Clark's Essential Physics in Imaging for Radiographers (Clark's Companion Essential Guides) [2 ed.] 0367511983, 9780367511982

Table of contents :
Cover
Half Title
Title Page
Copyright Page
Contents
Preface
The authors
Preface to second edition
Chapter 1. Overview of image production
Introduction
General principles
X-ray beam characteristics
Scattered radiation
Field Size
Geometry of Image Production
Magnification
Unsharpness
Bucky/grids
Resolution/definition
X-ray detectors
Ionisation
Display System and Viewing Conditions
Radiation dose
Practitioner's Skill and Perception
MCQs
Chapter 2. Mathematics for medical imaging
Introduction
Basic Mathematics
Exposure calculations
International System of Units
Measurement prefixes (powers)
Multiplication and division of powers
Logarithms (LOGS)
Graphs
Line focus principle
Similar Triangles
Inverse square law
Statistics
MCQs
Chapter 3. Physics for medical imaging
Introduction
The Atom
Atomic structure
Atomic number
Mass number
Electrons and electron orbitals
Binding energy
Atomic symbols and the periodic table
Atomic balance
Ions
Isotopes
Elements
Compounds
Radioactivity
Principles of radioactive decay
Alpha (α) emission
Beta particle (β) emission
Gamma (γ) emission
Gamma emission and X-rays
Examples of radionuclides used extensively in medical imaging are:
γ - emission and X-rays
Penetrating power of the emissions
Force, Work, Energy and Power
Heat
Transfer of heat
Conduction
Convection
Radiation
Waves
Sound
Magnetism
Electricity and Electric Charge
Electrical circuit
Electromagnetic Radiation
MCQs
chapter 4. X-rays, X-ray tube and X-ray Circuit
Introduction
X-rays
X-ray Tube
Parts of the X-ray tube
Insert (envelope)
Anode assembly
Filament circuit
Cathode assembly (filament)
Focusing cup
X-ray Circuit
The interaction of high-energy electrons with matter
Interactions between Incoming Electrons and Outer-Shell Electrons in tungsten
Interactions Producing Heat
Interactions producing X-rays
Interactions between incoming electrons and inner-shell electrons in tungsten (characteristic X-ray production)
Interactions between incoming electrons and the nucleus of the atom (Bremsstrahlung X-ray production)
X-ray spectra and factors affecting the quality and intensity of the X-ray beam
Impact of changing the mA
Impact of Changing the kV
The Impact of Filtration on the X-ray Beam
MCQs
Chapter 5. X-ray Interactions in Matter
Introduction
Interactions of X-rays in Matter
Attenuation
Linear attenuation coefficient (µ)
The Processes of Attenuation in Diagnostic Radiography
Elastic scatter
Pair production
Photoelectric absorption
Photoelectric absorption coefficients
Compton scatter
Compton scatter coefficients
MCQs
Chapter 6. Principles of Radiation Detection and Image Formation
Introduction
Desirable Characteristics of Radiation Detectors
Detective Quantum Efficiency
Ionisation chambers
Ionisation Chambers Used for Automatic Exposure Control Circuits
Scintillation crystals/photocathode multiplier
Scintillation crystal/photocathode X-ray image intensifier
Scintillation crystals/silicon photodiode multiplier
Large Field Detectors (Overview)
Indirect, Direct, Computed and Digital Radiography
Computed radiography in detail
CR plate construction
Production of the latent image
Processing or reading the latent CR impression
Indirect digital radiography technology in detail
Indirect digital radiography using thin film transistor technology
Active matrix array in detail
Indirect digital radiography using charged coupled device technology
Charged coupled device coupling via optical fibre
Slot scan chest radiography
Charged coupled devices optically coupled by a mirror and high quality lens
Direct digital radiography
Digital Fluoroscopic Systems
Image intensifier linked to charged coupled device
Fluoroscopic flat panel detectors
Solid-state X-ray image intensifier
Reference
MCQs
Chapter 7. Image Quality
Introduction
Geometry of Imaging
Magnification and distortion
Signal-to-noise ratio
Unsharpness
Movement unsharpness
Resolution of the imaging system
Spatial resolution
Measurement of unsharpness in an image
Viewing digital images
Brightness and contrast
Effect of scatter on contrast
MCQs
Chapter 8. Radiation Dose and Exposure Indicators
Introduction
Radiation Dose
Detection and measurement of radiation
Ionisation of air
Exposure
SI Units
Absorbed dose
Equivalent dose
Effective dose
Lumbar spine
Which is equal to 2.1 mSv
CT dose and DRL's
Linear energy transfer and relative biological effectiveness
Quality factor for radiation
Radiation monitors and personal monitoring
Optically stimulated luminescent dosimeter (OSLD)
Thermoluminescent Dosimeters (TLD)
Exposure Indicators
MCQs
Chapter 9. Image Display And Manipulation In Medical Imaging
Introduction
Image Production Pathway
Raw data image matrix
Display of reconstructed image matrix
Image quality and matrix size
Image Interpolation
Image Manipulation
Manufacturer-defined manipulation tools
Operator-defined manipulation tools
Windowing
Zooming and enlarging
Interpolated zoom
Reconstructed zoom
Geometric zoom
Noise Reduction by Background Subtraction
Simple Edge Enhancement
Spatial Domain Filtering for Smoothing and Sharpening
Noise Reduction by 'Low-Pass Spatial Filtering'
Image Sharpening and Edge Enhancement 'High-Pass Spatial Filtering'
Standards
MCQs
Chapter 10. Computed tomography
Introduction
General Principles of Computed Tomography
Defining areas within the body
Hounsfield scale
Creating raw data attenuation values
Creation of Data in the xy Dimension
Scan field of view (SOV)
Image matrix
Pixel size
Number of projections
Simple back projection
Image reconstruction using filtered back projection (FBP) or iterative filtered back projection (IFBP)
Creating Data in the z Dimension
z axis data from directly scanned sequential slices
z axis data from spiral acquisition
Multi-spiral scanning
How does all this relate to scan considerations / protocols
Spatial Resolution (SR)
Slice thickness (controls z axis SR)
Matrix size (controls xy axis SR)
Number of projections
High resolution filter convolution
Contrast resolution (CR)
Slice thickness and matrix size
mAs
The Link Between Slice Thickness, Dose, Photon Density and Signal Value
The principles still apply when we consider multi-slice scanners
Temporal Resolution and Control of Movement
Artefacts and Other Scan Considerations
Beam hardening
Partial volume artefact
Cone beam artefact
MCQs
Chapter 11. Radiation protection and safety
Introduction
Legislation
IRR17
IR(ME)R 2017
Justification
Diagnostic reference levels (DRL)
Optimisation
Training
Local rules and systems of work
Responsibilities
Radiation protection of staff
Radiation risk assessment
Dose limits
Physical, chemical and biological effects of ionising radiation
Cancer Effects (Stochastic)
Deterministic (tissue reactions) effects
Standard Operating Procedure for Women of Reproductive Capacity
Pregnancy enquiry procedure
Minimising radiation risks in pregnancy
Inadvertent fetal exposures
Reporting of Radiation Incidents
Significant accidental or unintended exposures (SAUE)
Under-exposures
Over-exposure
MCQs
Chapter 12. Benefit-Risk
Introduction
Benefit-Risk Analysis
Benefits of Examinations Using X-rays and Gamma Photons
Risks from X- and γ-radiation"+"-Radiation";?>
Possible Effects of Ionising Radiations on Human Cells
MCQs
Answers to MCQ's
Chapter 1
Chapter 2
Chapter 3
Chapter 4
Chapter 5
Chapter 6
Chapter 7
Chapter 8
Chapter 9
Chapter 10
Chapter 11
Chapter 12
Chapter formulas
Chapter 1
Chapter 2
Chapter 3
Chapter 5
Chapter 7
Chapter 10
Index

Citation preview

ESSENTIAL CLARK'S

PHYSICS IN IMAGING FOR

RADIOGRAPHERS

ESSENTIAL CLARK'S

PHYSICS IN IMAGING FOR

RADIOGRAPHERS Second edition Ken Holmes Senior Lecturer School of Medical Imaging Sciences University of Cumbria Carlisle and Lancaster, UK Marcus Elkington Senior Lecturer in Diagnostic Imaging College of Health, Wellbeing and Life Science Sheffield Hallam University Sheffield, UK Phil Harris Senior Lecturer in Medical Imaging University of Cumbria Carlisle and Lancaster, UK Series Editor: A. Stewart Whitley International Society of Radiological Technologists

First edition published 2021 by CRC Press 6000 Broken Sound Parkway NW, Suite 300, Boca Raton, FL 33487-2742 and by CRC Press 2 Park Square, Milton Park, Abingdon, Oxon, OX14 4RN © 2021 Taylor & Francis Group, LLC CRC Press is an imprint of Taylor & Francis Group, LLC The right of Ken Holmes, Marcus Elkington and Phil Harris to be identified as the author of the editorial material, and of the authors for their individual chapters, has been asserted in accordance with sections 77 and 78 of the Copyright, Designs and Patents Act 1988. This book contains information obtained from authentic and highly regarded sources. While all reasonable efforts have been made to publish reliable data and information, neither the author[s] nor the publisher can accept any legal responsibility or liability for any errors or omissions that may be made. The publishers wish to make clear that any views or opinions expressed in this book by individual editors, authors or contributors are personal to them and do not necessarily reflect the views/opinions of the publishers. The information or guidance contained in this book is intended for use by medical, scientific or health-care professionals and is provided strictly as a supplement to the medical or other professional’s own judgement, their knowledge of the patient’s medical history, relevant manufacturer’s instructions and the appropriate best practice guidelines. Because of the rapid advances in medical science, any information or advice on dosages, procedures or diagnoses should be independently verified. The reader is strongly urged to consult the relevant national drug formulary and the drug companies’ and device or material manufacturers’ printed instructions, and their websites, before administering or utilizing any of the drugs, devices or materials mentioned in this book. This book does not indicate whether a particular treatment is appropriate or suitable for a particular individual. Ultimately it is the sole responsibility of the medical professional to make his or her own professional judgements, so as to advise and treat patients appropriately. The authors and publishers have also attempted to trace the copyright holders of all material reproduced in this publication and apologize to copyright holders if permission to publish in this form has not been obtained. If any copyright material has not been acknowledged please write and let us know so we may rectify in any future reprint. Except as permitted under U.S. Copyright Law, no part of this book may be reprinted, reproduced, transmitted, or utilized in any form by any electronic, mechanical, or other means, now known or hereafter invented, including photocopying, microfilming, and recording, or in any information storage or retrieval system, without written permission from the publishers. For permission to photocopy or use material electronically from this work, access www.copyright.com or contact the Copyright Clearance Center, Inc. (CCC), 222 Rosewood Drive, Danvers, MA 01923, 978-750-8400. For works that are not available on CCC please contact [email protected] Trademark notice: Product or corporate names may be trademarks or registered trademarks and are used only for identification and explanation without intent to infringe. ISBN: 978-0-367-51198-2 (hbk) ISBN: 978-0-367-51197-5 (pbk) ISBN: 978-1-003-05279-1 (ebk) Typeset in Berling by MPS Limited, Dehradun

CONTENTS Preface The Authors Preface to Second Edition

Chapter 1 Overview of Image Production Introduction General principles X-ray beam characteristics Scattered radiation Field size Geometry of image production X-ray detectors Ionisation Display system and viewing conditions Radiation dose Practitioner‘s skill and perception MCQs

ix xi xiii

1 1 1 2 3 3 4 9 9 10 11 11 12

Chapter 2 Mathematics for Medical Imaging 15 Introduction Basic mathematics International system of units Measurement prefixes (powers) Logarithms (logs) Line focus principle Similar triangles MCQs

15 15 17 17 19 20 21 25

Chapter 3 Physics for Medical Imaging

27

Introduction The atom

27 27

Radioactivity Force, work, energy and power Heat Waves Sound Magnetism Electricity and electric charge Electromagnetic radiation MCQs

39 39 41 42 42 43 44 46

Chapter 4 X-rays, X-ray Tube and X-ray Circuit

49

Introduction X-rays X-ray tube X-ray circuit The Interaction of high-energy electrons with matter Interactions between incoming electrons and outer-shell electrons in tungsten Interactions producing heat Interactions producing X-rays Interactions between incoming electrons and inner-shell electrons in tungsten (characteristic X-ray production) Interactions between incoming electrons and the nucleus of the Atom (Bremsstrahlung X-ray production) X-ray spectra and factors

34

49 50 50 55

56

56 57 58

58

62 v

affecting the quality and intensity of the X-ray beam Impact of changing the mA Impact of changing the kV The impact of filtration on the X-ray beam MCQs

Chapter 5 X-ray Interactions in Matter Introduction Interactions of X-rays in matter The processes of attenuation in diagnostic radiography Elastic scatter Pair production Photoelectric absorption Compton scatter MCQs

65 66 67 68 71

73 73 73 77 78 78 78 83 87

Chapter 6 Principles of Radiation Detection and Image Formation 89 Introduction Desirable characteristics of radiation detectors Detective quantum efficiency Ionisation chambers used for automatic exposure control circuits Large field detectors (overview) Indirect, direct, computed and digital radiography Digital fluoroscopic systems MCQs vi

89 90 90

93 99 100 118 121

Chapter 7 Image Quality Introduction Geometry of imaging Magnification and distortion Signal-to-noise ratio Unsharpness Movement unsharpness Resolution of the imaging system Spatial resolution Measurement of unsharpness in an image Viewing digital images Brightness and contrast Effect of scatter on contrast MCQs

Chapter 8 Radiation Dose and Exposure Indicators Introduction Radiation dose Exposure Radiation monitors and personal monitoring Exposure indicators MCQs

125 125 126 128 129 130 131 131 131 132 132 132 135 136

139 139 140 141 148 150 153

Chapter 9 Image Display and Manipulation in Medical Imaging 155 Introduction Image production pathway Image interpolation

155 156 159

Image manipulation Noise reduction by background subtraction Simple edge enhancement Spatial domain filtering for smoothing and sharpening Noise reduction by ‘low-pass spatial filtering’ Image sharpening and edge enhancement ‘high-pass spatial filtering’ Standards MCQs

159 163 165 165 167

169 171 173

Chapter 10 Computed Tomography 175 Introduction 175 General principles of computed tomography 176 Creating raw data attenuation values 179 Creation of data in the xy dimension 179 Creating data in the z dimension 190 How does all this relate to scan considerations/protocols 195 Spatial resolution (SR) 195 Contrast resolution (CR) 198 The link between slice thickness, dose, photon density and signal value 199 Temporal resolution and control of movement 203

Artefacts and other scan considerations MCQs

204 207

Chapter 11 Radiation Protection and Safety 209 Introduction 209 Legislation 210 Dose limits 215 Physical, chemical and biological effects of ionising radiation 215 Cancer effects (stochastic) 216 Standard operating procedure for women of reproductive capacity 218 Reporting of radiation incidents 222 MCQs 224

Chapter 12 BenefitRisk

227

Introduction Benefit-risk analysis Benefits of examinations using X-rays and gamma photons Risks from X- and γ-Radiation Possible effects of ionising radiations on human cells MCQs

227 227

Answers to MCQs Chapter formulas Index

235 241 245

229 230 232 233

vii

PREFACE The second edition of Clark’s Essential Physics in Imaging for Radiographers is a revised and updated version of the first edition. The authors have considered feedback from students, lecturers and practitioners using the book and have taken the opportunity in revision to improve and enhance existing text, update information where appropriate and include some additional material. These changes make the book more useful to students and will better inform clinical practice. The overall aims, however, remain the same – to give the reader an understanding of the basic physics underpinning diagnostic radiography, CT scanning and imaging science. It is essential that any practitioner working in an imaging department and using ionising radiation has a sound knowledge base. In order to understand the various factors affecting the production of diagnostic images, there is a requirement to demonstrate an understanding of the fundamental definitions of physics and how these principles may be applied to radiography. The idea is to provide a clear guide to the subject using clear text and diagrams/ photographs to support the text. Each chapter has clear ‘learning objectives’ and a series of multiple choice questions (MCQs) to test these learning outcomes. The content has been updated throughout and includes a new chapter on CT imaging, updated regulations on ionising Radiation Regulations 2017 (IRR17) and The Ionising Radiation (Medical Exposure) (Amendment) Regulations 2018. There is additional material on radioactivity, dosimetry, and imaging display and manipulation. The book has chapters that give an overview of image production, basic mathematics and physics relevant to medical imaging followed by detailed chapters on the physics relevant to producing diagnostic images using X-rays. Diagnostic radiography involves the safe use of ionising radiation and the production of diagnostic images. The process by which images are produced involves the conversion of energy from one form to another and this underpins the fundamentals of imaging. It requires knowledge of specialised equipment, such as the Xray tube, image detectors, computers and image processors. Understanding the fundamental principles of this equipment is the basic knowledge base of any practitioner.

ix

The final chapters review the factors affecting image quality and radiation dose: how correct exposures are indicated by the equipment together with how to manipulate images, data management and display parameters. Discussion of risk benefit, safety and radiation protection conclude the book as these are necessary requirements of health-care practitioners using ionising radiation.

x

THE AUTHORS Ken Holmes is now retired but was the programme leader for the BSc (Hons) Diagnostic Imaging at the University of Cumbria (formerly St Martins College). He is one of the co-authors of Clark’s Pocket Handbook for Radiographers and believes the time is right to develop a pocket physics book to use alongside the technique one. Ken started education as a clinical tutor and has worked at several higher education institutes in the UK and has taught physics and imaging principles for 30 years. Marcus Elkington is a senior lecturer in Diagnostic Imaging at Sheffield Hallam University. He has a great interest in imaging and physics related to diagnostic radiography and has been helping students understand physics for many years. Marcus feels there is a place for a pocket physics book produced in a student-friendly format that is aimed specifically at the core topic areas surrounding general radiographic imaging. Phil Harris was head of school at Medical Imaging Science at the University of Cumbria for many years and has always taken the greatest pleasure in passing on a basic understanding of radiation science to radiography students, many of whom enter into this subject with some considerable trepidation. This book has been written especially with these students in mind.

xi

PREFACE TO SECOND EDITION The second edition of Clark’s Essential Physics in Imaging for Radiographers is a corrected and updated version of the first edition. The authors have considered comments from students using the book and reviewers to correct mistakes, rewrite confusing text, update information and include some additional material. This should make the book more suited to students and better inform. practitioners The aims are still the same to give the reader an understanding of the basic physics underpinning diagnostic radiography, CT scanning and imaging science. It is essential that any practitioner working in an imaging department and using ionising radiation has a sound knowledge base. In order to understand the various factors affecting the production of diagnostic images, there is a requirement to demonstrate an understanding of the fundamental definitions of physics and how these principles may be applied to radiography. The idea is to provide a clear guide to the subject using clear text and diagrams/photographs to support the text. Each chapter has clear ‘learning objectives’ and a series of multiple choice questions (MCQs) to test these learning outcomes. The content has been updated throughout and includes a new chapter on CT imaging, updated regulations on Ionising Radiation Regulations 2017 (IRR17) and The Ionising Radiation (Medical Exposure) (Amendment) Regulations 2018. There is additional material on radioactivity, dosimetry, and imaging display and manipulation. The book has chapters that give an overview of image production, basic mathematics and physics relevant to medical imaging and then detailed chapters on the physics relevant to producing diagnostic images using X-rays. Diagnostic radiography involves the safe use of ionising radiation and the production of diagnostic images. The process by which images are produced involves the conversion of energy from one form to another and this underpins the fundamentals of imaging. It requires knowledge of specialised equipment, such as the X-ray tube, image detectors, computers and image processors. Understanding the fundamental principles of this equipment is the basic knowledge base of any practitioner.

xiii

The final chapters review the factors affecting image quality and radiation dose: how correct exposures are indicated by the equipment together with how to manipulate images, data management and display parameters. Discussion of benefit-risk, safety and radiation protection conclude the book as these are necessary requirements of health-care practitioners using ionising radiation.

xiv

CHAPTER 1 OVERVIEW OF IMAGE PRODUCTION INTRODUCTION The aim of this chapter is to give the practitioner an understanding of the basic principles of image production. It is essential that any prac­ titioner understands the principles involved in obtaining diagnostic images. Images must be produced using the lowest radiation dose consistent with diagnostic quality. The practitioner therefore needs to understand how to adjust the factors affecting dose and image quality. Learning objectives The student should be able to: ◾ Understand and explain the principles of producing images using X-radiation ◾ Explain the terms magnification, unsharpness, scatter, contrast, definition and resolution

GENERAL PRINCIPLES The objective of diagnostic imaging is to produce images of optimum quality for diagnosis and to aid in the management/treatment of the patient. There must be a valid reason for the examination. The proce­ dure must also affect the clinical management of the patient. The procedure should produce images with limited magnification, minimum 1

Overview of Image Production

unsharpness and a radiation dose as low as reasonably practicable (ALARP). The ideal set-up is to have the body part being imaged parallel to and in contact with the image detector. The X-ray beam should be at right angles to the detector and not angled across it as this produces a distorted image. However, there are situations where the patient or X-ray beam is angled to deliberately distort/elongate the image, e.g. 30° angled elongated scaphoid projection. There are a number of factors which affect the quality of the image and/or radiation dose to the patient when producing diagnostic images using X-rays. These are: ◾ The X-ray beam characteristics: ◾ Focal spot size ◾ Filtration of the beam ◾ Exposure factors ◾ Field size ◾ The production and management/reduction of scatter ◾ The geometry of image production. ◾ The patient: ◾ Ability to keep still ◾ Thickness and density of the body parts ◾ The detector and imaging system: ◾ Using computed radiography (CR) and digital radiography (DR) technology ◾ Quantum Detection Efficiency (QDE) ◾ The display system ◾ Viewing conditions

X-RAY BEAM CHARACTERISTICS The production of X-rays will be described in a later chapter however, in terms of image production there are a number of requirements of the Xray beam. ◾ The beam needs to be filtered to preferentially remove low energy photons which will not penetrate the patient. This reduces radiation dose and changes the energy range of the X-rays in 2

Field Size

the beam, which hardens the beam (makes the beam more homogenous, i.e. there are a smaller range of intensities). ◾ The source of radiation (focus) from the X-ray tube is small (typically from 0.3 mm2 fine focus to 2 mm2 broad focus). ◾ The size of the radiation beam can be collimated to the body part to reduce scatter and intensity. ◾ The energy of the beam needs to be adjustable to enable a range of exposures: ◾ Kilovoltage from 40 to 125 ◾ Milliamperage from 50 to 1000 ◾ Exposure times from 0.001 to several seconds.

SCATTERED RADIATION The primary beam of radiation leaving the X-ray tube interacts with the patient. There are only three possibilities for the X-ray photons leaving the X-ray tube: 1. The photons are absorbed by the patient and cease to exist (this may cause radiation damage). This gives information about the density and thickness of the patient and help create an image (signal). 2. The photons pass through the patient and produce a point of information in the detector and also help create an image (signal). 3. The photons are scattered within the patient or detector. a. This contributes to noise if they interact with the detector. b. Absorbed photons in the patient again may cause radiation damage with no benefit to the image. The image on the detector can therefore been seen as an attenuation ‘map’ of radiation which has passed through the patient.

FIELD SIZE The area of the patient irradiated can be controlled by collimation of the X-ray beam. The maximum field size at 100 cm focus receptor distance (FRD) is 43 cm2. However, it is critical that the beam of 3

Overview of Image Production

radiation is limited only to the area of interest. This can improve image quality and reduce the radiation dose to the patient and therefore staff by minimising the amount of scattered radiation produced.

GEOMETRY OF IMAGE PRODUCTION All radiographic images produced using an X-ray source and detector will be larger than the object being imaged. However, this is not always apparent from the image on the monitor as the image is optimised for image viewing by the computing system and may appear ‘life-size’. There are some important aspects determined by the geometry of the imaging system which are relevant when producing the image. ◾ Magnification ◾ Unsharpness ◾ Resolution/definition All of these factors are altered when selecting the equipment and factors for the set up of the X-ray examination. Table 1.1 states terms related to geometry.

Magnification As stated above, all images produced are larger than the body part being X-rayed. One key skill of the practitioner is to produce images with minimal magnification and unsharpness. Any unsharpness produced is magnified by the object receptor distance. Magnification is reduced by close contact between the patient’s body part and the image receptor. In practice, a standardised FRD should be used. A FRD of 100 cm for table work and 180 cm for erect chest images is used (Figure 1.1). Distances must be standardised within departments to standardise magnification. Table 1.1 Terms related to geometry. (See Figure 1.1)

4

Distances

Abbreviation

Focus to receptor distance Focus to object distance Object to receptor distance

FRD FOD ORD

Geometry of Image Production

Focus

FOD FRD

Object OR D Image Figure 1.1 Distances used in radiographic image production.

The magnification in an image may be represented by the formula,

Magnification =

FRD FOD

Unsharpness All images produced in radiography have a level of unsharpness, which should not be visible to the practitioner. Image unsharpness becomes ap­ parent at approximately 0.3 mm and is determined by a number of factors: ◾ Movement of the patient ◾ The patient may need to be immobilised or asked to arrest respiration to avoid movement unsharpness ◾ Geometry of imaging ◾ Related to distances of the patient, focus and detector ◾ How the data are displayed ◾ Type of monitor and its characteristics

5

Overview of Image Production ◾ Brightness and contrast of the monitor ◾ Viewing conditions ◾ Background conditions, e.g. light intensity in the room. ◾ Resolution and quality of the monitor ◾ Perception of the practitioner ◾ Affected by the contrast, resolution of the image and their

experience of viewing images. The unsharpness (penumbra) of the image can be calculated by the equation:

Degree of unsharpness =

Focus × ORD FOD

To minimise geometric unsharpness (Figure 1.2) ◾ Fine focus should be used where possible. ◾ The object should be as close to the detector as possible (ideally in contact).

Focus

FOD

FRD

Subject

ORD

Detector Area of unsharpness (penumbra) Figure 1.2 Diagram to demonstrate factors affecting geometric unsharpness.

6

Geometry of Image Production ◾ The FRD should be as long as practicable as this minimises the

penumbra. n.b. Images with visible unsharpness are not diagnostic and need to be repeated.

Bucky/grids Bucky assemblies/antiscatter grids are the most commonly used method to reduce noise, and thus improve contrast in an X-ray image in large body parts, e.g. spine and pelvis. This is achieved by absorbing scattered radiation, which is produced in all images. The grid allows a majority of the primary X-rays to be transmitted and a majority of the scattered X-rays to be absorbed. It is constructed from lead strips (which absorb most of the scatter) that are separated by aluminium, carbon fibre or fibrous material (Figure 1.3). The grid however, also absorbs some of the primary beam and the radiation exposure to the patient must be increased by a factor of 2–3 to compensate for the loss of primary and scattered X-rays removed by the grid. Grids can be stationary and simply placed between the patient and the image receptor or inserted into a bucky mechanism where they move in a reciprocating manner to absorb most of the scattered pho­ tons. Grids therefore increase the radiation dose to the patient and the exposure is increased according to the grid factor. The bucky is located under the X-ray table or vertical stand. The specifications of different grids can vary based on:

Patient Photons Grid

Detector Figure 1.3 Application of a grid showing preferential absorptive of scatter.

7

Overview of Image Production ◾ The number of strips over the length of the grid (numbers of

strips per centimetre). ◾ The grid ratio (height of the lead strips to the interspace distance). ◾ Whether the grid is focused or parallel (alignment of the lead

strips). A focused grid allows more useful photons to reach the image detector, but has restrictions for the FRD used. ◾ The pattern of the lead strips or orientation. The grid frequency can range from 30 to 80 lines/cm. The ratio can range from a ratio of 4:1 to 16:1. The higher the frequency or the ratio, the more the noise can be reduced by the effective removal of scatter. The grid pattern can be linear (strips running in one direction) or crossed (strips running perpendicular to one-another), where crossed grids are associated with higher noise reduction and higher patient doses. The higher the grid capability to absorb scattered radiation (i.e. higher frequency or ratio or grid with crossed pattern), the more it absorbs useful X-ray photons and the more increase in exposure is needed. It is also important to note that grid misalignment may result in grid ‘cut off’ where a large number of useful X-ray photons are absorbed by the grid, thus causing a substantial loss of image density and the necessity repeat the X-ray procedure. Modern Direct Radiography (DR) systems may have a software application of a virtual grid which can enhance image contrast and remove noise.

Resolution/definition The resolution of any radiographic image can be measured objectively using a test tool. It is normally stated in line pairs per millimetre and, for a matrix size of 10 pixels/mm2, there are approximately 5 lp/mm. If the image is assessed visually (subjectively), the ability to determine anatomy is referred to as ‘definition’. Resolution and definition are both affected by all of the components of the digital image chain, as well as the geometry of the patient positioning. Definition (spatial resolution) is also affected by the character­ istics of the detector and monitor, the pixel size and depth, the processing and display of the image (see Chapter 7, Image quality). When assessing the diagnostic quality of an image, it may be better to measure definition, as the person viewing the image can assess if they can see structures within the image, e.g. joint space, bone trabeculae.

8

Ionisation

X-RAY DETECTORS Ideally, all of the unscattered radiation leaving the patient should be ab­ sorbed by the imaging plate, and the scatter ignored by the detector. Unfortunately, this is not achievable. Digital detectors (photostimulable storage phosphor (PSP)) absorb up to 35% of the transmitted beam and this may increase to 60% with direct conversion digital systems. The remaining radiation passes through the detector and may again be scattered. There are two main types of detector for conventional projection X-ray imaging: computed radiography (CR) and digital radiography (DR). Both use photostimulable phosphors and will be explained in more detail in Chapter 6, detective quantum efficiency (DQE) is often a measure that is quoted in order to make comparisons between various imaging systems.

IONISATION This is the process of removing one or more of the electrons in an atom leaving the atom in an excited state or ionised. The remaining atom is then called an ‘ion’ and is positively charged as the electron is ejected. Ionisation is significant in a number of processes for image production. Figure 1.4 demonstrates the process of ionisation. Charged particles and photons of radiation are all able to ionise other atoms and the process features in the following instances: ◾ Production of X-rays in a tungsten target ◾ Thermionic emission at the filament ◾ Production of heat in a tungsten target ◾ Detection of radiation ◾ Radiation measurement ◾ Dosimetry of radiation effects ◾ Fluoroscopy ◾ Image production ◾ Radiation protection

9

Overview of Image Production

Energy

Ejected electron

Ion

Electron orbitals Nucleus

Figure 1.4 The process of ionisation.

DISPLAY SYSTEM AND VIEWING CONDITIONS Images are viewed on visual display units (VDU). The image is a display of pixels (these are the smallest element of the image). The matrix size affects the spatial resolution of the image. All pixels within the digital images have been processed before they are dis­ played. A look-up table (LUT) is applied to each pixel and this en­ hances the contrast and dynamic range of the image. Data may be enhanced differently by applying a different LUT and an almost in­ finite number of variables can be used to manipulate the data to provide an optimum image. The images can also be manipulated postacquisition by the operator to enhance various factors, e.g. contrast, brightness, or suppress factors, like noise.

