineated, should this be necessary, with targeted pre-contrast or post-contrast

MRA IN CLINICAL PRACTICE MRA is an excellent technique for imaging most vascular territories but is generally avoided in unstable and/or ventilated patients and those with severe trauma because of the hazards of the MR environment and the difficulties in monitoring patients within the MR room. Standard contra-indications to MRI (pacemakers, intracranial uneurysm clips) also preclude use of MRA.

Thoracic aorta and great arteries Because of the relatively large size of these vessels, they can be demonstrated with a wide variety of techniques but CEMRA is favoured in most instances owing to its rapid speed of acquisition and the quality of the images generated33,34.

Pulmonary arteries Although several studies have established high accuracy for MRA compared with pulmonary angiography for the evaluation of suspected pulmonary embolism, it is not widely used clinically35,36. Reasons for this include a reluctance to refer potentially unstable patients to MRI; the availability of CT pul-

Figure 6.9

Normal pulmonary MRA.

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images as only the lumen is demonstrated with CEMRA40. Calcium within the wall is not demonstrated, however, and aneurysm assessment for planning endovascular stenting is usually performed with CTA.

Carotid arteries Because of the requirement accurately to differentiate stenosis at a 70 per cent cut-off within a relatively small (internal carotid) artery, there are stringent spatial resolution requirements for carotid MRA41. As the carotid bifurcation does not move with respiration a relatively long data acquisition that generates images with sufficiently high (isotropic) spatial resolution is recommended. Despite the fact that the resultant data acquisition is substantially longer than the arterio-venous transit

A

D

G

B

E

H

time of 8–12 s (the blood–brain barrier prevents the parenchymal extraction of gadolinium and therefore facilitates rapid transit from artery to vein), this does not lead to unacceptable venous contamination of the images as there is sufficient time within this arterio-venous window for acquisition of the contrast-defining central lines of k-space. Clearly, acquisition of these central lines must be synchronized with the peak arterial bolus by combining some form of bolus detection as described previously with a ‘centric’ order of k-space filling. In comparison with cathether angiography CEMRA has demonstrated high accuracy in evaluating the carotid artery, e.g. for differentiating between significant and insignificant stenoses, differentiating between critical stenoses and occlusions, and for depicting carotid and vertebral dissections (Fig. 6.10).

C

F Figure 6.10 T2, FLAIR and diffusion-weighted images in a patient with right-sided weakness and aphasia. All the images were acquired in the axial plane at the same level. (A–D) Note the high signal on T2 (A), FLAIR (B), and Diffusion-weighted images (C) (arrowed) with a corresponding low-signal intensity on ADC map (D), ringed indicating an acute cerebral infarct. (E, F) Whole-volume MIP images from a TOF MRA in coronal (E) and axial (F) orientation demonstrate reduced signal intensity within the left internal carotid (short arrows) and left middle cerebral arteries. Targeted MIP images of the left-sided carotid and vertebral arteries demonstrate a severe stenosis of the internal carotid artery just beyond the bulb (not shown). Axial T1 (G) and T2w (H) images with fat-saturation reveal crescentic high-signal intensity within the left ICA (long arrows) at the site of the stenosis demonstrated on CEMRA, diagnostic of acute dissection.

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• ANGIOGRAPHY: PRINCIPLES, TECHNIQUES (INCLUDING CTA AND MRA) AND COMPLICATIONS

Peripheral arteries In the peripheral arteries, as in most other areas, TOF MRA has been superseded by CEMRA. Although the spatial coverage offered by single-field-of-view imaging is insufficient to address all of the relevant anatomy, the introduction of moving-table MRA has opened the way for routine noninvasive MRA of the entire run-off arteries in a short timeframe (<20-minute examination)42,43 (Fig. 6.11). The major limitation of moving-table MRA is venous contamination in the third location. Methods to eliminate this include the use of tourniquets inflated to sub-systolic pressures to delay onset of venous enhancement, careful attention to detail in setting up the examination (to avoid any redundancy in anatomical coverage), the use of high parallel-imaging factors and the exploitation of ultra-short TR imaging42,43.

Figure 6.11 Peripheral MRA performed using a threestation, moving-table approach. The study demonstrates excellent visualization of the entire run-off arteries from the mid-abdominal aorta to the level of the pedal arch. Note extensive vascular stenoses and occlusions and a bypass graft from the right popliteal artery to the right anterior tibial artery.

‘SUPPLEMENTAL’ IMAGING: WHEN IS IMAGING OF THE LUMEN NOT ENOUGH? Currently, virtually all clinical decision making is based on the appearance of the lumen, though clearly the relevant pathology (atheroma) lies within the wall. Efforts to delineate the various components of plaque, in particular the ‘unstable’ plaque, have met with mixed results. A particular situation exists in the imaging of carotid and vertebral artery dissections, where imaging of the lumen alone simply demonstrates narrowing indistinguishable from other causes, but where cross-sectional imaging of the lumen with spin-echo imaging is essential to demonstrate the typically bright appearance of the usually crescent-shaped intramural haematoma. Therefore, in all cases of suspected carotid or vertebral dissection pre-contrast T1 and T2w fat images saturated should be performed. Failure to do so results in the identification of a focal stenosis only, and robs the clinician of the opportunity to prescribe anticoagulation therapy, the treatment of choice for dissection. Other instances in which imaging additional to that provided by the ‘luminogram’ is helpful include arteritides (e.g Takayasu’s and Behçets disease), aneurysms and the assessment of vessels narrowed or invaded by tumour.

FUTURE DIRECTIONS Improvements in MRA are aimed at increasing both spatial and temporal resolution whilst minimizing contrast-to-noise loss.

Improved coil efficiency and new parallel imaging methods that increase acquisition speed and/or spatial resolution by a factor of 2–16 are becoming routine in clinical practice44,45 and 3T imaging opens new horizons due to its improved signalto-noise capability44,46. New contrast agents (1.0 M and bloodpool agents) and the potential for routine functional imaging of end-organ function offer exciting prospects for future improvements. Currently, new applications are being added constantly, and isotropic submillimetre resolution studies for all regions of interest combined in a single examination with functional evaluation are anticipated within the near future.

CATHETER ARTERIOGRAPHY TECHNIQUE The risks associated with modern arteriography are extremely small. Arteriography is still nevertheless an ‘invasive’ procedure and it should never be undertaken unless the radiologist is satisfied that the likely benefits justify the potential risks. An arteriogram should never be carried out simply because it has been scheduled or ‘routinely requested’ by a clinical team;

mistakes inevitably occur and the radiologist responsible for the procedure should be satisfied in every case that proper indications exist for the particular study requested. The arteriographer should also be quite clear before starting as to what information is required from the procedure; this ensures that the correct studies and projections are obtained and allows for rational decision making during the procedure if something unexpected is shown or a problem arises. Most of these

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comments are clearly applicable to any imaging technique but are arguably more important and perhaps easier to enforce than is the case for non-invasive studies.

PREPARATION OF THE PATIENT Informed consent should be obtained for arteriography. A doctor, preferably the responsible radiologist or a member of the radiology department, should see the patient before the procedure to explain what is to be done, check that no contraindications to the study exist, check the appropriate pulses and ensure that adequate premedication is arranged.The groin should be shaved if a femoral approach is to be used. It has been the usual practice for patients to be ‘nil by mouth’ for an appropriate period before the procedure to avoid the risk of aspiration during a possible reaction to contrast medium or other serious accident. It is now the policy in many departments, however, only to stop solid foods and to permit free oral fluids unless general anaesthesia or heavy premedication is being used. Whatever regimen is adopted, adequate measures should be taken to avoid dehydration during the procedure and the recovery period.

CONTRAINDICATIONS There are very few absolute contraindications to arteriography but there are many factors that considerably increase the hazards of the technique. Always check that a patient is not pregnant before arteriography, as the radiation dose may be considerable. If arteriography is essential in a pregnant patient, the dose to the fetus should be minimized by protection, field collimation and careful choice of filming sequences. Caution should be exercised in patients on anticoagulant therapy or with other bleeding diatheses.Arteriography should be avoided if possible in such cases; if it is essential then all possible steps should be taken to correct or improve the coagulation defect before and during the procedure if this is clinically acceptable. Other factors that increase the risk of bleeding from an arterial puncture site include systemic hypertension and disorders predisposing to increased fragility of the vessel wall such as Cushing’s syndrome, prolonged steroid treatment and rare connective tissue disorders such as certain types (especially type IV) of the Ehlers–Danlos syndrome. Arteriography may be necessary in a patient with a suspected or known previous adverse reaction to contrast medium and this problem is discussed in Chapter 2. Arteriography can require larger doses of contrast medium than any other radiological procedure and particular care must be exercised in infants and in dehydrated or shocked patients, patients with serious cardiac or respiratory disease, patients in hepatic or renal failure, and other patients with serious metabolic abnormalities.

ANAESTHESIA Most arteriography is performed under local anaesthesia, though general anaesthesia is necessary for babies and young

children, for confused, difficult or very nervous patients, and in some complex and interventional procedures. Although general anaesthesia can be more pleasant for the patient than local anaesthesia and reduces motion artefact on the radiographs, it nevertheless adds to the risks of arteriography. This is not only because of the (small) risks inherent in general anaesthesia but also because it masks the patient’s subjective symptoms and reactions.These may provide the radiologist with immediate warning of a mishap such as the subintimal injection of contrast medium or the inadvertent wedging of the catheter tip in a small artery: a warning that may well prevent more serious injury.

ARTERIAL PUNCTURE It is often said that the most important part of an angiographic procedure is the initial vessel puncture and there is no doubt that a good technique is likely to make the subsequent angiogram not only more comfortable for the patient but also easier for the angiographer. It should go without saying that the most suitable route of access in order to achieve the aims of the study should have been chosen before the patient enters the angiographic suite; this decision will often depend upon a number of factors including previous imaging studies, operation notes and clinical examination. For example, when performing a diagnostic arteriogram (or angioplasty) of the main artery supplying a renal transplant it is worthwhile reviewing the operative notes to see if the arterial anastomosis was made end-to-end with the internal iliac artery or end-to-side with the external iliac artery; the former is often easier to selectively catheterize from a contralateral femoral arterial approach, whilst the latter is best studied via an ipsilateral puncture. The most common vessel punctured for diagnostic and therapeutic angiography is the common femoral artery. Axillary, brachial or radial approaches can be used but these routes are usually reserved for those patients in whom a femoral approach is not possible due to iliac occlusive disease.There has, in recent years, been a trend towards the greater use of upper extremity access even in those individuals with patent femoral vessels.This is partly because DSA enables adequate studies to be obtained by injection through smaller calibre catheters (presumed to be associated with a lower morbidity than larger catheters) and partly because of the move towards the performance of angiography as an outpatient or day-case investigation. The latter trend, motivated by economic and logistic considerations as well as those related to the patient’s convenience, favours the upper extremity approach because of the rapidity with which a patient who has undergone an axillary, brachial, or radial procedure can be mobilized in comparison with one who has had a groin puncture and the greater ease with which the upper extremity puncture or arteriotomy site can be monitored and controlled during the hours following the investigation.A popliteal artery puncture may be useful in certain instances such as angioplasty of the superficial femoral artery when it is not possible to catheterize this vessel via an antegrade approach. The technique of arterial puncture and catheterization of those vessels most commonly used for arterial access will be described.