10

Practitioner’s Skill and Perception

RADIATION DOSE It is a legal requirement (Ionising Radiation Regulation 2017 (IRR17)) to record the radiation dose delivered to the patient. This will be discussed in Chapter 8. There are a number of ways used in practice, the simplest being noting down the exposures used and the room used. Alternatively, the diamentor(DAP) reading may be noted, alternatively some DR and fluoroscopy units state a dose reading. These methods give sufficient data to allow the radiation dose to the patient to be calculated at a later date using more sophisticated methods.

PRACTITIONER’S SKILL AND PERCEPTION The concept and science of image perception is beyond the scope of this book, however, it should be noted that the practitioners viewing the image displayed on the monitor may see very different images. There is research which clearly demonstrates that the experience and skill of the practitioner affects their ability to perceive pathology or abnormalities with the image.

11

Overview of Image Production

MCQs 1. Which of the following statements best describes ‘image definition’? a. A subjective measurement b. Ability to resolve 5 lp/mm c. Ability to see a specific anatomical structure d. Diagnostic quality image. 2. Scattered radiation can be described as: a. X-rays produced by the X-ray tube b. Radiation absorbed by the patient c. X-rays which improve image quality d. Noise which does not convey useful information from the patient. 3. You are producing an image of the spine. The focus receptor distance (FRD) is 100 cm and the spine is 20 cm from the imaging plane. The spine is 40 cm long. What is the length of the spine in the image? a. 50 cm b. 48 cm c. 40 cm d. 20 cm. 4. Which of the following factors will not enhance image quality? a. Short ORD b. Long FRD c. Large focal spot d. Small focal spot. 5. Which of the following abbreviations best describes the principle of radiation protection? a. IR(ME)R b. ALARA c. ALARP d. IRR17. 12

MCQs

6. If the focus of the X-ray tube is 1 mm, the focus is 100 cm from the detector and the object is 1 cm from the detector calculate the unsharpness: a. 1 mm b. 0.1 mm c. 0.02 mm d. 0.01 mm. 7. You are producing an image of the finger. The FRD is 100 cm and the finger is 1 cm from the imaging plane. What is the magnification in the image? a. 0.9 times b. 0.5 times c. 1.1 times d. 1.01 times. 8. Which of the following statements best describes the function of a grid? a. The grid only removes scattered radiation b. The grid only removes the primary beam radiation c. The grid removes scattered radiation more efficiently than the primary beam d. The grid removes both scattered and the primary beam as efficiently as each other. 9. The grid ratio is: a. The number of lead strips per centimetre b. The height of the lead strip to the height of the interspace c. The height of the lead strip to the thickness of the interspace d. The direction of the lead strips in relation to the primary beam. 10. Ionisation of atoms produces: a. Free electrons b. Ionised atoms c. Positive atoms d. All of the above.

13

CHAPTER 2 MATHEMATICS FOR MEDICAL IMAGING INTRODUCTION The aim of this chapter is to give the student an understanding of the basic principles of the mathematics underpinning diagnostic radio­ graphy. It is essential that any practitioner operating within an imaging department and using ionising radiation has a sound base for their knowledge. You need to comprehend the maths and be able to explain the factors affecting the production of diagnostic images, the principles of exposure manipulation and safety within X-ray departments. Learning objectives The student should be able to: ◾ State the base International System of Units (SI) units ◾ Explain the applied SI units for radiography ◾ Understand and explain the basic mathematical concepts used in radiography

BASIC MATHEMATICS There are a number of basic tasks which all radiographers should be able to perform. Simple addition, subtraction, multiplication and division, for example, enable the student to manipulate exposure factors at dif­ ferent distances and calculate radiation dose. 15

Mathematics for Medical Imaging

Exposure calculations The radiation output from an X-ray tube is a product of the current applied to the X-ray tube (measured in milliamps [mA]), the duration of the exposure (measured in seconds [s]) and also the voltage applied to the X-ray tube (measured in kilovoltage [kV]). Radiation output is normally known as the intensity of the X-ray beam and needs to be varied to enable different body parts to be im­ aged. Other factors affect the intensity of the radiation beam reaching the detector. These are: ◾ The energy of the beam ◾ The medium through which the beam passes ◾ The distances between the X-ray source and the detector ◾ If a grid or Bucky is used to eliminate scattered radiation ◾ The filtration applied to the X-ray tube When the radiographic technique needs to be modified either to change the distances between the elements or if the exposure time needs to be reduced, then a new exposure may need to be calculated. This can be done by using the following formula:

mAs × kV4 mAs × kV4 = 2 grid factor × FRD grid factor × FRD2 The mAs and kV, use of a grid and distance for the initial examination, are put into the left-hand side of the equation and the new ones into the right-hand side. Changing the mAs is straightforward as the intensity of the beam is directly proportional to the mAs. Changing the kV is more complex and the calculation should be performed with the kV to the fourth power, i.e. kV4. If no grid is used, the grid factor is 1, e.g. initial exposure: 1 mAs, 70 kV 150 cms FRD and grid factor 1. Question: What is the mAs at 180 cms FRD, grid factor 1 and kV 70?

1 = 1 × 1502 1 × 1802 The new exposure needs to be 1.4 mAs to account for the in­ creased FRD.

16

Measurement prefixes (powers)

INTERNATIONAL SYSTEM OF UNITS To standardise the units of measurement used in science, SI units are used within the scientific community. There are seven standard base units and these are listed in Table 2.1. From these standard base units, other SI units may be derived which are more applicable to radiography for example. These derived SI units are defined in Table 2.2 with their applications.

MEASUREMENT PREFIXES (POWERS) There are a number of times in radiography when we use numbers which are either multiplications or fractions of the base units. These can be expressed as indices, i.e. 102, which is shorthand for 10 × 10 or 100. If numbers are divided by 10 s, the indices are minus numbers, i.e. 10−2 or one hundredth, e.g. kilovoltage and milliamps. ◾ The voltage which drives the electrons across the X-ray tube is measured in kilovolts (103 volts). ◾ The current used to produce a stream of electrons to the filament of the X-ray tube is measured in milliamps (10−3 of an amp).

Table 2.1 SI base units.

Base quantity

Name

Symbol

Length Mass Time Electric current Temperature Amount of substance Luminous intensity

metre kilogram second ampere kelvin mole candela

m kg s A K mol cd

17

Mathematics for Medical Imaging Table 2.2 Units used in radiography.

Unit

Symbol

Definition of unit

Application

Electric current

A (ampere)

X-ray tube current

Electrical potential

V (voltage)

Quantity of electrons flowing in an X-ray circuit. Usually expressed as the milliamperage (mA) Force which moves electrons around the X-ray circuit. Usually expressed as the maximum voltage applied across the X-ray tube (keV) mA multiplied by the duration of the exposure (exposure time) Ability to do work Resistance of an electrical conductor

Milliamperage mAs second Joule (energy) J Resistance Ω (ohm)

Power

J s-1

Rate of doing work

Gray

Gy

Energy imparted to a body

Sievert

Sv

Dose in Grays × quality factor

X-ray tube voltage

Quantity and quality of the exposure Production of X-rays Needs to be minimised in electrical circuits Output of X-ray generator Measurement of the absorbed radiation dose Measures the biological effect of ionising radiation

The scale of values needed in radiography ranges from: ◾ Tera as in ‘terabytes’ (TB) (1012 bytes or 1 billion bytes) to ◾ Nano as in ‘nanometre’ (nm) (10−9 of a metre or 1 thousand millionth of a metre) Other useful powers are: ◾ Giga as in ‘gigabecquerels’ (Gbq) (109 becquerels or one thousand million becquerels) ◾ Mega as in ‘megahertz’ (MHz) (106 hertz or 1 million hertz) ◾ Centi as in ‘centigray’ (cGy) (10−2 Grays or 1/100th of a Gray) ◾ Micro as in ‘microgram’ (μg) (10−6 gram or 1 millionth of a gram)

18

Logarithms (LOGS)

Multiplication and division of powers If we need to multiply indices together, as long as the base is the same, we simply add the powers together, i.e. 102 × 102 = 104 or 10,000; to divide, we turn the lower indices to a negative number and simply add again, i.e.

10,000 100 This becomes 104 × 10−2, which is 102 (100).

LOGARITHMS (LOGS) Before the invention of the calculator, there was a simple method of multiplying and dividing complex numbers. These were called common logarithms. Tables were used to convert numbers into indices and then the numbers were simply added or subtracted as above. There are two important types of logs used in science: 1. Common logs or logs to the base 10 2. Natural logs (ln) or logarithms to the base e Natural logs are associated with exponential functions, such as the half value thickness or radioactive decay. Logarithmic scales are useful when displaying data graphically with a large range of values, e.g. 1 to 1 million. Exponential data displayed on a log scale will product a straight line rather than a curve. Some computed radiography systems express the exposure index(EI) logrithmically and you need to be aware of the magnitude of change on this scale, e.g. if the expected EI is 2, ◾ A variation of +1 is 10 times the intended exposure. ◾ A variation of +0.3 is twice the intended exposure. ◾ A variation of −0.3 is half the intended exposure.

Graphs A graph is a way of displaying data in a diagram. In its simplest sense, a graph can be used to display one set of data and its relationship against another. Alternatively, different data sets can be displayed to make visual comparisons. For example, see Chapter 4 and Chapter 5, where the graphs are used to display X-ray spectra. There are occasions when a logarithmic scale is used to display scientific 19

Mathematics for Medical Imaging

Activity

Exponential decay 450,000 400,000 350,000 300,000 250,000 200,000 150,000 100,000 50,000 0

0

2

4

6 Time

8

10

Log/linear display 1,000,000

Activity (log)

100,000 10,000 1,000 100 10 1 1

2

3

4

5 6 7 8 Time Figure 2.1 The effect of different scales on data display.

9

10

data in radiography, such as radioactive decay. Examples of the effect of a logarithmic scale when presenting the same data are shown in Figure 2.1. Having the data represented by a straight line allows more accurate estimation of values between data points (interpolation) and points before or after the data measured (extrapolation).

LINE FOCUS PRINCIPLE X-rays are produced when a fast-moving stream of electrons are de­ celerated by the target of the X-ray tube. The area bombarded by the 20

Similar Triangles

Real focus

Target

Apparent focus Figure 2.2 Line focus principle.

X-rays is known as the focus. There are two conflicting variables when producing X-rays in an X-ray tube. These are: ◾ The focal area should be as large as possible to dissipate the heat produced. ◾ The apparent focus should be as small as possible by to produce sharp images. These contradicting requirements are resolved as much as possible by the line focus principle. The focal track of the anode disc is angled at about 16–20° and forms the outer diameter of a large disc which may be up to 200 mm in diameter, whereas the apparent focus may be as small as 0.3 mm2. The relationship between the real focus and the apparent focus can be given by the equation:

a = r sin . where a is the apparent focus, r is the real (actual) focus, θ is the angle of the anode. Figure 2.2 show the context of the anode in relation to the X-ray tube.

SIMILAR TRIANGLES It may be useful to practitioners to use similar triangles to calculate values used in radiography. Similar triangles can be used to calculate: 21

Mathematics for Medical Imaging Focus

FOD

B

FRD

A b

a

c

Patient

ORD Detector C

Figure 2.3 Similar triangles can be used to calculate ratios/distances in radiography.

◾ The magnification of the image ◾ The geometric unsharpness of the image

If we consider the set up for imaging from Chapter 1 to demonstrate magnification, you can see there are two triangles of different sizes. Both triangles have the same internal angles, but one is bigger than the other. To calculate the length of any side of similar triangles, the ratio of the lengths is used (Figure 2.3). For example, if the lengths of two of the three sides are known, the size of the third can be calculated using ratios. The two triangles, from the focus to the patient and from the focus to the detector, are similar, i.e. they are the same shape with the same internal angles. It is possible to calculate the length of any of the side of the triangles by using their ratios to each other. Capital letters are for the big triangle. Lower case letters are for the small triangle.

A B C = = a b c

or

size of the image FRD = FOD size of the patient

Example: the ratio of the FRD (A) to the FOD (a) (similar sides of the triangle) can be used to calculate magnification. If a finger is imaged, the FRD is 100 cms and the finger is 1 cm from the detector. The magni­ fication would be: 22

Similar Triangles

100 divided by 99 (FRD-ORD) = 1.01 The finger is only slightly magnified (one hundredth times bigger). If the distance between the detector and a lumbar vertebra was im­ aged, the FRD would still be 100 cms but the spine would be 20 cms from the detector. 100 divided by 80 (FRD-ORD) = 1.25. The spine is one and one quarter times bigger.

Inverse square law The inverse square law is a mathematical way of calculating the in­ tensity of an X-ray beam at differing distances from the X-ray tube output. It has important consequences for radiation protection and calculating the exposure factors needed when modifying radiographic techniques at different distances. The inverse square law for radiation states: The intensity of an X-ray beam is inversely proportional to the square of the distance. The law only applies if the radiation is from a point source, the radiation beam is homogenous and there is no attenuation between the source of radiation and the detector. None of these three conditions apply to an X-ray beam used for radiography. However, for practical purposes, the inverse square law may be generally applied to radiographic practice. Practically, therefore, if the beam is measured at distances from a source of X-rays, the following applies: ◾ If the distance is doubled, the intensity falls to one-quarter of its original value. ◾ If it is trebled, the intensity falls to one-ninth. ◾ At four times the distance, it is 1/16, etc. The formula is therefore:

I

1 d2

where I is intensity and d is distance. This may be represented as de­ picted in Figure 2.4. In terms of radiation protection, the inverse square law demonstrates the effect of distance from a radiation source, e.g. X-ray tube or 23

Mathematics for Medical Imaging Distance from source 3D

2D Area D

4A

A

Intensity

L

L/4

9A

L/9

Figure 2.4 Inverse square law.

radioactive patient. Application of the inverse square law may be helpful in reducing radiation dose to staff.

Statistics It may be necessary to collect, state, analyse and display data from groups (populations) in radiography. The simple expression of the mean (average), mode (most common), median (middle number) and range of values may be useful to calculate. Standard deviation may be used to demonstrate the variance from the mean. A low standard deviation indicates that the data points tend to be very close to the mean, whereas high standard deviation indicates that the data points are spread out over a wide range of values. Descriptive statistics are used to summarise the population data by describing what was observed in the sample numerically or graphically. Inferential statistics uses patterns in the sample data to draw inferences about the population represented. Statistical analysis may again be useful, and there are a number of common tests which may be used to analyse data: ◾ Analysis of variance (ANOVA) ◾ Chi-squared test ◾ Mann–Whitney U test ◾ t-test 24

MCQs

MCQs 1. The voltage of an X-ray beam is conventionally measured in: a. Joules per second b. kV c. Joules d. mAs. 2. One joule is equivalent to which of the following quantities? a. 1 m2 s b. 1 N c. 1 kg m2 s−2 d. 1 K m2. 3. Which of the following is not an SI base unit? a. Metre b. Second c. Kelvin d. Gray. 4. If a radiographic image requires 20 mAs to produce the required density and the mAs was set at 200 mAs, what is the time setting? a. 4 seconds b. 0.1 seconds c. 10 seconds d. 0.01 seconds. 5. If the output of an X-ray tube is measured at 20 mGy at 100 cm, what will be the approximate output at 3 m? a. 8 mGy b. 2.2 mGy c. 22 mGy d. 0.22 mGy.

25

Mathematics for Medical Imaging

6. Using the formula a = r sin θ, calculate the size of the apparent focus if the anode angle is 17° and the real focus is 2 mm (sin 17° is 0.3) a. 0.6 mm b. 2 mm c. 1.2 mm d. 0.22 mm. 7. The SI unit of absorbed dose is the: a. Milligray b. Gray c. Sievert d. Megagray. 8. The SI unit of dose equivalent is the: a. Milligray b. Gray c. Sievert d. millisievert. 9. The SI unit of radioactivity is the: a. Curie b. Rad c. Becquerel d. Coulomb. 10. Ten to the power of 2 (102) is equivalent to: a. 10 b. 100 c. 1000 d. 10,000.

26

CHAPTER 3 PHYSICS FOR MEDICAL IMAGING INTRODUCTION The aim of this chapter is to introduce aspects of physics that are im­ portant within imaging science. These include; atomic structure, atomic number, mass number, electrons, elements, compounds, radioactivity, isotopes, principles of radioactive decay, work, energy, heat, transfer of heat, waves, sound, magnetism, electricity, electromagnetic radiation. Learning objectives The student should be able to: ◾ Explain atomic structure. ◾ Explain the principles of radioactivity and radioactive decay. ◾ Explain concepts of work, heat, waves and different forms of energy. ◾ Explain electromagnetic radiation and understand energy characteristics in respect of the electromagnetic spectrum.

THE ATOM All matter within the universe is made from atoms.

Atomic structure The simplest atom known is that of hydrogen, it consists of 2 subatomic particles; a proton and an electron. All other atoms have a 3rd sub-atomic particle known as a neutron. 27

Physics for Medical Imaging

Protons and neutrons form the nucleus of the atom and are known as nucleons. The electrons orbit the nucleus and are not attached directly to it. They are held in place primarily by the electrostatic force produced by the positively charged protons attracting the negatively charged electrons. Protons: have a relative positive charge of +1 and relative atomic weight of 1. Although it is possible to examine the sub-structure of a proton in much more detail we will just mention quarks. The proton is composed of 2 up quarks and one down-quark. Neutrons: DO NOT have a charge but have the same relative mass of 1. Apart from not having a net electric charge, neutrons do have a similar structure to protons and also contain quarks but this time the neutron is composed of one up-quark and two down-quarks. Why talk about quarks? The reason we mention quarks is that they play a very important role in holding the nucleus together. As the protons in a nucleus all have a positive charge they naturally repel each other and try to separate the nucleus but it is thought that the ar­ rangements of quarks produce very strong bonds to hold the nucleus together. These are referred to as the short range nuclear binding energy Electrons: have a relative negative charge of –1 but have a mass 1840 times less than both protons and neutrons. Electrons are known as ‘elementary particles,’ which mean they do not have a sub-structure in the same way as protons and neutrons. Figure 3.1 illustrates the arrangements of the sub-atomic particles, while the Table 3.1 summarises them.

Atomic number All atoms have a specific atomic number and this is based on the number of protons in the nucleus. For example, naturally occurring carbon has 6 protons forming part of its nucleus and therefore has an atomic number of 6.

Mass number The mass number is the total number of protons and neutrons that form the nucleus. As electrons have negligible weight they are not included in the mass number. The commonest form of carbon consists of 6 protons and 6 neutrons and therefore has a mass number of 12. This natural form of carbon is referred to as ‘carbon 12’. 28

The Atom

Electrons

Protons

Neutrons

Figure 3.1 Arrangement of sub-atomic particles.

Table 3.1 SI characteristics of the fundamental particles.

Sub-atomic particle Proton Neutron Electrons

Relative charge

Relative mass

+1 0 −1

1 1 1

Location Part of the nucleus Part of the nucleus Orbits the nucleus

Electrons and electron orbitals The electrons in any atom ‘orbit’ the nucleus in energy bands called electron shells. If this did not happen the electrons would exist remotely from the nucleus and randomly within matter. Many aspects of modern physics, including the way X-rays are produced and interact with matter provide us with a strong indication that electron orbitals provide the most likely structural framework for those electrons which are bound to atoms. Everything about modern physics points to the ex­ istence of discrete energy bands each of which can contain a predictable maximum number of electrons. 29

Physics for Medical Imaging ◾ It is important to say at this stage that inner energy bands are the

ones which fill up first and that if an electron is forcibly removed from an inner shell then another from a shell further away from the nucleus will move in to take its place. ◾ In making this transition however, the electron loses energy as it requires less energy to exist in a shell closer to the nucleus of an atom that in one further out. ◾ Each shell has within it, a whole series of sub-shells which have very slightly different energy values from each other, thus allowing the electrons to move about within their shells but without undergoing repeated collisions with each other. When considering electron shells it is always easiest in the early stages to consider the balanced atom where there are an equal number of electrons as there are protons (i.e. where the atomic number of the atom gives the number of electrons circulating in its energy shell or shells). ◾ The shells of the atom are commonly identified using letters, with the inner-most shell (that’s the one nearest the nucleus) being the K shell and subsequent shells, (working away from the nucleus), being called L, M, N shells etc. When it comes to determining the maximum number of electrons in each shell then a different identification system has to be used. Now the K shell becomes shell number 1, the L shell is shell number 2 etc (see the table below). We can now apply a formula to determine the maximum electron capacity for each shell: If we take the n-value of the shell to be its identifying number then the maximum electron capacity per shell is given by the formula 2n2. Table 3.2 shows the shells and electron capacity. The mean energy value of each shell is determined by the atomic number of the atom and this energy value is greatest for the innermost (K)

Table 3.2 Maximum number of electrons in shells.

Shell K L M N

30

Shell number

2n2 value

Maximum no of electrons

1 2 3 4

2 8 18 32

2 8 18 32

The Atom

shell, diminishing sequentially in shells further away from the nucleus. The level of attraction between the positive nucleus and the negatively charged electrons will determine how securely the electrons are bound into the atom. (See page 51 for tungsten atom with its orbitals.)

Binding energy The binding energy of an electron is that amount of energy which is required to remove the electron from its atom and such an event will leave an atom which has been ionised. It is worth noting that an electron so affected must, in the first place, have been bound into an atom (i.e. existed within an electron shell) and that this electron must be removed from the atom, not simply removed from its shell (which is an alter­ native process leading commonly to fluorescence). The binding energy of an electron is determined by: a. The number of protons in the nucleus of the atom (i.e. its atomic number) b. The proximity of any electron to its nucleus (i.e. the orbit it sits in) Essentially the nucleus can be considered as a strong attractive force for electrons, especially in high atomic number materials where the nucleus contains larger numbers of positively charged protons. This should help to explain both a) & b) above. ◾ You also need to understand though, that it is not possible for electrons to come to rest within the atom anywhere other than in the electron orbitals which explains why the electrons are not simply found ‘glued’ to the surface of the nucleus by the electrostatic forces which exist between the positively and negatively charged particles. Every electron in any given shell of a specific atom will have the same binding energy. For example, the binding energy of K shell electrons in tungsten (W) is always 69.5 keV and L shell electrons in tungsten al­ ways have a binding energy of 10.2 keV. In short, the binding energy of any electron is not determined by the electron itself but rather, it is a function of the atomic number of the atom and the energy orbital (or shell) in which the electron exists.

31

Physics for Medical Imaging

Atomic symbols and the periodic table Every atom known is included on a table known as the periodic table. Each type of atom has its own letters (symbol) accompanied by the atomic number and mass number. Carbon should therefore be written as: (mass number) 12 (atomic number) 6C

This is also known as a ‘nuclide’.

Atomic balance You should think of the nucleus and the electron shells as being related, but separate from each other because it is very rare for natural events in daily life here on earth to cause the electrons to enter the nucleus. Likewise, the nucleons (the collective name for protons and neutrons because they exist in the nucleus) and rarely move out of the nucleus. In any given atom it is common for there to be the same number of electrons, collectively, within the electron shells as there are protons in the nucleus. This is known as an electrically balanced atom. But what happens if the nucleus and the electron shells become imbalanced electrically? ◾ If an atom gains an electron (which is a relatively common occurrence for some atoms), it is referred to as a negatively charged atom, also known as a negative ion or an anion (because it is attracted to a positive electrode or anode). ◾ If an atom loses an electron (again, relatively common) it leaves a positively charged atom, known as a positive ion or a cation (pronounced cat-ion), because it is attracted to a cathode. An ion therefore is an atom in which there is an electrical imbalance between its nucleus and electron orbits. In this atom, the nucleus is actually about 3,700 times heavier than the collective electrons circulating around it and therefore that nucleus contains far more than 99% of the mass of the atom. Let’s imagine that the atom was the size of an international rugby stadium. In such a scenario the nucleus of the atom, on the same scale, would be no bigger than a cherry (and the electrons would be smaller 32

The Atom

than pin heads). The nucleus is extremely dense (because density = mass/volume) Within physics most but not all stable atoms try to be in balance.

Ions We mentioned before that atoms are normally electrically neutral with equal numbers of protons and electrons. It is possible for atoms to lose this balance and have a different number of electrons than protons re­ sulting in a net positive or negative charge; it is then known as an ion.

Isotopes Some atoms of the same element which have the same numbers of protons and electrons may have differing number of neutrons. These variations are called isotopes and have the same atomic number but a different atomic mass (Total of nucleons). A good example of an atom with a number of isotopes is carbon, which has a total of 15 known isotopes. Carbon 12 is by far the com­ monest accounting for over 99% of all the carbon on earth. Three other well known isotopes of carbon are carbon 11, carbon 13 and carbon 14. These and other isotopes of carbon are unstable and undergo radioactive decay.

Elements An element is quite simply a substance that only contains a single type of atom.

Compounds A compound on the other hand is composed of more than one type of atom. Compounds can be quite different in the way they react when compared to the parent atoms and can have completely different properties. For example; common table salt is actually sodium chloride, we use this everyday in cooking and food preservation but if we look at the parent atoms they are very different. Sodium is actually explosive in water and highly volatile, while chlorine is a dense green poisonous gas. 33

Physics for Medical Imaging

RADIOACTIVITY This is a random process associated with the nucleus of an atom. In its simplest term, it is the break down in nuclear coherence and does NOT involve the electrons. As such it is not influenced by external environ­ mental factors such as heat or pressure. This also means it cannot be controlled or even predicted when considering individual atoms. Activity is the number of disintegrations per second of a radionuclide. The SI unit of radioactivity is the Becquerel and is defined as: 1 disin­ tegration per second. This unit is too small to be of practical value and the common unit used in radionuclide imaging (RNI) is the Megabecquerel (1 million disintegrations per second, i.e. 106). The old unit of radioactivity was the Curie which is 3.7 × 1010 disintegrations per second and was too big for common usage in radionuclide imaging (RNI).

Principles of radioactive decay As we said earlier, atoms naturally try to be in balance. Nuclear decay involves an unstable atom trying to become stable and balanced. In order for the atom to try and stabilise, the nucleus undergoes some form of emission of particles and/or charge, so that it can move to a more stable state. There are three types of emission that occur: 1. Alpha particle emission (α-emission) 2. Beta particle emission (β-emission) 3. Gamma ray emission (γ-emission)

Alpha (α) emission We have already said that atoms are balanced when the number of neutrons and protons is the same. Protons naturally try to repel each other as they all have a positive charge but the neutrons act almost like a type of glue and this is thought to be because of the arrangement of quarks. However, in certain atoms that have high atomic numbers such as uranium, the arrangement of quarks is not strong enough to hold the nucleus together. There are too many protons with positive charges randomly distributed in the nucleus and there is more chance of 34

Radioactivity

neutrons not being in the ideal place for the quarks to line up in the ideal pattern resulting in weak bonds. There is a tendency to emit equal numbers of protons and neutrons until the nucleus is smaller and the likelihood of quarks being in a good place to produce a stronger bond increases. Each group of two protons and two neutrons emitted together is known as an α particle and is identical to a helium-4 nucleus. The remaining parts of the atom now form a new atom known as a daughter nucleotide. This will be a different element as the number of protons has changed. This daughter may decay again and again until a stable daughter is achieved. An example is the decay of radium to radon (Figure 3.2).

Alpha-decay α particle 226

Ra

88

α particle 222

Rn

86

Key: Ra -- Radium (solid) Rn -- Radon (gas)

Figure 3.2 Alpha decay of radium.

Beta particle (β) emission This time the instability is because of an imbalance between the neu­ trons and protons. As the neutrons and protons are held tightly by the arrangements of quarks then a particle cannot be directly ejected. What happens is that the neutron or proton splits and ejects a very small part of itself known as a β particle. This is effectively a small particle with the same mass and charge as an electron. There are two types of β particles depending on whether the proton or neutron splits. These are known as beta minus β− and beta plus β+. Beta particles ejected only have a tiny mass of 1836th that of a proton or neutron and therefore the atomic mass changes by such a small 35

Physics for Medical Imaging

amount that it does not affect the mass number which remains the same as the parent atom. However, the relative charge of the emitted β particle changes the balance of the nucleons. This results in either the conversion of: ◾ Neutrons to protons or ◾ Protons to neutrons Figure 3.3 illustrates the two processes of beta decay The net result is that in order to balance the unstable atom the β− and + β particles decay effectively converts nucleons by emitting a charged particles. ◾ For beta minus emission the resultant atom has the same mass number but the atomic number increases by 1 as it now has more protons. ◾ For beta plus emission the resultant atom has the same mass number but the atomic number is decreased by 1 as it now has 1 less proton. The β particle itself is NOT radioactive but the energy it contains can cause harm to living tissues as it is capable of breaking certain chemical bonds which may subsequently form ions which can cause damage.

n

p + β– + υ–

p

n + β+ + υ

Key n is a neutron p is a proton β are beta particles υ is a neutrino Figure 3.3 Beta decay.

Gamma (γ) emission Gamma (γ) rays are produced as a by product of either α or β decay. The nucleus changes as a result these emissions in the atom have excess energy. As the atom tries to become stable and balanced it has to 36

Radioactivity

dispose of this excess energy. The result is that monoenergetic photons of energy are emitted in the form of gamma radiation.

Gamma emission and X-rays X-rays and γ radiation with the same energy and wavelength have identical properties ie an X-ray generated at 140 kV will be identical to γ radiation emitted by the decay of technetium 99mTc. The only way to distinguish them is their origins. γ radiation comes from the nucleus of the atom and X-ray are produced within the orbits of the atom.

Examples of radionuclides used extensively in medical imaging are: ◾ The β− decay of metastable technetium-99m (99mTc).

The radionuclide 99mTc is created by the ejection of a beta particle from Molybdnum99 (99Mo) which has too many neutrons to be stable. This produces a gamma emitting radionuclide with a half life of about 6hrs. 99mTc ejects a gamma photon of 140 KeV which is ideal for imaging (Figure 3.4)

99mTc 43

γ radiation 140 keV 6hr half life Key: Tc -- Technetium m -- metastable

99Tc 43 Figure 3.4 Decay scheme for Technetium

99m

Tc.

◾ The β+ decay of Fluorine 18 FDG (18Fl FDG) to oxygen 18 (18O).

The radionuclide 18Fl has too many protons to be stable. During the decay process a proton is converted into a neutron with the production 37

Physics for Medical Imaging

of a β+ particle. The reduction of a proton in the nucleus converts the Fluorine to Oxygen. The process allows the production of β+ used in Positron Emission Tomography (PET) (Figure 3.5).

18F 9

β+ particle 18B 8

Key: C -- Carbon B -- Boron

Figure 3.5 Decay scheme for Fluorine 18.

γ – emission and X-rays γ rays are produced as a by product of either α or β decay. As the nu­ cleus changes as a result these emissions it can result in the atom having excess energy. As the atom tries to become stable and balanced it has to dispose of this excess energy. The result is that monoenergetic photons of energy are emitted in the form of gamma radiation.

Penetrating power of the emissions Radioactive emissions have different power to penetrate materials. Alpha particles are highly ionising radiation and very destructive. Fortunately they are relatively big and heavy in atomic terms and can only travel a few centimetres in air and easily stopped even by a piece of paper. Beta particles have a lower charge than alpha particles. The charge is not really high enough to directly cause ionisation but capable of breaking certain chemical bonds which may subsequently form ions. Gamma radiation is a highly penetrating form of ionising radiation. They have a specific wavelength of γ radiation which can only have come from a specific atom undergoing radioactive decay. 38

Heat

FORCE, WORK, ENERGY AND POWER Force: is defined as the ability to move a stationary body OR to increase/ decrease the speed of a moving body. It’s unit of measure is the newton (N).

1 newton = 1 kg × m/s2 The Newton is therefore the amount of force required to move a mass of 1 kg at an acceleration rate of I meter per second, per second. Work: Is effectively done if the object is moved. If a force of 1 newton was needed to move an object over 1 metre then this would equate to 1 joule of work done. The unit of measure is the joule (J).

1 joule = 1 newton × 1 metre Energy: Is effectively the same as work in that it equates to the ability to do work and uses the same unit of measure the joule. However, it is quoted in two ways, potential and kinetic energy. ◾ Potential Energy (PE) is the amount of energy which a body is capable of emitting. ◾ Kinetic Energy (KE) is the amount of energy actually being used at the time, ie the work being done during the activity. Power: Is derived by the rate at which energy is used. It is measured in two ways either as joules per second or as watts, 1 joules per second (J/s) is the same as 1 watt (W).

HEAT All atoms are in constant motion and because of this they possess kinetic energy. Heat as a form of energy is also measured in joules. Different materials such as solids, liquids or gases have slightly different amounts of motion depending on their chemical structure. This inherent motion is also 39

Physics for Medical Imaging

affected by certain extrinsic factors. Anything that increases the amount of motion will lead to an increase in kinetic energy and effectively to an in­ crease in temperature.

Transfer of heat There are 3 forms of heat transfer; conduction, convection and radiation.

Conduction Requires the atoms to be very tightly packed together and as such is the dominant form of heat transfer for solids but conduction does occur in liquids and to a lesser extent even gases. As a material is heated the atoms within it start to vibrate more due to their increased energy state. Conduction occurs when this vibration starts to spread and pass heat energy to adjacent atoms. It is affected by certain physical character­ istics such as the material density, its cross section and length as well as the temperature difference between adjacent points.

Convection Heat transfer occurs when a current or flow of heat is created inside the material which is why it is restricted to liquids and gases. As a liquid or gas is heated the atoms tend to spread out due to the increase in their energy state this makes them less dense and they tend to rise and move away from the source of heat. As they move further away from the heat source they start to cool down become more dense and start to fall back down again. This forms a circular current called convection.

Radiation This form of heat transfer is related to vibration of atoms in a similar way to the other forms of heat transfer. As the atoms are continuously heated and given energy they vibrate and as they do they give off pure energy as a wave which has both magnetic and electrical properties. This type of energy is part of the electro-magnetic spectrum known as infra-red ra­ diation. As it is an electro-magnetic wave heat transfer does not require a medium and can even occur through a vacuum or void in space. All atoms absorb and radiate heat by this method to some extent as all atoms are in constant motion, unless they are at absolute zero (0 Kelvin or −273°C) where all motion stops. Heat transfer is influenced by such things as the surface texture and colour of the material involved and at its most efficient is called black body radiation. 40

Waves

WAVES Waves allow energy, such as electrical, magnetic or sound, to travel or transfer from one point to another and so on through any form of medium. A good analogy is to imagine how a wave is created by lots of people in football stadium. Everyone in the ground sits down to begin with then one person stands up, then the person next to them stands up, the first person sits back down again and the 3rd person now stands up, the second sits down the 4th stands up and so on, before long this movement has travelled all the way around the stadium. The people themselves are still in the same seats but energy has travelled all the way around the stadium. This is essentially how energy is transferred. Waves have three main components; amplitude, frequency and speed. The amplitude is the height of the wave while the frequency relates to how quick it switches between its lowest and highest points, while its speed is obviously how far it travels in a certain time. In our analogy the amplitude is how high people in the crowd stand up, the quicker they stand up and down the higher the frequency of the wave. While the speed is how quickly this movement travels around the stadium. The length of the wave is also important (wavelength) and may be defined as the distance over which the wave repeats itself (Figure 3.6). wavelength

Amplitude Time

• Frequency is the number of wavelengths passing a point • Velocity is frequency x wavelength Figure 3.6 Shows the parts of a sine wave.

41

Physics for Medical Imaging

Waves generally travel in straight lines but they can be reflected, deflected, amplified or absorbed. Waves of energy can also interfere with each other. If we consider two sound waves heading towards each other if they cross at the point where both waves happen to be at their peak the waves will be added together and effectively amplified.

SOUND Sound is created by a vibrating object that causes compression and decompression pressure to build within a medium in sync with the vibration. This creates a longitudinal pressure wave which flows away from the vibrating object. If the pressure wave travels towards the human ear it causes a vibration of the ear drum at the same frequency as the source and if the frequency falls somewhere between 20 Hertz (20 cycles per second) and 20 Kilohertz KHz) we perceive this as sound. Vibrations still occur outside this fre­ quency range and strike the ear drum in the same way, it is simply that it falls outside the frequency range we are sensitive to. Ultrasound is used in diagnostic imaging and uses sound above the audible range of 20 KHz and are usually between 2 and 18 Megahertz (MHz).

MAGNETISM We are all familiar with a basic magnet, one that has a north and south pole. The basic law of magnetism states that like poles repel each other and opposite poles attract. Inside the magnet there are a series of much smaller bits known as domains which have a north and south pole. In a non magnetic material these domains are randomly arranged so the material does not have a net magnetic force as all the magnetic forces cancel each other out. If on the other hand we do have a magnetic material it’s domains are arranged so they line up with their north and south poles in the same direction. Around this magnet are what is known as ‘lines of magnetic flux energy’, which form relatively large loops flowing between the north and south poles (Figure 3.7). 42

Electricity and Electric Charge

Magnetic Resonance Imaging MRI makes use of the property of nu­ clear magnetic resonance (NMR) to image nuclei of atoms inside the body. This provides excellent soft tissue images of the body.

S

N

Figure 3.7 Lines of magnetic flux.

ELECTRICITY AND ELECTRIC CHARGE Electrical charge is related to the number of electrons in relation to the number of protons in a material. If these are the same then there is no net electrical charge. If we have more protons than electrons we ef­ fectively get a net positive charge overall and if we have more electrons than protons we have a small net negative charge. Like charges repel each other and unlike charges attract each other. Electricity is a flow of electrons within a conductor.

Electrical circuit Electricity is a flow of electrons around a circuit and is called a current. If a potential difference is applied in an electrical circuit it will cause a current to flow. The positive electrode in a circuit is known as the anode and the negative the electrode the cathode. Electrons will flow from the 43

Physics for Medical Imaging

negative electrode to the positive one providing there is a potential difference and the circuit is closed. Electrical conductors have a liberal supply of electrons loosely bound in their outer conduction shell. Silver is the best metal electrical conductor but copper is commonly used due to the expense. Both atoms have 1 electron in their conduction band. Electrical insulators have all their electrons firmly bound to its mo­ lecules. In an insulating material it is much more difficult to disrupt the electrons and they usually break down before the electrons can flow. Good insulators are oil, plastics and rubber. The SI unit of electric current is the Ampere (A). The ampere (amp) may be defined in a number of ways but the most suitable is: 1 amp is 1 Coulomb of charge flowing per second. Power in an electrical sense can be determined from the current (amps) and voltage (volts). The power of an electrical circuit for X-ray production is measured in kilowatts (kW). Electricity flows through the circuit in a number of forms which are relevant in radiography. These are: ◾ Direct current (flows in one direction) ◾ Alternating current flows in either directions depending on the potential difference

ELECTROMAGNETIC RADIATION This is a form of energy that is composed of both electrical and mag­ netic fields and travels through space as a wave. As such it can be measured and classified by its wavelength and fre­ quency. The electromagnetic spectrum includes radio waves, micro­ waves, infrared radiation, visible light, ultraviolet radiation, X-rays and gamma rays (Figure 3.8). Different types of electromagnetic radiation (EMR) carry energy at different frequencies and wavelengths and this dictates their properties (see spectrum above). All EMR travel in a vacuum at the speed of light (3 × 108 m/s) and travel in straight lines. They comprise of transverse variations of electrical & magnetic fields. (i.e. a transverse wave) All EMR carry energy and are not affected by electrical/magnetic fields. As the wavelengths shorten this leads to an increasing likelihood that they will interact with the atom causing ionisation. 44

Electromagnetic Radiation Electromagnetic Waves Electric field Magnetic field

Wavelength λ The wave is travelling in this direction

Figure 3.8 Electromagnetic spectrum.

These EMR are known as ionising radiation and this is particularly im­ portant within radiography with both X and Gamma rays directly inter­ acting with the atoms within the body and causing ionisation (Figure 3.9).

Non-Ionising Radiation Low induced currents No proven effect at environmental levels

6

10

1 km 103

Low Freq 0

50

Power Lines

High induced currents

1m 11

106 1 MHz AM

Electronic excitation Photochemical effects

Heating

Radio waves

103 1 kHz

Ionising Radiation

1 mm 0–3 µ wave

109 1 GHz

1 µm 10–6 Infrared

3001012

Broken bonds DNA damage

1 nm 10–9 1Å UV

1015 Visible

Wavelength (metres) 10–12

10–15

X-, Gamma, & Cosmic Rays 1018

1021

1024

Frequency (Hertz)

Traffic Tanning Medical X-rays radars booths TV/FM Laser Radioactive µ wave Heating X-ray radars sensors Cell phones ovens lamps machines

Figure 3.9 Electromagnetic spectrum.

45

Physics for Medical Imaging

MCQs 1. If we consider carbon written as follows: 12 6C

Which statement is true? The atoms have: a. 12 protons and 6 neutrons and 6 electrons b. 6 protons, 6 neutrons and 6 electrons c. 6 protons and 6 neutrons and 12 electrons d. 12 protons and 6 neutrons and 6 electrons. 2. Which statement best describes electrons a. Negative charge with an atomic weight of 1 b. No charge with an atomic weight of 1 c. Positive charge with an atomic weight of 1 d. Negative charge with an atomic weight of 1/1840th that of a proton e. positive charge with an atomic weight 1/1840th that of a proton. 3. Which statement best describes elements? a. Elements are made of one type of atom b. Water H2O is an example of an element c. They consist of different types of atoms d. All compounds are included on the periodic table. 4. Which statement best describes an isotope? a. Are forms of an element that have the same number of protons but different numbers of neutrons b. Is an atom with less or more electrons than protons c. Is an atom with either a net negative or positive charge d. Always undergoes radioactive decay. 5. How would you define radioactivity? a. The breakdown in nuclear coherence causing the ejection of an electron b. A process directly influenced by heat and temperature

46

MCQs

c. A process not influenced by external environ-mental factors such as heat or pressure and does NOT involve the electrons d. A process not influenced by external environ-mental factors such as heat or pressure and involves the ejection of electrons. 6. Radioactive emissions have different ability to penetrate materials. Alpha particles are? a. Relatively big and heavy in atomic terms and a highly ionising form of radiation able to penetrate the skin b. Relatively big and heavy in atomic terms and a highly ionising form of radiation but not able to penetrate the skin c. Relatively small and light but able to penetrate the skin d. Relatively small and light but unable to penetrate skin. 7. Waves have three main components; amplitude, frequency and speed. The amplitude is: a. Is the height of the wave b. How quick it switches between it’s lowest and highest points c. Its velocity d. The number of waves per second. 8. Electrical charge is present if: a. There are more neutrons than electrons resulting in a net positive charge overall b. There are more protons than neutrons resulting in a net positive charge overall c. There are more protons than electrons resulting in a net positive charge overall d. There are more electrons than protons resulting in a net positive charge overall. 9. Electromagnetic radiation is: a. Radiation that has relatively long wavelengths much larger than an atom b. Radiation that has relatively short wavelengths, much smaller than an atom 47

Physics for Medical Imaging

c. Radiation that has a wavelengths that is of a similar size to that of the atom d. Radiation that has a constant velocity in a vacuum of 3 × 108 metres per second. 10. Electromagnetic radiation is: a. Classified only by its wave length b. Classified only by its frequency c. Classified by both its frequency and wavelength d. Always visible.

48

CHAPTER 4 X-RAYS, X-RAY TUBE AND X-RAY CIRCUIT INTRODUCTION It is essential that any practitioner understands the properties of X-rays and the way in which they are produced. The ability to generate a varied X-ray beam is necessary to produce diagnostic images with optimum quality and minimum dose. The circuitry of the X-ray tube coverts a mains supply of 415 volts and 13 amps to the range of kilovolts (kV) and tube currents (mA) required to generate a range of X-rays. The circuitry allows an electrically safe process and facilitates exposure times from milliseconds to several seconds. The X-ray tube not only produces a beam of X-rays, but also dissipates the heat produced as a byproduct of the process efficiently, so repeated exposures can be made. It also needs to be electrically safe at the high voltages used. Learning objectives The student should be able to: ◾ Explain the properties of X-rays, the basic X-ray circuit and the range of exposure values generated. ◾ Understand and explain how electron interactions generate X-rays and the spectra produced. ◾ Explain and illustrate how changes to exposure factors and X-ray tube settings will affect the spectral output of X-rays. ◾ Explain how a significant amount of heat is dissipated following X-ray production. 49

X-rays, X-ray Tube and X-ray Circuit

X-RAYS X-rays are invisible, cannot be heard, have no odour and are not affected by electric or magnetic fields. They are a form of electromagnetic radiation and have wavelengths in the range of 0.01 to 10 nanometres. X-rays are commonly referred to as ‘photons’ and have the ability to ionise other substances, i.e. they cause the atoms through which they pass to eject electrons from their electron shells. This ability accounts for imaging properties and their potential harmful effect. The ejected electrons can be absorbed and scattered in different media. X-rays can be detected by their ability to ionise other substances, cause fluorescence, give rise to colour changes in several substances and produce changes which can be made visible in photographic film.

X-RAY TUBE This book will only outline rotating anode X-ray tubes as this arrangement accounts for the majority of X-ray equipment. Stationary anode X-ray tubes have low ratings (heat capacity) and are now only found in dental equipment and small portable machines. An X-ray tube consists of two components: 1. The insert which is evacuated and is where the X-rays are produced 2. The tube shield which supports the insert and is responsible for electrical and radiation safety The function of the X-ray tube is to: ◾ Provide a beam of X-rays from as near a point source as possible (focus) ◾ Dissipate the heat produced effectively to prevent damage to the X-ray tube (approximately 99 per cent of the energy conversions produce heat) ◾ Provide a consistent quality (kV) and quantity (mAs) of radiation ◾ Allow X-rays to emerge only from the window (port) of the housing of the tube and exclude emissions from elsewhere in the housing, which is lined with lead sheet 50

X-ray Tube ◾ Provide an electrically safe environment for the practitioner ◾ The tube is securely supported, but capable of easy move-

ment into any position and then being maintained in that position There are numerous materials used in the construction of the X-ray tube and tube shield. These include: ◾ Tube housing – steel construction lined with lead (except port) ◾ Port – plastic or beryllium Anode Heat shield Stator Sleeve bearing

Ceramic insulator Bearing + 75 kV

Anode stem

(a)

Cathode

– 75 kV

Filament

(b) Figure 4.1 (a) X-ray tube and (b) housing.

51

X-rays, X-ray Tube and X-ray Circuit ◾ Insulation between the housing and insert – mineral oil ◾ Insert – nowadays these are made from metal/ceramics, but

historically were made from borosilicate glass ◾ Filament (cathode) assembly/focusing cup – nickel or stainless

steel ◾ Filament – tungsten ◾ Anode disc – molybdenum alloy or graphite disc or tungsten

(90 per cent) and rhenium (10 per cent) alloy focal track with graphite backing ◾ Anode stem – molybdenum ◾ Stator windings – copper ◾ Additional filtration – aluminium and sometimes copper Figure 4.1(a, b) shows a schematic diagram of the X-ray tube and a photograph showing the X-ray tube in its housing.

Parts of the X-ray tube Insert (envelope) This insert maintains a vacuum for X-ray production and contains the anode assembly and cathode assembly. Nowadays, metal/ceramics have replaced the borosilicate glass envelope as the metal component can Cathode assembly

Vacuum

Anode disc Anode stem

Rotor

HT connection

MT cable Filament leads Broad Common fine

HT cable

Bearings Glass envelope

Figure 4.2 Tube insert.

52

Filament set in focusing

Focal track

Rotor support

X-ray Tube

be earthed so that there is no build up of static. The cathode and anode assemblies are fixed within the envelope and the envelope also supports these two electrodes in correct alignment at the correct distance. All seals and metal poles are carefully chosen to match the expansion coef­ ficients of the different parts which will reduce the risk of damage to the insert during operation. Having the potential of +75 kv and –75 kv allows a maximum of 150 kV to be used. Figure 4.2 illustrates the alignment and constituents of the tube insert.

Anode assembly The anode assembly consists of: ◾ Anode disc and focal track ◾ Anode stem ◾ Rotor assembly ◾ HT connection for the positive side of the tube circuit The rotating anode consists of a tungsten rhenium disc which is ty­ pically 90–150 mm in diameter. A large rotating disc increases tube rating and thermal capacity. It may be a composite disc with a tungsten/ rhenium focal track and a graphite or molybdenum backing which re­ duces the overall mass and acts as a heat sink.

–

High voltage

+

Anode

Cathode Electron beam Filament

e–

X-ray photons Figure 4.3 Cathode focusing the electron beam on to the anode angle where X-rays are produced.

53

X-rays, X-ray Tube and X-ray Circuit

The disc has a bevelled edge which forms the anode angle and focal track. The cathode focuses the electron beam on to the focal track where X-rays are produced (Figure 4.3). Typical angles for rotating X-ray tubes are between 7°and 20°. The anode is attached to a molybdenum stem which connects it to the rotors. The anode stem has a small cross-section and is as long as possible in order to restrict the conduction of heat to the bearing as­ sembly (Note: Molybdenum is a metal, so is an excellent conductor of heat). Heat conduction is restricted by the size and shape of the stem, not the material. The stem is connected to the copper rotor assembly. The rotors use induction via the stator windings to rotate the anode at speeds of 3000–9000 r.p.m. during exposures.

Filament circuit The filament circuit provides a circuit to the filament to facilitate thermionic emission of electrons. Changing the mA varies the current so that the temperature varies. As the mA increases, the number of electrons increases proportionally to the temperature.

Cathode assembly (filament) ◾ ◾ ◾ ◾

The cathode assembly consists of: The filaments (fine and broad) Focusing cup which is negatively charged Electrical supply and connections: ◾ Filament supply ◾ HT supply to the negative side of the X-ray tube circuit The purpose of the cathode is to produce thermionic emission of electrons which can be focused on and attracted to the anode. Thermionic emission is achieved by supplying a variable mA which heats the filament. The filament is a tightly coiled wire and increasing the mA causes increases in the tem­ perature of the filament and hence the number of electrons in the electron cloud. Doubling the mA from 100 to 200 doubles the number of electrons and hence is directly proportional to the number of X-rays produced.

Focusing cup The filament is encased in a nickel housing which focuses the electrons on to the focal track of the anode. This is achieved by having sharp edges to the focusing cup, which become negatively charged when the 54

X-ray Circuit

Figure 4.4 Focusing cup and foci.

tube voltage is applied. The sharp edge concentrates the negative charge and this narrows the electron beam directed towards the anode. Figure 4.4 shows the arrangement for fine and broad focus.

X-RAY CIRCUIT The function of the X-ray circuit is to: ◾ Provide an electrically safe environment for the production of X-rays: ◾ The housing is connected to earth via the high tension cables. ◾ Live wires are electronically insulated. ◾ The high tension and filament circuits are isolated from the mains supply. ◾ Provide a stable voltage to the autotransformer which is electronically isolated from the mains supply. This allows for a non-fluctuating supply of electricity to the components of the circuit. ◾ Modify the mains supply of 415 volts alternating current (ac) to a unidirectional current at voltages ranging from 50–150 kV for radiography and 25–40 kV for mammography (high tension (HT) supply). ◾ Provide heat to the filament circuit. This creates thermionic emission to produce electrons at the anode of in the X-ray tube. ◾ Supply accurate and consistent control of the duration of the exposure from 0.001 seconds to several seconds (timer circuit). 55

X-rays, X-ray Tube and X-ray Circuit

THE INTERACTION OF HIGH-ENERGY ELECTRONS WITH MATTER Electrons released from the filament of the X-ray tube by thermionic emission are accelerated across the X-ray tube towards the target (anode) by the potential difference between the cathode and the anode. ◾ The number of electrons released in each exposure is determined by the selected mA. ◾ The kinetic energy acquired by the electrons is determined by the selected kV. As the electrons reach the tungsten target of the X-ray tube, they will start to undergo sub-atomic interactions with the atoms in the target. Essentially three types of interaction will occur: 1. Incoming electrons with outer-shell electrons of the target atoms 2. Incoming electrons with inner-shell electrons of the target atoms 3. Incoming electrons with nuclei of the target atoms Incoming electrons will undertake many interactions (probably around 1000) with the target atoms before giving up all their kinetic energy. These interactions are most likely to occur within 0.5 mm of the surface of the target.

INTERACTIONS BETWEEN INCOMING ELECTRONS AND OUTER-SHELL ELECTRONS IN TUNGSTEN Tungsten has an atomic number of 74, and therefore there are (usually) 74 electrons in each atom of the target material available for interaction. The majority of those electrons will be in outer orbitals of each of the tungsten atoms and therefore the incoming electrons will require rela­ tively small amounts of energy to interact with the shell-bound elec­ trons as they are less tightly bound to the atom. Figure 4.5 shows the atomic arrangement of tungsten. 56

Interactions Producing Heat

P 74 N 110

Figure 4.5 Tungsten atom demonstrating the number of electrons in the outer shells. P, protons; N, neutrons; ●, electron; electron shells 2, 8, 18, 32, 12, 2, Corresponding to the K,L,M,N,O,P orbits.

INTERACTIONS PRODUCING HEAT As the incoming electrons interact with the electrons in tungsten atoms (millions of times during each X-ray exposure), small amounts of en­ ergy are given up as electromagnetic radiations (EMR) due to the electrostatic repulsion which occurs between the two negatively charged sub-atomic particles. The small amounts of energy released during each interaction are far too small to produce X-rays and virtually all of the energy is released as heat. As the greater proportion of the target material (in terms of volume) occurs in the outer orbitals of each tungsten atom, it follows that large proportions of the energy provided by the incoming electrons will be converted into EMR energies at the lower end of the spectrum – most commonly infrared energy (heat). Each of the incoming electrons can undergo many thousands of these low energy-release interactions before coming to rest, making it very easy to understand how up to 99 per cent of the energy produced at the target of the X-ray tube occurs in the form of heat. 57

X-rays, X-ray Tube and X-ray Circuit

INTERACTIONS PRODUCING X-RAYS X-rays are produced in the X-ray tube by two interactive processes between incoming electrons and the atoms of the target: 1. Characteristic radiation 2. Bremsstrahlung

INTERACTIONS BETWEEN INCOMING ELECTRONS AND INNER-SHELL ELECTRONS IN TUNGSTEN (CHARACTERISTIC X-RAY PRODUCTION) It is normal to consider K and L shells as inner shells or orbitals for the purposes of this description. These are the only two shells in tungsten where the ejection of an electron could lead to the emission of a photon of EMR which would both be X-rays in nature and also have sufficient energy to escape from the X-ray tube. Two interactions are possible: 1. Excitation: where sufficient energy is given to the bound electron to raise it to the next orbital 2. Ionisation: where the binding energy of the orbital electron is overcome and it is released from the atom Where excitation occurs, the electron which was removed from its shell has only been removed temporarily. It still remains within the structure of the atom and will ultimately return to its original position, giving out a photon of EMR with an energy equivalent to that acquired by the electron in raising itself to a different shell (or valence band). If the electron transition away from its shell and back again takes less than 10−14 seconds, then the process is known as ‘fluorescence’; if it takes longer, it is known as ‘phosphorescence’. In addition to the possible production of characteristic radiation, both ionisation and excitation will also give rise to some heat production as 58

Interactions between incoming electrons and inner-shell electrons

energy is transferred to the target material during both processes but no X-rays will be produced. In the case of ionisation, the released electron is known as a delta-ray and carries with it kinetic energy donated during the interaction. The delta-ray will go on to interact with other atoms until it has lost its acquired kinetic energy at which point it becomes indistinguishable from other electrons in the material. It must be noted that ionisation can only occur in this way if the kinetic energy carried by the incoming electron (i.e. the keV applied across the X-ray tube) is equal to or greater than the specific binding energy of the orbital-bound electron. In the case of tungsten, the K-orbital binding energy is 69.5 keV and the L-orbital binding energy is given as 10.2 keV. Ionisation of a K-shell electron will leave the atom with a vacancy in this orbital which cannot be sustained in nature. Immediately an elec­ tron from a shell further from the nucleus will drop into the K-shell and in doing so a photon of EMR will be released whose value equals the difference in binding energies between the receiving shell (the K-orbital here) and the ‘donating’ shell (Figure 4.6).

Ejected K-shell electron (delta ray)

Target atom

Nucleus K L M Rebounding incident electron

Figure 4.6 Diagram of the process of characteristic radiation.

59

X-rays, X-ray Tube and X-ray Circuit

If we consider this process in tungsten, then the removal of a K-orbital electron would have arisen from the interaction of one of the shell-bound electrons with an incoming electron which carried with it a kinetic energy of at least 69.5 keV. There would then be a ‘gap’ in the K-orbital which, if filled by an L-orbital electron from within the atom, would give rise to a photon of EMR with energy of 59.3 keV (K-orbital binding energy (69.5 keV) minus L-orbital binding energy (10.2 keV)). A photon of EMR with energy of 59.3 keV would fall well within the X-ray region of the electromagnetic spectrum and would be known as a ‘K-alpha emission’. It is entirely possible that the vacancy in the K-orbital in the example above could be filled with an electron arising from the M-orbital. In such a case, the energy of the emission arising from the interaction would equal the difference in binding energies between the K-orbital electron (69.5 keV) and the M-orbital electron (2.5 keV), i.e. approximately 67 keV. This emission would be known as a K-beta X-ray photon. It should be clear at this stage that irrespective of the energy of the incoming electron, provided that energy is greater than the binding energy of the shell-bound electron, removal of that electron will always lead to the production of X-rays which fall into well-defined energy bands. These radiations are characteristic of the material in which they are produced and their energy bands depend on three variables: 1. The atom type, i.e. the atomic number of the material which, in turn, determines the binding energies of the shells. 2. The shell from which the electron was ejected (e.g. K, L, M) for any given atom. 3. To a lesser extent, the shell from which the ‘replacement’ electron comes. Characteristic line spectra are always produced by this process alongside the continuous spectra, as demonstrated in Figure 4.7. From an analysis point of view, the identity of every material can be de­ termined by identifying its emission spectra (X-ray spectroscopy). More to the point, useful X-rays can be produced using this process, pro­ viding the target material has a high enough atomic number and the tube voltage is above 70 KeV. It should be noted that characteristic radiation in any inner shell will not be produced alone. Wherever a K-alpha series of characteristic radiations is produced in tungsten, there will always be an L-series of emissions and an 60

Interactions between incoming electrons and inner-shell electrons

Photon abundance

(a)

K-alpha K-beta

0

10

20

30

40

50

60

70

80

Voltage (kV)

Photon abundance

(b)

0

10

20

30

40

50

60

70

80

Voltage (kV)

Figure 4.7 (a) Characteristic radiation (b) Bremsstrahlung with characteristic radiation.

61

X-rays, X-ray Tube and X-ray Circuit

M-series of emissions (through to the last shell). The process works on a cascade basis, so that any gaps which appear in any shell are filled im­ mediately to maintain the orbital stability of the atom. A beam from characteristic X-ray production is called a homogenous beam. For tungsten K-shell emission, it gives a: ◾ 59.3 keV energy line (K-alpha) ◾ 67 keV energy line (K-beta)

INTERACTIONS BETWEEN INCOMING ELECTRONS AND THE NUCLEUS OF THE ATOM (BREMSSTRAHLUNG X-RAY PRODUCTION) This interaction process occurs when one electrically charged particle (in this case the incoming electron from the filament with its small negative charge) and large kinetic energy is deflected by a second charged mass (the nucleus of the tungsten atom which has a mass many thousands of times greater than the mass of the incoming electron, and a much greater (positive) charge also) (Figure 4.8). The deflection will cause a change in momentum for the incoming electron which, in turn, gives rise to the emission of a photon of EMR. The loss of energy causes the incoming electron to slow down and therefore the common name for this process of X-ray production is the Bremsstrahlung process (meaning ‘braking radiation’ in German). The amount of kinetic energy given over to the photon of EMR (which is in the X-ray part of the electromagnetic spectrum) from the incoming electron is determined by how close the electron passes to the nucleus. The closer the electron passes to the nucleus, the greater will be the deflection experienced by it and the greater will be the loss of energy from the electron as a quantum of radiation. The exact energies of the various quanta produced in each X-ray exposure may be many and various from very low energy photons (which are completely ab­ sorbed by the X-ray tube) right through to interactions where the total energy of the incoming electron is given up. Therefore, as each 62

Interactions between Incoming Electrons and the Nucleus

Figure 4.8 Diagram of the process of Bremsstrahlung.

interaction is likely to produce slightly different results from those of neighbouring interactions, the graphical representation, which is produced as a result of the activation of an X-ray tube, will demonstrate a range of X-ray energies being produced rather than the single-energy spectra found with characteristic X-ray production. The closer the in­ teraction to the nucleus the greater the loss of energy and the shorter the wavelength of the X-ray photon produced. The graph in Figure 4.7b is effectively a line chart which delineates an area underneath the graph. It is this whole area (underneath the line) which represents the total X-ray energy found in the beam and this enables the graph or X-ray spectrum to be used to compare beam outputs when exposure factors and other parameters are changed. ◾ A beam from Bremsstrahlung which is composed of a range of energies is known as a ‘heterogeneous beam’. ◾ The only definite fixed point to be found on a graph representing the X-ray spectrum for the Bremsstrahlung process is the upper energy point (Emax). This is determined by the generating voltage 63

X-rays, X-ray Tube and X-ray Circuit

◾

◾

◾

◾

◾

64

applied across the X-ray tube (in keV) and therefore the kinetic energy carried by each of the incoming electrons. If the generating voltage is set at 100 keV, for example, then no electron can give up more than that amount of energy as it interacts with the nucleus of the target atom. It is found that some (though very, very few) incoming electrons do give up all of their energy in a single interaction with the charged nucleus. The energy of the photon of radiation produced would equal that amount of energy acquired by the electron as it accelerated towards the target of the X-ray tube and, which was then released in the interaction with the target atom – in our example above that would be 100 keV. Many incoming electrons will give up a proportion of their energy in an interaction with the nucleus of a target atom and, having done so, carry on through the target material to undergo further interactions (some of which will produce X-rays, but many more of which will produce yet more unwanted heat). Figure 4.7 demonstrates that many of the Bremsstrahlung interactions produce X-rays which, as a proportion of the maximum possible photon energy, occur at about one-third to one-half of that energy. This is represented as the peak intensity of the beam of X-ray energies. The peak intensity of X-ray energies is also known as the kVp (kV-peak) and the relationship between keV and kVp can be confusing for students. The kVp for a given beam is numerically the same as the keV applied as the generating voltage across the X-ray tube during any given exposure. As an example, if the X-ray beam is generated with a tube voltage of 60 keV, then the peak energy output from the resulting X-ray beam will be 60 kVp. If you check this out on a 60 keV spectrum, it will be clear that the keV and the kVp do not occur at the same point on the energy axis of the graph. As stated above, the kVp will have an energy which is between onethird and one-half of the keV. Thus, although the two values are the same numerically, they are not the same. It is convenient to consider the beam output from the X-ray tube as being comparable to the input voltage used to generate the beam and this can be done easily if the numbers are the same. Perhaps the easy way to understand what is happening is to consider the

X-ray Spectra and Factors

keV as the energy which is put into the X-ray tube to produce a similar kVp output value. The difference in actual energy between the two values is utilised in the inevitable production of heat, which always accompanies the production of X-rays.