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• ANGIOGRAPHY: PRINCIPLES, TECHNIQUES (INCLUDING CTA AND MRA) AND COMPLICATIONS

Retrograde Femoral Artery Puncture (Fig. 6.12) The artery is palpated to select the site of puncture before local anaesthetic is injected. Various anatomical descriptions are given regarding the site at which to puncture but the aim should be to enter the common femoral artery a short distance below the inguinal ligament. This is best achieved by puncturing the vessel where it can be most easily palpated, irrespective of the relationship of this point to the inguinal skin crease. Since this is the point where the artery crosses the pubic bone it is also the easiest point at which to achieve haemostasis afterwards. It is important that the arterial puncture site should be adequately anaesthetized. After cleansing the skin with a suitable preparation, 5–10 ml of 0.5–1 per cent lignocaine is infiltrated around the artery. It is important to inject the local anaesthetic agent posterior to the artery as well as anterior to it, even if a single-wall puncture of the vessel is subsequently made, as it makes the arterial puncture and subsequent catheter manipulation much more comfortable for the patient. In addition, if the puncture site is inadequately anaesthetized arterial spasm is more likely, which may make selective catheterization very difficult because of the lack of free catheter movement. After local anaesthesia a small scalpel incision is made in the skin (large enough to accommodate the anticipated catheter), the skin being temporarily drawn laterally during the incision to avert the risk of a scalpel injury to the arterial wall. A pair of fine artery forceps is inserted into the incision and used to create a tunnel through the subcutaneous tissues down to the artery. This manoeuvre is particularly

Figure 6.12 Percutaneous catheterization. One of the commonly used techniques of percutaneous arterial catheterization. The artery (1) is transfixed (2). The needle is partially withdrawn and re-angled (3). A guidewire is passed into the needle during free backflow of blood (3, 4), the needle removed and a catheter or introducer inserted over the wire (5, 6). When the catheter is safely within the arterial lumen the wire is withdrawn (7).

important in large or obese patients, for it not only facilitates catheterization but it reduces the risk of a postoperative haematoma: any escaping blood emerges through the incision and is immediately apparent, instead of collecting subcutaneously. The technique employed for puncturing the artery is a matter of personal preference. A reliable way is to feel the artery with the middle and index fingers of the left hand and insert the needle (held in the right hand) between the two palpating fingers. Gentle pressure with the needle will alter the pulse as detected by the lower (index) finger, giving reassurance as to its correct alignment. The needle is held angled forwards and passed through the anterior wall alone or right through the artery depending upon the angiographer’s preference (see later). A single-wall puncture technique is most commonly performed with a one-part needle, in which case free pulsatile blood flow will occur as soon as the needle enters the vessel lumen. When a double-wall puncture technique is performed a two-part needle is used; after passing through the vessel, the central stylet is removed, the needle angled slightly more towards the horizontal and then withdrawn at an even rate, assisted by gentle rotatory movements to avoid any sudden jerking. When the tip of the needle is safely in the arterial lumen there will be a free, spurting backflow of blood from the hub. While the needle is held steady with one hand the soft tip of a guidewire is threaded through the needle into the artery. When a sufficient length of wire is inside, the needle is removed and firm manual pressure maintained on the puncture site until the needle has been exchanged for a catheter or dilator. The guidewire is removed when the tip of the catheter is in a satisfactory position and the catheter is then flushed free of blood with heparinized saline. At the end of a correctly conducted insertion procedure there should be no bleeding around the catheter, which will move freely and painlessly through the puncture site when manipulated. There are arguments in favour of both single- and doublewall arterial punctures; some radiologists prefer to puncture only the anterior wall of the artery to minimize the trauma to the vessel, although puncture to both vessel walls does not increase the risk of complications as these are usually caused by the subsequent manipulations with guidewires and catheters. Proponents of a double-wall puncture technique maintain that this method is safer, particularly when first learning angiography, with a lesser risk of intimal dissection when introducing the guidewire through the needle when compared with a single-wall technique. Whichever method is employed, if a guidewire does not pass freely into the arterial lumen it should not be forced; this technique is never successful and usually causes an intimal dissection. If there is good backflow of blood the reason for failure of the wire to pass freely may be either that the needle is angled sharply towards one wall of the blood vessel or that the wire is passing into a small branch vessel such as the deep circumflex iliac artery. A number of manoeuvres may help, including fluoroscopy, cautious repositioning of the angle of the needle or changing to a J-wire. A gentle injection of contrast medium is possible to help ascertain the nature

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of the problem but only if good backflow is present. If there is poor backflow then the needle may be positioned near an atheromatous plaque or a stenosis, may be only partially in the lumen, or may have caused an intimal dissection. In these circumstances discretion is advised and it is wise to start again with a fresh puncture. In difficult cases remember that there is usually a femoral artery in the opposite leg! Better two groin punctures than a dissection or a large haematoma. The catheter is flushed with a heparinized saline solution throughout the procedure to prevent clotting.Various concentrations of heparin are used in different centres; the authors use 2000 IU heparin L-1 normal saline. During an average procedure it would be unusual for a patient to receive much more than about 500–1000 IU in this way. It is almost always preferable to give a firm hand flush intermittently, rather than maintain a continuous slow flush infusion. This technique not only leaves the proximal end of the catheter free for manipulation but also is more effective, since a slow infusion may only clear the proximal holes of a catheter with multiple side-ports; clot forms in the end-port and the more distal side-ports, and is then blown out into the vascular system when a pressure contrast injection is performed.

Antegrade femoral artery puncture The most common indication for an antegrade puncture of the femoral artery is when performing ipsilateral superficial femoral, popliteal or infragenicular artery angioplasty. Once again, the aim should be to puncture the common femoral artery just below the inguinal ligament. The most common problem associated with this procedure is catheterization of the profunda femoris and many methods have been described in order to manipulate the catheter out of this vessel into the superficial femoral artery when this has happened. With a good technique, however, this complication should rarely occur. Preparation of the patient is important; obese patients may have an abdominal ‘apron’ that hangs over the groin and this should be lifted superiorly and strapped out of the way with heavy-duty tape. It is well worthwhile when performing an antegrade femoral puncture to screen over the femoral head to determine the correct site for vessel puncture, as this is often considerably higher than anticipated when using palpation alone. The fovea of the femoral head is a useful landmark for the correct site of arterial access and this can be marked on the skin before local anaesthetic is infiltrated in the same way as for a retrograde puncture. A skin incision is made with a scalpel blade and blunt dissection is again performed of the soft tissues over the femoral artery. The vessel is then punctured with the needle angled slightly towards the feet. The best wire to use is a wide-angled J-wire with a curvature of radius 7.5 mm. This should be introduced through the needle with the tip of the curve directed anteriorly so that, as it exits the needle tip within the vessel lumen, it is directed into the superficial femoral artery rather than the profunda femoris. A straight guidewire is much more likely to pass directly into the latter vessel because of the posterolateral orientation of this artery.

Popliteal artery puncture The patient is placed prone on the angiography table and US is used to visualize and puncture the popliteal artery, which lies deep and lateral to the popliteal vein in the popliteal fossa. Intra-arterial vasodilator agents are commonly employed routinely to reduce the risk of spasm.

Axillary artery puncture When performed correctly this is arguably the safest puncture to perform in the upper limb47–49. The term axillary puncture is in fact incorrect as the most proximal portion of the brachial artery is punctured rather than the axillary artery itself and the description of ‘high brachial’ is preferred by some. With the arm abducted (the left arm is preferred) to approximately 135 degrees and the hand placed under the patient’s head, the artery is palpated over the superior portion of the humerus, where it is punctured after local anaesthetic infiltration and skin puncture as previously described. Complications of haematoma and pseudoaneurysm formation are seen in less than 1 per cent of patients and the feared complication of brachial plexus injury is rare but if a haematoma should develop there should be low threshold for performing surgical decompression to prevent a neurological deficit. Arterial spasm is less common than when low brachial and radial artery punctures are performed and intra-arterial vasodilator agents are not administered routinely.

Brachial artery puncture The brachial artery may be punctured anywhere along its course but the most common site is the antecubital fossa above the elbow joint, where it lies medial to the biceps tendon50. Arterial spasm is more common than when an ‘axillary’ artery puncture is performed51,52 and intra-arterial vasodilators are administered routinely by some practitioners once the vessel has been catheterized. If a sheath of greater than 7 Fr in size is likely to be necessary then a formal surgical cut-down is preferred by some.

Radial artery puncture This vessel is cannulated using a micropuncture set. A 4 Fr sheath is then introduced through which a 120 cm diagnostic catheter is manipulated over a guidewire into the aorta. Intra-arterial vasodilator agents are routinely administered to reduce the risk of arterial spasm. Before employing this route it is important to confirm that there is a good collateral supply to the hand through the ulnar artery in case of subsequent radial artery occlusion.

Selective catheterization By manipulating a catheter under fluoroscopic control it is possible to insert the catheter selectively into various branches of the vascular system such as the renal artery, coeliac axis, axillary artery, etc. Different catheter shapes are available (Fig. 6.13), each of which is suitable for a particular manoeuvre or for catheterizing certain arterial branches. Superselective catheterization (also known as subselective embolization) is the term used for the catheterization of small subsidiary arteries that themselves arise from named branch arteries and is most frequently performed during embolization procedures. A co-axial

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are taken so as to embrace all the different phases of the cardiac cycle, then appropriate mask selection will subsequently permit virtually any contrast-filled image to be presented free of artefact.The same technique can be used in those individuals who are unable to hold their breath for any length of time. Such patients should be asked to breathe normally throughout the run and multiple images are obtained before the injection of contrast medium so that a suitable mask for subtraction is available for every phase of respiration. This technique is particularly suited to visceral angiography (Fig. 6.15).

INTRAVENOUS DIGITAL SUBTRACTION ANGIOGRAPHY (IVDSA) Figure 6.13 Some examples of catheters in common use. Some of the many different catheter shapes available are illustrated. From left to right: straight flush; pigtail; cobra; and sidewinder. Note the side-ports in some of the catheters.

catheter (one that passes through the lumen of a diagnostic or guiding catheter) is often used for the catheterization of these small vessels (Fig. 6.14).

DIGITAL SUBTRACTION ANGIOGRAPHY With any DSA examination particular attention must be paid to the elimination of movement artefact. Respiratory motion needs to be controlled (a nose-clip is often helpful) and in the case of abdominal and pelvic examinations the effects of bowel movement can be minimized by the use of paralytic agents. Multiple mask acquisition is a simple and muchneglected technique for counteracting the effect of inevitable movement such as, say, the artefacts produced in pulmonary arteriograms by cardiac motion. If several precontrast images

The introduction of IVDSA was hailed as a major breakthrough and enthusiasts initially predicted that it would eventually replace intra-arterial angiography. Experience, however, has proved that IVDSA often produces nondiagnostic images because of patient movement, poor cardiac output and the superimposition of vessels. It may also require the use of very large volumes of contrast medium, which not only restricts the number of studies that can be obtained during any one procedure but also makes the technique unsuitable for patients in whom a significant load of contrast medium would be particularly undesirable; the indications for its use are now limited.