X-RAY SPECTRA AND FACTORS AFFECTING THE QUALITY AND INTENSITY OF THE X-RAY BEAM This section will consider the factors which influence the intensity and quality of the X-ray beam produced by the Bremsstrahlung process. First, there are a number of factors associated with a Bremsstrahlung curve which we must consider. The key thing to remember here is that: ◾ The quality of the radiation measures the beam’s overall energy. ◾ The intensity is a measure of the number of X-ray photons. The intensity of the beam represents the quantity of radiation pro­ duced at a given energy and the maximum beam intensity for the full range of energies which make up the heterogeneous beam is given by the curve on the Bremsstrahlung graph. Radiation intensity is a measure of the number of photons in a beam of a given cross-sectional area. The height of the Bremsstrahlung curve will be directly proportional to the intensity of the X-ray beam. This can be referred to as the size of the curve. The quality of the X-ray beam measures how readily the beam will penetrate any given material (often measured using thicknesses of alu­ minium for a diagnostic X-ray beam). The half-value thickness of a beam of radiation is that amount of a given material which will at­ tenuate 50 per cent of the intensity of the X-ray beam. Although the measure of half-value thickness can only strictly be used with homo­ genous (monochromatic) X-ray beams, it will provide a useful guide to the penetrating power of a heterogeneous (Bremsstrahlung) beam also. The quality of any beam of X-rays is proportional to its half-value thickness for any given material. With reference to the Bremsstrahlung curve, the energy of the beam is directly related to the position of the curve along the energy axis, i.e. a higher energy beam will be 65

X-rays, X-ray Tube and X-ray Circuit

represented by the peak of the Bremsstrahlung curve being moved to the right along the energy (x) axis. This can be referred to as the shape of the curve. The maximum beam energy is often referred to as E-max and is re­ presented by the point at which the Bremsstrahlung curve’s high energy point crosses the x-axis. For the purposes of the diagnostic X-ray beam, E-max will always be equal to the generating voltage (keV) of the particular X-ray beam.

IMPACT OF CHANGING THE mA The mA is a measure of the current flowing across the X-ray tube, often called the ‘tube current’. Current in electrical terms is the flow of electrons (caused by a potential difference between the cathode and the anode) and its value in mA is determined by the number of electrons flowing per unit time. If the tube current is doubled from 200 to

Photon abundance/relative intensity

Eeff

0

20

40

60

80

100

X-ray photon energy (keV) Figure 4.9 Variation of the intensity of the X-ray beam with mA. The graph shows that the only feature to change is the height (or size) of the curve and that change will represent a doubling of the area under the curve for a doubling of the mA.

66

Impact of Changing the kV

400 mA, then there will be twice as many electrons making up the tube current and flowing across the X-ray tube. Since each electron will be subjected to exactly the same potential difference, it will have exactly the same chance of creating an X-ray photon. Therefore, doubling the mA will double the number of X-rays produced but will not affect the energy range of the X-rays. The op­ posite would happen if the mA were halved (Figure 4.9). (E-max, the minimum energy value and the position of the peak of the curve across the graph will all remain the same.) It can therefore be deduced that: ◾ Beam intensity is proportional to mA. ◾ Beam quality is unaffected by changes in mA.

IMPACT OF CHANGING THE kV The X-ray tube voltage (measured in keV) determines the potential difference between the cathode and the anode of the X-ray tube. This in turn will determine the kinetic energy acquired by each of the electrons as the current flows across the X-ray tube. As the keV is increased, the speed at which the electrons impact on the target anode is also increased, which means that there are more opportunities for the conversion of the energy into X-rays (and also heat unfortunately). As more energy is available in the interaction process at the anode, the following events will all occur: ◾ The maximum photon energy achievable will increase and so E-max will increase to match the higher keV applied across the tube. ◾ There will be an increase in the average energy of each photon of X-rays comprising the X-ray beam and therefore the peak energy (kVp) will shift towards a higher energy value (i.e. to the right along the energy axis). It also follows that if the peak energy of the X-ray beam is one-third to one-half of the maximum energy then as the E-max increases, the kVp must shift proportionately. ◾ There will be more opportunities for individual X-ray photons to be produced and therefore as the total number of photons increases, the intensity of the X-ray beam must increase. 67

X-rays, X-ray Tube and X-ray Circuit

Photon abundance

Eeff

0

20

40 60 80 100 X-ray photon energy (keV)

120

Figure 4.10 The change of the X-ray spectrum with applied kilovoltage.

All of these changes are represented in Figure 4.10. Logic will tell you that each of the changes mentioned above will be reversed if there is a reduction in the applied tube voltage. We can therefore determine that: ◾ Beam quality ∝ keV2 ◾ Beam intensity ∝ keV

THE IMPACT OF FILTRATION ON THE X-RAY BEAM The X-ray beam, in the process of leaving the X-ray tube, will pass through certain components of the tube. These are collectively called ‘inherent filtration’: ◾ The glass or ceramic envelope which maintains the vacuum ◾ The cooling oil 68

The Impact of Filtration on the X-ray Beam ◾ The window of the X-ray tube, which is usually made of

beryllium due its low atomic number (which means it will not absorb too many of the X-rays) Additionally, every X-ray tube you use will have external or added filtration added to preferentially remove the lower energy, ‘softer’ X-rays from the beam. Usually these filters will be made of aluminium for use with the diagnostic range of X-ray energies. Aluminium is used because it has a low atomic number and will therefore be more likely to absorb (or scatter) low energy photons and more likely to leave higher energy (useful) photons to form the radiographic image. The lower energy X-radiation would have sufficient energy to in­ crease the radiation dose to the superficial tissues of the patient, but it would not be able to penetrate through the patient and contribute to the process of radiographic image formation (or tumour treatment in the case of a radiotherapy beam therefore only add to patient dose). The amount of filtration added will depend on the maximum gen­ erating voltage of the tube (i.e. the maximum kV available to the radiographer). The total filtration should be 2.5 mmAl equivalent for diagnostic X-ray tubes. There is increasing evidence that an additional 1mm Copper filter will reduce radiation dose to patient without any loss of image quality. In order that we can understand the impact of the X-ray beam on the patient, we must first understand what effect filtration will have on the beam and the best way to do that is to consider what happens when inherent filtration and added filtration is allowed to impact on the X-ray beam. In Figure 4.11, you can see that the higher energy components of the beam remain broadly the same, however, the lower energy components will vary. ◾ Spectrum A represents the beam without any filtration (i.e. it is still within the X-ray tube insert). ◾ Spectrum B represents the beam after it has passed through the envelope, the oil and the tube window (the inherent filtration) and it can be seen that some of the lower energy photons have been removed from the beam. ◾ Spectrum C represents the beam after it has passed through the added filtration at which point most of the lower energy, less useful X-ray photons have been removed. 69

Intensity

X-rays, X-ray Tube and X-ray Circuit

A

B

C

Energy Figure 4.11 Continuous spectrum demonstrating the effect of filtration on effective energy.

The graph shows us that: ◾ The process of filtration will not affect E-max (because the higher energy photons are largely unaffected by the process of beam filtration). ◾ The intensity of the beam will be decreased as the beam passes through increasing thicknesses of the filter (due to the lower energy photons being absorbed or scattered). ◾ The average energy of the beam (represented by the position of the peak of the graph along the energy axis) will increase with additional thicknesses of filter material. We can therefore determine that beam quality is proportional to filtration and beam intensity and is inversely proportional to filtration.

70

MCQs

MCQs 1. With reference to the interaction of electrons from the cathode with atoms of the anode, what percentage of heat typically occurs? a. 1 percent b. 10 percent c. 2 percent d. 99 percent. 2. The inner envelope of an X-ray tube is usually made from: a. Ceramics b. Lead c. Copper d. Aluminium. 3. Typical anode angles in general diagnostic X-ray tubes (excluding mammography) tend to be between: a. 4° and 6° b. 15° and 20° c. 25° and 30° d. 30° and 45°. 4. The X-ray tube incorporates an angled target in order to: a. Decrease the size of real focal spot b. Decrease the size of apparent focal spot c. Increase the size of apparent focal spot d. Increase length of target track. 5. The following material is added to the anode disc of a rotating X-ray tube to prevent the crazing effect: a. Molybdenum b. Carbon c. Rhenium d. Copper.

71

X-rays, X-ray Tube and X-ray Circuit

6. Modern anode discs, which contain more than one material in their construction, may be referred to as a a. Bi-anode b. Double anode c. Compound anode d. Rare-earth anode. 7. The added filtration of a diagnostic X-ray tube typically consists of: a. Aluminium or copper b. Aluminium and beryllium c. Copper or tin d. Tin or lead. 8. The filtration of an X-ray beam has the effect of: a. Improving the quality of the transmitted X-ray beam b. Increasing the quantity of the transmitted X-ray beam c. Reducing the quantity and decreasing quality of the transmitted X-ray beam d. Improving the quality and increasing quantity of the transmitted X-ray beam. 9. When X-rays are produced, the maximum energy of an X-ray photon is determined by the: a. mA b. mAs c. Temperature of the filament d. keV. 10. When X-rays are emitted from the X-ray tube, the minimum energy of the radiation beam is determined by the: a. Added filtration b. mAs c. mA d. Exposure time.

72

CHAPTER 5 X-RAY INTERACTIONS IN MATTER INTRODUCTION The aim of this chapter is to give the practitioner an understanding of the interactions of X-rays with matter. It is important to understand the principles of attenuation processes and be able to adjust exposure and other factors to enhance the image and reduce scatter, or remove it before it affects the image. Learning objectives The student should be able to: ◾ Understand and explain the principles of attenuation. ◾ Understand and explain the processes of Compton scatter and the photoelectric effect. ◾ Explain the terms ‘linear’ and ‘mass’ attenuation coefficient and their significance to radiography.

INTERACTIONS OF X-RAYS IN MATTER X-rays will interact with all manner of things both inside the X-ray tube and after they have left it, but as practitioners we are primarily inter­ ested in how they interact with patients and detectors, in terms of producing the image and managing the radiation dose to the patient. The principles of interaction are similar for all materials and they are 73

X-ray Interactions in Matter

primarily affected by the atomic number of the medium and the energy of the X-rays themselves. It is a simple fact that the interactions we are discussing occur be­ tween photons of X-rays and atoms (or to be more exact, the electrons contained within atoms) of the medium which is irradiated. If we want to reduce the number of interactions therefore, we must either reduce the number of photons in the beam or reduce the number of atoms exposed (or both). The use of tissue displacement bands (often called ‘compression bands’) will reduce the number of atoms in the beam and collimation of the beam will reduce both the number of photons re­ leased from the tube and the number of atoms irradiated. All radiography practice should be based on an understanding of these basic facts; they are the very essence of radiation dose reduction. It must be clearly understood that a patient will only receive a ‘radiation dose’ if energy from the X-ray beam is deposited in their tissues. If all photons were to go straight through tissues then radiography would be a process which was risk free in terms of radiation risk. However, it is probably clear to you that if all of the X-rays were to pass straight through tissues without any interaction there would be no radiographic image either!

Attenuation In a clinical environment, all X-ray beams (although not all X-ray photons) will interact with the medium through which they pass and therefore, to a greater or lesser extent, the beam will be attenuated. To attenuate is to weaken or reduce something and in the case of the diagnostic X-ray beam, that reduction occurs through two processes: absorption or scattering of the photons. It therefore follows that as a beam of radiation passes through a medium, any one of three things can happen to its photons (Figure 5.1). They can: A. Pass straight through the medium; B. Be absorbed by the medium and therefore cease to exist; C. Undergo scattering where, following the interaction event, the photon of radiation will travel in an entirely new and different direction. It can be stated then that for all diagnostic X-ray beams: attenuation = absorption + scatter. 74

Interactions of X-rays in Matter

A No interaction Photoelectron B

Absorption K-characteristic X-ray (L →K transition)

Nucleus

C

φ θ

Incident X-ray with incident energy, E0 photons

Recoil electron Compton scattering Es < E0

Figure 5.1 Summary of interactions.

Linear attenuation coefficient (µ) The chances of any particular X-ray photon interacting with a given atom are not great at all, however, every time a patient undergoes an X-ray examination, so many millions of atoms are exposed to many X-ray photons that it is inevitable that some absorption and scattering events will take place. An estimation of the probability of absorption and scattering events occurring in any given material can be gained by considering the linear attenuation coefficient (represented by µ [the Greek letter mu]) for that material. A coefficient is simply a number or symbol which, in this case, re­ presents the ability of a medium or material to attenuate radiation per 75

X-ray Interactions in Matter

unit thickness of the material (thickness being a linear measurement, hence the term ‘linear attenuation coefficient’). In order to fully understand the mathematical concept of linear at­ tenuation for X- and γ-radiation, it is necessary to study the exponential process and the role of linear attenuation within it. However, this book is intended to be a practical handbook and further background detail in this area is left to other authors. It is probably sufficient to say here that: The total linear attenuation coefficient (µ) is the fraction of X-rays removed from the beam per unit thickness of the irradiated medium.

The total linear attenuation coefficient represents the sum of all of the absorption and scattering events which occur during the passage of an X-ray beam through the material. ◾ τ (tau) is the symbol used to determine the linear absorption coefficient for the photoelectric absorption process and ◾ σ (sigma) is used to represent the linear scattering coefficient for Compton scatter. Therefore, for diagnostic X-ray beams:

µ=t+ i.e. total attenuation = photoelectric absorption + Compton scatter. The main problem with using linear attenuation coefficient as the guide to how much attenuation will occur in a given thickness of ma­ terial is that as the physical conditions of the material change (a material could be a solid, a liquid or a gas), then the accuracy of that coefficient can change. If, for example, a material was heated and expanded, it is extremely likely that its thickness would change (i.e. its volume would increase, its density would decrease and the number of atoms per unit volume would decrease proportionately). This would mean that the linear attenuation coefficient would also change. However, if the linear attenuation coefficient were divided by the density of the medium, the numerical value would remain pro­ portionally consistent. In this context, density is represented by ρ (the Greek letter rho): (total) mass attenuation coefficient (µ/ρ). 76

The Processes of Attenuation in Diagnostic Radiography

The mass attenuation coefficient is, therefore, the linear at­ tenuation coefficient (represented by µ) divided by the density of the medium (represented by ρ) and this ratio removes the anomalies caused by any physical changes which may occur as a result of environmental conditions. The mass attenuation coefficient is therefore a more reliable measure of attenuation under a variety of circumstances and will be referred to during the remainder of this section. Total mass attenuation coefficient can be defined as: The fractional reduction of X-rays per unit mass of the medium.

It will comprise the individual mass absorption and scattering coef­ ficients: τ/ρ is the mass absorption coefficient (for the photoelectric process) and σ/ρ is the mass scattering coefficient (for Compton scatter). Therefore for diagnostic X-ray beams:

µ

=

+

i.e. total mass attenuation = mass absorption + mass scattering.

THE PROCESSES OF ATTENUATION IN DIAGNOSTIC RADIOGRAPHY Attenuation coefficients can help us to understand how much at­ tenuation will occur based on a given set of circumstances, but we also need to understand how the individual processes occur. The remainder of this chapter will consider the attenuation processes which contribute to the radiographic image. Consideration will be given to the impact of the processes on image formation and on patient radiation doses. Just to remind you:

Attenuation = Absorption + Scatter 77

X-ray Interactions in Matter

There are four processes of attenuation which need to be mentioned but, from a diagnostic perspective, only two of these need to be de­ scribed and thoroughly understood. The four processes are: 1. Elastic (unmodified) scatter 2. Pair production 3. Photoelectric absorption 4. Compton (inelastic or modified) scatter.

Elastic scatter Elastic scatter occurs between low energy photons and electrons which are bound into atoms of the target material. This process has little significance in diagnostic imaging, largely because many of the low energy photons have been removed from the beam prior to reaching the patient.

Pair production This process does not occur in diagnostic X-ray beams. The process only occurs with a photon energy of 1.02 MeV and above. The photon in­ teracts with the nucleus of an atom and produces a positron and an electron. Positron annihilation is the basis of positron emission tomo­ graphy (PET) scanning.

Photoelectric absorption The previous processes are not really relevant to diagnostic radiography, however, the photoelectric process described here is very relevant because it occurs predominantly in the X-ray energy range used in diagnostic radiography. The photoelectric effect is the main source of the data for digital image production in diagnostic radiography. This process can be best described by annotating Figure 5.2: 1. The incoming photon must have an energy which is equal to or slightly greater than the binding (ionisation) energy of the electron. (As the beam energy increases the process is much less likely to occur.)

78

The Processes of Attenuation in Diagnostic Radiography Incident X-ray photon (1)

Outer-shell electron fills gap (3) Photoelectron Plus KE (2)

Photon of EMR (4)

Figure 5.2 Photoelectric effect.

2. The bound electron will be ejected from the atom leaving it ionised. Any excess energy over and above the binding energy used by the photon to remove the electron is donated to the electron as kinetic energy and therefore the photon has donated all of its energy and ceases to exist (i.e. its energy has been absorbed). The electron will travel through the material until the donated kinetic energy has been used, when it comes to rest and acts like any other electron. 3. The vacancy in the electron shell occurring as a result of the ionisation process in (2) will be filled by an electron from a shell further away from the nucleus of the atom. 4. A photon of electromagnetic radiation will be released as a result of the electron transition in (3) above. The energy of this photon will be dependent entirely on the atomic number of the material (as it is the attractive force of the nucleus due to the number of protons it contains which directly determines the binding energy of the electrons in each electron orbital or shell). If this process occurs in patient tissues, the energy of the transition photon will 79

X-ray Interactions in Matter

be very low and it is likely that it will fall in the infrared part of the electromagnetic spectrum as heat. This interaction process will therefore see the X-ray energy carried in some of the incoming photons absorbed by human tissue with the associated potential for cellular and chromosomal damage.

Photoelectric absorption coefficients The mass attenuation coefficient (which, in this process, is effectively an absorption coefficient as there is no scatter involved) for the photoelectric process is dependent on two factors: 1. The atomic number of the medium (Z), and 2. The energy of the X-ray beam (E). In broad terms, we find that:

Z3 ◾ This means that relatively small changes in the atomic number

◾

◾ ◾

◾

◾

80

of the medium give rise to reasonably significant changes in the absorption of X-rays. Many radiographic examinations (including all extremity examinations) rely on the ability of the radiographer to distinguish between bone and soft tissues. As an estimation, it can be assumed that soft tissue has an atomic number of 7.4, while the atomic number of bone is around 13. Therefore, if those numbers are both cubed (as per the formula above) we can see that rather than bone absorbing slightly more Xrays than soft tissue, it actually absorbs more than five times as much. This will have a significant impact on the level of image contrast between bone and soft tissue that we see when viewing extremity images, for example, and it will improve our ability to make bony diagnoses from plain radiographic images. The process is also put to good use in mammographic imaging where a primary aim is to demonstrate ‘microcalcifications’ (often the first sign of breast tumours) against the soft tissue background of the otherwise normal breast.

The Processes of Attenuation in Diagnostic Radiography

Additionally, we find that:

1 E3 ◾ Essentially, this means that the chance of the photoelectric

◾

◾

◾

◾

process occurring decreases very rapidly as the radiation beam energy increases. It is therefore clear that if we wish to utilize the photoelectric effect to enhance image contrast, then we shall need to ensure that relatively low beam energies are used. It is no coincidence that much of our work as diagnostic radiographers is undertaken in the 60–150 kV range as that is where the photoelectric absorption process predominates when soft tissues and bone are being X-rayed. We would tend to use the upper part of this kV bracket when a greater degree of beam penetration is required, e.g. for a thicker patient body part, but in extremity work for example, the tendency is to use an image-enhancing lower kV of around 60. The downside to this practice is that by using lower beam energies to enhance the probability of the photoelectric process occurring, we are also increasing the amount of energy absorption and therefore increasing the dose to the patient. It does, however, affirm the point that, after the need for a radiographic examination has been determined, the production of a high quality image becomes the priority to fully justify the inevitable (if often small) radiation dose. If the two equations above are combined, then we can see clearly that there are two major influences on the ability to utilize the photoelectric effect: one which can be controlled by the radiographer (the beam energy) and one which cannot (the atomic composition of the patient).

Z3 E3 ◾ The graph in Figure 5.3 confirms the point that the probability

of the photoelectric absorption process decreases sharply with 81

X-ray Interactions in Matter

L Edges Attenuation coefficient µ (m–1)

105

K Edge

104

103

102 10

100 Photon energy (keV)

1000

Figure 5.3 Graph to show a representative mass absorption coefficient (τ/ρ) in human soft tissue.

increasing energy, however, it also appears to show a couple of anomalies where the probability spikes at certain energies. ◾ This happens because each electron orbital has a different binding energy for its electrons depending on how far that orbital is from the nucleus. ◾ The graph shows the probability of the photoelectric process decreasing inversely with respect to the beam energy, however, at a certain energy (depending on the nature of the atom) the binding or ionisation energy of the electrons in the next orbital of the atom will be reached. ◾ Up to that point, the electrons in that orbital could not be involved in the process, but as soon as that binding energy value is achieved, additional electrons can be involved and as the photoelectric process occurs as a result of a photon of X-ray energy interacting with an electron, whenever additional electrons become available, the probability of the process occurring must increase. 82

The Processes of Attenuation in Diagnostic Radiography ◾ Hence, there is a sharp upward spike in the graph at the exact

binding energy of the electrons in the next orbital and then the probability immediately begins to fall away again as the beam energy continues to increase. ◾ A further spike may be achieved, but if we assume that this corresponds to the inclusion of the K-orbital electrons being involved, then there will be no additional spikes as the K-orbital is the innermost orbital (and its electrons have the highest binding energy of any in that atom) and so the probability line for that process in that atom falls away towards zero. We now need to consider what will happen if the incoming photon has an energy which is not equal to or slightly greater than the binding energy of the bound electrons (as was the case with the photoelectric absorption process), but has one which is much greater.

Compton scatter This process will meet the parameters set out in the last paragraph above. It occurs when the energy of the incoming photon is considerably greater than the binding energy of the electron involved. In the case of the Compton scattering process, the photon energy is so much greater than the electron binding energy that some texts describe the electron as being a ‘free’ electron. This somewhat confusing term can lead you to think that the electron is not bound into an atom at all. This is not the case, it is a term which is used to try and indicate the mega-mismatch between the two energies involved i.e. the X-ray photon and the weekly-bound electron. Once again, we will use annotations in Figure 5.4 to explain the process which leads to the production of Compton scatter. 1. An incoming photon with an energy much in excess of the binding energy of the electron, interacts with and ejects the electron from the atom. 2. The ejected electron will receive a proportion of the photon’s energy as kinetic energy and, as with the photoelectric process, will go on to interact with other atoms until all of its energy has been dispersed. This deposition of energy into tissues clearly means that there is some absorption of radiation energy and that the patient receives a small radiation dose as a result of this process occurring. 3. The photon undergoes a change in direction as a direct result of the collision with the electron and is therefore scattered. It also 83

X-ray Interactions in Matter

Scattered photon (3) Incident photon (1)

Angle of scatter

Ejected electron (2)

Outer orbital

Figure 5.4 Compton scatter.

experiences a reduction in energy which is equivalent to the binding energy of the electron and the kinetic energy acquired by it as it left the atom. ◾ The angle of scatter can be directly associated with the energy lost by the photon and is greatest when the photon is backscattered along its original path (i.e. it is scattered through 180°). ◾ The angle of scattering is greatest at lower photon energies and smallest at higher energies. As the beam energy increases, there is a reduction in the amount of energy donated to the electron and therefore the overall amount of scatter is reduced. ◾ This may be a difficult concept to understand as you are likely to be aware that anti-scatter grids tend to be used for thicker patient parts and with higher kVs. ◾ However, what must be remembered is that although there will be an overall reduction in the amount of scatter at higher beam energies, it is also the case that more of that scatter will be travelling at a reduced scattering angle (i.e. in a forward direction, towards the image receptor); therefore, the anti-scatter grid is more likely to be needed at these higher energies. 84

The Processes of Attenuation in Diagnostic Radiography

Compton scatter coefficients ◾ The mass attenuation coefficient (which, in this process, is

effectively a scatter coefficient) for the Compton scatter process is dependent on two factors: (1) the energy of the beam and (2) the electron density of the medium. ◾ For this process, we find that:

1 E ◾ This reflects the statement above indicating that as the energy of

the beam increases, so the probability of Compton scatter occurring is reduced. The probability of the Compton scattering process occurring is also seen to be directly proportional to the electron density of the attenuator:

/

electron density

The electron density of a material represents the number of electrons per unit mass of the medium. ◾ As you will know from your study of atomic structure, the great majority of the mass is found in the nucleus of the atom with each proton and neutron accounting for approximately one atomic mass unit each. ◾ In most atoms, there will normally be an equal number of protons and electrons (to maintain the electrical balance). ◾ It therefore follows that all atoms with equal numbers of protons and neutrons (which accounts for virtually all atoms at the lower end of the periodic table), will have a mass:electron ratio of 2:1. ◾ Further up the periodic table (where atoms have a tendency to become more unstable), we find that the ratio is closer to 2.5:1, as there are more neutrons than protons. ◾ However, there is one atom which is different in this context because there is just one that does not have any neutrons at all. ◾ Hydrogen, as you will recall, has only one nucleon (a proton) in its nucleus and therefore it has an atomic mass of 1. It also has one 85

X-ray Interactions in Matter

◾ ◾

◾

◾

◾

86

electron and so it follows that the mass: electron ratio (electron density) for hydrogen is 1:1. This means that hydrogen has an electron density which is at least twice as great as any other atom. As stated above, the probability of Compton scatter occurring is proportional to the electron density of the material and therefore any medium which contains large amounts of hydrogen will produce greater amounts of Compton scatter. Probably the most commonly encountered material with a high hydrogen content will be water, with two hydrogen atoms out of a total of three atoms per molecule. In clinical radiography, the materials we consider to have a high water content are soft tissues and fat and it follows therefore, that the largest amounts of Compton scatter will be encountered anywhere where there are significant amounts of these materials. As radiographers, you should think about the X-ray examinations and patient types where secondary radiation grids are needed – these are probably areas with larger amounts of soft tissue present. Don’t forget, you are likely to be able to remove large amounts of soft tissue (including fatty tissues) from the image by using displacement bands which will really reduce the amount of scatter produced. This is an effective and much underused dose reduction measure.

MCQs

MCQs 1. In diagnostic radiography, photoelectric absorption occurs most in: a. Air b. Bone c. Tissue d. Fat. 2. Attenuation of an X-ray beam within matter is not affected by: a. Compton scatter b. Photoelectric absorption c. Transmission of X-rays d. Atomic number and electron density. 3. Mass attenuation increases with: a. Decreasing mass number b. Increasing temperature of the material c. Increasing beam energy d. Electron density. 4. The linear attenuation coefficient: a. Defines the probability of absorption or scattering process taking place b. Is higher for fat than soft tissue for the same photon energy c. Decreases attenuation per cm of the attenuating medium d. Defines the fractional reduction in X-rays per unit mass of the attenuator. 5. Which interaction process does not take place in the range of intensities of a diagnostic beam? a. Compton scatter b. Photoelectric absorption c. Pair production d. Coherent scatter.

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X-ray Interactions in Matter

6. The probability of photoelectric absorption occurring is greatest when: a. The energy of the incoming photon is equal to or just above the ionisation energy of the atom with which it is interacting b. The energy of the incoming photon is much greater than the ionisation energy of the atom with which it is interacting c. The energy of the incoming photon is less than the ionisation energy of the atom with which it is interacting d. The energy of the incoming photon is much less than the ionisation energy of the atom with which it is interacting. 7. As a. b. c.

the photon energy of an X-ray beam increases: The incidence of Compton scattering increases The incidence of photoelectric absorption increases The incidence of Compton scattering and photoelectric absorption both decrease d. The incidence of Compton scattering increases and photoelectric absorption decreases.

8. The mass attenuation coefficient is: a. Equal to the linear attenuation process multiplied by the density b. Equal to the linear attenuation process divided by the density c. Independent of atomic number d. Different for ice and water. 9. The probability of a Compton interaction is: a. Proportional to electron density of the medium b. Inversely proportional to electron density c. Proportional to atomic number d. Proportional to the beam energy. 10. In the photoelectric absorption process: a. All energy of the photon is passed to the free electron b. No ionisation of the atom takes place c. The vacancy from the photoelectron is filled from an inner shell electron d. The vacancy from the photoelectron is filled by an electron from an orbital (shell) further out in the atom. 88

CHAPTER 6 PRINCIPLES OF RADIATION DETECTION AND IMAGE FORMATION INTRODUCTION The aim of this chapter is to explore how radiation is detected, mea­ sured, quantified and used in order to produce images. There are various types of radiation detector which are designed for different purposes within medical imaging. There are automatic exposure devices and Computed Tomography (CT) detectors, as well as those used within general radiographic and fluoroscopic imaging. This chapter will begin by looking generally at the types of detector we may come across in the radiography department, but the focus and bias later in the chapter revolves specifically around large field detectors used in general radiography. Learning objectives The students should be able to: ◾ Discuss how radiation is detected, measured, quantified and used in order to control exposure, as well as produce images. ◾ Discuss various detectors and how they are used for different clinical purposes. ◾ Discuss the benefits and limitations of various detector types used within different imaging systems.

89

Radiation Detection and Image Formation

DESIRABLE CHARACTERISTICS OF RADIATION DETECTORS There are a number of characteristics which are considered for any kind of radiation detector. The main ones include: ◾ Absorption efficiency is clearly desirable that a detector is able to absorb as many of the incident X-rays as possible. The overall absorption is dependent on the physical density (atomic number, size, thickness). ◾ Conversion efficiency is essentially the ability of a detector to convert absorbed X-ray energy into a usable electronic signal. ◾ Capture efficiency is dependent on the physical area of the face plate minus the interspace between individual detectors and side and end walls. ◾ Dose efficiency is influenced by both conversion and capture efficiency. Typical dose efficiency is anywhere between 50 and 80% for individual detector designs, but nearer 30–60% for flat panel detectors. ◾ Temporal response should be as fast as possible and is the time it takes the detector to absorb the radiation, send a signal and prepare for the next reading. ◾ Phosphorescence or afterglow affects temporal response; until the detector has stopped giving off a signal, it cannot detect another signal. ◾ Wide dynamic range, in its simplest terms, is the range of radiation intensities the detectors are sensitive to. ◾ High reproducibility and stability help avoid drift and resultant detector fluctuation or noise variation.

DETECTIVE QUANTUM EFFICIENCY Detective quantum efficiency (DQE) is often a measure that is quoted in order to make comparisons between various imaging systems and takes account of all the characteristics mentioned above. The DQE describes how well an imaging system performs, essentially based on its overall signal-to-noise ratio (SNR) when compared against 90

Detective Quantum Efficiency

a theoretical ideal detector. It is essentially a measure of how much of the available signal is degraded by the imaging system. A very simplistic way of looking at it is that the DQE value represents the probability of a signal being produced by the detector system. A DQE of 50% means that approximately 50% of the available quanta is used by the system (compared to an ideal system) to produce a signal. If we consider two imaging systems with different DQEs, but the same SNR, the one with the higher DQE would require less signal and con­ sequently less radiation exposure for the same eventual image quality. So, in some ways, it can almost be used as a measure of dose efficiency. The actual measures of true DQEs are a little more complex as DQE is also affected by spatial frequency. The DQE of a particular system can also vary as signal values change; the signal is effectively produced by the exposure (especially the kV value), as well as the detector’s

C ontrast

MTF

SN R

D QE Resolution

N oise

WS

Figure 6.1 Factors affecting the DQE: Modulation transfer function (MTF) takes account of the combined effects of resolution and contrast and how they influence each other; signal-to-noise ratio (SNR) takes account of the combined effects of contrast and noise and how they influence each other; Weiner spectra (WS) is es­ sentially the combined effects of noise and resolution and how they influence each other (see Lança and Silva, 2009).

91

Radiation Detection and Image Formation

internal structure. The same system will probably have a slightly dif­ ferent DQE for different kV values. As such, manufacturers often supply a series of graphs of DQE plotted against spatial frequency and kV. Figure 6.1 illustrates the complex relationships involved in assessing DQE. The main reason it is often quoted is that it is a helpful measure of detector performance but, if taken at face value, can mislead without careful consideration of how it is derived.

Ionisation chambers In their simplest configuration, ionisation chambers consist of a positive (anode) and a negative (cathode) electrode plate which are placed at opposite ends of a sealed chamber (Figure 6.2). The material used to construct the chamber is an electrical insulator. The space in between the electrodes forms the sensitive volume and this is filled with a gas, such as air. The electrodes are supplied with a voltage, but as the chamber is made of an insulating material and the air in between the electrode plates is also naturally a good insulator; then a current will not flow between the electrodes. However, when X-rays pass through the chamber, some of them interact with the outer shell electrons of the atoms that make up

G as filling C athode Ionizing radiation

Positive ions + – – + – + + – + + – +

V0

– + + – – – – + – – + – Electrons

G as-tight window

Figure 6.2 Ionisation chamber.

92

–

+ Anode

Ionisation Chambers Used for Automatic Exposure Control Circuits

air inside the chamber. This causes the ejection of the electron from its orbit. This results in a free negatively charged electron (negative ion) and a positively charged ion. This process is known as ‘ionisation’. The negative ions flow to the positive electrode and the positive ions flow to the negative electrode. This causes a current to flow between the positive and negative electrode plates. The electrons move much faster as they have much less mass than the positive ions so the charge is usually measured from the anode. The amount of current that flows is directly related to how much of the air is ionised, which in turn, is dependent on the amount of radiation passing through the sensitive volume. Air ionisation chambers are not used in clinical practice to form images due to their relatively large size, but they were widely used by engineers to calibrate other radiation detectors in clinical departments. They are still used by standards laboratories to provide reference values against which all other detectors are measured. They do have an important clinical role to play and that is in auto­ matic exposure control (AEC) circuits which exploit the desirable characteristics of this type of detector.

IONISATION CHAMBERS USED FOR AUTOMATIC EXPOSURE CONTROL CIRCUITS The sensitive volume can be made very thin allowing it to be positioned between the patient and image receptor and is constructed of radi­ olucent materials so it is not visible on the resultant image. The X-rays emerging from the patient pass through the automatic exposure control (AEC) on to the imaging system (Figure 6.3). As the detector is very thin and contains gas, relatively few interactions take place so only a tiny amount of the primary beam is absorbed, but it is enough to cause ionisation within the detector and produce a small signal in proportion to the X-ray energy passing through it.

93

Radiation Detection and Image Formation Patient

A EC detectors

Inc ident X-rays X-ray generator and tube

Once a pre-set amount has been measured it sends a signal to the generator to stop the exposure

ontrol circuit accumulates this signal

Imaging receptor

EC sends a signal proportionate to the emergent X-rays passing through it

Figure 6.3 Demonstrates the set up for an automatic exposure control (AEC).

The circuitry is preprogrammed to measure the size of this signal and once it reaches a predetermined level terminate the exposure. The chambers are typically around 5 or 6 cm long by 3–4 cm wide but only a few millimetres deep. The device is crude in some respects as it is influenced by all the incident radiation that passes through its area. In other words, it cannot take account of variations in X-ray intensity within its 6 × 4 cm dimensions; it simply measures the total amount passing through that area. As such, it is important that the radiographer takes into account the patient’s anatomy that overlies the AEC area. In general radiography, we use a system of three or five chambers: Correct exposure can only be achieved if we select an appropriate chamber for the anatomy overlying it or we deliberately increase or decrease the sensitivity of the chamber to account for an area we know will result in an over exposure. 94

Ionisation Chambers Used for Automatic Exposure Control Circuits

Good collimation is essential when using AEC’s to reduce scatter in an over- or underexposed image. Figure 6.4 indicates where the AEC chambers may be positioned on an abdominal X-ray with 3 chambers. Whilst ionisation chambers have not been phased out entirely, modern digital systems tend to use ‘virtual’ chambers more and more. We simply select the relevant area as before but the digital system samples the signal in real time as it is recorded by the detector. A feedback loop still triggers the termination of the exposure. We still tend to use the traditional chamber locations but we can actually select how many chambers and whereabouts we want to measure/sample on some sys­ tems. So we can have 5 or 7 area sampling if we desire or even more and we can place them anywhere within the imaging frame.

Figure 6.4 Position of the automatic exposure control (AEC) chambers on an abdominal X-ray, where R is right AEC; L, left AEC; C, central AEC.

95

Radiation Detection and Image Formation

Scintillation crystals/photocathode multiplier Scintillation crystals/photocathode multipliers have a role as scintilla­ tion counters within nuclear medicine and the gamma camera is an extensively modified scintillation counter (Figure 6.5). They were also used as an early type of detector primarily with first and second gen­ eration CT scanners. X-ray and gamma radiation detection is essentially a three-stage process: 1. A solid scintillation crystal captures and converts X-rays into light. 2. Light is then converted into a small electrical signal by the photocathode. 3. Finally, a photomultiplier is used to amplify the signal into a much larger useful electronic signal. This type of detector is used in medical imaging but no longer to produce images from X-ray systems. It was notorious for drifting and afterglow, resulting in image degradation and inaccuracies.

X-rays or gamma rays Solid scintillation crystal Light Photocathode surface (converts light into an electrical signal)

+– +–

Photomultiplier (amplifies the electronic signal)

+– +– Electrical output

+–

Figure 6.5 Scintillation crystal and photocathode arrangement.

96

Ionisation Chambers Used for Automatic Exposure Control Circuits

Scintillation crystal/photocathode X-ray image intensifier One technology that is very similar and is still being used clinically is the X-ray image intensifier (Figure 6.6). It only merits a brief description as this technology is slowly being phased out of production. Image production is a four-stage process with the whole system en­ cased in a vacuum tube: 1. A solid scintillation crystal coats the inside of the vacuum tube face plate. It captures and converts X-rays into light. 2. Light is then converted into a small photoelectrical signal by the photocathode. 3. The photoelectrical signal is accelerated and focused by high kV electrodes arranged around the inside diameter of the tube. The electrodes decreasing in circumference along the length of the tube towards the output phosphor. 4. The highly focused and energetic photoelectric signal strikes the output phosphor which subsequently converts the signal to light. The light output can then be recorded using a solid-state charged couple device (CCD)-based camera. The diagram above shows optical

Input phosphor

Electrodes (30 kV ) Ceramic/metal construction O utput phosphor (30 mm)

X-ray beam Fibreoptic plate Z oom B

Z oom A Input window (170 to 400 mm) Photocathode Figure 6.6 Image intensifier.

97

Radiation Detection and Image Formation

fibres connecting the output phosphor to the CCD system which will be digitised. This technology is currently in clinical use, but its days are numbered as it is slowly be phased out in favour of flat panel technology.

Scintillation crystals/silicon photodiode multiplier Solid-state type of detectors (Figure 6.7) are used extensively within CT scanners, but their principles of operation are also very similar to some of the large field detectors that are discussed as the next topic. The latest solid crystal detector materials have many advantages over their predecessors, including high stability and relatively small size, together with possible cost savings. Radiation detection is essentially a two-stage process this time: 1. A solid scintillation crystal captures and converts X-rays into light. 2. The light is then converted into a useable electrical signal by the photodiode. The signal is proportional to the quantity and quality of the incident X-rays. The principles of signal creation are similar to indirect digital radio­ graphy based on thin film transistor technology which is described later X-ray

Solid scintillation crystal Light

Silicon photodiode multiplier

+–

Electrical output Figure 6.7 Solid state detector.

98

+–

Large Field Detectors (Overview)

in this chapter. The latest solid-state detectors of this type have virtually zero afterglow and are used almost exclusively in spiral scanners, cer­ tainly the case for multi-slice/spiral scanners and those capable of CT fluoroscopy. Individually they have a face plate of 0.5 mm2 but are housed in large detector arrays, with a 64 slice scanner having 64,000 individual detectors, typically 64 in the z axis by 1000 in the xy axis. Please see chapter 12 for more information about CT detector arrays.

LARGE FIELD DETECTORS (OVERVIEW) Large field detectors (LFDs) are specifically designed to produce fullsize radiographic images and have replaced film screen technology. In order to produce an image, we need to detect the radiation that emerges from the patient’s body being examined. Traditionally, this was performed using photographic emulsions which were either directly exposed to the emergent radiation or more commonly the emergent radiation was used to excite intensifying screen phosphors which caused them to produce light. This in turn was then used to produce a latent image in the photographic film emulsion that was subsequently pro­ cessed using photographic chemicals. The discrete individual detectors just mentioned would be simply too big physically to replace film screen technology, recording unacceptably large pixel sizes for general radiography. For example, modern multi­ slice CT detectors are about as small as it gets in terms of the size of individual detectors, with the smallest currently available individual solid-state detectors being around 0.25 mm2 across the face plate. Even if we could get the detectors into an imaging plate thin en­ ough to fit inside the bucky trays of conventional X-ray equipment, 0.25 mm2 would only equate to two line pairs per millimetre (as a line pair is one black line and one white line). This is because we could only squeeze in four detectors in the x-axis by four detectors in the y-axis giving a total detector density of 16 detectors per mm2. The information or data measured by each detector is used to form the picture elements (pixels) in the resultant image, so this system would also give a pixel density of 16 pixels per mm2. In reality, it would 99

Radiation Detection and Image Formation

be even less than this when we account for the interspace material re­ quired to separate the individual detectors. In order to get close to the resolution available with photographic emulsions, we need to use detector technology in a slightly dif­ ferent way. A typical resolution of an older fast film screen technology used for larger body areas, such as the spine or abdomen, would be around five line pairs per millimetre. This is equivalent to a resultant image having ten pixels in each axis giving a pixel density of 100 pixels/mm2, which also equates to a pixel resolution of 100 μm. A traditional fine or detailed film screen combination used to image extremities would have to have an even greater resolution of at least ten line pairs per mm resulting in the equivalent of a pixel density of 400 pixels/mm2, which equates to a pixel resolution of 50 μm. In order to achieve these resolution values, then any digital detector system needs to have at least 400 individual areas per mm2. In radiography, rather than using individual detectors as we would with say CT, we use a large flat panel detector which produces a signal covering the whole area of the panel. We then need to put this into a grid known as the image matrix to make sense of it and this is where technology varies. There are several manufacturers employing different technologies to detect/capture emergent radiation and subsequently produce an image. There are a few terms used when discussing these technologies, but the main categorisations are usually indirect and direct systems.

INDIRECT, DIRECT, COMPUTED AND DIGITAL RADIOGRAPHY Indirect systems may either be computed radiography (CR) or in­ direct digital radiography (IDR), but in both cases X-rays are first absorbed and converted into light before being converted to an electrical signal. Direct digital radiography (DDR) does not use an intermediate stage. The emergent X-rays directly cause the system to produce an electrical signal with no intermediate conversion of X-rays to light. 100

Indirect, Direct, Computed and Digital Radiography

Computed radiography in detail CR is a system that produces digital radiographic images utilising ima­ ging plates. From a user’s perspective, it is very similar to film screen technology and was introduced because it does not generally require modifications to the X-ray equipment itself. Following an exposure, the CR imaging plate retains a latent image in a similar way to previous film screen technology. The differences occur when we process the latent image. Rather than being processed chemically, the latent CR image is scanned using a laser beam and digitised in a CR reader. The data are then sent to a computer for display, manipulation and archive. Computed radiography using imaging plates (photostimuable phosphors (PSP)) is currently in widespread clinical use to absorb X-rays and convert them to light (Figure 6.8).

X-rays Photostimuable phosphor plate

T rapped electrons forming latent image Figure 6.8 Trapped electrons in a photostimuable phosphor (PSP) plate.

CR plate construction The imaging plates of CR systems are actually very similar to X-ray intensifier screens used in film screen technology in that their function is to absorb X-rays and convert them to light (Figure 6.9). The main difference is that the phosphor material allows a delay to occur as part of the process which will be discussed in more detail shortly, but first we will look at the structure of the imaging plate itself. There are some alternatives, but for our purposes we will use the principles associated with a PSP, such as barium fluoro-halide activated or doped by europium (Ba F Brx I 1–x:Eu). 101

Radiation Detection and Image Formation

C onductive layer

Reflective layer Protective layer (soft backing)

M agnetic bar code ID Figure 6.9 A cross-sectional representation of a computer radiography (CR) ima­ ging plate. Please note the reflective layer is missing on a higher resolution version.

Production of the latent image The emerging X-rays from the patient pass through the surface of the cassette on to the PSP. The X-rays interact with the electrons of the atoms within the PSP’s conductive layer and transfer some energy. This

X-rays

Figure 6.10 The triangles represent the crystalline structure of the photostimuable phosphor (PSP). The dot represents atoms within the crystalline structure that contain electrons in higher energy bands due to them capturing the energy from the X-ray beam.

102

Indirect, Direct, Computed and Digital Radiography

causes the energised electrons of the PSP to move to a higher energy band within the atom’s structure through the process of excitation (Figure 6.10). The phosphor crystals are ‘activated or doped’ and this forms electron traps that hold on to the energised electron in this higher energy band. This forms the latent image as an analogue impression across the surface of the imaging plate. The plate will retain this impression until it is processed by the CR reader.

Processing or reading the latent CR impression The energised electrons require additional energy in order to escape this energy band and return to their original/natural energy band. Refer to Chapter 3, for a further explanation of band theory. The CR reader does this using a laser, with the light being the energy source. This gives the energised electrons enough extra energy to escape the trap. These electrons then fall back to their original energy band/orbit and, as they fall, they give off the excess energy in the form of different coloured light to that of the laser, usually blue light. This is measured by a moving scanning blue light detector (Figure 6.11).

Detector

Signal

Figure 6.11 A diagram illustrating the construction of a computer radiography (CR) imaging plate being read.

103

Radiation Detection and Image Formation

Rotating mirror L aser Detector Light released by laser light guide (rapid scan across) Imaging plate Plate slowly moving lengthwise

Figure 6.12 A simplistic diagram illustrating the main principles surrounding the reading of a photostimuable phosphor (PSP).

The scanning detector measures light output in both the x- and y-axis. It does this by moving across the width of the imaging plate while at the same time, the imaging plate passes through the reader (Figure 6.12). Electronics track the x and y coordinates of the laser and detector as well as the quantity of light emitted at each point. The reader even­ tually builds up a map of the light output across the whole of the imaging plate in the form of a grid which becomes the image matrix. The data for individual squares within the grid form the picture ele­ ments or pixels in the final image. The data are sent to a computer workstation for display, manipulation and storage. The rate at which the laser, detector and plate move through the reader can be slowed down allowing a greater number of measurements to be taken per mm2 and this subsequently results in a finer matrix being formed. Typical CR resolutions range from 100 to 200 μm, so spatial resolution is lower than fine or detailed film screen technology. Fortunately, it benefits as from higher contrast resolution so is similar and generally regarded equivalent in terms of overall image quality. The latest systems can employ multiple parallel lasers and light de­ tectors/scanners which has significantly reduced the time it takes to read an imaging plate to under 10 seconds, which is a similar time to digital radiography (DR) technology. 104

Indirect, Direct, Computed and Digital Radiography

Another criticism relates to processing of the imaging plates during mobile and theatre cases or other areas that did not have CR readers available nearby. This has been overcome lately with portable imaging plate readers being incorporated into mobile X-ray equipment or po­ sitioned around the hospital. The mobile CR readers can either be di­ rectly plugged into a wired network or even sent wirelessly to the main system. Criticism of early systems was that they required relatively high ex­ posures when compared to DR and fast film screen technology. Some manufacturers put two PSP layers one on either side of a transparent substrate to form the imaging plate, while others incorporated reflective layers. The result is a near doubling of sensitivity and lower noise, allowing almost a halving of exposure and subsequent dose to the patient. However, these techniques also cause spatial resolution (detail) to reduce. Modern CR systems are at a point where processing times are within a few seconds of DR. Exposures and system sensitivity and therefore dose is very similar. The cassettes tend to be more compact and easier to use in challenging examinations, they are also far more robust than most DR systems and with the introduction of mobile readers, similar ben­ efits are enjoyed away from the main imaging department.

Indirect digital radiography technology in detail When the scintillator is exposed to X-rays, it immediately produces fluorescent light in proportion to the quantity and quality of X-rays interacting with it. The X-rays give electrons enough energy to move to a higher orbit, but unlike the phosphors used with CR, there are no electron traps so the energised electrons fall immediately back to their natural orbit releasing their excess energy as light. There are two main systems that come into this category: those based on thin film transistor (TFT) technology and those based on charged coupled device (CCD) technology. Both designs use phosphors/scintillators that produce light when exposed to X-radiation. The differences in the systems revolve around how this light is detected and converted into a useful electrical signal that represents the quantity and quality of X-rays that fell on a particular area of the scintillator. There are a few phosphors/scintillators that may be used by manu­ facturers, such as gadolinium oxisulphide (Gd2O2S) or caesium iodide 105

Radiation Detection and Image Formation

(CsI). There are advantages and disadvantages with any scintillator, but generally speaking the materials are subclassified as being either struc­ tured or non-structured. Nonstructured scintillator crystals, such as Gd2O2S, are arranged randomly throughout the scintillator. The crystals themselves have si­ milar dimensions in all planes, i.e. the face plate may be no larger or smaller than one of the side or oblique walls. Although not always the case, they tend to have a relatively high light output or conversion ef­ ficiency as the face plate is similar in dimension to any other surface of the crystal and there is, therefore, a relatively high chance of interaction with the X-ray beam in comparison to a structured crystal with a re­ latively small face plate. However, the light output has quite a large spread due to the shape of the crystal and does not produce a focused light in one direction. Structured scintillator crystals, such as CsI, have their crystals arranged more formally and tend to lie in parallel lines. This is due to the crystals being produced as long thin rods. As the end of the rod is the part of the crystal that faces the X-ray beam, it has a relatively small face plate area and therefore less chance of X-rays interacting with it. This results in a far more focused emission of light from the other side of the crystal facing the photodiode or CCD array, but the overall amount of light produced tends to be much lower than with unstructured scintillators. A secondary advantage of using structured crystals is that the shape of the crystal means incident X-rays have to fall directly onto the face plate. Any oblique rays (scattered radiation) are unlikely to cause the crystal to emit light reducing the need for a secondary radiation grid enabling the radiation dose to be reduced (Figure 6.13). Fortunately, the characteristics of both types of crystal are carefully matched to the recording systems. Generally speaking, TFT systems work better with unstructured crys­ tals and can utilise the high light output because they are closely coupled in a sandwich to the back of the scintillator crystal, so light is captured before it spreads out too much. Whereas CCD systems tend to use structured crystals as the more directional light, output is optically cou­ pled through a mirror and lens (or optical fibres) and has to travel a relatively long distance to the CCD array. But this does not have to be the case and many manufacturers use CsI with their TFT designs. Whereas CT detectors tend to use Gd202S but there are alternates. 106

Indirect, Direct, Computed and Digital Radiography Primary X-rays = solid arrows Secondary X-rays = dotted lines G d2O 2S C sI Scintillation crystal Secondary light production

Solid arrows represent primary light output. is is greater with G d2O 2S, but spread more widely. Light output of C sI is lower but far more focused. Figure 6.13 Features of a scintillation crystal.

Indirect digital radiography using thin film transistor technology The scintillator forms the top layer of a sandwich, the next layer being the photodiode layer. The light produced by the scintillator interacts with the photodiode which produces an electrical charge, which is proportional to the amount of light interacting with it. The principles are similar to the single scintillation crystals/silicon photodiode multiplier detector described earlier in the chapter. The difference is that this is one large flat plate currently approaching 43 cm2 (Figure 6.14). How then do we get pixel densities of up to 400 mm2 with IDR technology? In contrast to CR technology, there is no latent image phase in the conventional sense (see below). The emitted electrical charge passes directly to the last part of the sandwich, the active matrix array (AMA) formed by the TFT charge collector layer covering the entire surface area of the photodiode. This layer is divided into an extremely fine 107

Radiation Detection and Image Formation

Scintillator crystal

X-rays

Photodiode

Light

Q

Q Q

Q Q

Q

Electrical charge

T FT charge

Y a xis co-ordinate grid reference X axis co-ordinate grid reference Figure 6.14 Diagram of an indirect radiography system.

grid of minute areas where the charge is collected and measured (Figure 6.14). The grid itself forms the raw data matrix with each area of the grid being given a co-ordinate reference in both the x and y axis.

Active matrix array in detail The active matrix is essentially a very fine grid of transistors and capa­ citors held together in a thin layer (Figure 6.15). It is the same size (in the x and y axis) as the scintillator crystal and photodiode layers that sit above it. The matrix itself will directly form the pixels in the resultant image. The grid will contain as many areas, known as detector elements (dels), as required for adequate resolution. Earlier we considered a resolution of ten line pairs per millimetre 108

Indirect, Direct, Computed and Digital Radiography

C omponents of detector element (Del)

T FT array

C harge area (capacitor) also acts as the electrode

T ransistor which acts as a switch

Drain lines

G ate lines/drivers

Figure 6.15 Diagram of an active matrix array.

equating to a pixel density of 400 mm2. This means that for full re­ solution images to be produced, the TFT charge collector will need to have at least 400 dels (each containing a transistor and capacitor) for every mm2 as well. We also said earlier that no latent image is formed. This is true in the conventional sense as no relatively long-term latent image is formed. However, the electrical charge from the photodiode is connected to the electrode of the TFT and creates a short-term latent charge in the ca­ pacitor of the individual TFT dels to be stored, but it is only held for a fraction of a second. This is because a very short time later, the gate of the transistor for a particular del is switched on allowing the charge to be released and read from the TFT’s drain line. In reality, this is not done 1 del at a time, many dels are turned on in a co-ordinated sequence with multiple readings being taken and digitised simultaneously, something known as ‘multiplexing’. 109

Radiation Detection and Image Formation

In Figure 6.15, the columns and rows of the array form the gate and drain lines. Following an exposure, electronic circuits energise the gates of the transistors in the entire column. This causes a charge to be re­ leased from every del in the column with their charges flowing down individual drain lines (rows). This results in a specific amount of charge for every del in that column which is equivalent to the radiation that interacted with it. The next column is then energised and another set of charges flow down the drain lines and so on. This is done extremely quickly as it only requires the circuits to be switched electronically enabling us to obtain all the information from the entire matrix in just over a second. Every del in the entire array will have an individual value attributed to it which is representative of the amount of radiation that interacted with it during exposure. This is digitised and displayed on a monitor in less than 10 seconds. Information quoted about a particular TFT array often refers to what is known as the ‘fill factor’. Essentially this is the proportion of sensitive area (charge collection area) against the dead areas of the array which includes the tiny electronic circuits (gate, drain, transistor and capacitor electronics) between the collection areas. A fill factor of 1 means the entire area is sensitive, but such a system cannot exist with current technology as we will always need the associated electronics to send the information, creating the dead area. There is a manufacturing limit on how small we can make the electronics which are similar sized, regardless of the resolution of the array. This means that the sensitive area is relatively large in com­ parison to the electronics for a low resolution array, in the region of 0.8 but is more likely to be around 0.5, for high resolution systems. This means that only 50% of the array is sensitive to radiation in the higher resolution system, the rest is filled with electronic circuits. This obviously reduces the DQE with the higher resolution system, it also effectively limits the spatial resolution that TFT systems can achieve.

Indirect digital radiography using charged coupled device technology In many respects, CCD technology produces very similar results to TFT technology. The difference lies in the physical size of the CCD array 110

Indirect, Direct, Computed and Digital Radiography

which is not big enough to cover the planar dimension of the scintil­ lator. In some respects, even though CCDs are solid-state detectors, they actually work in a similar way to ionisation chambers except they are designed to work with photons of light rather than be directly exposed to X-radiation. However, X-rays will also affect them, which is why they are carefully positioned to avoid X-rays interacting with them. Their main advantages include high spatial resolution, wide dynamic range, low electronic noise and linear response. As they are sensitive to light, it also means they do not need the photodiode layer of the TFT system, as they can produce a signal di­ rectly from the light output of the scintillator. The photon of light from the scintillator strikes the surface of the CCD del and is enough to eject an electron from its orbit. A potential difference is applied across the individual CCD del which causes the ions to move to different areas of the del where they are collected. This is where the technology varies to other designs. Rather than the signal being read here it transfers its charge to its neighbouring del. While at the same time its neighbour also transfers its charge, and so on, across the whole array, hence the name ‘charged coupled devices’. This happens in both the x and y axis at the same time producing a serial signal which is collected at one corner of the array. If this serial signal is then calculated against a time line, it is possible to determine exactly from where the signal originated within the array. One reason for designing the array in this way is that we do not need signal wires similar to the gates and drains of TFT technology running throughout the array and therefore the space between the dels is much smaller, meaning they have far better resolution with a CCD pixel size of around 0.10–0.14 μm. Typical high resolution CCD systems have 4000 × 4000 dels, giving a total of 16,000,000 dels (effectively a 16 mega pixel system), a similar number of dels to those in TFT arrays. However, in order to maintain high charge coupling ratios required for serial transmission, CCD arrays are limited to relatively small sizes, with typical overall dimensions of only 4 × 4 cm. This means the light output covering an area of 43 cm2 from the scintillator has to be reduced to cover the photosensitive areas of the CCD which is only 4 cm2. Consequently, even though the CCD pixel is 111

Radiation Detection and Image Formation

around 0.10–0.14 μm, this is the demagnified pixel size of CCD array; the actual raw data pixel represents an area in the region of between 100 and 200 μm and this is the true resolution of the system. One of the design considerations of these systems is how to connect and where to place the CCD array. In addition to light, CCDs are also sensitive to X-rays. Any X-rays interacting with the CCD could affect the charge coupling and create a false signal. Therefore, the array cannot be positioned in line with the scintillator crystal where X-rays may interact with it. There are two ways to achieve this: one way is to use fibre optical tapers; the other way is to use a mirror and optical lens arrangement.

Charged coupled device coupling via optical fibre Figure 6.16 shows six tapering optical fibres. In reality, if we wanted to manufacture such a system to cover the full 43 cm2 area of the scin­ tillator and match it to the CCD array, we would need 4000 × 4000, a total of 16,000,000 tapers. Due to the relative expense of having this many fibres, it would be prohibitively expensive for use in a general X-ray room.

Scintillator crystal

X-rays

Light is focused within the tapering optical fibres CCD charge collectors

Figure 6.16 Diagram of a charged couple device (CCD) coupling via optical fibre.

112

Indirect, Direct, Computed and Digital Radiography

However, there are some systems that use a modified scaled down version of this technology. Two of the most common systems will be briefly outlined. Both of the systems use what is known as ‘slot scan techniques’ to produce images.

Slot scan chest radiography The first system is a dedicated chest X-ray system where the X-ray beam is tightly collimated to an area of around 1 cm high × 43 cm wide. The corresponding imaging system uses a 1 cm high scintillating crystal, again 43 cm wide. The scintillator is attached to a series of CCD arrays connected by optical fibres, but as the coverage is only 1 cm × 43 cm, the number of fibres required is much less than the number required to cover 43 cm2 making such a system now finan­ cially viable (Figure 6.17). How then do they cover the full 43 cm2 required for a chest ex­ amination? These systems essentially scan the chest to produce the image. The X-ray beam and detectors move from the top of the chest to the bottom during exposure with a 1 cm moving strip covering the width

Scanning motion

Start

Translating X-ray tube

Scatter (not detected)

Linear CCDs t=0s

Primary X-rays (differential attenuation) 43 cm

1 cm Fibreoptic taper

t=5s End Pre-patient collimation

Post-patient collimation

Fibreoptical CCD array

Figure 6.17 Linear charged couple device (CCD) slot scan array large image acquisition.

113

Radiation Detection and Image Formation

of the chest as it moves down the chest; this is all done in a single breath hold. The advantages of this technique include a two- to four-fold reduc­ tion in dose over conventional CR and DR systems, very high resolution and also benefits from low levels of noise and little scatter due to rodshaped CsI scintillator crystals acting a little like a secondary radiation grid. It is currently regarded as one of the most effective systems for radiography of the chest. The only real disadvantage is that the equipment is only really capable of this one job and would only warrant being installed dedicated chest X-ray facility. The second system is a dedicated digital mammography system which uses a batch of fibreoptic tapers linked to a CCD detector array with 8192 × 400 channels. It operates in a very similar way to the chest slot scan systems, but everything is scaled down to produce a very fine scanning strip of 1 cm × 22 cm. Due to the small size of CCD it actually needs four CCD arrays to cover the 22 cm width. Each CCD has a matrix of 2048 wide × 400 high, by adding the four arrays together. This give us a matrix which is 8192 channels wide × 400 channels high. It takes about 6 seconds for the beam to scan the breast, but the images produced exhibit exceptional image quality together with a relatively low radiation dose. However, for the reasons previously discussed, this technology for all its benefits does not suit large area single exposure techniques required for most body areas. This means the optical coupling method discussed below is the only viable option for general radiography.