AFTERCARE When an arteriographic study is completed the catheter is withdrawn and firm manual pressure applied to the puncture site for 5–10 min. The radiologist should be absolutely satisfied that bleeding has stopped before the patient leaves the angiography suite. The wound site is then checked at regular intervals by the nursing staff, who should also record pulse and

Figure 6.14 Co-axial catheterization. (A) A coeliac axis arteriogram performed through a 7 Fr catheter shows an aneurysm at the splenic hilum. This needs selective embolization but the long tortuous splenic artery presents a formidable technical obstacle to catheterization with a conventional catheter. (B) A co-axial (tracker) catheter has been passed to the site of the aneurysm, which was successfully occluded without compromising the splenic circulation.

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Figure 6.15 Example of use of breathing technique during visceral angiography. Arterial (A) and venous (B) images from a superior mesenteric arteriogram acquired during patient respiration. Intestinal peristalsis has been obliterated with hyoscine butylbromide. The arterial and venous images were acquired at different phases of respiration but the use of different masks (mask 2 for the arterial image and mask 10 for the venous image), acquired before the injection of contrast medium, has resulted in excellent quality images.

blood pressure observations for a reasonable period following the procedure and check that distal pulses remain palpable. Pressure pads, sand bags and other accoutrements are generally a waste of time. It is much better to be able to see the puncture site than to cover it up. If bleeding does not stop from a puncture site, press for longer! Almost all post-catheterization bleeding can ultimately be controlled by local pressure unless the artery has been torn or there is a coagulation abnormality. Several devices are now available that are used to seal the puncture point in the vessel wall at the time of catheter withdrawal in order to allow rapid mobilization and discharge of the patient, often within 1–2 h53–55. These have become popular in many centres following cardiac catheterization and are being utilized more commonly in peripheral angiography. Care should be exercised in their use, however, as a least one meta-analysis of published data has demonstrated little evidence that they are effective and concluded that they may be associated with an increased risk of haematoma and pseudoaneurysm formation56. An adequate record of the procedure should be entered in the patient’s case notes. This should include: • date • name of the operator • puncture site • catheter size • studies performed • names and doses of anaesthetic agents • volumes and concentrations of contrast medium and other drugs administered • preliminary findings • any complications during the procedure • integrity or otherwise of the pulses peripheral to the puncture site at the end of the procedure • post-procedural nursing instructions.

These notes are important not only for patient care but also as a medico-legal record and they should be comprehensive and accurate.

COMPLICATIONS The principal complications of arteriography are listed in Table 6.1.

Contrast medium reactions The many possible adverse effects of contrast media, together with details of their prevention and treatment, are discussed in Chapter 2.

Adverse drug reactions Apart from contrast medium reactions, local anaesthetic agents or other drugs given during the procedure may cause adverse or idiosyncratic effects.

Puncture site complications Haemorrhage Haemorrhage may occur from the puncture site and may cause external blood loss or a subcutaneous haematoma, which can result in extensive bruising (Fig. 6.16). Retroperitoneal bleeding is uncommon but may occur if the arterial puncture is performed above the inguinal ligament. Perhaps the commonest reason for this is the introduction of the needle during retrograde femoral artery catheterization at too acute an angle so that it passes through the back wall of the external iliac artery above the inguinal ligament. It is less well recognized, however, that this complication may also occur when the femoral artery is correctly punctured; this is probably due to a downward extension of the pelvic and abdominal wall fascial layers around the femoral artery and vein, the so-called femo-

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Table 6.1

• ANGIOGRAPHY: PRINCIPLES, TECHNIQUES (INCLUDING CTA AND MRA) AND COMPLICATIONS

COMPLICATIONS OF ARTERIOGRAPHY

Contrast medium-related complications Minor adverse reactions Major adverse reactions and death Local vascular changes (effects on blood cells, viscosity, vascular tone; results of extravasation, etc.) Systemic vascular changes (effects on blood volume, osmolality, etc.) Individual organ toxicity (heart, kidney, brain, etc.) Adverse reactions to local anaesthetic or other drugs Puncture site complications Haemorrhage (external bleeding or haematoma) Intramural or perivascular injection of contrast medium Vascular thrombosis (dissection, local trauma) Peripheral embolization from puncture site Vascular stenosis or occlusion Aneurysm or pseudoaneurysm formation AV fistula Local sepsis Damage to nerves Damage to other local structures Catheter-related and general complications Catheter thrombus embolism Air embolism Gauze embolism Dissection, perforation or rupture of vessels Organ ischaemia or infarction secondary to spasm, dissection or embolism Interventional accidents Fracture and loss of guidewire or catheter fragments Knot formation in catheters Inadvertent injection of toxic material (e.g. skin-cleansing lotion) Inadvertent overheparinization Vasovagal reaction

Figure 6.16 Extensive groin haematomas in a patient following bilateral femoral artery catheterization.

ral sheath. If this sheath is transgressed at the time of vessel puncture, and bleeding subsequently occurs from the puncture site after catheter or sheath removal for any reason, then blood may spread along the fascial planes continuous with the femoral sheath into the retroperitoneum or, indeed, into the anterior abdominal wall. Four types of haematoma may occur after femoral artery puncture: (A) abdominal wall, (B) retroperitoneal, (C) groin and thigh and (D) intraperitoneal. The first three may all result from puncture of the common femoral artery (i.e. below the inguinal ligament), with (A) and (B) resulting when there is bleeding into the femoral sheath and (C), the commonest, when there is spread into the femoral triangle. Intraperitoneal haemorrhage is even less frequent than retroperitoneal bleeding and generally requires transgression of the peritoneum itself, which is more likely to occur if vessel puncture is performed above the level of the inguinal ligament. Some intraperitoneal bleeding has also been described, however, when the arterial puncture is below the inguinal ligament; it has been postulated that this may be due to the presence of defects in the parietal peritoneum. Excessive bleeding is usually the result of bad technique. Particular caution is necessary in patients with a bleeding diathesis or hypertension and following the use of balloon catheters in transluminal angioplasty. If inexplicable bleeding continues, check that inadvertent over-heparinization has not occurred.This can be corrected if necessary by the administration of protamine sulphate (10 mg of which counteracts the effects of approximately 1000 IU heparin).

Intramural and perivascular contrast medium injection Contrast medium may be inadvertently injected into the wall of a vessel or outside the vessel (perivascular). In most cases little harm (apart from pain) results from a perivascular injection but it is possible to dissect and occlude an artery with a subintimal injection of contrast medium. Never inject into a needle or catheter that does not exhibit free backflow. Vascular thrombosis can result from severe trauma to the vessel at the puncture site or from a subintimal contrast injection. It is also possible that thrombus wiped off the outside of the arterial catheter during its extraction forms a nidus for thrombus at the puncture site.Vascular trauma at the puncture site can be minimized by good technique. Never use force to introduce a wire into a vessel; if it does not pass easily something is wrong! It is often better for the inexperienced operator to start again with a fresh needle puncture than to persist with one that is causing problems. A 10-minute delay with a successful outcome is always preferable to a dissection and/or a groin haematoma. Peripheral embolization from the puncture site probably occurs to a minor degree in many cases but clinically obvious embolization is rare. Local vascular complications such as false aneurysm (pseudo-aneurysm) or AV fistula formation and late stenosis or occlusion can all result from arterial procedures. Good technique is the best preventive measure. Femoral artery pseudoaneurysms are usually due to a combination of a low puncture

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– the superficial femoral or profunda femoris arteries have often been cannulated instead of the common femoral artery when this complication occurs – and inadequate compression of the vessel at the end of the procedure. Local sepsis may occur following an arterial puncture (although it is extremely rare) and this factor is particularly important if early surgery is contemplated. The utmost care should be taken to observe sterile precautions and when a study is performed in a patient with local skin contamination (e.g. open wound, ileostomy, etc.) a protective adhesive sheet helps to keep the operative field uncontaminated. Injury to local structures such as nerves, joints and bones is rarely of clinical significance. Occasionally damage to branches of the femoral nerve gives rise to areas of cutaneous anaesthesia or paraesthesia in the thigh57. These normally recover completely, although this may take several months in some instances.

Catheter-related and general complications 1 Thrombi can form in or on a catheter (Fig. 6.17) and be ejected into the vascular system. This is always undesirable and in areas such as the cerebral or coronary circulation is extremely dangerous. Catheters should be flushed assiduously during all arteriographic procedures to prevent thrombus formation. Other types of embolism that may occur are air embolism (sometimes from incorrectly loaded pressure injectors) and thread embolism from fragments of gauze swab (Fig. 6.17). 2 Vascular injuries distant from the puncture site may be produced by the catheter or guidewire, or by the intramural injection of contrast medium or saline. If there is ever any doubt about whether a needle or catheter tip is in a satisfactory position, contrast medium should always be injected in preference to saline. Under fluoroscopic control, it is then possible to stop the injection immediately if any extravasation or other mishap is apparent. The most common injury is dissection of the tunica intima from the tunica media and this complication is far more likely to occur in previously diseased vessels than in normal vessels. The intima forms a raised flap (Fig. 6.18) that may completely occlude the ves-

Figure 6.18 Dissection of the aorta and iliac vessels. A catheter has been passed subintimally from the right groin and created an extensive arterial dissection involving the aorta and both iliac systems. Note the suspiciously straight edge of the aortic opacification and the ‘twisted tape’ appearance of the left common iliac artery as the dissection spirals down the vessel (arrow).

sel. It is also possible to perforate vessels with a guidewire or to rupture them when a forced injection is made through a catheter wedged into a vessel of the same calibre. 3 Injuries to organs are normally caused by ischaemia during arteriographic procedures (other than those related to the effects of contrast medium. This may occur through wedging of the catheter so that the normal flow to an organ is obstructed; dissection of the feeding artery; spasm, thrombosis or rupture of the feeding artery; or embolism. The ischaemia resulting from one or more of the earlier mentioned events may have no observable clinical sequelae, may result

Figure 6.17 Catheter thrombus. During a renal arteriogram (A) a filling defect was noted at the catheter tip (arrow). On cautious withdrawal of the catheter a thrombus was found protruding from the side ports. (B) Note the fragment of gauze thread on the catheter tip; this was incorporated in the clot and presumably caused it. Care should always be taken when handling swabs in association with wires or catheters.

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• ANGIOGRAPHY: PRINCIPLES, TECHNIQUES (INCLUDING CTA AND MRA) AND COMPLICATIONS

in temporary or permanent functional abnormalities in the affected organ or system, or may cause infarction of the organ. The clinical importance of these accidents depends very much on the vascular territory in which they occur; complete occlusion of a carotid, coronary or renal artery is likely to have disastrous consequences whereas occlusion of, say, a hepatic artery may not necessarily produce any adverse effects. 4 Guidewire fracture. Occasionally fragments of guidewire or catheter may become detached within the vascular system. Catheters may also become knotted during overenthusiastic manipulation procedures.With good technique they should not occur in the first place. 5 Injection accidents. Tragedies have occurred when toxic substances have been inadvertently injected into blood vessels.