Charged coupled devices optically coupled by a mirror and high quality lens The light output from the scintillator is bent through 90° by a high quality mirror, it is then reduced in size and focused by a very high quality optical lens on to the 4 cm CCD array (Figure 6.18). This is potentially a very effective system, but it does suffer from a few disadvantages. One issue relates to the size of optics and mirror which require a relatively large space within the X-ray equipment, meaning it is only possible to incorporate this type of technology into specifically designed X-ray couches and upright imaging systems. 114

Indirect, Direct, Computed and Digital Radiography X-rays Scintillator C rystal

OPTICAL LENS

Light is focused onto the smaller surface of the C C D by the lens

Light

Mirror

Figure 6.18 Charged couple device (CCD) system and optical mirror.

Other issues are related to demagnification and optical main­ tenance. Demagnification is inherent in the system design and is basically the effect of taking a relatively large scintillator light output and minimising it to fit the size of the CCD and then en­ larging it again to view on a workstation, which subsequently re­ duces image quality. The other issue of optical maintenance is related to anything that degrades the effectiveness of the optics, such as dust or alignment problems interfering with the fidelity of the transmission of light.

Direct digital radiography One system that serves as an example is provided in Figure 6.19. These detectors work in a similar way to ionisation chambers. When incident radiation passes into the sensitive volume, it causes electrons to 115

Radiation Detection and Image Formation X-rays H igh voltage Top electrode + –

---

+ +- ++ ++

C harge

+

+-

Amorphous selenium C harge electrode

+

in-film transistor Storage capacitor Glass substrate

Figure 6.19 Direct digital radiography (DDR) system.

be liberated from their orbits forming positive and negative ions to carry the charge from one electrode to the other. These are attracted to their respective electrodes creating a current which the image acquisition system, in contrast to IDR, converts the X-rays to an electrical signal without the need for first converting it to light. With solid-state semiconductor materials, the incident radiation produces electrons and holes in pairs that carry the charge. The top surface of the imaging system is made of a thin metal coating and forms one of the electrodes. X-rays pass through this first layer quite easily into the layer below the dielectric layer (electrical insulator), again relatively unhindered. They then pass into the next layer, which is a solid-state semi-conductor material where the inter­ actions take place. The X-rays interact with the atoms in the semiconductor layer and form electron hole pairs in proportion to the quantity and quality of incident X-rays falling on it. The last layer is a TFT array that also forms the other electrode. This TFT electrode is not a single area, but is actually composed of many minute electrodes formed by the sensitive areas of every del in the TFT which directly form the image matrix. The electrodes of the face plate and those forming the TFT have an opposite charge applied to them in the order of a few kilovolts, so one will be positively charged and the other negatively charged. It does not necessarily matter which of the plates is positively charged and which is 116

Indirect, Direct, Computed and Digital Radiography

negatively charged, but different polarities do cause the systems to have slightly different properties which are exploited differently by different manufacturers. It is beyond the scope of this book to explore this in more detail. Electron hole pair refers to the electron and the remaining positive atom which both have an equal and opposite charge Following exposure, electron hole pairs are created in the semiconductor layer by interactions of the X-ray beam. Due to the high kilovoltage, the electrons immediately flow towards the positive elec­ trode while the positive atom flows towards the negative electrode. The high kilovoltage also inhibits the recombination of the electrons and holes. The electrodes are carefully aligned making use of the high kV to create field lines that will funnel the ions pairs directly towards their respective electrode plates with very little lateral spread of information. This ensures few if any ions fall outside the del on to the interspace (or dead space created by the associated electronics, the gates and drain lines), resulting in excellent spatial resolution. The electrodes on the underneath of the semi-conductor layer are connected directly to or from part of a matching TFT array which essentially has the same role as with IDR using TFT technology. As with IDR technology, the size of the dels in these TFT systems is again the main limiting factor in resolution. However, in addition to this being limited by the dead space, the minute electrodes in this appli­ cation of TFT technology also have to be a certain size in order to cope with a relatively high kV. The main semi-conductor currently in use with these systems is amorphous selenium (a-Se). This substance can have issues where the charges become trapped at the electrode interface making it difficult to fully clear these charges before the next exposure resulting in remnants of the previous exposure affecting the latest image. Some systems perform detector ‘relaxation’ following an exposure to release trapped charges within the substrate. This phenomenon gets worse with time as the a-Se naturally tries to crystallise with ever increasing effects on its semi-conducting properties. It is worth noting that newer manufacturing techniques, such as doping a-Se with arsenic, have reduced these effects significantly. Another substance, amorphous silicon (a-Si), is also being developed and is used in a similar way and exhibits very similar properties. 117

Radiation Detection and Image Formation

So which is the better system: CR/DR or Dl? The main supporters of CR would argue it has the greatest versatility of any system and can be used with unmodified equipment as well as being available in different cassette sizes, while proponents of DR argue for the speed of the system. In terms of resolution, it is possible to alter this with CR, but most current research suggests that DR has the advantage. Recent studies also suggest that if we require resolutions less than 200 μm, then a-Se performs better, but for resolutions of more than 200 μm, then a-Si panels may perform better.

DIGITAL FLUOROSCOPIC SYSTEMS Image intensifier linked to charged coupled device Early digital fluoroscopy systems tended to revolve around an analogue image intensifier (II) linked to a CCD camera. The CCD output was converted to a digital output via an analogue to digital converter. Images need to be produced at a rate of 30 or more frames per second in order to produce a smooth real-time moving image. As CCDs use fast serial data collection together with extremely fast transmission speeds, these frame rates are easily achieved. CCDs also benefit from extremely low inherent noise and can produce good image quality even from the relatively low exposure used during the fluoroscopy mode, as well as extra low-dose pulse techniques. For these reasons, CCD using an image intensifier became the dominant system and this type of technology is still widespread in clinical use. Even so, there are some issues: it requires a large housing and its image is distorted to some degree by the signal amplification that takes place inside the image intensifier. As a result of these inherent limitations, this technology is now being replaced by the flat panel technologies.

Fluoroscopic flat panel detectors Initially, there were issues associated with using flat panel technology to produce real-time images. The systems suffered from lag and slow 118

Digital Fluoroscopic Systems

refresh rates, which although not an issue for still radiographic images, is a real issue in fluoroscopy mode. The system must collect all the signals from all detector elements, for each frame, within a thirtieth of a second; this is extremely difficult to achieve and places high demands on the switching characteristics of the components that make up the TFTs, as well as the speed of the charge amplifiers and digitisers of the output stage. The result is that early systems were not able to achieve the 30+ frames per second required for smooth motion realtime imaging. The second issue relates to radiation dose during fluoroscopic ex­ aminations. While flat panel detectors are comparable in terms of dose with other systems for still images they initially gave relatively high doses when used for moving fluoroscopic images. Traditional fluoro­ scopic systems including digital image intensifier systems using CCD technology could produce images of good quality using relatively low mA. Image quality was still acceptable with even more aggressive dose reduction techniques where the beam was pulsed during fluoroscopic mode. This allowed a significant lowering of the overall dose the patient received for an examination. Unfortunately, early flat panel technology did not respond very well to the low quantity of incident radiation during standard fluoroscopy mode and consequently suffered very poor signal-to-noise ratios. The signal-to-noise situation was even worse with pulsed low-dose techniques. Recent developments in flat panel technology using modified for­ mulations of scintillator and photodiode materials, as well as in­ troducing overcharging protection circuits, help in significantly reducing lag, speeding up refresh rates, as well as responding better to low-dose fluoroscopic techniques, enabling much better overall image quality. They do not suffer from geometric distortion, such as pincushion and S-shaped distortion, offering excellent image uni­ formity. Current systems also benefit from an image area of up to 43 cm2 and are therefore able to cover the commonly used techni­ ques and examinations of the entire body, something early systems also struggled with. This has led to more widespread uptake of flat panel technology for fluoroscopic purposes. One of the biggest de­ velopments though, relates to improvements in image processing. There has been much research revolving around software 119

Radiation Detection and Image Formation

enhancements and how images are processed. Using techniques such as artificial intelligence, model based and iterative image pro-cessing, usable images are possible with very low signal/dose techniques whilst effectively managing relatively high noise associated with these techniques. Please see Chapter 9 where image processing techniques are dis-cussed in much more detail.

Solid-state X-ray image intensifier There has also been much research into the technology of the solid-state X-ray image intensifier (SSXII) which is based on a technology called electron multiplying CCD (EMCCD). SSXII is essentially a series of modified CCD arrays which are butted together forming a much larger array (43 cm2). The modifications in­ clude an on-chip amplifier which removes the need for a traditional image intensifier. A CsI scintillator is still used with a traditional ana­ logue image intensifier, but instead of the light being amplified by the intensifier it is used directly by the EMCCD. The system is very similar in design to the fibreoptic coupled CCD systems used in slot scan techniques discussed earlier in this chapter. The light output of the CsI scintillator in response to an exposure is coupled to the EMCCD. By using optical fibres, this system has a much higher resolution than the three line pairs per millimetre (lp/mm) available with both the CCDII and flat panel detectors (FPD) and is able to provide an effective pixel size of 32 μm. It has all the advantages of the flat panel systems and also benefits from no lag or ghosting. Signal to noise ratios are excep­ tional; all CCDs benefit from extremely low noise levels, but in addition EMCCDs also benefit from a built-in amplifier, enabling them to detect tiny signals.

Reference Lança L, Silva A. Digital radiography detectors – a technical overview: part 2. Radiography 2009; 15: 134–138.

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Mcqs

MCQs 1. The DQE is: a. Essentially based on its overall signal-to-noise ratio (SNR) when compared against a theoretical ideal detector b. Not affected by spatial frequency c. Not a variable as signal values change in response to changes in exposure d. Essentially based on the modular transfer function (MTF) when compared against a theoretical ideal detector. 2. Indirect systems: a. May be CR or IDR b. Are just CR c. Are just IDR d. May be CR, IDR or DDR. 3. The following crystals: a. They have scintillation b. They have scintillation c. They have scintillation d. They have scintillation

are characteristics of structured scintillation high output in crystals and light low output in crystals and light low output in crystals and light high output in crystals and light

comparison to unstructured is in a more forward direction comparison to unstructured is in a more forward direction comparison to unstructured is in a less forward direction comparison to unstructured is in a less forward direction.

4. TFT systems work better with: a. Unstructured crystals due to their higher light output b. Structured crystals as the light output is more directional c. Unstructured crystals as the light output is more directional d. Unstructured crystals due to their lower light output.

121

Radiation Detection and Image Formation

5. TFT arrays are: a. Usually connected to the scintillator crystals by optical fibres b. Optically connected to the scintillator by a lens and mirror system c. Directly connected to the back of the scintillator d. Attached to the output phosphor of a digital fluoroscopy image intensifier 6. Which of the following materials is used in systems that directly convert X-rays to electrical signal without the intermediate conversion to light? a. Gadolinium oxisulphide (Gd2O2S) b. Caesium iodide (CsI) c. Amorphous selenium (a-Se) d. Barium fluoro-halide:europium activated/doped (Ba F Brx I 1–x:Eu). 7. The electrodes in a DDR system are formed by the face plate and those forming the TFT. Why do they have an opposite charge applied to them in the order of a few kilovolts, so one will be positively charged and the other negatively charged? a. To create field lines in which the light can travel between electrodes b. To create field lines which inhibit the recombination of the electrons and holes c. To create field lines to increase signal, but they also create a small amount of interference which causes the signal to spread d. To create field lines in which the light can travel to the electrodes. 8. Define the fill factor as related to the TFT array. a. This is the ratio of the thickness of the sensitive layer in comparison to the whole thickness of the array b. This is the proportion of electronic circuits to the thickness of the sensitive volume c. Ratio of the number of live detector elements against the number of dead detector element in an array d. Is the proportion of sensitive area (charge collection area) against the dead areas of the array which includes the tiny electronic circuits. 122

Mcqs

9. What is the minimum number of frames in order to produce a smooth real-time moving image? a. 30 frames per second b. 10 frames per second c. 20 frames per second d. 50 frames per second. 10. Early designs of flat panel technology used to produce real-time images were associated with: a. Exceptional temporal resolution, but relatively low signal output b. Exceptional temporal resolution, but required relatively high exposures c. Lag and slow refresh rates d. Poor temporal resolution, but benefited from relatively low exposure requirements.

123

CHAPTER 7 IMAGE QUALITY INTRODUCTION It is essential that any practitioner understands the principles involved in producing and assessing diagnostic images. Images must be produced with the lowest radiation dose consistent with diagnostic quality (ALARP [as low as reasonably practicable] principle). The practitioner therefore needs to understand how to adjust the factors affecting image quality to ensure the images answer the diagnostic question. Image quality is subjective (it depends on the skills of the observer) and may be difficult to define, however, an optimum quality image enables the observer to make an accurate diagnosis. Poor quality images are easier to define as they have a poor signal-to-noise ratio, poor spatial resolution and detract from the process of extracting information. There are characteristics of an image which may be evaluated and this enables the practitioner to determine the diagnostic quality of an image. These characteristics include: ◾ The positioning of the patient, X-ray beam and detector ◾ Collimating and centring of the beam to the area of interest ◾ Precise patient positioning with the area under examination parallel to the detector ◾ Ensuring the patient is comfortable and still to minimise movement. ◾ The data acquired by the detector ◾ Quantity and quality of photons which pass through the patient without attenuation (brightness and contrast) ◾ Scattered photons (noise) ◾ The display system used to view the image ◾ Monitor size and matrix size 125

Image Quality ◾ Software processing applied to the raw data ◾ Viewing conditions (background illumination)

Learning objectives The student should be able to: ◾ Understand and explain the principles of producing and assessing images for their image quality. ◾ Explain the factors which contribute to radiation dose and image quality.

GEOMETRY OF IMAGING All radiographic images are larger than the object being X-rayed. This mag­ nification is due to the geometry of imaging. The ideal situation is to have: ◾ The object being imaged parallel to the X-ray beam and the image receptor ◾ The radiation beam at right angles to the object ◾ A long focus to receptor distance and small object to receptor distance This minimises the distortion of the image and the magnification of the unsharpness in the image. The ideal conditions to produce radiographic images are shown in Figure 7.1. X-ray tube

Body part Image receptor Resultant image Figure 7.1 Ideal positioning for X-ray imaging.

126

Geometry of Imaging

The object should be as close to the image receptor as possible. As the object moves away from the image receptor the magnification increases. This makes the object bigger but also magnifies any unsharpness in the image. The positioning of the patient (geometry) to produce the image has a direct relationship on the quality of that image. Figure 1.2 in Chapter 1, Overview of image production, is a diagram­ matic representation of image production and shows that a penumbra (unsharpness) is formed with any image that is produced from a finite source (focal spot). The diagram uses a large distance between the object and image receptor to illustrate the principle of penumbra. In practice, the amount of geometric unsharpness (Ug) in an image is small. The point at which we begin to perceive unsharpness in an otherwise optimum image is approximately 0.3–0.4 mm. Measurement of the penumbra (Ug) is a straightforward calculation using similar triangles. Figure 1.1 in Chapter 1, for example, demon­ strates the diagrammatic representation of similar triangles. It is possible to calculate the unsharpness in an image of a finger due to geometric unsharpness. The formula is:

Ug =

Focal spot size × ORD FOD

To calculate the penumbra you need to know all the factors on the right side of the equation. For example, if the: ◾ Focal spot size is 0.3 mm; ◾ The focus receptor distance (FRD) is 110 cm; ◾ The object receptor distance (ORD) is 1 cm; ◾ This makes the FOD 109 cms. The unsharpness is calculated as only 0.003 mm, which is negligible and looks sharp to the observer. The values used here are typical in radiography. When undertaking an X-ray of the lumbar spine, the geometric unsharpness is much larger. For example, if the: ◾ Focal spot size is 1 mm. ◾ FRD is 110 cm. ◾ FOD is 80 cm. The unsharpness is calculated at 0.37 mm, which is approaching an image which the observer would perceive as blurred. Other factors, 127

Image Quality

such as movement, and the resolution of the monitor will also increase this level of unsharpness.

MAGNIFICATION AND DISTORTION If the object being imaged is not parallel to the image receptor, it will be magnified, however, different aspects will be magnified differently and this will produce distortion. This may be elongation or foreshortening of the image. Figure 7.2 shows the set up for producing a distorted image and Figure 7.3 shows a deliberately elongated image of the scaphoid which aids in the diagnosis of a fracture.

X-ray tube

Central ray

FRD

Body part

{ORD}

Image receptor

Resultant image is distorted Figure 7.2 Set-up with variable ORD, which will give a distorted image.

The distance between the patient and the image receptor (ORD) should be as short as possible. For practical reasons, the FRD is usually 110 cm for techniques on the X-ray detector and 180 cm for erect chest and cervical spine work. If possible, the object is in contact with the image receptor, however, using a Bucky assembly increases magnifica­ tion of the image, but may be necessary to reduce scatter and improve 128

Signal-to-noise ratio

Figure 7.3 Elongated scaphoid projection.

the contrast of the image. Practically, the mechanism which moves the grid and houses the Potter–Bucky is kept as small as possible.

SIGNAL-TO-NOISE RATIO Image quality may be assessed by the signal-to-noise ratio. The signal is the useful information from the patient being imaged and noise is anything which detracts from accessing the information. Useful information is derived from photoelectric interactions within the patient (absorption by dense body structures and no minimal absorption in air or instead of low density structures). Noise is derived from Compton scatter. The image 129

Image Quality

receptor does not have the ability to determine the origin of scattered photons and there is also electrical noise from the system. Radiographic images that have a signal level which is high compared to the noise will enable structures to be clearly seen, but if the signal level is similar to or less than the noise level the structure will become obliterated. Images produced by X-rays are often a compromise between ob­ taining a perfect signal and reducing the noise. The production of images is constrained by a number of factors including the radiation dose level and each component of the imaging chain, e.g. when imaging radiosensitive areas gonads, areas with red bone marrow, there is a constraint of minimising the radiation dose to that area. Extremities of the human body are less radiosensitive and images with higher dose and better definition may be reasonably acquired. Use of a grid will also enhance contrast by removing scattered pho­ tons before they reaches the detector. The increased contrast by the use of a grid must be justified by the value of the practitioner, as using a grid will increase the dose proportionally by the grid factor.

UNSHARPNESS Unsharpness of an image is related to: ◾ The geometry of the imaging system ◾ Focal spot size ◾ Relationship (distance) between, focus, patient and detector. ◾ Intrinsic sharpness of the detector employed ◾ Subject contrast ◾ The quality and resolution of imaging system (film screen combination or video display unit [VDU]) ◾ Beam quality (kV) ◾ Scatter (obscures the ability to observe sharpness) ◾ Movement unsharpness (usually patient movement) In order to determine the image quality the image must be reasonably sharp. Blurring will reduce the image quality and also reduce the diagnostic quality of the image. The sharpness of a system is best characterised in terms of its modulation transfer function (MTF).

130

Spatial resolution

MOVEMENT UNSHARPNESS Movement unsharpness should be kept to a minimum by careful radiographic technique. Making the patient comfortable, giving them precise instructions on breath-holding and practising the procedure may all help reduce movement unsharpness. Using short exposure times also helps reduce blur in the image. Exposure times as low as 0.001 s for a chest radiograph enables the practitioner to ensure minimum movement even in paediatrics.

RESOLUTION OF THE IMAGING SYSTEM This will be covered in more detail in other chapters of this book. However, it should be noted that the smallest exposure compatible with diagnostic quality should be used. When determining the quality required in an image, the practitioner must be aware of what structures they need to define. An image to determine the position of bones in a plaster cast following an orthopaedic reduction needs less resolution than the original image to diagnose the fracture. Ideally, the smaller of the foci of the X-ray tube should be used (fine focus), but if this does not enable a short exposure time to be used on a patient likely to move, the practitioner may need to use a broad focus. This is another example of where the practitioner needs to make a decision which is a com­ promise between the ideal conditions and getting a diagnostic image.

SPATIAL RESOLUTION Spatial resolution refers to the ability of the imaging system to represent distinct anatomic features within the object being imaged. It may be defined as the ability of the system to distinguish neighbouring features of an image from each other and is related to sharpness. The maximum spatial resolution of an image is defined by pixel size and spacing. Pixel 131

Image Quality

size affects the system resolution and varies between systems. Detective quantum efficiency (DQE) combines spatial resolution (MTF) and image noise to provide a measure of the signal-to-noise ratio. The sharpness of an imaging detector or system is best characterised in terms of its MTF.

MEASUREMENT OF UNSHARPNESS IN AN IMAGE The sharpness of the image can be measured using a test tool (resolution) or by viewing the image and determining if fine structures can be visualized (definition). The resolution can be determined by measuring the ability of the imaging system to resolve the smallest object in the test tool and is expressed in line pairs per millimetre (lp/mm). This is perhaps an abstract concept to illustrate sharpness and it would be more sensible to determine the definition within the image, e.g. can I see the bony trabecular? Again however, definition within the image is subjective and depends on a number of factors including the resolution of the monitor, viewing conditions and the experience of the practitioner viewing the image.

VIEWING DIGITAL IMAGES All digital images have a look-up table (LUT) applied to the raw data. The computer software enhances the image and displays the data with an ‘optimum’ brightness and contrast.

BRIGHTNESS AND CONTRAST The brightness may be defined as the intensity of light that represents the individual pixels in the image. Brightness is controlled by the pro­ cessing software and can be adjusted following image processing. Digital systems generally have a linear response to exposure and a wide dynamic range. All digital images are ‘auto-windowed’ so the 132

Brightness and contrast

Optimum exposure

Overexposed 40× optimum

Underexposed 1/10th optimum

(a)

(b)

(c)

Figure 7.4 Underexposed image (a); optimum image (b); and overexposed image (c).

computer will automatically apply an algorithm to the data detected and provide the ‘best’ range of densities (contrast) and brightness. Overand underexposure is therefore not visually apparent and other means are necessary to determine if the optimum exposure has been used. These checks include: ◾ Evaluation of the exposure indicator (EI). The number allocated to the image should be in the range expected by the system, e.g. ◾ Computed radiography (CR) system acceptable range is given as 1700–2300. The ideal exposure is 2000 (may vary slightly for different body parts). ◾ Digital radiography (DR) system range is given as 200–800. The ideal exposure is 400. ◾ Evaluation of noise within the image due to underexposure (Figure 7.4a). ◾ Evaluation of ‘burn through’ due to gross overexposure (Figure 7.4c). If the EI is consistently higher than recommended, the patient is being overexposed and receiving to high a radiation dose. Figure 7.4 shows a range of images from a CR system. The images range from an optimum quality image (Figure 7.4b), an underexposed image (Figure 7.4a) and a grossly overexposed image (Figure 7.4c) as measured using the EI. 133

Image Quality

1 mAs 50 kV

1 mAs 100 kV

Figure 7.5 Images at 50 and 100 kV auto windowed by computer system.

Optimum images can be obtained from a wide range of exposures and Figures 7.4a and 7.4c show the extremes where the computer failed to adjust the image into the diagnostic range due to the extremes of over-/underexposure. Contrast may be defined as a measure of the relative brightness dif­ ference between two locations (area or pixels) in an image. The contrast of an imaging system is described by the characteristic response curve of the system. The X-ray beam is attenuated in the patient depending on the energy of the X-ray beam (quality) and the total exposure (mA and kV). These factors were critical using film screen technology, however, when the image is ‘auto-windowed’. In digital systems, these factors are less important and images taken at 60 and 110 kV may have similar brightness and contrast levels when displayed and viewed on a monitor (especially with the mA adjusted to compensate for the change in kV). Figure 7.5 demonstrates images with differing kilovoltages autowindowed by CR system to produce similar contrast and brightness. Note: If a range of exposures exists for producing a diagnostic image, the practitioner should always select the total exposure with the lowest radiation dose. When imaging variations within the human body, there is a natural subject contrast. This varies depending on thickness of the patient and the area of the body, i.e. 134

Effect of scatter on contrast ◾ The thorax has high inherent contrast because adjacent areas

have large differences in density of structures (dense heart muscle surrounded by air). ◾ The abdomen has low inherent contrast with adjacent areas having a similar atomic number (muscle, fat and organs). The subject contrast cannot be altered, however, the auto-windowing can modify the image to improve the brightness and contrast before the image is displayed and the practitioner can also ‘window’ the image during post-processing. Both these processes may be able to ‘create’ a diagnostic image from poor raw data. Generally, if the image appears ‘optimum’ for both 50 & 100 kV, the high kV should be used as it delivers a lower radiation dose.

EFFECT OF SCATTER ON CONTRAST Scattered radiation reaching the detects reduces the contrast of any image as it does not carry useful information, but does create a signal on the detector. The effect can be minimised by: ◾ Reducing the production of scatter ◾ Close collimation to the area of interest ◾ Displacement of the body part (used in mammography) ◾ Preventing the scatter reaching the detector ◾ Using a grid or Bucky assembly ◾ Using an air gap Scattered radiation should therefore be minimised by the practitioner.

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MCQs 1. You are producing an image of the spine. The focus receptor distance (FRD) is 100 cm and the spine is 20 cm from the imaging plane. The spine is 40 cm long. What is the length of the spine in the image? a. 50 cm b. 5 cm c. 44 cm d. 40 cm. 2. Which of the following factors do not directly affect the geometric sharpness of an image? a. Focus receptor distance (FRD) b. Focal spot size c. Scattered radiation d. Object receptor distance (ORD). 3. Which of the following factors produces an optimum image? a. Longest FRD practicable b. Smallest focus available c. Close contact between the patient and detector d. All of the above. 4. Calculate the penumbra if focus is 0.2 × 0.2 mm, the FRD is 110 and ORD is 2 cm. a. 4.2 mm b. 0.42 mm c. 0.042 mm d. 0.0042 mm. 5. Optimum images have: a. High signal-to-noise ratio b. Low signal-to-noise ratio c. Equal signal-to-noise ratio d. High electrical noise.

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MCQs

6. Which of the following arrangements produces an image with minimal unsharpness? a. Short FRD and long ORD b. Short FRD and short ORD c. Long FRD and short ORD d. Object parallel to the detector. 7. Which of the following factors cause noise in an image? a. Photoelectric absorption b. Transmitted photons c. Compton scatter d. Using a grid. 8. Spatial resolution is the ability of the system to: a. Minimize radiation dose b. Represent distinct anatomical detail c. Reduce blurring d. Reduce noise. 9. Which anatomical area has the highest inherent contrast? a. Thorax b. Abdomen c. Hand d. Pelvis. 10. If a a. b. c. d.

range of exposures produces a diagnostic image you should: Use the lowest kV Use the highest kV Use the exposure which gives the lowest radiation dose Use the lowest mA.

137

CHAPTER 8 RADIATION DOSE AND EXPOSURE INDICATORS INTRODUCTION The radiation exposure from diagnostic X-rays is the largest manmade source of radiation exposure to the general population worldwide. Man-made exposure contributes about 14% of the total annual exposure from all sources. Although diagnostic X-rays pro­ vide great benefits, their use involves some small risk of developing cancer. The aim of this chapter is to give the practitioner an under­ standing of the basic ways radiation dose can be measured and ex­ pressed to the patient and other staff. It will also explain exposure indicators used with digital systems. It is essential that any practi­ tioner operating within an imaging department and using ionising radiation has a sound base for their knowledge. You need to com­ prehend and be able to explain the factors affecting radiation dose, methods of measuring dose and how dose may be expressed using exposure indicators. Learning objectives The student should be able to: ◾ State and explain the factors affecting radiation dose and the methods of measuring dose ◾ Understand and explain the principles of exposure indicators

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RADIATION DOSE Radiation may be measured and expressed in several ways. These ac­ count for the different sensitivities of the body tissues. For the purposes of this book, we will only consider X-rays and γ-radiation which have identical properties.

Detection and measurement of radiation There are several methods used to detect and measure radiation. Some are more sensitive and accurate than others, but they all work on the principle of ionisation of a material and then measuring the effect. The methods include: ◾ Direct ionisation ◾ An ionisation chamber ◾ Free air ionisation chamber ◾ Thimble dosimeter ◾ Dose area product (DAP) metres ◾ Automatic exposure devices (AED) used to terminate exposures rather than measure the dose ◾ Geiger–Muller tube ◾ Gold leaf electroscope (Quartz fibre dosimeter) ◾ Luminescence ◾ Photostimulated luminescence (PSL) resulting in the emission of blue light in an amount proportional to the original X-ray irradiation ◾ Phosphors in screens used with film ◾ Scintillation ◾ Scintillation probe and gamma camera ◾ Optical stimulated luminescence dosimeter (OSL) ◾ Thermoluminescenct dosimeters (TLD) ◾ Dosimeters for whole body doses and separate monitors for small body parts ◾ Semi-conductor detection ◾ Gamma camera ◾ Flat panel detectors ◾ Fluoroscopy units 140

Exposure

Ionisation of air Air in its normal state is considered to be a good electrical insulator as it does not contain any conduction electrons. However, if air is exposed to X- or γ-radiation, some of the photons of radiation will release electrons from the atoms in the air. This causes ionisation enabling the air to conduct electricity. The more radiation the air is exposed to, the more electrons emitted and the better able it is to conduct electric current. By measuring the electrons in the sample of air, the quantity of radiation causing the ionisation may be calculated/estimated. The ionisation in air by the beam of radiation is proportional to the energy absorbed. The energy required to produce one ionisation in air is approximately 33 electron volts (eV). Therefore, if we have a homogenous beam X-rays at 60 kVp, this will produce about 1800 electrons. (This is a tiny amount compared to an electrical current of 1 picoamp, which is about 1 million electrons flowing through the conductor per second.) Atomic numbers of air (7.6) and soft tissue (muscle) (7.4) are similar and therefore absorb about the same amount of energy. This means the mass absorption coefficients of air and soft tissue are very similar and by measuring the value in air, the value in tissue can be accurately estimated.

EXPOSURE The traditional method of measuring the amount of ionisation in air was exposure. This measured the ratio of the total charge produced (of one sign, usually electrons) in a small volume of air. The unit of exposure for air is Coulombs per kilogram of air (C/kg). The unit of exposure only applies to X- and γ-radiation. The intensity (exposure rate) of a beam of X-rays can also be measured as the energy passing through unit area in unit time.

SI Units There is an important distinction between the absorbed dose Gray (Gy) in matter usually air and the sievert (Sv) which is the absorbed dose in Grays multiplied by a quality factor to represent the health risk. 141

Radiation Dose and Exposure Indicators

There are three SI units to remember: ◾ Absorbed dose Gray. Defined as: one Gray deposits 1 joule of energy per kilogram of matter. It is used as a measure of absorbed dose, specific energy (imparted), and kerma (an acronym for kinetic energy released per unit mass). The Gray is defined independently of any target material. However, when measuring kerma the reference target material must be defined explicitly, usually as dry air at standard temperature and pressure. ◾ Equivalent dose is determined by multiplying the absorbed dose by a quality factor depending on the type of radiation and related to human tissue. ◾ Effective dose is determined by equivalent dose multiplied by the sensitivity of the tissue irradiated.

Absorbed dose The Gray (Gy) is the SI unit of absorbed dose and is equivalent to the absorption of one joule of energy in a kilogram of a substance by ionising radiation. For X-rays and gamma radiation, the Gray equates to the Sievert (equivalent dose) as the quality factor (QF) for X- and γ-radiation is one. The Gray is a large unit and, for normal radiation protection pur­ poses, it is more common to use subunits like the: ◾ Microgray (μGy) is one millionth of a Gray (1 × 10−6) ◾ Milligray (mGy) is one thousandth of a Gray (1 × 10−3) ◾ Centigray (cGy) is one hundredth of a Gray (1 × 10−2) Kerma dose is different from absorbed dose especially at high en­ ergies (up to 1 MeV) and roughly equal at low energies. The unit for kerma is joule per kilogram (Gy), which is the same as for absorbed dose. When a photon beam interacts with a medium, the photon interactions release electrons with kinetic energy into the medium (kerma) which then move on to deposit energy along ionisation tracks. The energy deposited by these electrons per unit mass is the absorbed dose. So ‘kerma’ is energy released and ‘absorbed dose’ is energy absorbed. The total absorbed energy delivered by the sec­ ondary electrons produced by a photon beam is the ‘integral dose’. The integral dose attempts to describe energy deposition within the whole body. 142

Exposure

When determining the absorbed dose in tissue you need to know the effective atomic number of tissues. Tissues contain a mixture of ele­ ments so their absorption depends on the sum of the elements and their density. The 2 most common effective atomic numbers (Zeff) used in dosimetry are: ◾ Soft tissue Zeff is 7.4 ◾ Bone Zeff is 13.8.

Equivalent dose Equivalent dose allows the effect of radiation exposure on human tissue to be determined. It relates the absorbed dose in human tissue to the effective biological damage of the radiation. Not all radiation has the same biological effect, even for the same amount of ab­ sorbed dose. The SI unit of equivalent dose is the Sievert (Sv) and represents the stochastic biological effect. The Sievert is a large unit and for normal radiation protection levels a series of pre-fixes are used: ◾ Microsievert (μSv) is one millionth of a Sievert (1 × 10−6) ◾ Millisievert (mSv) is one thousandth of a Sievert (1 × 10−3) To determine equivalent dose (Sv), you multiply absorbed dose (Gy) by a radiation weighting factor (WR) that is unique to the type of radiation. The WR takes into account that some kinds of radiation are inherently more dangerous to biological tissue, even if their ‘energy deposition’ levels are the same. For X-rays and gamma radiation, and electrons absorbed by human tissue, the WR is 1. To determine the dose in Sieverts from the dose in Grays, simply multiply by the WR. This is obviously a simplification. The WR ap­ proximates what otherwise would be a very complicated calculation. The values for WR change periodically as new research refines the risks associated with radiation exposure. Exposure occurs over time, of course. The more Sieverts absorbed in a unit of time, the more intense the exposure. So we express exposure as an amount over a specific time period, e.g. 5 mSv per year. This is called the ‘dosage rate’. In the UK, the dose rate from background radiation, the sum of all natural radiation, is about 2.5 mSv per year. A table of radiation weighting factors is in Figure 8.1 143

Radiation Dose and Exposure Indicators Radiation type X-radiation g-radiation Beta particles Fast neutrons Alpha particles

Quality factor WR 1

20 20

Figure 8.1 Radiation weighting factors WR.

Effective dose The probability of a harmful effect from radiation exposure depends on the part or parts of the body exposed. Some organs/tissues are more sensitive to radiation than others. A tissue weighting factor (WT) is used to take this into account. When an equivalent dose to an organ is multiplied by the tissue weighting factor for that organ, the result is the effective dose to that organ. If more than one organ is exposed, then the effective dose, EfD, is the sum of the effective doses to all exposed organs (Figure 8.2). Organ or tissue Gonads Red bone marrow Colon Lungs Stomach Bladder Breast Liver Oesophagus Thyroid Skin Bone surfaces Remainder Sum Figure 8.2 Tissue weighting factors.

144

Weighting factor WT 0.2 0.12 0.12 0.12 0.12 0.05 0.05 0.05 0.05 0.05 0.01 0.01 0.05 1

Exposure

Therefore whole body scan like CT and PET have a tissue weighting factor of 1 ie the sum of all the tissue weighting factors is 1. Using the effective dose provides a dose relevant to the whole body, This can then be used to compare the effective dose for different radiological examinations and natural sources. It is used to measure the stochastic risk of non-uniform radiation Figure 8.3 demonstrates dose quantities and their relationship

Dose Quantities Absorbed dose Energy “deposited” in a kilogram of a substance by the radiation Equivalent dose Absorbed dose weighted for harmful effects of different radiations (radiation weighting factor wR) Effective dose Equivalent dose weighted for susceptibility to harm of different tissues (tissue weighting factor wT)

Figure 8.3 Dose quantities and their relationship.

The effective dose can be calculated using the following equation: EfD = D × WR × WT Key: D: Absorbed dose WR: Radiation weighting factor WT: Tissue weighting factor Typical effective doses for X-ray examinations are in Figure 8.4

145

Radiation Dose and Exposure Indicators Examination Dental X-ray Chest X-ray CT scan of the head CT scan of the chest CT scan of the whole spine

Effective dose (mSv) 0.005 0.02 1.4 6.6 10

Figure 8.4 Typical effective doses for X-ray examinations.

Example of a dose calculation for a radiographic X-ray examination:

Lumbar spine If the absorbed Dose to skin is 6.4 mGy ◾ Effective dose. ◾ Calculated from various tissue weighting factors and organ absorbed doses in field of view of exam. ◾ Absorbed dose (6.4) × WR X-rays (1) × WT (Colon 0.12 + ovaries 0.2 + skin 0.01 = (0.33) ◾ 6.4 × 1 × 0.33

Which is equal to 2.1 mSv CT dose and DRL’s Absorbed dose in gray’s is defined as the energy absorbed by tissue per unit mass. The radiation output from a CT procedure is described by using the volume CT dose index (CTDIvol) and the dose length product (DLP). The CTDIvol is calculated using standard phantoms and is an estimate of the average X-ray output for any given CT examination. The DLP is the total X-ray output integrated throughout the entire ex­ amination and is calculated by multiplying the CTDIvol by the scanned length. This calculation does not take a patient’s body habitus into ac­ count therefore the size specific dose estimate (SSDE) is calculated by multiplying the CTDIvol by a correction factor based on patient size. Effective dose is a measure of the uniform whole body equivalent dose and is calculated by multiplying the DLP by a conversion factor based on the region of the body that is scanned. CT protocols should be developed in accordance with the ALARP principle considering the 146

Exposure

patient size and age, and must also take into account justification and the clinical question. DRLs are reported by CTDIvol and DLP values for a single CT ex­ amination for a ‘standard’ patient. They provide a diagnostic DRL for different CT examinations. The purpose of DRLs is to establish the optimised dose level to avoid unnecessary high doses. CT departments can compare their local DRLs with the National DRLs to develop protocols that result in optimal examinations at the lowest practicable patient dose (ALARP). Table 8.27c compares common CT examinations accompanied by the associated DRLs. National and local DRLs are listed from a survey of CT examinations in the North West of the UK. Local DRLs must be subject to annual review for each scanner, to ensure ongoing dose optimization.

DRl’s for common CT examinations

Examination Chest Chest-abdomen-pelvis Head

National reference level DLP (mGycm)

Local dose reference level DLP (mGycm)

610 1000 970

508 730 844

Linear energy transfer and relative biological effectiveness When ionising radiation passes through tissue (a medium) it may in­ teract with it and deposit energy along its path of travel. The average energy deposited per unit length is called the linear energy transfer (LET). The energy absorbed in tissue depends on the type of charged ionising particle which is travelling through the medium and the type of medium. LET is measured in kiloelectron volts (KeV) per micron (10−6 m) and is an important factor in assessing potential tissue damage. For diagnostic beams the LET is low. When diagnostic beams interact with tissue it causes damage by the production of free radicals which may cause damage to the DNA. High LET radiation eg alpha particles

147

Radiation Dose and Exposure Indicators

are more destructive to biological tissue as they lose their energy in a shorter length of tissue and are more ionising. Another consideration in dosimetry is the relative biological effectiveness (RBE) of radiation. This describes the relative ability of ionising radiation to produce a biological reaction. RBE is not suitable for determining the biological effects in humans for all types of radiation but is similar for X-rays in the diagnostic range.

Quality factor for radiation The same absorbed doses of different types of radiation cause a different biological damage in tissue. X-rays, gamma radiation and beta particles all produce virtually the same biological effect on tissue for equal absorbed doses and are given a quality factor of 1. A quality factor is used to adjust the dose equivalence of other types of radiation eg alpha particles have a quality factor of 20 and have a much greater effect on tissue.

RADIATION MONITORS AND PERSONAL MONITORING The principles of ionisation chambers are described in Chapter 6, Ionisation monitors may be used for dosimetry. Other types of personal dosimeters used to monitor staff doses are: ◾ Optically stimulated luminescent dosimeter (OSLD), and ◾ Thermoluminescen dosimeters, which are used for whole body doses or small body parts.

Optically stimulated luminescent dosimeter (OSLD) This is the most commonly used device used for monitoring of occu­ pational exposure and has largely replaced film badges and TLD’s. They are normally worn for a period of 1-3 months and either sent to man­ ufacturer or read in-house. They consist of three different filters, alu­ minium (Al), tin (Sn) and copper (Cu). The device allows for energy discrimination as it attenuates radiation proportionally.

148

Radiation monitors and personal monitoring

Higher energy radiation can penetrate filters easier than lower energy radiation. ◾ High energy reading is taken behind all three filters ◾ Low energy reading is behind Al filter and cannot penetrate other filters. Energy ranges classified as (i) deep, (ii) eye and (iii) shallow and gives a correlation between energy and penetration depth. The dosimeter has a wide range of sensitivity ◾ X-ray and gamma radiation has an accurate reading as low as 10 µSv in energy range of 5 keV to 40 MeV. The maximum reading is 10 Sv. ◾ Beta particles has a range of 100 µSv to 10 Sv for energies 150 keV to 10 MeV. ◾ Neutron have a range of 200–250 µSv for energies greater than 40 keV OSLD’s are more sensitive than a TLD so ideal for monitoring employees in low radiation environments and pregnant workers. They are lightweight, durable, self contained and resistant to heat moisture and pressure. They can be used for long periods of time and perform repeat readings.

Thermoluminescent Dosimeters (TLD) The TLD card comprises two elements of lithium fluoride (LiF) mounted in an aluminium frame and has a uniquely numbered barcode that can be read automatically and the radiation dose can be measured. After readout, the card can be annealed to get rid of any residual reading and then reissued. A TLD card can go through this readout/anneal cycle more than 500 times. The badge is comprised of a TLD card (Figure 8.5), which is placed in a holder that incorporates a filter system. This allows the radiation type

Figure 8.5 Thermoluminescence dosimeter badge.

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Radiation Dose and Exposure Indicators

and energy to be determined. It is used to determine the whole body exposure of people who may be exposed to beta-, gamma- or X-rays. It can be worn for 4 to 12 week period wearing period depending on the work carried out and the risk to the operator. The extremity (or finger) TLD is used by anyone who may be exposed to significant doses to the fingers. The monitor consists of a small plastic sachet containing a TLD which can be chemically disinfected if necessary. The doses are determined by the measurement of the light output from the TLD card. Thermoluminescent materials store energy inside their structure when they are irradiated, as electrons and holes are trapped in trapping centres due to crystaline defects. When that ma­ terial is heated, electrons and the positive atom recombine, at lumi­ nescence centres, and thus light is emitted. The light is measured using a PMT (photomultiplier tube) inside the reader device. The photons which are emitted are in the visible region and they comprise the thermoluminescent (TL) signal.

Exposure Indicators This is probably the most complex aspect of producing diagnostic images at the optimum radiation dose (ALARP). Developed concurrently by the International Electrotechnical Commission (IEC) and the American Association of Physicists in Medicine (AAPM), in cooperation with digital radiography system manufacturers, the exposure index has been im­ plemented as an international standard. It’s known as the IEC exposure index. The IEC exposure index is unique to the receptor type being used and to the exam performed. This will allow audit of and between systems to further optimise exposures in diagnostic imaging. Exposure indicators are sometimes known as detector dose indicators (DDI), exposure indicators (EI) are used to provide feedback, in the form of a standard index, to operators of digital radiographic systems. This reflect the adequacy of the exposure that has reached the detector after every exposure event. Due to the very wide dynamic range of digital imaging systems, there is now little visual indication of exposure variation in the images displayed on the monitor. A digital image pro­ cessing technique called ‘histogram auto-ranging’ is used in all systems to stop ‘over- or underexposure’ from changing the lightness or darkness of the image. The International Commission for Radiation Protection 150

Radiation monitors and personal monitoring

(ICRP) decreed that all computed radiography (CR) and digital radio­ graphy (DR) systems must incorporate some form of exposure in­ dicator. Their main fear was that dramatic overexposure might occur without the radiographer being aware of it, if they simply relied on the image as they were used to doing with film/screen systems. The direct connection between the level of detector exposure and optical density is well established in film-screen radiology. This is not the case in digital radiography, where almost always, a constant image characteristic is achieved using automatic image processing. Consequently, deviations from the intended exposure, i.e. over- and underexposure, are not noticeable by a corresponding deviation in image brightness. While considerable under­ exposure results in an increased level of noise, the more alarming aspect (from a radiation protection point of view) is that overexposure cannot be recognised easily in the displayed image. The final brightness of the image is controlled not by the exposure to the detector, but by automatic image processing applied to the acquired raw data. Consequently, overexposed images may not necessarily be dark, and underexposed images may not appear light. This may be a new and confusing concept for operators of digital radiographic systems who are accustomed to screen-film imaging. For more than a decade, the phenomenon of ‘exposure creep’ in photostimulable storage phosphor imaging has been reported. This is attributed to the fact that digital imaging systems can produce adequate image contrast over a much broader range of exposure levels than screen-film imaging systems. Average exposure levels tend to creep up over time if a clear indicator of exposure is not provided and routinely monitored. Techniques required to achieve optimal radiographic ima­ ging in DR may be different from those used for film/screen imaging. This broad dynamic range is one of the benefits of digital detectors. However, if the detector is underexposed higher noise levels may ob­ scure the presence of subtle details in the image. Excessive detector exposures may produce high quality images with improved noise characteristics, but at the expense of increased patient dose. The dynamic range for the chest X-ray is from 1 mAs and 70 kV to 20 mAs and 110 kV with all images of diagnostic quality (Figure 8.6). The American Association of Physicists therefore recommended an exposure index for every image taken and a deviation index which in­ dicates the amount the exposure varies from a designated aim point. 151

Radiation Dose and Exposure Indicators

1 mAs

70 kVp

1 mAs

110 kVp

20 mAs

110 kVp

Figure 8.6 Dynamic range of digital radiography for a chest image.

This will identify: ◾ Under- and over exposure ◾ Exposure distribution and exposure drift ◾ A basis for evaluation of new technology.

An index of detector exposure is appropriate because it reflects the noise content, and thus the signal-to-noise ratio in the image. For DR systems, the appropriate incident exposure is variable based on the de­ sired signal-to-noise ratio rather than on the resulting optical density of a radiograph. Different digital detectors may require more or less radiation exposure to achieve the same noise content depending upon the detec­ tive quantum efficiency (DQE) of the detector technology in use. The concept of an exposure indicator was adopted, however, there are currently over a dozen systems in place. Some manufacturers have different systems for their CR and DR systems. Image evaluation for exposure should perhaps be conducted using EIT ie the EI for a given radiograph and the target exposure. The 3-point scale for producing diagnostic images is: 1 = Unacceptable, must be repeated 2 = Marginal 3 = Acceptable Images should be repeated if relevant anatomical or clinical details cannot be distinguished. It’s the responsibility of the practitioner to select a technique that provides enough exposure to reduce the amount of noise while also adhering to ALARA standards. 152

MCQs

MCQs 1. Which of the following contributes the most to the background dose of radiation to which the population is exposed? a. Flying in a plane b. Radiation from nuclear discharges c. Medical radiation exposure d. Gamma radiation from buildings. 2. The SI unit of absorbed dose is the: a. Milligray b. Gray c. Sievert d. MilliSievert. 3. A device which may be used to measure the X-radiation delivered to a patient is called: a. Densitometer b. Geiger-Muller tube c. Photocathode d. Diamentor. 4. The SI unit of dose equivalent is the: a. Milligray b. Gray c. Sievert d. MilliSievert. 5. An a. b. c. d.

atom which loses an electron: Is called an ion Changes its position in the periodic table Becomes negatively charged Becomes radioactive.

6. One Gray is the absorption of: a. 1 calorie of energy per kilogram of matter b. 1 joule of energy per kilogram of matter

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c. 1 joule per gram of matter d. 1 calorie per gram of matter. 7. The equivalent dose is calculated by: a. Multiplying the absorbed dose by a weighting factor which takes account of the type of radiation being used b. Multiplying the equivalent dose by a weighting factor which takes account of the type of radiation being used c. Multiplying the equivalent dose by a weighting factor which takes account of the radiosensitivity of the tissues being irradiated d. Multiplying the radiation exposure by a weighting factor which takes account of the radiosensitivity of the tissues being irradiated. 8. The quality factor for X-rays is: a. 10 b. 20 c. 1 d. 100. 9. Approximately how many ionisation events will be released by a 70 keV beam: a. 21 b. 210 c. 2100 d. 21,000 10. Which of the following is not an exposure indicator? a. Exposure index b. Sensitivity number c. Log median value d. Exposure mean.

154

CHAPTER 9 IMAGE DISPLAY AND MANIPULATION IN MEDICAL IMAGING INTRODUCTION This aim of this chapter is to explore digital imaging display and manip­ ulation in medical imaging. All the raw data from digital medical images is manipulated by the computer before the practitioner can view them. The chapter explores the various stages in the image production pathway from the formation of the raw data matrix through to the final displayed image matrix. The computer will display an image following certain processes in a format which has been optimized either by the processes themselves or preferences determined by a practitioner. You therefore need to understand and explain relevant processes and the terminology used in the department, how a digital image is displayed on the monitor and how the image can be manipulated to enhance the quality. It finally gives a brief introduction into relevant terms and standards that are widely used within imaging depart­ ments associated with storage and transmission of images and data. Learning objectives The student should be able to: ◾ Explain terms associated with image display and manipulation, such as data matrix, spatial and contrast resolution, image interpolation filter algorithms ◾ Understand and explain the principles of image manipulation, enhancement and optimisation and how the operator can modify image quality 155

Medical Imaging

IMAGE PRODUCTION PATHWAY All digital images are recorded in the form of a grid known as the ‘raw data image matrix’. This grid of raw data is then used to form the pixels of the displayed or reconstructed image matrix. A great deal happens to the raw data before an image is actually displayed. Figure 9.1 lists the main stages involved and diagrams at the beginning and end of the process.

Raw data image matrix This is the grid of fixed actual values collected by the computed radiography (CR)/digital radiography (DR) detector plate following exposure. Consider a chest X-ray recorded with film screen technology which has a resolution of 10 line pairs per millimetre. We have already shown how this would equate to 20 pixels per millimetre if it were a digital system. A standard chest x-ray film screen combination enclosed in a cassette of 43 × 35 cm therefore has the equivalent of a matrix of 8600 × 7000 pixels, a total of 60,200,000 pixels (or 60 mega pixels). In Chapter 6, we also discussed that spatial resolution was not the only parameter involved in image quality and that digital systems whether they are CR or DR based generally have much greater contrast resolution. For general non-specialised radiographs, it is currently accepted that a digital system with a resolution of 10 pixels/mm2 (equivalent to 5 lp/mm) will give similar overall image quality to film screen technology, having double the spatial resolution. A digital system would therefore require a smaller matrix of 4300 × 3500 pixels, a total of 15,050,000 (or 15 mega pixels) to offer arguably similar overall image quality. Every pixel in the raw data matrix will have its own data value based on the incident radiation that interacted with it, effectively meaning the raw data contain a grid reference and signal value for all 15,050,000 pixels.

156

Image Production Pathway Raw data matrix is produced Image interpolation applied Manufacturer applied system filter algorithms Manufacturer algorithms applied as part of pre-set protocols Operator/user-defined manipulation tools (LUT) Reconstructed/displayed image

(a)

(b) Figure 9.1 (a) Stages involved in image manipulation; (b) Raw data and processed image.

Display of reconstructed image matrix The raw matrix is used to form the display image matrix. If all detail is to be displayed on the monitor, then the reconstructed image matrix has to have at least as many pixels as the original raw data matrix. It is also important that a monitor used for image interpretation is able to display the full resolution of the raw data matrix and as such there are typically at least 4300 × 3500 pixel displays. Lower resolution monitors are usually employed to check the technical quality in the viewing areas of most imaging departments. These are typically what is known as a 2K monitor (2048 × 2048 pixels). 157

Medical Imaging

Image quality and matrix size There is a considerable effect on resolution by altering matrix size. If we increase/double the pixels in both directions from say 4 × 4 to 8 × 8, we actually have four times as many pixels. The more we

2 × 2 matrix (4 pixels)

8 × 8 matrix (64 pixels)

4 × 4 matrix (16 pixels)

16 × 16 matrix (256 pixels)

Figure 9.2 The effect of the matrix size on image display.

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Image Manipulation

increase the matrix, the better the representation of the original object. If we consider a random shape which may represent a small body structure, we can see the effect of matrix size and how it will be recorded or displayed (Figure 9.2). Although we can clearly see an improvement in how well the larger matrix represents the original object, we have to be aware that every doubling of matrix size, quadruples the number of data areas forming the matrix, which also means we require greater computer power for subsequent image manipulation and greater archive storage requirements as we have four times the data.

IMAGE INTERPOLATION Many different types of interpolation are used simultaneously in all digital imaging modalities, e.g. linear exponential interpolation. All interpolation algorithms fill in the gaps in the data. They are basically a mathematical filter that alters the value of data either before or after the data has been processed. No matter how expensive, all digital systems will have minor manufacturing faults. A thin film transistor (TFT) used in a DR system, for example, will contain a few dead detector elements. Consider a small area of a TFT-based DR system that contains a dead detector element. As this particular detector element (del) is dead, it will not send a signal so the system puts in an interpolated (educated guess) value of what it might have been. It does this by taking into account what the surrounding live dels have recorded.

IMAGE MANIPULATION There are many ways to manipulate images, but essentially all techni­ ques alter either the values of the pixels in the raw data matrix or re­ constructed data matrix. Before displaying the stored image data from the system memory, the data values at each address can be processed and manipulated to improve perceived image quality. 159

Medical Imaging

Manufacturer-defined manipulation tools Manufacturers will incorporate some form of data manipulation which may be applied immediately to the raw data in order to maximise system efficiency and reduce noise in the system. The operator does not usually have any control over this process. These manufacturer-based algorithms are applied to the raw data matrix to improve image quality. However, their effectiveness depends on how good the original mathematics and mathematical formulae are, which is itself dependent on the accuracy of the original mathematics. A manufacturer may also incorporate some manipulation techniques as part of pre-set anatomical protocols. A CR or DR system pro­ grammed for a chest X-ray examination will have a different filter algorithm applied to that of an abdominal X-ray examination to take account of the very different subject contrast. These filters are applied to make the most of the data collected for particular body areas and are generally very effective. There is a small caveat in that they are generally based on typical patient models and may not be as quite effective with atypical patients.

Operator-defined manipulation tools Finally, the operator will be able to apply various manipulation tech­ niques. The user applies these techniques if they think the original image can be improved in terms of subjective choice. They can be and are often part of a pre-set protocol and based usually on a typical patient model. These processes are operator-dependent and may not be seen as an improvement by all viewers. Automatic or pre-programmed image reconstruction processes in­ clude: ◾ Windowing ◾ Zooming and enlarging ◾ Noise reduction by background subtraction ◾ Noise reduction by ‘low-pass spatial filtering’ ◾ Edge enhancement by ‘high-pass spatial filtering’

Windowing Every del of the raw data matrix has its own value following exposure. Potentially, we could allocate a shade of grey on the display matrix that 160

Image Manipulation

is directly related to the amount of radiation collected by the individual dels. However, the human eye is only really able to distinguish around 32 shades of grey on an image. Hence, we choose which signal values we wish to display by selecting a range of signal values for each shade of grey. This is a process called ‘windowing’. The window width can be increased to include a greater range of tissue types, but it also means there is less difference between various displayed tissue densities. On the other hand, window width can be narrowed to a particular tissue type, for example soft tissue, but this may mean more dense and less dense tissues, such as bone or cystic areas, may be outside the range displayed.

Zooming and enlarging Three types are commonly available: 1. Interpolated zoom (IZ) 2. Reconstructed zoom (RZ) 3. Geometric zoom (GZ) Manufacturers use a variety of terms and technical jargons in the lit­ erature they produce, but all the techniques will be based around the methods listed here.

Interpolated zoom Interpolated zoom is undertaken on a workstation and involves ma­ nipulation of the raw data matrix. Interpolation algorithms of one sort or another are used under a variety of manufacturers’ trade names to produce a zoomed image. The raw data pixels are enlarged to produce a bigger image by simple magnification. However, the pixels in the new image are now much larger, so a second process is applied by dividing the larger ‘real’ pixels into smaller but interpolated pixels. The result looks reasonable as pixel size is reduced relative to image enlargement, therefore potentially in­ creasing spatial resolution (SR). However, the data values of the new smaller tissue volumes tend to be an average of the original larger tissue volume. So while pixels re­ present smaller volumes they do not represent true patient data, but rather arithmetic averages (interpolations).

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Medical Imaging

Reconstructed zoom Reconstructed zoom is undertaken on a workstation and involves ma­ nipulation of the raw data matrix. It allows reconstruction of smaller areas and uses the original data set. It is very effective if the raw data matrix is smaller than the display matrix size, as we effectively have more information to reconstruct than is actually being used. However, once the display matrix is greater than the raw data matrix, it effectively requires interpolation and is no longer true image zooming. As such it may or may not use zooming inter­ polation algorithms, dependent on original data sampling and data set size in relation to display matrix.

Geometric zoom True geometric zoom requires the movement of the detector plate and tube relative to the object, thereby altering the geometry (Figure 9.3). It means that a small object is projected over many more dels with greater

Number of detectors used to form raw data matrix

Figure 9.3 Geometric zoom.

162

Noise Reduction by Background Subtraction

coverage of the detector plate, in comparison with the other two tech­ niques. This technique is more commonly known as ‘macroradiography’.

NOISE REDUCTION BY BACKGROUND SUBTRACTION In its simplest form, the signal values are all reduced by the same amount, reducing not only the level of noise, but also the useable signal. As such the raw signal-to-noise ratio is actually very similar. If we consider a data set of 16 pixels, the height of the bars represents the actual signal collected by the relevant del of the imaging system. All dels give off a small amount of noise with any real signal being added to this background noise (Figure 9.4). If these values were displayed, the image would look something like the grid of pixels shown in Figure 9.5, with the low signal values in pixels 1, 2, 10 and 16 being just above background noise. If we apply background subtraction, all pixels are reduced equally, reducing the background noise of the system. However, it also reduces the signal values by the same amount (Figure 9.6).

Detector signal

Original data values 16 14 12 10 8 6 4 2 0

1

4

10 7 Pixel number

13

16

Figure 9.4 Actual signal values (raw data).

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Medical Imaging

1

2

3

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5

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9

10

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14

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16

Figure 9.5 How this raw data might reconstruct.

Subtracted signal

Signal values following background subtraction 16 14 12 10 8 6 4 2 0 1

3

5

7 9 11 Pixel number

13

How it might look following background subtraction

How original data set might look 1

2

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Background subtraction applied

Figure 9.6 New signal values following background subtraction.

164

Spatial Domain Filtering for Smoothing and Sharpening No edge enhancement

With edge enhancement

Figure 9.7 Edge enhancement.

SIMPLE EDGE ENHANCEMENT This technique examines the data set and detects where the data change from a high attenuating area to a low attenuating area. The pixels nearest to the edge of this change are given an artificially high signal value effectively outlining the structure (Figure 9.7).

SPATIAL DOMAIN FILTERING FOR SMOOTHING AND SHARPENING Spatial domain filtering is one of the commonest more advanced enhancement techniques used in digital imaging. The idea is we have a target pixel that will be influenced by the values in the surrounding pixels. We select the amount of surrounding pixels we want to consider, typically a 3 × 3 and this is called the mask. We now need to consider the pixels making up the image matrix in groups of nine. The grey scale 165

Medical Imaging S

S

S

S

C

S

S

S

S

Figure 9.8 Shows a central pixel surrounded by 8 pixels used in this process.

value of the central pixel of this group of nine (as indicated by C in Figure 9.8) is added as a proportion value of the eight surrounding pixels (S in Figure 9.8). These proportion values can be changed depending on how much we want to smooth or sharpen the image. We decide how strong an effect we want to apply by putting different mathematical weightings into the mask, these weightings are what is known as the filter kernel. We can put different numbers in for different effects so either smoothing or sharpening can be done in exactly the same way, simply by having a different kernel inside the mask.

filter product kernel (mask) Original matrix with target group of pixels

modified output pixel data value

modified matrix with new pixel data values

Figure 9.9 Demonstrates the target pixel and its surrounding neighbors in a typical 3 x 3 mask.

166

Noise Reduction by ‘Low-Pass Spatial Filtering’

The mask and its kernel are actually applied to every pixel value in the image and as such every pixel value will be influenced by what signal was recorded by its neighbours. The other way we can influence the effect is by increasing the mask to say 5 × 5 or 7 × 7 and it could be even larger but the bigger the mask then the more general the effect. The mask does not even have to be symmetrical and we can ex­ aggerate the effect vertically or horizontally. The ‘Sobel’ gradient filter is a good example and actually works diagonally across an image.

NOISE REDUCTION BY ‘LOW-PASS SPATIAL FILTERING’ Low pass filters are a form of spatial domain filtering and are generally used for smoothing, as well as noise reduction. They are more sophis­ ticated than simple background subtraction techniques and whilst they are very effective at reducing noise, they do also result in a slight re­ duction in spatial and contrast resolution. This essentially leads to very subtle areas of light or dark being removed from the data, whether they are noise or real data. The small signal values relating to pixels 1, 2, 10 and 16 could be smoothed out of the data if this technique is applied too aggressively (Figure 9.10).