Skin-cleansing fluids should always be removed from the instrument trolley immediately after the puncture site has been prepared. Drugs should always be double-checked before injection through a catheter and particular care is necessary with heparin, which is available in solutions that differ considerably in their concentration. 6 Vasovagal reactions. Vagally mediated reactions may occur during arteriography in response either to the injection of contrast medium or to the discomfort and psychological effects of the procedure. Bradycardia is a prominent feature of such reactions, which must be distinguished from acute allergic responses to the contrast medium or local anaesthetic. The incidence of vasovagal reactions is considerably reduced if proper premedication is employed.

CATHETER VENOGRAPHY TECHNIQUE Direct venography This is the most common means of opacifying the venous system. A needle or catheter is placed directly into the vein to be imaged and contrast medium injected. This can be done using an antegrade approach, for example a needle inserted into an antecubital fossa vein to demonstrate the brachial, axillary, subclavian and brachiocephalic veins and superior vena cava, or a catheter placed in the iliac veins to demonstrate the inferior vena cava. Alternatively, a catheter may be directed retrogradely into the venous system, as in selective hepatic venography, which is now most commonly performed as part of a transjugular liver biopsy procedure or when performing a transjugular intrahepatic portosystemic stent shunt.

Indirect venography In this approach contrast medium is injected into the arterial system of the area being studied and delayed films are obtained to image the venous return. This technique is used to great advantage for examining the portal venous system, for while it is possible to gain direct access to the portal venous system, the latter approach is more invasive and requires greater technical expertise than the indirect approach. Other areas where indirect venography is used include the cerebral and renal veins.

COMPLICATIONS The complications that may occur following venography can be divided into three main groups: those occurring locally at the site of injection, those occurring within the vascular system, and systemic complications. Many of the complications of arteriography (Table 6.1) also apply to venography.

Local complications Pain may occur at the site of contrast injection. Hypertonic contrast agents in particular may cause local pain due to irritation of the vessel. Extravasation of contrast medium will also cause pain due to both the irritant nature and volume of the medium injected. Extravasation of contrast media into the skin may cause tissue necrosis and sloughing of skin, although this is less likely with low-osmolality agents. When injecting any substance careful attention should be paid to the patient’s comments. Severe pain at the injection site should never be ignored, as it does not occur if the needle is correctly positioned within the vein.

Vascular complications Dissection of the vein being cannulated/catheterized may occur at the puncture site or distally, due to catheter and/or guidewire manipulation.

Intraosseous venography

Systemic complications

This technique is now of historical interest only; it involves the injection of contrast medium into the bone marrow cavity at a site distal to the venous system that is to be imaged. Contrast medium is injected rapidly into the marrow cavity from which it drains via sinusoids and small veins into the deep venous system of the region. The technique was most widely used for leg venography with injections made into the malleoli or calcaneum for the calf, into the tibial tubercle or femoral condyle for the thigh, and into the greater trochanter to show the pelvic veins and inferior vena cava.

1 Pulmonary embolism. Ultrasound is now most commonly used to confirm or exclude lower limb deep vein thrombosis and ascending venography is performed infrequently. If venography is considered necessary it should be remembered that calf and femoral vein clots, if loose, can be dislodged by manipulation of the leg and the forceful injection of contrast medium. This complication is rare but when deep vein thrombosis is demonstrated, vigorous massage of the calf is inadvisable. Clots in the iliac vein, renal vein, hepatic vein and inferior vena cava can all be dislodged by the manipulation

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of catheters and guidewires, as well as by pressure injections of contrast medium. 2 Cardiac arrhythmias. The manipulation of catheters in the right side of the heart for any reason can cause arrhythmias. Injections of large volumes of contrast medium in the right heart, especially for pulmonary angiography, may also

cause dysrhythmias and continuous electrocardiographic monitoring is essential. 3 Air embolism. Because of the danger of air embolism in the venous system, care should be taken neither to inadvertently inject air into a vein, nor to leave the lumen of any large-bore cannula or catheter open to the atmosphere.

EMBOLIZATION TECHNIQUES The deliberate occlusion of arteries, veins or abnormal vascular spaces by embolic material injected through a selectively positioned catheter or needle is one of the major therapeutic applications of interventional radiology. Several of the indications for embolization are covered in other chapters in this book and for this reason only the important basic principles and techniques of the procedure will be discussed here.

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•

EMBOLIC MATERIALS AND TECHNIQUES Many different materials have been used to embolize blood vessels but those most commonly used are metallic coils, absorbable gelatin sponge, polyvinyl alcohol foam, sodium tetradechyl sulphate, absolute ethanol and N-butyl-2-cyanoacrylate. The choice of agent in any particular setting will depend upon the indication for embolization and the angiographic findings. Detailed descriptions of the radiological techniques employed in embolization are not included here but a few important points should be noted: • Embolization should always be preceded by high-quality angiography to define precisely the vascular territory under consideration and its anastomotic communications. • Embolizing catheters should be sited as selectively as possible to avoid the unintentional embolization of adjacent vascular territories. Coaxial catheters may be necessary to achieve the necessary degree of selectivity and failure to obtain a satisfactory catheter position will require the procedure to be abandoned. • The likely route of any ‘overspill’ of emboli should be noted during preliminary angiography. Reflux of emboli into other vessels occurs very easily as flow progressively diminishes in the embolized bed and this hazard is best avoided by injecting emboli in small quantities at a time, under constant fluoroscopic control, or by the use of a balloon catheter (see later). It is important to remember that a selective preliminary arteriogram may not reveal all routes of collateral flow, even if they communicate directly with the vessel into which contrast medium is being injected. This is because the blood flow in collateral vessels may be towards the lesion under investigation (especially if it is very vascular or is shunting into a vein) and therefore they may not necessarily be opacified when contrast medium is injected into a principal vessel feeding the lesion. When the flow haemodynamics are changed during the embolization procedure, flow may reverse along such collaterals

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and carry embolic material to an adjacent bed that was not visualized on the preliminary arteriograms and, therefore, fallaciously regarded as being unconnected with the lesion being embolized. Non-opaque emboli should always be injected as a suspension in contrast medium so that they are visible as filling defects during the injection sequence. Particular care is necessary in certain areas (e.g. CNS, lower GI tract, peripheral limb vessels); where there has been previous surgery (which may have reduced any potential collateral circulation to a region); and where there is preexisting vascular insufficiency (e.g. atheroma). It may be useful to employ a balloon catheter in some situations, either to reduce flow through the bed to be embolized or as a temporary obstruction to flow in nearby branches supplying normal structures when there seems a potential danger of emboli entering such vessels. It is prudent to cover some embolization procedures with antibiotic prophylaxis.

INDICATIONS FOR THERAPEUTIC ARTERIAL EMBOLIZATION There are numerous indications for arterial embolization but the most common are: • the control of severe bleeding (e.g. gastrointestinal or posttraumatic haemorrhage from the liver, kidney, or spleen) • the reduction of blood flow to certain tumours as a definitive treatment (e.g. uterine fibroids, certain aneurysmal bone cysts) • as a preoperative measure (e.g. juvenile nasal angiofibroma, meningioma) • as a palliative measure (e.g. hepatocellular carcinoma, neuroendocrine hepatic metastases) • in the management of vascular malformations. The most common indication for venous embolization is the treatment of male varicoceles. Interested readers are referred to other chapters in this book for more details of some of these techniques. The incidence of serious complications following embolization is low provided that the procedure is conducted with patience and scrupulous care and that only suitable patients are selected for treatment by the technique.

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• ANGIOGRAPHY: PRINCIPLES, TECHNIQUES (INCLUDING CTA AND MRA) AND COMPLICATIONS

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50.

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aorta and lower extremity arteries: preliminary experience. Radiology 211:59–67 van den Brink J S, Watanabe Y, Kuhl C K et al 2003 Implications of SENSE MR in routine clinical practice. Eur J Radiol 46:3–27 Herborn C U 2004 On the difference between 0.5-M and 1.0-M gadolinium chelates for 3D MR angiography. Eur Radiol 14(Suppl 5):M21–M22 Lin W, An H, Chen Y et al 2003 Practical consideration for 3T imaging. Magn Reson Imaging Clin N Am 11:615–639 McIvor J, Rhymer J C 1992 245 transaxillary arteriograms in arteriopathic patients: success rate and complications. Clin Radiol 45:390–394 Inglese L, Lupattelli T, Carbone G L et al 2004 Axillary artery access for interventional procedures. J Endovasc Ther 11:414–418 Ray C E, Kaufman J A1996 Complications of diagnostic angiography. In: Ansell G, Bettman M A, Kaufman J A, Wilkins R A (eds) Complications in Diagnostic Imaging and Interventional Radiology, 3rd edn. Blackwell Science, Cambridge, chapter 14, pp 303–316 Heenan S D, Grubnic S, Buckenham T M, Belli A M 1996 Transbrachial arteriography: indications and complications. Clin Radiol 51: 205–209 Hildick-Smith D J, Khan Z I, Shapiro L M, Petch M C 2002 Occasionaloperator percutaneous brachial coronary angiography: first, do no arm. Catheter Cardiovasc Interv 57:161–165

52. Chatziioannou A, Ladopoulos C, Mourikis D, Katsenis K, Spanomihos G, Vlachos L 2004 Complications of lower-extremity outpatient arteriography via low brachial artery. Cardiovasc Intervent Radiol 27:31–34 53. Boccalandro F, Assali A, Fujise K, Smalling R W, Sdringola S 2004 Vascular access site complications with the use of closure devices in patients treated with platelet glycoprotein IIb/IIIa inhibitors during rescue angioplasty. Catheter Cardiovasc Interv 63:284–289 54. Nikolsky E, Mehran R, Halkin A et al 2004 Vascular complications associated with arteriotomy closure devices in patients undergoing percutaneous coronary procedures: a meta-analysis. J Am Coll Cardiol 44:1200–1209 55. Applegate R J, Sacrinty M, Kutcher M A et al 2006 Vascular complications with newer generations of angioseal vascular closure devices. J Interv Cardiol 19:67–74 56. Koreny M, Riedmuller E, Nikfardjam M, Siostrzonek P, Mullner M 2004 Arterial puncture closing devices compared with standard manual compression after cardiac catheterization: systematic review and metaanalysis. JAMA 291:350–357 57. Hallett J W Jr, Wolk S W, Cherry K J Jr, Gloviczki P, Pairolero P C 1990 The femoral neuralgia syndrome after arterial catheter trauma. J Vasc Surg 11:702–706

CHAPTER

Radionuclide Imaging: General Principles

7

Djilda Segerman and Kenneth A. Miles

• • • • • • • • •

Introduction and general principles Basic physics The gamma camera Principles of positron emission tomography Hybrid systems: SPECT-CT and PET-CT Quantification and image processing Image interpretation Nonimaging applications Radiopharmaceuticals

INTRODUCTION AND GENERAL PRINCIPLES Frequently, the first impression of radionuclide images is one of poor anatomical definition compared with other imaging techniques. If the aim of diagnostic imaging were merely the demonstration of human anatomy, the lower spatial resolution of radionuclide imaging would indeed be a disadvantage. However, the intent of diagnostic imaging is not illustration of anatomy, rather the improvement of health outcomes for patients through more accurate diagnosis. The increasing incorporation of radionuclide imaging into national clinical management guidelines, for example the National Institute for Clinical Excellence (NICE) guidance on the use of myocardial perfusion scintigraphy in the UK, testifies to the ability of radionuclide imaging to impact positively on the care of patients. The aim of this chapter is to describe how the ‘fuzzy’ images produced using radionuclides achieve this clinical impact. Radionuclide images depict organ function rather than structure.The images are obtained by mapping the distribution of an administered radiopharmaceutical within the body.Thus, as for X-ray radiography and computed tomography, radionuclide imaging uses ionizing radiation. However, the radiation is emitted from within the patient and subsequently detected in the imaging device, rather than transmitted through the patient from an external X-ray source. The specific organ

function depicted is determined by the biological behaviour of the radiopharmaceutical. For example, a radiopharmaceutical excreted by the kidney will evaluate renal function whilst one excreted in the bile will depict the biliary tree. The biological behaviour of radiopharmaceuticals can be traced to particular cellular or molecular mechanisms. Thus, by using different radiopharmaceuticals, a range of cell types can be imaged within the same organ (Fig. 7.1), or different molecular mechanisms with the same cell type (Fig. 7.2). It is perhaps ironic that, despite limited spatial resolution, radionuclide images depict structures (i.e. cells and molecules) that are too small to be visualized by anatomical imaging techniques. Another advantage of radionuclide imaging is that quantification of the amount of radiopharmaceutical within a particular organ can be achieved through counting the amount of radiation emitted within a specific time. By obtaining repeated images, the change in concentration of radiopharmaceutical over time can be measured and displayed as time–activity curves, giving additional insights into the functional status of tissues (Fig. 7.3). With appropriate mathematical modelling, these changes in concentration can be used to quantify particular physiological processes, such as glomerular filtration rate (GFR) or left ventricular ejection fraction. As repeated images can be obtained after a single administration of radiopharmaceutical, these temporal data are obtained with no additional radiation exposure to the patient. It is the ability of radionuclide imaging to depict function that confers the ability to impact upon health outcomes. The magnitude of change in functional activity, rather than the size of an abnormality, determines whether a disease can be detected by radionuclide imaging. Radionuclide imaging devices are highly sensitive and can detect relatively small changes in functional activity. Thus, a change in tissue function is frequently detectable before any change in structure has occurred or can be detected confidently by anatomical imaging. For example, increased glucose metabolism of cancer cells within lymph nodes can be detected using positron emission tomography (PET) before the size of the lymph

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• IMAGING TECHNIQUES AND GENERAL ISSUES

Figure 7.1 Radionuclide images depicting three different cell types within the liver obtained using different radiopharmaceuticals. (A) 99mTc sulphur colloid taken up by Küpfer cells. (B) 99mTc HIDA taken up and excreted by hepatocytes. (C) Leucocytes labelled (ex-vivo) using 99mTc HMPAO. (The crosses mark the costal margins.)

node exceeds the threshold to be diagnosed as abnormal on computed tomography (CT) (Fig. 7.4). Similarly, successful treatment may return functional activity to normal before anatomical abnormalities have resolved and thus radionuclide imaging can be advantageous in monitoring treatment (e.g. for patients with cancer). In other circumstances, the functional status of an organ is more important for determining clinical management than its structural appearance. For instance, the anatomical finding of a dilated renal collecting system does not necessarily indicate renal tract obstruction. It is the presence of obstruction and its impact on renal function, as demonstrated by renal radionuclide imaging, that more accurately determines whether clinical intervention is required. More recently, the limited anatomical definition of radionuclide imaging has been addressed to some degree by the development of hybrid imaging in which radionuclide imaging devices are combined with computed tomography in a single imaging system. The resulting images display the functional data obtained from the radionuclide distribution (usually in colour), overlaid on the anatomical information from CT (usually in grey scale) (Fig. 7.5). Because the two image data sets are acquired using the same imaging system within a short time of each other, the two image data sets can be co-registered more accurately than when obtained on separate imaging systems. Not only do these hybrid systems allow abnormalities

Figure 7.2 Radionuclide images of breast cancer (arrows) demonstrating different uptake mechanisms in the same cell type (adenocarcinoma). Uptake of 18F-FDG on PET (A) is due to increased expression of glucose transporters and hexokinase. Uptake and retention of 99mTc methoxyisobutylisonitrile on gamma scintigraphy (B) are dependent on perfusion and expression of multidrug resistance pglycoprotein. (Note uptake in nodal metastases in the axilla in case B, as indicated by the arrowhead.) Figure 7.4 A mediastinal lymph node (arrow) is not sufficiently enlarged to confidently diagnose nodal metastasis on CT (A) but demonstrates uptake of 18F-FDG on PET (B), implying the presence of tumour cells. (* = primary tumour.) MAG3 Renogram (LASIX AHEAD) Right peak

4000 3600 3200

Left kidney Right kidney Right uptake region Left uptake region Aorta Bladder

2800 2400 2000 1600 1200

Left peak

Activity (counts)

130

800 400 0 0

4

8

12

16 20 Minutes

24

28

32

Figure 7.3 Time–activity curves from the kidneys, bladder and aorta following injection of a radiopharmaceutical (MAG-3) that is excreted by the kidney.

Figure 7.5 PET-CT images from a patient with active lymphoma. The 18 F-FDG uptake has been displayed in colour over grey scale CT images in (A) transaxial, (B) coronal and (C) sagittal formats.

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seen on radionuclide images to be assigned to precise anatomical structures, they also enable the morphological appearances of disease processes depicted by CT to be assimilated into the interpretation of the findings on radionuclide images. For example, such combined interpretation can aid the distinction of malignant and inflammatory causes of uptake of the positron emitting radiopharmaceutical 18F-fluorodeoxyglucose (FDG). Further advantages of hybrid systems include the use of the CT data to correct radionuclide images for artefacts resulting from attenuation of photons within the body and the ability to incorporate radionuclide image data into CT-based radiotherapy planning systems.

• RADIONUCLIDE IMAGING: GENERAL PRINCIPLES

In 1910 Rutherford established experimentally that each atom had to have a dense, positively charged nucleus which was surrounded by a cloud of negative electrons. Bohr’s model (1913) included electrons in discrete orbits circling about the dense nucleus. This is the model of the atom generally accepted today. Atoms are composed of three types of particle: protons, neutrons and electrons. The protons and neutrons form a central nucleus. Each proton has a positive electric charge (+); neutrons have no charge; electrons have a negative electric charge (−) and orbit around the nucleus, attracted by its positive charge. Protons and neutrons are nearly 2000 times heavier than electrons, so most of the mass of an atom is in the nucleus. A nucleus is described by its mass number, A, which is the sum of the number of neutrons (N) and protons. The atomic number, Z, is equal to the number of protons. The mass number is therefore given by:

it will emit radiation either as photons (X or gamma rays) or as high-velocity particles.The remaining protons and neutrons will form a daughter nucleus, which may itself be radioactive, or may be stable. The particular nature and energy of emissions are characteristic of each radionuclide. The radiation generally is ionizing, meaning that it is highly energetic and damaging to living tissue. The rate of radioactive decay is characteristic of each radionuclide. Decay is governed by probability, which means that it is impossible to predict when a particular nucleus will disintegrate. However, with large numbers of nuclei the overall rate of decay is entirely predictable. As the nuclei in a sample of radionuclide disintegrate, the number of radioactive nuclei left is reduced, and the rate of decay reduces in proportion. This effect is described mathematically by the equations of exponential decay. An important parameter is the half-life (T1/2) which is defined as the time period during which half the nuclei will decay. For example, the half-life of 131I is 8 d. In the course of 8 d, half the 131I nuclei in a sample will have decayed, and the rate of decay will have fallen to half its original value. The relationship is extremely accurate for large numbers of nuclei; with smaller numbers it is true but only on average. This can be observed in the erratic clicks of a Geiger counter that is picking up just a few counts per second. The rate of radioactive decay of a sample is called activity. The SI unit of activity is the becquerel (Bq), defined as one nuclear disintegration per second. Doses of radioisotopes given to patients for diagnostic purposes are usually measured in kBq or MBq. Originally the unit of activity was the curie (Ci) defined as 3.7 × 1010 disintegrations per second, thought to be the number of disintegrations found in 1 g of radium. Conversion factors are:

A=Z+N

1 mCi = 37 MBq; 1 microCi = 0.037 MBq.

BASIC PHYSICS Structure of the atom

The atomic number Z is constant for a given element. The same element can have varying numbers of neutrons; these different forms of an element are called isotopes. The chemical properties of the different isotopes of an element are the same. Isotopes with more neutrons will be heavier, so some physical properties may differ. To distinguish the isotopes of an element, the mass number A is normally written as a superscript. For example, the isotope of iodine with A = 131 is written 131I. Other conventions are: I131 and 53131I. In the latter example the subscript signifies Z, the atomic number of the element. A nuclide is the term to denote a type of atom characterized by both Z and N. A radionuclide is a nuclide that is radioactive.

Radioactivity A radioactive nucleus is one that is unstable and will spontaneously disintegrate or decay at some time. The reason for the instability is either that the nucleus is too large or because there is an imbalance between the protons and neutrons. Unstable nuclei have an excess of energy. When a nucleus disintegrates,

Table 7.1 lists some medically important radionuclides.

Alpha radiation Alpha decay occurs only when a heavy nucleus ejects an alpha particle. The alpha particle (α) consists of two neutrons and two protons and is identical to a helium nucleus. Alpha particles interact strongly with matter, so have a very short range, typically 1 mm or less in tissue.Within this range alpha particles repeatedly collide with atoms, disrupting their chemistry, which is extremely damaging in living tissue. Historically, radium and radon were the principal alpha emitters of medical interest.These are no longer used in medicine. Other alpha emitters are being researched for therapeutic approaches using radiopharmaceuticals that can target the delivery of short half-life alpha emitters into cancerous cells2. Due to their very short range, alpha particles have the potential to deliver a lethal radiation dose to small metastatic cell clusters, while mostly sparing the surrounding tissue. All work with alpha emitters must be conducted under very strictly controlled conditions. A significant health hazard would exist

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Table 7.1 PHYSICAL PROPERTIES OF SOME COMMONLY USED RADIONUCLIDES Main emissions (keV) Radionuclide

Half-life

Main use

Gamma or X

Beta (Emax)

Carbon-11

11

C

20 min

p

511

960

Carbon-14

14

C

5730 years

n

–

157

27.7 d

n

320

–

51

Cr

Chromium-51 Cobalt-57

57

270 d

n

122

–

Cobalt-58

58

Co

71 d

n

811

475

Fluorine-18

18

109 min

p

511

634

Gallium-67

67

Indium-111

111

Co F

Ga In

78 h

g

93, 185, 300

–

67 h

g

171, 245

–

Iodine-123

123

13.2 h

g

159

–

Iodine-125

125

60 d

n

27, 31

–

Iodine-131

131

8d

t (g)

365

606

I I I 81m

Kr

Krypton-81m Nitrogen-13 Oxygen-15 Selenium-75

13

N

15

O

75

Se

89

Strontium-89

Sr 99m

Tc

Technetium-99m 201

Tl

Thallium-201

13.5 s

g

191

–

10 min

p

511

1199

2 min

p

511

1732

120 d

g

136, 265, 401

–

50.7 d

t

–

1492 –

6h

g

140

3d

g

71

–

Xenon-133

133

5.2 d

g

31, 81

346

Yttrium-90

90

2.7 d

t

–

2284

Xe

Y

Main use: g = gamma camera, n = nonimaging diagnosis, p = PET, t = therapy. Beta (Emax): maximum energy of continuous spectrum; mean energy = approximately one-third of Emax. Main source of decay data1.

if these radionuclides were inadvertently ingested or inhaled, as cells lining the tracheobronchial tree or gastrointestinal tract would then receive high doses of radiation.

Beta radiation A beta particle (beta− or β−) is an electron emitted by a nucleus. Beta decay occurs in nuclei that contain an excess of neutrons. One of the neutrons (n) will convert to a proton (p+), an electron (e−) and a neutrino (ν). This can be visualized as: n → p+ + β− + ν + energy The proton remains in the nucleus, and the beta and neutrino are emitted with high velocity. The electron and the neutrino share a set quantity of energy in variable proportions, so the electrons will have a continuous spectrum of energies. An example of beta decay is that of 89Sr: 89 Sr 38

→ 3989 Y + β− + ν + energy

One of the neutrons has converted to a proton, increasing the atomic number from 38 to 39; strontium is converted to yttrium.The neutrino has an extremely small mass, no electric charge and interacts extremely weakly with matter. It will have no effect on the patient. Beta particles cause damage to tissue in a similar manner to alphas, but less severely. They follow a zig-zag path, undergoing many collisions with atoms. At each collision the beta gives up some of its energy, causing excitation and

ionization of the atoms. The range of betas depends on their initial energy and the electron density of the material.Typical ranges in tissue are a few millimetres; in air the range may be up to several metres. Beta emitters are routinely used to deliver therapeutic radiation doses to tumours.

Positron emission A positron (beta+, β+ or e+) is a particle similar to an electron except that it has a positive electric charge. The decay process is a type of beta decay, except in this case the nucleus has an excess of protons. p+ → n + β+ + ν + energy Positron emitters are used in PET imaging. An example is fluorine-18: 18 F 9

→ 818O + β+ + ν + energy

The behaviour of positrons in tissue is very similar to beta particles, with one important difference: once a positron has been slowed down by atomic collisions, it is annihilated by interaction with an electron from a nearby atom. The combined mass of the β+ and e− is converted into two annihilation photons, each with energy 511 keV, in accordance with Einstein’s famous equation, E = mc2: β+ + e− → γ + γ

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The two photons are emitted at 180 degrees to each other, a property exploited by PET.

X and gamma radiation Definitions Gamma rays and X-rays are both a form of electromagnetic radiation, and have the same properties. The two names arose historically. Roentgen discovered and named X-rays; these were produced by a beam of electrons striking a heavy object, the same mechanism as in a modern X-ray tube. A few years later, the names alpha, beta and gamma were given to the different rays emitted by radioactive uranium. It was only later realized that X-rays and gamma rays are essentially the same. The modern definitions are that gamma rays are emitted by an atomic nucleus in an excited state, whereas X-rays are the result of changes involving electrons. The properties exhibited are both wave-like and particle-like. The particles are called photons. Each photon has a definite energy and zero mass. Its wave-like properties are a frequency which is in proportion to its energy, and a wavelength which is in inverse proportion.

Reactions that produce X and gamma radiation Some beta emitters produce gamma photons in the course of beta decay.The decay is a two-stage process, with beta emission leaving the nucleus in an excited state. The excited nucleus then decays to its stable daughter, emitting the excess energy as a gamma photon. There may be several possible excited states, each with a particular probability of occurrence. The result will be a range of gamma photons, each with its specific energy and probability. An example is the beta emitter 131I, whose principal gamma ray is used in imaging. Certain excited nuclei do not decay immediately but remain in their excited state for some time. These nuclei are termed metastable.This type of decay is known as an isomeric transition (IT) because the parent and daughter nuclei are isomers, i.e. they contain the same number of protons and nuclei. An example is the metastable radionuclide of technetium, 99mTc, which emits a single gamma photon of energy 140 keV, and is the most commonly used radionuclide for imaging. In many decay schemes, nuclear energy is transferred to the atom’s orbital electrons. When the orbital electrons revert to their ground state they emit characteristic X-rays. In some cases, instead of producing an X-ray, the energy will eject an outer orbital electron as an Auger electron. In their properties, and in their interactions with matter, X-rays are the same as gamma rays, and Auger electrons are the same as beta particles. Electron capture (EC) is another radioactive decay process that can occur in a nucleus with excess protons. One of the orbital electrons is captured by a proton in the nucleus and changes the proton into a neutron. If the original nucleus has enough excess energy to produce two 511 keV photons, it may decay by either positron emission or electron capture. If the excess energy is less than 1022 keV, then positron emission cannot occur and electron capture is the only decay mode. Electron capture can be visualized as: p+ + e−→ n + ν + energy

• RADIONUCLIDE IMAGING: GENERAL PRINCIPLES

Some of the energy is carried off by the neutrino and some is transferred to the orbital electrons of the atom, resulting in characteristic X-rays and possibly Auger electrons. In some cases EC will leave the nucleus in an excited state, which will decay to its stable daughter by gamma emission. Beta particles and Auger electrons can give rise to another form of radiation called Bremsstrahlung (German for braking radiation).When a high-energy beta particle passes close to the nucleus of another atom, electrostatic attraction deflects and decelerates the beta particle, and its lost energy is converted to an X-ray photon (Fig. 7.6). This Bremsstrahlung radiation has a continuous spectrum. The energy tends to be higher when the betas have high energy and the material they are passing through has a high atomic number. For this reason radiation shields for high-energy beta emitters are made of plastic which has a low atomic number.

Interaction of X and gamma radiation with matter (Fig 7.7) Two types of interaction are important in nuclear medicine: • In a photoelectric absorption, an atom absorbs all the energy of an incident photon. The photon ceases to exist but the energy is transferred to a photoelectron which is ejected from the atom. • In a Compton scattering, the photon transfers only some of its energy to a recoil electron. As a result the photon is ‘scattered’ or deflected off in a new direction with a lower energy. In both interactions an orbital electron is ejected. It will lose its energy locally in the same way as a beta particle. The Compton-scattered photon will subsequently undergo either another Compton scattering or a photoelectric absorption. Unlike alphas and betas, gamma rays and X-rays do not have definite ranges, but are attenuated on passing through matter. A certain fraction of the photons will pass through, the fraction depending on the material, its thickness, and the energy of the photons.

Figure 7.6 A fast-moving beta particle is deflected by a nucleus, producing a Bremsstrahlung photon. (From Hart G, Armes F 1992 Medical physics for advanced level. Simon & Schuster Education, London.)

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Figure 7.7 (A) Photoelectric absorption. (B) Compton scattering. (From Hart G, Armes F 1992 Medical physics for advanced level. Simon & Schuster Education, London.)

THE GAMMA CAMERA A gamma camera (Fig. 7.8) is a system of one or more detectors, linked to an acquisition computer. In front of each detector is a collimator, which projects an image of the radioactive distribution onto the crystal. These images are ‘shadow pictures’, analogous to planar X-rays. The crystal converts the incoming gamma-ray signal to a light signal, which is detected by an array of photomultiplier tubes (PMTs).The PMT signals are processed to form the image which is stored on an acquisition computer. Images are downloaded to a processing computer. Depending on the clinical information required, this computer may, for example, track changes with time, or create a tomographic reconstruction similar to CT or MRI.

Detector and acquisition computer The crystal is a single crystal of sodium iodide. It is typically large and rectangular, perhaps 40 × 55 cm, and between 9.5 and 19 mm thick. (Older gamma cameras tended to have circular crystals.) When a gamma photon enters the crystal, electrons are ejected from their atomic orbits in photoelectric or

Compton interactions (see above). The ejected electrons collide with nearby atoms in the crystal and energize them (see above). The atoms of the crystal release their excess energy by scintillation, giving off a characteristic blue light. In most cases all the energy of the incoming gamma photon is converted to scintillation photons3. Attached to the back of the crystal is an array of perhaps 90 or so PMTs. These extremely sensitive light detectors are coupled to the crystal with special grease to give maximum transmission of light from the crystal. Each PMT consists of an evacuated glass envelope, containing a series of electrodes. A scintillation photon striking the first electrode causes emission of an electron. This electron is accelerated towards the next electrode, which emits two or three electrons for every one striking it. This amplification process is repeated at each electrode in the cascade, until at the last electrode the signal may be amplified by a factor of about 1 million. Image formation is accomplished by combining the signals from several PMTs. In the crystal, the scintillation photons resulting from one gamma photon are given off in all directions, and are picked up by the nearby PMTs. From the relative size of signals from all the PMTs, the detector determines the position of the gamma photon’s interaction in the crystal to within a few millimetres, even though each PMT is several centimetres across. It is important to exclude from the image radiation that has been scattered within the patient or gamma camera before interacting with the crystal. Only those gamma rays that trace back to the distribution of radioactive material in the patient are used for image formation. Exclusion of scattered photons is made possible because the scattered rays have reduced energy (see above, Compton scattering). Figure 7.9 shows a typical energy spectrum for 99mTc displaying the number of events detected by the gamma camera across a range of photon energies. The tall peak, known as the photopeak, represents the unscattered gammas. These have their original energy, which is 140 keV in the case of 99m Tc. The scattered gammas appear as the broad distribution to the left of the photopeak. The detector identifies the energy of each incoming gamma photon by summing the signals from all the PMTs. A suitable energy window is set at the photopeak as shown in Figure 7.9. Each gamma within the energy window will register one count, and is saved in the pixel corresponding to the X and Y coordinates that denote the position of the photon interaction within the crystal. As most of the scattered gammas have energies outside the window, they are excluded from the image.

Collimators Up to this point, only the physical steps involved in the detection of a gamma-ray interaction within the crystal have been discussed. To form a meaningful image of the distribution of radioactive material, a collimator is placed in front of the crystal to restrict the direction of gamma rays entering the crystal. Without a collimator, gamma rays arising from radioactivity in all parts of the body can strike all parts of the crystal and so no meaningful image is formed.

CHAPTER 7

HV supply

• RADIONUCLIDE IMAGING: GENERAL PRINCIPLES

Computer memory Digital circuitry A/D converters Amplifiers Photomultiplier tubes Crystal Collimator

B Figure 7.8 (A) A gamma camera. This model has two detectors mounted on the circular gantry ring to the left of the picture. (Image supplied by Siemens Medical Solutions.) (B) Components of a modern ‘digital’ gamma camera with one analogue-to-digital (A/D) converter for each photomultiplier tube.

The usual form of collimator is a block of lead with an array of holes orientated perpendicular to the face of the crystal. Gamma rays that originate directly in front of a hole can pass through to the crystal, but rays arriving at an angle are stopped by the lead septa between the holes. The performance of the collimator is determined by the number of holes, the diameter and length of each hole and the septal thickness. Collimators with differing performance can be chosen depending on the energy of the gamma rays being detected and whether high resolution or high sensitivity is needed. Because their septa need to be thicker, the resolution of collimators designed for high-energy isotopes is generally poorer than that of those used for low-energy isotopes. Several types of magnifying and minifying collimators have been used in the past.These have included converging, diverging and fan-beam collimators, all of which employ nonparallel arrangements of holes.The pinhole collimator is a hollow metal cone that fits over a circular detector face, and has a single hole at its point. As all of these collimators introduce distortions, the scale of an anatomical feature on the image would depend on its distance from the detector.These types of collimator have largely fallen into disuse.

Count rate

Energy window

In a modern gamma camera, magnification is controlled by choice of a zoom factor on the acquisition computer. Typically this may range from 1.1 to 8.0. The zoomed image is magnified by the zoom factor, but covers only a portion of the detector’s field of view. The images are easier to interpret because zoom magnification is not affected by distance from the detector.

Temporal resolution The speed at which a gamma camera can count incident photons is limited by the properties of the crystal and the electronics of the camera.The detector can only process one event at a time. After a photon is absorbed in the crystal it will scintillate for approximately 230 nanoseconds4. If a second photon hits the crystal during this ‘integration time’, the two photons will not be detected as separate events, but rather as one with excess energy and a misplaced position. The capacity of the camera to record many scintillations rapidly is also a function of the time for the electronics to process each event. The overall ‘dead time’ of a detector is that time during which the system cannot respond to a second incident photon. Most types of imaging camera operate in a mode with a dead time of 1–2 ms.The count rate performance is described by the count rate at which there is a 20% loss of recorded scintillations due to dead time, which may then typically be around 75 000 counts per second. Most imaging is performed at count rates considerably less than this. Sometimes cameras have a high count rate mode with a slightly reduced integration time and hence a reduced dead time which can be used for quantitative dynamic studies with high count rates (e.g. radionuclide angiocardiography).

Techniques of image acquisition and processing 0

140 Gamma energy keV

Figure 7.9

99m

Tc spectrum and energy window.

A major strength of nuclear medicine imaging is its ability to show and quantify physiological function.This capability is supported by a range of acquisition protocols and image processing software. This section introduces the main types of study.

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• IMAGING TECHNIQUES AND GENERAL ISSUES

Static image This is the basic image; other acquisitions consist of a series of such images. A matrix size is chosen, usually 64 × 64, 128 × 128, 256 × 256, or 512 × 512. This determines the number of pixels there will be in the image. Each gamma ray that is accepted by the camera is saved as a count in the pixel corresponding to its location. A typical image may contain between 100 000 and 1 million counts, and take up to several minutes to acquire. Dynamic imaging consists of a sequence of basic images or frames acquired over time. The frame rate is chosen according to the rate of change of the radiodistribution, and may be as rapid as 40 frames per second. The number of counts acquired in a short time can be increased by the use of a high-sensitivity collimator which has larger holes and therefore poorer resolution than a general purpose collimator. Processing can generate time–activity curves of selected organs (e.g. the kidneys in a renogram). Parametric or functional images display particular aspects of a physiological process in each individual pixel (picture element) making up the image; for example, in a renogram, the colour scale might show time-to-peak activity in different regions across the kidney. (For an introduction to parametric images, see Chandler & Thomson 2003, in Suggested Further Reading.) Whole body imaging Most modern cameras allow whole body imaging in which either the camera or the patient is moved slowly longitudinally during acquisition. The result is an extended image that can include the whole patient. With a dual-headed camera it is possible to acquire anterior and posterior whole body images simultaneously and to allow, for example, a whole body survey for bone metastases to be acquired in less than 20 min. For quantitative measurements, a series of static images is preferable. Gated imaging It is possible to use a physiological signal from a patient to ‘gate’ images as they are acquired on a gamma camera. The commonest application is to image the blood pool or the myocardium gated with an electrocardiogram (ECG) signal. The result is a series of images of the heart at points throughout the cardiac cycle that is built up over several hundred cardiac cycles. Dual isotope imaging involves two radionuclides which are imaged simultaneously by setting two different energy windows on a detector; for example, a lung study using 81mKr gas for imaging ventilation and 99mTc microspheres for perfusion. The 81mKr energy window at 190 keV is well separated from the 140 keV 99mTc window. The camera acquires two separate images. Care needs to be taken with technique to avoid artefacts due to activity from one or other isotope being erroneously detected in the wrong window. SPECT acquisition (single photon emission computed tomography) Most current gamma cameras can be used for both planar and tomographic imaging. Tomographic images are obtained by rotating the detectors around the patient.

A series of projection images is acquired from many angular views around the patient from which the tomographic slices are reconstructed using computer processing. The sensitivity of a SPECT system is proportional to the number of its detectors which typically range from one to three. Gated SPECT acquisition is commonly used in myocardial perfusion imaging. SPECT reconstruction produces a set of axial slices, perpendicular to the camera’s rotation axis (and the patient’s spine), and can be performed using either filtered back-projection (FBP) or iterative algorithms such as ‘OSEM’. Mathematical filters reduce noise and enhance the images. In essence, the counts in each pixel of a slice correspond to the radioactive concentration at that point within the patient, though this is not strictly true because of the various ‘camera limitations’ discussed in the next section. These data can also be rearranged into coronal and sagittal slices, with all three sets displayed simultaneously. New slices can be defined to align with anatomical axes (e.g. of the heart or brain). In reconstruction it is essential to use the same algorithm and filter parameters for a particular investigation, and not to alter these for individual patients5.

Gamma camera limitations The effects described here are not faults, but are inherent to the workings of gamma camera imaging. The resultant distortions of the image need to be understood and allowed for. • The spatial resolution of a gamma camera is limited primarily by the size of the collimator holes and statistical variations in the distribution of light photons detected by the PMTs. A small radioactive source will appear as a broad peak. Manufacturers generally specify resolution in terms of full width half maximum (FWHM). This parameter is a measure of the spread of the image of a very small source (Fig 7.10A). A typical value near the surface of a high-resolution collimator is 6–8 mm, and, for a high-sensitivity or a high-energy collimator, 10–12 mm. Resolution is degraded when the source is at a distance from the collimator (Fig 7.10B). • A consequence of limited resolution is the partial volume effect. When an object is smaller than the resolution of the camera, the measured signal is reduced in amplitude. For example, a tumour that is smaller than approximately 2 × FWHM will appear larger than its true size and fainter than its true uptake of radionuclide. • Attenuation is the reduction in intensity of a beam of gamma rays, due to absorption and scatter. Patient tissue between the radioactive material and the detector can very significantly reduce the radiation reaching the detector. For 99m Tc, a soft tissue thickness of approximately 4.5 cm will reduce the intensity to one half 6. To correct for attenuation in SPECT, some cameras offer attenuation correction. The correction algorithm needs, as input, a map of the distribution of scattering material within the patient. This information may be obtained from previously acquired CT or by using a hybrid SPECT-CT system. Several conventional gamma cameras use a radioactive line source, usually gadolinium

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• RADIONUCLIDE IMAGING: GENERAL PRINCIPLES

Crystal Collimator Hole Lead

Maximum height

Source A

Full width half maximum

A

B

Source B

Figure 7.10 (A) Full width half maximum is a measure of the width of a peak, taken at half its maximum height. (B) The effect of distance on resolution. Gamma rays from source B will pass through a larger number of holes in the collimator, and activate a larger area of the crystal. Resolution is therefore poorer for the more distant source B.

(153Gd), to obtain transmission data. As with all methods of attenuation correction, national guidelines emphasize the need for proper validation and quality control, and for comparison of the images with the noncorrected reconstruction5,7. • For distant sources, the number of counts acquired is not reduced, but image quality is compromised. This can be understood from Figure 7.10B. In both cases, the cone of rays that passes through the collimator has the same angle, which means that the same number of rays is transmitted. However, those from the more distant source are spread over a larger zone of the crystal. This reduces both the sharpness (spatial resolution) and the contrast of the image. • Scattered radiation (Compton scattering) contributes random noise that tends to obscure the signal. Most of the scattered radiation is eliminated by the energy window. • Count rate limitation. Because of the finite time the detector takes to process each gamma photon, acquisitions carried out at high count rates should not be used for quantitative estimations of absolute or relative organ uptake.

PRINCIPLES OF POSITRON EMISSION TOMOGRAPHY PET uses the coincidence detection of paired 511 keV annihilation photons from positron-emitting radionuclides such as 11 C, 13N, 15O and 18F. A PET camera (Fig. 7.11A) contains several rings of block detectors as shown in Figure 7.11B. The principle of PET is that an event is only registered if two detectors record a photon interaction at almost exactly the same time. Because the photons are emitted 180 degrees apart, it can be inferred that the annihilation occurred somewhere along a line of response (LOR) connecting the two detectors. This inherent spatial information makes it possible to acquire and reconstruct a tomographic image without the need for collimators. PET can therefore potentially offer much higher sensitivity and better spatial resolution than SPECT. Although sodium iodide detectors similar to those used in gamma cameras can be employed, the physical properties are less than ideal for the high-energy photons involved in PET. PET detectors typically use crystals such as bismuth germinate (BGO),

B Figure 7.11 (A) A combined PET/CT system. (Image supplied by Siemens Medical Solutions.) (B) Coincidence detection of paired photons in a PET imaging system. (From Sharp P F, Gemmell H G, Smith F W 1998 Practical nuclear medicine, 2nd edn. Oxford University Press, Oxford.)

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lutetium orthosilicate (LSO) and gadolinium orthosilicate (GSO). PET detectors surround the patient and acquire data from all angles simultaneously, which allows dynamic acquisitions with frame times as short as 1 s 8. Gated acquisitions are also possible with PET. In 2D mode, each ring of detectors works independently, and only registers annihilations that occur within its own plane. Lead septa between the rings absorb any oblique photons and prevent them from registering. Using algorithms similar to SPECT, the data from each ring are reconstructed to form an axial image.The stack of axial images forms the complete tomographic map. To increase sensitivity, the septa may be designed so that each axial is formed from a set of adjacent detector rings; this causes some loss of spatial resolution in the axial direction. In 3D mode the septa are removed, and sensitivity is greatly increased by acquiring coincidence events from any pair of detectors. Special 3D algorithms are used for image reconstruction. PET images are degraded by scatter, attenuation, random coincidences, dead time and noise. Attenuation correction is particularly important for PET because absorption or scatter of either of the pair of photons will result in loss of detection of the annihilation event. Spatial resolution of PET is limited by the length of the positron path (a few millimetres), noncolinearity of annihilation photons (about quarter of a degree) and the size of the detector blocks. Gamma cameras have been developed that include coincident detection for use in PET imaging9. Their use is limited, partly because gamma cameras are not optimized for PET, so their performance is inferior. Another reason is that it is not economic for a department to conduct occasional PET studies, because the high radiation doses necessitate substantial expense and reorganization of a department. As well as the physical advantages, PET also benefits from the types of radiopharmaceutical that can be used. Most physiological molecules are made of carbon, nitrogen and oxygen, enabling them to be labelled with positron emitters 11C, 13N and 15O. Cellular uptake of the glucose analogue 18F-FDG reflects the rate of transmembranous exchange of glucose. PET provides the clinician and researcher with a unique tool to study and quantify physiological and pathological functions of human tissues and organs. Examples of how PET radiopharmaceuticals can be used to measure glucose and oxygen utilization in tumours, or to quantify cerebral or myocardial blood flow, are given in subsequent chapters.

The CT image can also be used for attenuation correction of the nuclear medicine image. By using a single gantry, errors due to movement of the patient between the CT and nuclear medicine acquisitions are significantly reduced. However, respiratory motion can lead to a mismatch of organs near the diaphragm, particularly if patients hold their breath for the CT acquisition, but breathe naturally during the nuclear medicine image (Fig. 7.12). Motion correction by respiratory gating of the nuclear medicine acquisition may lead to improvements10. PET-CT has now replaced dedicated PET for the systems currently in production. The early systems that used CT merely for attenuation correction and localization of abnormalities are now being replaced by diagnostic-quality CT. A greater throughput of patients is achievable with PET-CT due to the shorter acquisition time of the CT as compared to the transmission images performed using dedicated PET systems. This results in an increasing need for clinicians proficient in interpreting both nuclear medicine and CT11.

QUANTIFICATION AND IMAGE PROCESSING Processing computers provide a range of software to extract quantitative data reflecting clinically relevant physiological parameters from data acquired either with gamma cameras or PET. Dynamic and gated acquisitions generate particularly large quantities of data. Basic processing functions are as follows: • Measurement of activity within regions of interest (ROIs) outlining specific organs or areas. ROIs may be drawn manually or determined by various computer algorithms. • Background subtraction, to compensate for activity in overlying tissues, when estimating the activity within an

HYBRID SYSTEMS: SPECT-CT AND PET-CT The development of SPECT-CT and PET-CT systems allows the superimposition of nuclear medicine and CT images, a technique known as functional–anatomical mapping (Fig. 7.5). In a range of clinical scenarios, interpretation of images obtained using these hybrid devices has proved to be more diagnostically accurate than evaluation of the two sets of images separately. This combined imaging approach also allows functional data to be incorporated into the selection of radiation fields during radiotherapy planning.

Figure 7.12 PET-CT study in which the CT data (right) has been used for attenuation correction of the PET image (left). A crescent of lung just cranial to the left hemidiaphragm (arrow) demonstrates reduced activity—an artefact due to differences in position of the left hemidiaphragm between acquisition of CT and PET data.

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• •

organ. Various algorithms are available, or the background may be defined in terms of a background ROI. Consistency in the estimation of background can be extremely difficult to attain, and has led to significant differences in the numerical results from different centres. Generation of time–activity curves (TACs) for ROIs applied to dynamic image series. Smoothing and filtering, to reduce statistical noise, both spatial and temporal. Curve fitting, to reduce complex TACs to a few parameters that can be compared against normal values (e.g. gastric emptying half-times). Mathematical modelling to quantify discrete physiological processes. Statistical analysis to highlight image regions that vary significantly from a reference set of normal images.

For a discussion of these functions, and an introduction to more sophisticated methods, see Lawson (1999), in Suggested Further Reading. When the counts in an image are low, random variations become significant. These variations result in a grainy appearance that can be removed by smoothing algorithms, often automatically included in processing protocols. Low counts are a particular problem in applications involving subtraction of tissue background from counts within an organ. The background region chosen will be a compromise between a small ROI with a large percentage error and a larger ROI that includes areas less representative of background behind the organ. Measurements of activity within a ROI are often compared to activity within a reference region (e.g. target-to-background ratio). A more precise quantitative measure of uptake of radiotracer in tissue is the standard uptake value (SUV): SUV =

Activity per unit mass of tissue Injected activity per unit body mass

Estimation of SUVs may be affected by choice of ROI, time since injection of the radiopharmaceutical and by some of the camera limitations discussed above. (For a discussion of these factors, and of how uptake may be affected by clinical considerations, see Thie 2004, in Suggested Further Reading.) For tomographic images, accurate quantification is more readily achieved with PET than SPECT. SUV measurements are more commonly used in PET. Image registration is the geometric alignment of two images. There are two fundamentally different applications: • In dealing with images from one patient, registration may align different examinations (e.g. CT to nuclear medicine) or may be used to correct for patient motion. Software may operate automatically, or with shifts and rotations determined by the operator. • A patient’s images may also be compared against those of another individual by stereotactic normalization or warping. The images are deformed to make them the same size and shape. This technique has found application in brain

• RADIONUCLIDE IMAGING: GENERAL PRINCIPLES

imaging, where an individual’s image is compared to a ‘normal’ image, a composite formed by warping images obtained from several normal subjects. The technique is recommended in dopamine transporter SPET imaging, as outlined in the EANM guidelines12. Current research is assessing the usefulness of different methods13.

IMAGE INTERPRETATION The processes of acquiring radionuclide images, along with the focus on depiction on function rather than structure, result in the need for an approach to image interpretation that differs in some aspects from other imaging techniques. Image annotation is important for radionuclide images which may have few anatomical landmarks to help orientation. Unlike conventional radiography, planar images obtained with a gamma camera may differ considerably depending on whether they have been acquired from the anterior or posterior aspect of the patient (Fig. 7.13). This difference is due to attenuation and scatter of radiation emitted by structures further from the camera.The resulting images are displayed in the orientation in which they are acquired. Thus, a posterior view will be orientated with the patient’s right-hand side on the right side of the image. Tomographic views from SPECT and PET are displayed in accordance with the same conventions used for other cross-sectional imaging techniques. Image artefacts may be caused by the gamma camera being out of adjustment or by inappropriate selection of imaging conditions by the operator. Any geometrical or nonanatomical shape on an image is likely to be a camera artefact. The detection of artefacts is improved by: (A) being familiar with their typical appearance; (B) adjustment of the grey scale; and (C) varying one’s distance from the computer screen. The examples below are restricted to artefacts that occur in patient images, i.e. those that are of concern to clinicians. Technical staff will be concerned also with the artefacts of quality-control images. Busemann Sokole’s ‘Atlas’ (see Suggested Further Reading) illustrates and explains a comprehensive range of gamma camera artefacts. • A common artefact of older gamma cameras is a regular pattern of dark or light spots that mimic the arrangement of the PMTs behind the crystal.The example shown in Figure 7.14 is a faulty uniformity flood test, acquired using a uniform sheet of radioactive material known as a flood phantom. With a patient on the camera, the image would show the organs containing radioactivity, but with the pattern of spots superimposed. This pattern can have various causes, for instance a shift in the high voltage supply to the PMTs, a fault in one of the camera’s correction matrices, or selection of the wrong energy window for the radionuclide being scanned. • A single cold spot of similar size would indicate a single faulty PMT. • A straight line, or lines, across the field of view may indicate a fault in the circuits that determine the spatial coordinates.

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Anterior

Posterior

Anterior

Tissue phase

A

Posterior Delay

B

C

D

Figure 7.13 Bone scintigram in a child with osteomyelitis of the distal left femur acquired using 99mTc methylene diphosphonate. The distribution of the radiopharmaceutical changes between the early tissue phase images (A, B) and the delayed images (C, D). The bones are visualized more clearly on later images. The epiphyseal plates show increased uptake, a normal finding in children. There are differences between corresponding anterior and posterior images and right–left (R–L) annotation on anterior and posterior images.

• Star-like rays emanating from hot spots can be produced by septal penetration if a low-energy collimator is used to image high-energy gamma rays. • Irregular lines or sharp-edged patches usually mean the crystal has cracked, or humidity may have leaked in. The

NaI crystal is hygroscopic and very delicate. Its front and sides are protected by thin aluminium sheeting, and the back is sealed to the PMTs with special light-transmitting grease. • If a SPECT reconstruction shows a circular ring or arc, this usually means that detector uniformity is at fault, and may need to be recalibrated. A ray-like artefact may be produced when a region of intense radioactivity lies within the reconstructed volume. The detection and understanding of artefacts is an essential aspect of quality assurance. Firstly, artefacts can distort images and therefore an understanding of them is essential to the interpretation of patients’ images. Secondly, certain artefacts indicate camera faults or poor working practice, and these will need to be corrected.

Figure 7.14 phantom.

Faulty uniformity flood test, acquired using a flood

Knowledge of the normal biodistribution of the injected radiopharmaceutical is essential for accurate interpretation, just as knowledge of normal anatomy is required for other imaging techniques. It is also important to be aware of how the distribution changes with time and with the age of the patient. For example, skeletal scintigraphy in a child will show normal epiphyseal activity related to growth, an appearance that is lost in the mature skeleton (Fig. 7.13). An unusual distribution of radioactivity may not necessarily indicate disease. Some normal tissues will show variable uptake of a radiopharmaceutical; for example, breast uptake of

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Ga-gallium citrate, or uptake of 18F-FDG in brown fat within the neck and superior mediastinum. The radiochemical purity may be poor; for example, an excess of unbound technetium, not combined in the intended form, would accumulate in the thyroid. The possibility that the patient may have been given the wrong radiopharmaceutical also needs to be considered. Radiodistribution is affected by patient preparation. Factors such as hydration, fasting and the patient’s medication are important in certain studies. Radioactive contamination will show up on images. Radiopharmaceuticals excreted in the urine may contaminate clothes or skin following micturition. Other examples include extravasation of radiopharmaceutical at a site of venous injection, a contaminated injection swab in a patient’s pocket and contamination transferred from the fingers to the patient’s skin or the camera. Motion of the patient during image acquisition will lead to loss of resolution and may produce other artefacts.

Abnormal biodistribution due to disease can comprise either increased or decreased uptake of radiopharmaceutical, depending on whether the pathological process concerned results in loss or upregulation of the mechanism of uptake. Some radiopharmaceuticals have highly specific uptake mechanisms and failure to demonstrate a site of disease may be due to absence of the uptake mechanism in that particular case. Such instances may not represent diagnostic failures because demonstrating the lack of uptake can have important implications for subsequent clinical management. For example, a tumour that fails to accumulate the somatostatin analogue 111In-octreotide will not be suitable for subsequent octreotide therapy. Abnormalities in nontarget organs can provide clinical information relevant to the disease under study or indicate a new concomitant diagnosis. For example, skeletal scintigraphy looking for metastatic cancer as a cause of back pain may reveal incidental renal obstruction or occasionally demonstrate uptake in additional hepatic metastases. Interpretation of images from hybrid devices requires skills that are additional to those involved in reporting each imaging investigation in isolation. Some image artefacts are specific to hybrid imaging devices (see above). Furthermore, the anatomical appearances of lesions can modify the interpretation of the radionuclide image data and vice-versa. The greater use of hybrid systems is likely to produce an increasing need for clinicians proficient in interpreting both nuclear medicine and CT11.

NONIMAGING APPLICATIONS