How it might look following low pass spatial filtering

How original data set might look 1

2

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Image smoothing applied

Figure 9.10 Effect of image smoothing.

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Medical Imaging

With smoothing filters they are based on averaging so product weighting within the kernel actually has to add up to 1. Otherwise if it was more or less than 1 the overall image would also be brighter or darker, as well as smoothed. If you refer to Figure 9.11 we have tabulated two examples. You can see how the kernel numbers are actually a portion of 1. In the first example we have a ‘blur’ kernel with a total count of 9 so individual kernel values are equivalent to 1/9. In our 2nd example, using the ‘Gaussian’ technique, which is more aggressive. The total values within the kernel add up to 16, so they really represent 16ths. We will focus on the ‘blur’ filter which you may see listed in the menus on your workstations. It is a relatively effective but gentle kernel that smooths and reduces some noise with little effect on contrast. It is classed as an averaging filter hence the equal weighting of 1/9ths in a 3 × 3 matrix. By studying the values in Figure 9.12 you can see how an original pixel value of 85 changes to 73. In isolation this pixel value change does not mean much but if you imagine the mask is applied to every pixel and that all pixels values will be re-written. The higher values such as the 88 will come down a little, whilst the lower value 57 will come up a little (depending on what pixels

Typical values and distribution for a ‘blur’ filter kernel Actual Filter kernel weighting

Typical values and distribution for a ‘Gaussian’ filter kernel

Effective multiplication factor

Actual filter kernel weighting

Effective multiplication factor

1

1

1

1/9

1/9

1/9

1

2

1

1/16 2/16 1/16

1

1

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1/9

1/9

1/9

2

4

2

2/16 4/16 2/16

1

1

1

1/9

1/9

1/9

1

2

1

1/16 2/16 1/16

Figure 9.11 Demonstrates the product weightings within two smoothing filter kernels using a typical 3 × 3 mask. Remember as both are types of averaging filter the total of the weighting is actually 1.

168

Image Sharpening and Edge Enhancement ‘High-Pass Spatial Filtering’ Original image

Product weight

Kernal

72

78

62

62

85

57

88

70

83

Total of all pixels above is 657

X

1

1

1

1

1

1

1

1

1

=

Note these add up to 1 so each one is effectively 1/9

Individual pixel values are multiplied separately by their corresponding kernal value

Modified pixel value

72

78

62

62

85

57

88

70

83

/ 9

Pixel values after being multiplied by corresponding kernel value

=

Sum of all values is 657 divided by 9

96

78

62

62

73

57

88

92

83

New pixel values total of product weights/9

=

73

Figure 9.12 Demonstrates how a ‘blur’ filter kernel influences the central pixel value.

lie next to them). The overall effect will mean a slightly narrower range of pixel values. If we consider a ‘Gaussian’ type kernel which you may also see listed on your menu, then this has a very strong effect and if used incorrectly can effectively over smooth to the point where there is very little contrast. The ‘Gaussian’ kernel is what we call a weighted averaging filter and you will notice the centre values as well as up, down and side to side are higher than the diagonal weightings. The total is still out of 1 but averaging takes place in proportion to the weighting, which is now in 1/16ths (Figure 9.13).

IMAGE SHARPENING AND EDGE ENHANCEMENT ‘HIGH-PASS SPATIAL FILTERING’ High pass filters are generally for sharpening and edge enhancement, named examples you might come across are ‘Laplacian’, ‘Laplacian-like’ or ‘Sobel gradient’ 169

Medical Imaging Original image

Kernal total of 1

72

78

62

62

85

57

88

70

83

Total of all pixels above is 572

X

Product weight

1

2

1

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=

Note the numbers represent portions of 1 so each one is effectively 1/16

Individual pixel values are multiplied separately by their corresponding kernal value

Modified pixel value

72

156

62

124

340

114

88

140

83

/ 1 6

Pixel values after being multiplied by corresponding kernel value 1179

=

Sum of all values is 340 + 839 = 1179 = divided by 16

96

78

62

62

74

57

88

92

83

New pixel values total of product weights/16

74

Figure 9.13 Demonstrates how a ‘Gaussian’ filter kernel influences the central pixel value.

These filters are often used for edge detection and should be applied after noise reduction otherwise you risk boosting the noise as well as the high signals leading to a very grainy image. As we are trying to detect edges rather than smoothing them out with this type of filter, we are no longer trying to average the pixel values. As such the total weighting has to equal 0. So you will notice our sharpening kernel has negative and positive values, which need to bal­ ance to 0 overall. In our example the centre kernel value is +8 whilst all 8 surrounding pixels are −1 giving a net weighting value of 0, these are typical values for a ‘Laplacian’ filter (Figure 9.14). With this type of filter if the pixels in a particular area are very similar it will not have much of an effect and no effect at all if they are the same. If on the other hand a pixel is distinct from its neighbors this difference will be amplified by a large amount.

170

Standards

Original image

Product weight

Kernal

72

78

62

62

85

57

88

70

83

Total of all pixels above is 657

X

-1

-1

-1

-1

8

-1

-1

-1

-1

=

Note the negative values BUT values are whole

individual pixel values are multiplied separately by their corresponding kernel value

Modified pixel value

-72

-78

-62

96

78

62

-62

+ 680

-57

62

108

57

-88

-70

-83

88

92

83

Pixel values after being multiplied by corresponding kernel value

=

+680 + -572 sum of all values = 108

Sum of original pixel

=

108

Figure 9.14 demonstrates the target pixel and its surrounding neighbors in a typical 3 × 3 mask ‘Laplacian’ filter kernel.

STANDARDS Traditionally, no or very few common standards were available for dif­ ferent pieces of equipment by different manufacturers. With modern equipment, it is important that an image produced by a particular equipment manufacturer might need to be manipulated, displayed and stored on a piece of equipment by a different manufacturer. A common standard was produced to allow this to happen and is known as ‘digital imaging and communications in medicine’ (DICOM). DICOM information contains a number of elements including in­ formation such as patient details and obviously the post-manipulated images themselves (as pixel values). A fully DICOM-compatible data set will contain additional information, such as the imaging protocols used, raw data values, as well as details of any manipulation filters already applied which are important for future manipulation of the image. In addition to DICOM, there are a number of other standards enabling picture archive and communications systems (PACS), as well as wider 171

Medical Imaging

teleradiolology systems, to operate. Although slightly beyond the remit of this book, some of these standards warrant a mention and include: 1. IEEE 802.3 (ethernet). This stems from computer networks and is associated with the physical layer on which the image data passes (network plugs, sockets, wires, switches and network cards). 2. IEEE 802.5 (token ring). Similar to ethernet standard, except that a ‘token’ is required to gain access to image data. A good example might be a dongle, which is a small device similar to a USB memory stick that has to be physically plugged into a computer in order to access certain images and data. 3. TCP/IP (transmission control protocol/internet protocol). This is a standard to ensure that data are transmitted over the internet without image or data degradation or corruption. 4. HL-7 (Health Level 7) is basically to ensure that a radiology information system (RIS) can talk to PACS and the hospital information system (HIS) systems. It also enables one hospital PACS, RIS and HIS to talk to another hospital’s set-up.

172

MCQs

MCQs 1. A digital system with a resolution of 10 pixels/mm2 is: a. Equivalent to 10 lp/mm b. Equivalent to 20 lp/mm c. Equivalent to 15 lp/mm d. Equivalent to 5 lp/mm. 2. If we double the image matrix size from 256 × 256 to 512 × 512, the data are: a. Twice the resolution b. Twice the resolution, but four times the data c. Four times the data and four times the resolution d. Four times the data, but half the resolution. 3. A 2K monitor will have approximately: a. 2000 pixels b. 4000 pixels c. 40,000 pixels d. 400,000 pixels. 4. A full resolution monitor needs to display at least a. 3500 × 3500 pixels b. 4300 × 3500 pixels c. 4300 × 4300 pixels d. 2000 × 2000 pixels. 5. The human eye is only really able to distinguish around: a. 60 shades of grey on an image b. 10 shades of grey on an image c. 4000 shades of grey on an image d. 32 shades of grey on an image. 6. Windowing: a. Allows images to be altered in size to fit the display b. Changes the contrast and brightness of the image

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Medical Imaging

c. Allows certain tissue densities to be divided into 32 discrete shades of grey for display d. Can be used to increase the number of pixels displayed. 7. Interpolated zoom: a. Requires the movement of the detector plate and tube relative to the object b. Is true image zooming c. Is a two-stage process that simply enlarges the pixels and then subdivides them to produce new smaller pixels d. Is an artificial image zoom. 8. Macro-radiography is another name for: a. Geometric zoom b. Interpolated zoom c. Reconstructed zoom d. Magnification. 9. Noise reduction by background subtraction: a. Is where the signal values are all reduced by the same amount b. Takes account of the average of the surrounding pixels c. Increases contrast and spatial resolution d. Decreases contrast and special resolution. 10. A common standard for transferring data files between manufacturers is: a. SITCOM b. DIDCOM c. DICOM d. OFCOM.

174

CHAPTER 10 COMPUTED TOMOGRAPHY INTRODUCTION The aim of this chapter is to explore how radiation is detected, measured, quantified and used in order to produce a CT scan. The focus in this chapter is to look at how we use and control the X-ray tube and detectors in the creation of CT data. The working prin­ ciples of these two components has already been covered in pre­ vious chapters, so a basic working knowledge of both is already assumed. The chapter will begin by looking at the principles of creating data and then how this is used to reconstruct CT images. There is some discussion surrounding scanner design variations and features, with a particular emphasis on multi-slice and wide field beam detector arrays to reflect modern scanner designs. The chapter culminates in a discus­ sion around scanning considerations within protocols and how these are employed to optimise CT scan examinations. Learning objectives The students should be able to: ◾ Discuss how an image is produced by a CT scanner ◾ Discuss various equipment design features necessary for the control and optimisation of the scanning process ◾ Discuss considerations taken into account during both data collection and reconstruction phase

175

Computed Tomography

GENERAL PRINCIPLES OF COMPUTED TOMOGRAPHY An X-ray beam passes through the body and is then measured by a detector array. The beam and detector array circle in unison with the beam being pulsed as it circles the body (Figure 10.1). Each time the tube is pulsed the amount of radiation that passed through the body is measured by the detector array. The array itself actually contains many individual detectors and each one sends a small signal proportional to the radiation that fell on it. This allows a number of individual at­ tenuation measurements from different angles to be taken at set inter­ vals as the tube and detectors circle the body. The individual signal values are digitised and sent to temporary sto­ rage. A typical CT scanner tube may be pulsed around 1000 times during a single 360 degree rotation, allowing us to collect 1000 sets of signals from its detectors. By applying many calculations to the 1000 sets of data, we are able to determine how much attenuation took place in particular volumes of the body.

X-ray tube producing a fan beam

A typical X-ray tube is pulsed around 1000 times as it circles the body through 360 degrees

The detector array is actually made up of many individual detectors

Figure 10.1 X-ray tube, collimated to a detector array circling the body.

176

General Principles of Computed Tomography

Defining areas within the body Any area within the body can be allocated an x, y and z co-ordinate and be formed into a cube/volume (Figure 10.2).

Figure 10.2 The trunk of the body is represented by the large cylinder whilst the small cube represents a single cube of data.

We can effectively divide the body up into a collection of 3 dimen­ sional cubes. The individual size and number of cubes are set by adjusting scan parameters within the protocol for a particular part of the body. After a series of pulses, also known as projections have taken place, we are able perform complex calculations using a process called Iterative Filtered Back Projection to provide information on the amount of attenuation that took place in each cube/volume of tissue. The cal­ culated attenuation values of these volumes, known as voxels form what is known as the raw data, they effectively indicate what attenuation took place throughout the body per unit volume. The raw data is then used for reconstruction where the attenuation value contained in the voxel is allocated a shade of grey for display as pixels. Different tissue types will attenuate certain amounts of radiation, so we are able to work out the likely tissue type for any particular voxel we are interested in. The amount of attenuation and hence raw data within the voxel is based on a CT or Hounsfield number (named after one of the inventors).

Hounsfield scale As water is the most prevalent substance in the human body it was selected as the reference value of 0 and the original Hounsfield scale went from −1000 to +1000. 177

Computed Tomography

Any tissue that attenuates more radiation than water is given a po­ sitive value, originally up to +1000 (cortical bone/calcifications). Any tissue less attenuating than water was given a negative value, down to −1000 (gas, air). A Hounsfield unit is calculated using this formula

HU =

1000 × water

water air

In reality the attenuation coefficient of air is effectively zero, so the formula above means that an increase or decrease of one HU actually represents the same amount of attenuation as 0.1% of the attenuation of water. The scale has since been extended to include metals as they may be present as implants or foreign bodies. One of the highest at­ tenuating materials known is gold at around +30,000 HU, which is much more dense (attenuating) than any natural body tissue (Figure 10.3).

Tissue type

Typical Hounsfield values

Water

0

Air

-1000

Bone

+700

White matter

+20 to +30

Grey Matter

+35 to +45

Kidney

+30 (+20-+40)

liver

+50 (+40 to +60)

Fat

-100 to -50

Muscle

+10 to +40

Cerebro Spinal Fluid

+15

Metal

above +1000

Figure 10.3 A table listing a few common tissue types and their respective Hounsfield Units.

178

Creation of Data in the xy Dimension

CREATING RAW DATA ATTENUATION VALUES We are now going to take a more detailed look at how we create the raw data volumes. Although we collect raw data in all 3 dimensions simulta­ neously, we are going to focus on what contributes to the data in the xy dimension first and then what contributes to the data in the z dimension.

CREATION OF DATA IN THE xy DIMENSION To help understand how we get data in the xy dimension, we need to consider the following; ◾ Scan field of view ◾ Image matrix ◾ Pixel size ◾ Number of projections (samples) ◾ Simple back projection ◾ Filtered & iterative back projection

Scan field of view (SOV) This is about selecting the size of the body part we are interested in. The X-ray tube of a typical scanner produces a fan beam set to the size 240 mm

X-ray fan beam

240 mm

Figure 10.4 A fan beam being collimated in the xy dimension to a width of 240 mm creating the SFOV.

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Computed Tomography

of the body part being scanned (Figure 10.4). Typically this would be 240 mm for a head scan and 450 mm for a thorax or abdomen and would be part of the scan protocol. This area is referred to as the Scan Field of View (SFOV).

Image matrix This is where we decide how many areas we want to populate with data. The SFOV is divided up into a virtual grid of areas by lines in the x and y axis. This grid is known as the image matrix and would typically be 512 × 512 or 1024 × 1024.

Pixel size Actual pixel size is calculated by dividing the SFOV in mm by the matrix (Figure 10.5). We now need to populate the image matrix with data so that every 0.23 mm2 area in our example has a data value (Hounsfield unit) which is based on the amount of radiation that has been attenuated by the voxel underlying that area.

÷

SFOV (240 mm)

=

Matrix size (1024)

Pixel size (0.23 mm2)

Figure 10.5 Demonstrates the matrix being applied over scan field of view. (The parameters in brackets represent typical values).

Number of projections The X-ray tube moves around the body producing a fan beam to match the SFOV. The whole fan is directed through the patient at a detector array which consists of many (typically around 800–1000) small dis­ crete detectors. Each detector simply measures how much radiation has 180

Creation of Data in the xy Dimension

X-rays produced as a fan beam

Attenuation measured by the detector bank

Producing a single Ray Sum Intensity profile

Figure 10.6 Represents the fan beam passing through the head onto the detector array producing a single ray sum intensity profile.

reached it. Collectively all the signals produce a Ray Sum Intensity Profile (Figure 10.6). Each part of the profile indicates how much at­ tenuation took place along a track from the tube focus to a specific detector. As the tube and detectors circle around the body the X-ray beam is pulsed to produce a series of very short exposures from different angles. On the second pulse we produce a second ray sum intensity profile from a slightly different angle. The beam continues to be pulsed producing a 3rd ray sum intensity profile, a fourth, fifth as it moves through 360 degree. In reality we typically collect around 1,000 ray sum in­ tensity profiles for a single rotation of the tube by pulsing the beam (turning it off and on) as it circles the object 1,000 times. Each of these pulses produces a ray sum intensity profile at a slightly different angle through the body and when combined with the ray sum intensity profiles from the other 999 projections it enables us to calculate how much attenuation took place throughout the volume.

Simple back projection The signal values that formed the ray sum intensity profile are digitized and put into temporary memory. They are then combined with the data from the other projections to form the raw data for a discrete slice or spiral through the body. 181

Computed Tomography

For the time being we are going consider a single pulse (exposure) of the beam as it passes through the body. To help simplify the process we are going to look at a very small area within the matrix which only has 9 pixels in it (Figure 10.7).

Figure 10.7 Demonstrates us focusing on a few pixels in the middle of the matrix and a small portion of the ray sum intensity profile.

As the single pulse of radiation passes through this small group of 9 pixels it creates a signal in the detectors (Figure 10.8).

20 X-ray beam

35 20 Matrix

Portion of ray sum intensity profile for these 9

Digitised to a value

Figure 10.8 Demonstrates how a small portion of the ray sum intensity profile is digitised by focusing on 9 data areas (pixels).

The signals forming the ray sum intensity profile are subsequently digitised to indicate numerically how much radiation was attenuated as the beam passed through the pixel rows. In our example both the upper and lower row of 3 pixels attenuated 20, with more attenuation taking place through the middle row of 3 pixels equating to 35, indicating slightly higher attenuating tissues were present in the middle row. What we do not know from a single projection though, is where along the row 182

Creation of Data in the xy Dimension

of pixels the attenuation took place. All we can say is that any one of the pixels along the upper or lower rows could have attenuated part of this 20 units. In a similar way we do not know the proportions of 35 that were attenuated by specific pixels on the middle row (Figure 10.9).

X-ray beam passing through matrix

Total of 20

0–20

0–20 0–20

Total of 35

0–35

0–35 0–35

Total of 20

0–20

0–20 0–20

Attenuation values actually collected

Possible data values

Figure 10.9 Demonstrates the totals collected along the rows and the possibilities of where attenuation took place.

We now collect another set of values on the next pulse of radiation as the tube moves around the body (Figure 10.10).

f5 lo 20 ta of To l 5 2 ta of To l 0 ta f2 To lo 5 ta of To tal To

X-ray beam passing through matrix from 2nd projection

Projection 2 attenuation values actually collected

0–25 0–20

5

0–20 0–25 0–20 5

0–20 0–25

Possible data values

Figure 10.10 The tube has moved around the object and a second pulse of ra­ diation results in a second set of values and pixel possibilities.

If we combine the two sets of data we can start to narrow down the possibilities of what area attenuated a certain amount of radiation. So if we consider the first projection in isolation all we knew was the top row added 183

Computed Tomography

up to total of 20. Now if we consider the data from the second projection we identified the top right pixel attenuated 5. This means the top left and top middle pixel can now only add up to a total of 15 (Figure 10.11). Data from 1st projection

Data from 2nd projection

0–20 0–20 0–20 0–35 0–35 0–35

0–25 0–25

+

0–20 0–20 0–20

Combined possibilities

5

0–15 0–15

0–20 0–25 0–20 5

=

0–20 0–25

5

0–20 0–25 0–20 5

0–15 0–15

Figure 10.11 Shows the two ray rum intensity values from projections 1 and 2 and by combining them to calculate the possibilities.

The tube continues to move around the patient and pulses for a third time and the detector array now measures the left and right columns which equal 20 whilst the middle column gives a total of 35 (Figure 10.12). If considered the 3rd projection alongside the other 2 projections we can narrow down the possibilities even further (Figure 10.13).

0–20 0–35 0–20 Total of 20

Total of 35

Total of 20

0–20 0–35 0–20 0–20 0–35 0–20

X-ray beam passing through matrix from 3rd angle

Projection 3 attenuation values actually collected

Possible data values

Figure 10.12 The tube has moved around the object and a third pulse of radiation results in a third set of values and data possibilities.

184

Creation of Data in the xy Dimension

Data from 1st projection 0–20

0–20 0–20

0–35

0–35 0–35

0–20

0–20 0–20

Data from 2nd projection

+

0–25

0–20

5

0–20

0–25

0–20

5

0–20

0–25

0–5

0–15

Data from 3rd projection

+

0–20

0–35

0–20

0–20

0–35

0–20

0–20

0–35

0–20

5

0–15 0–35 0–15 5

0–15 0–15

Possible data values

Figure 10.13 Shows the three ray rum intensity values from projections 1, 2 and 3. Combining these data sets allows us to calculate the possibilities even more.

We collect a fourth projection and this time it identifies the top left and bottom right are both 5 as well as slightly different attenuation ranges possible in the other pixels (Figure 10.14). If we combine all 4 data sets, it enables us to calculate precise values for every pixel in our example (Figure 10.15). If we focus on the top row of data areas we now have values for the top left of 5 (from projection 4) and top right also 5 (from projection 2) and as we know the total across all 3 adds up to 20 (from the 1st projection) then it must mean the middle data value of the top row is 10. In a similar way it is possible to calculate what all the data values are in our example creating the raw data for this 3 × 3 matrix. This is a very simplistic example and we have solved it by a process of elimination in a similar way to how we solve a ‘Sudoku puzzle’. In reality the raw data is calculated mathematically as we are probably interested in populating 1024 matrix requiring a little over a million bits of at­ tenuation data from around 1000 ray sum intensity profiles (projec­ tions) for a standard scan. 185

To ta To lo ta f5 To lo ta f2 l To 0 of ta 25 lo To f2 ta 0 lo f5

Computed Tomography

Projection 4 attenuation values actually collected

X-ray beam passing through pixels from 4th angle

5

0–20

0– 25

0–20

0–25

0–20

0–25

0–20

5

Possible pixel values

Figure 10.14 The tube has moved around the object and a fourth pulse of ra­ diation results in a fourth set of values and pixel possibilities.

Data from 1st projection

Data from 2nd projection

0–20 0–20 0–20

0–25 0–20

0–35 0–35 0–35

0–20 0–25 0–20

0–20 0–20 0–20

5

Data from 3rd projection 0–20 0–35 0–20

5

0–20 0–25

Data from 4th projection 5

0–20 0–25

0–20 0–35 0–20

0–20 0–25 0–20

0–20 0–35 0–20

0–25 0–20

5

10

5

10

15

10

5

10

5

5

Calculated answer Figure 10.15 By considering all the data possibilities from the 4 projections we arrive at the answer.

186

Creation of Data in the xy Dimension

There are issues in using pure back projection calculations as it would actually generate inaccurate raw data. So we are going to discuss why we have to modify or filter the data to correct for these inaccuracies. This is commonly achieved using what is known as the pipeline prin­ ciple of image reconstruction using filtered back projection or iterative filtered back projection.

Image reconstruction using filtered back projection (FBP) or iterative filtered back projection (IFBP) If we only consider pure signal values there are issues when we come to reconstruct. This is because the signal values are simply generated based on the total attenuation that took place between the tube and detector. So let us imagine an isodense object such as a single density phantom (a simple plastic uniform tissue equivalent disc). Every voxel within the phantom should have identical values whether it is at the edge or in the middle of the phantom, as the material density and hence attenuation in that area is identical. The problem is that the part of the beam going through the peripheral parts of the phantom are attenuated much less than central portions as it has to travel through less plastic. As the tube moves around to produce all the ray sum intensity values it will always record lower values around the periphery and yet the attenuation per unit volume should be identical (Figure 10.16).

Simple isodense object

Back projected attenuation profile

Simple unfiltered back projected reconstruction

Figure 10.16 Demonstrates how a reconstruction of an isodense object might look without filtered back projection.

187

Computed Tomography

This means the data has to be corrected and this is done mathema­ tically using FBP or IFBP. The mathematics involved are extremely complex and beyond the understanding required at this level of study but it is important to understand the process. Filtered back projection is essentially a corrective mathematical filter involving several mathematical equations being applied to overcome flaws in the data. They counter a number of effects including linear attenuation and beam hardening amongst other things. Basic iterative back projection was actually used very early on in CT development (1970s) but the computers of the time were not powerful enough to use this technique commercially. So Filtered Back Projection dominated from the 1970s for the next 40 or so years and is still in clinical use. Rather than look at the maths involved, Figures 10.17 and 10.18 help to illustrate this process using graphs and images. Iterative filtered back projection was introduced into clinical prac­ tice in some of the higher end scanners from around 2009. The iterative aspect means that once the filtered back projection calcula­ tions have been done the results are interrogated by an algorithm and then the calculations are redone. It is a process where data is reanalysed using a cycle of operations where each cycle (or iteration) brings us closer to the ideal answer, until there is ‘satisfaction’ in the answer.

+

Back projected attenuation profile

=

Mathematical correction filter

=

Attenuation profile New filtered and correction (corrected) filter combined attenuation profile

Figure 10.17 These graphs help illustrate how data is managed to reconstruct our isodense object using a correction filter (filtered back projection).

188

Creation of Data in the xy Dimension

+

Simple unfiltered back projected reconstruction

=

Corrective filter

New filtered reconstruction

Figure 10.18 Visually demonstrates the same filtered back projection tech­ nique and how data is managed to reconstruct our isodense object using a correction filter.

The benefit of this approach is that allows us to consider many more variables that could influence the fidelity of the reconstruction but one of the main reasons is its ability to deal with noise and very small signals. Traditional FBP includes high pass spatial filtering as one of the early stages in reconstruction, which is fine with higher S/N ratios but it does boost noise levels, which can be a big issue for lower S/N ratios, producing poor reconstructions. Modern multi-detector systems make use of very small detectors, producing very small slices and whilst they have excellent spatial resolution, they produce relatively low S/N ratio’s. Detailed discussion of the iterative process is beyond the scope of this book but it is enough to know that the iterative process includes a number of techniques such as interpolation and even ‘model’ based reconstruction, which are very effective at managing relatively low S/N ratios, much better than FBP alone. As the iterative aspect allows us to manage noise and low signals much better it also means that manufacturers are able to develop much lower dose techniques whilst maintaining image quality. It is interesting to note that much of the iterative process has to be tailored to particular machines. In the USA this even means that the algorithms used are subject to Food and Drug Administration (FDA) approval in a similar way to how a drug is approved for clinical use. 189

Computed Tomography

CREATING DATA IN THE z DIMENSION The process is different for a discrete directly scanned slice and one that is reconstructed from a spiral/helical scan and will be discussed separately.

z axis data from directly scanned sequential slices On early scanners the width of the beam in the z dimension was simply controlled by collimating the width of the fan beam as well as a second set of collimators over the detector array to the slice thickness we wanted to acquire (Figure 10.19).

Single 20 mm wide detectors

00 10 to 0 0 s y 8 tor all tec c i p de Ty

Collimators over detector array

Figure 10.19 Image of a portion of a single channel detector array, typically consisting of 800 to 1000 detector each being 20 mm in the z axis.

190

Creating Data in the z Dimension

Then came multi-slice scanners which were able to acquire multiple slices in a single rotation. Rather than a single detector in the z dimension, the width was split into several smaller detectors initially 2 and then 4. The example in Figure 10.20 below shows the original 20 mm single detector being divided into 4 individual × 5 mm detectors (in the z axis).

< 20 mm >

0

00

1 to 00 rs 8 o y all ect pic det y T 4 detectors each being 5 mm

Figure 10.20 A 4 slice detector array, based on 4 x 5 mm detectors replacing a large single 20 mm detector in the z axis (we still have 800–1000 in the xy dimension).

Later came 16 and then 64 slice scanners. A typical 64 slice scanner would have total detector array z axis width of 40 mm, divided up in to 64 detectors each one being 0.625 mm. If detector array also had 1000 around the x,y dimension then there would be a total of 1000 × 64 individual detectors (64,000). If we further consider a reconstructed 10 mm slice this would actually be made up of several very fine slices. If our detectors were 0.625 mm (typical for a 64 slice scanner) we could fit 16 detectors into 10 mm and therefore our 10 mm slice is actually made up of 16 very small slices (Figure 10.21).

191

Computed Tomography

40

m m

0 00

1 to 00 rs 8 y to all ec pic det y T

64 detectors total width of 40 mm Figure 10.21 Represents a 64 multi-channel detector array typically with 64 detectors spanning 40 mm wide (there are still the same 800–1000 in the xy dimension).

160 m

m

00

10 to 0 s 80 or ly tect l a ic de yp

T 320 detectors each being 0.5 mm2 total width 160 mm

Figure 10.22 Represents a wide multi-channel detector array with 320 detectors spanning 160 mm (there are still the same 800–1000 in the xy dimension).

192

Creating Data in the z Dimension

Some of the current scanners now have as many as 320 detectors across the z axis. These are based on a wide field detector array with a total z axis coverage of 160 mm (Figure 10.22). By using interpolations to create extra slices, this design actually allows as many as 640 very fine slices spanning a detector array width of 16 cm which can be acquired simultaneously with a single tube rotation. This is wide enough to image the whole heart or even a full brain scan in a single rotation.

z axis data from spiral acquisition If the scanner is in spiral mode, the tube moves around the patient as the table also moves along the z axis. Rather than producing discrete discs/slices of data, the data is produced as a spiral of continuous data through the volume. We effectively produce a 3 dimensional data set rather than a series of discrete slices. We do not view this spiral directly so we have to extract the in­ formation we need to reconstruct interpolated scan slices from this volume spiral data set. In order to do this we select a slice position and then use data from the spiral data that was recorded either side of the slice we want to reconstruct. Using a process known as interpolation, the computer calculates what data might be in this slice position. If you look at Figure 10.23 and were asked to make an educated guess at what number was in the slice position you would probably come up with 5 as it is half way between 1 and 10, the two actual values that were

Slice we are trying to interpolate

10 ?

1

Two measured data values for this portion of the spiral

Figure 10.23 Illustrates actual data values of 1 and 10 contained in the spiral data set and the disc with “?” mark in that requires data for reconstruction.

193

Computed Tomography

calculated and we know are accurate. As 1 and 10 are actual data points the process is known as 2 point linear interpolation.

Multi-spiral scanning On modern scanners we actually collect a series of very fine spirals ra­ ther than a single relatively thick spiral of data. This means we now have many more data points that we can use and rather than a 10 mm re­ construction being based on a single 10 mm spiral of collected data we may instead now have, lets say, 16 × 0.625 mm fine spirals (dependent on how the detector array is structured). So on modern multi-slice/multi-spiral scanners the interpolation is a little more complex and accurate than simple 2 point interpolation. Data is considered across multiple spirals before and after the re­ construction slice level. This way it is possible to judge how the tissues are changing from one area to the next either side of the in­ terpolated reconstruction. This results in a much more accurate in­ terpolation. This is known as multipoint or filtered interpolation (Figure 10.24).

Interpolated slice

XXXX----? ------XXXX X

Eight measured data values for this portion of the spiral

Figure 10.24 Demonstrates multipoint filtered interpolation with 4 spirals of actual data either side of the slice being reconstructed and therefore 8 data points.

194

Spatial Resolution (SR)

HOW DOES ALL THIS RELATE TO SCAN CONSIDERATIONS / PROTOCOLS As a simplistic overview, we generally start by defining the z axis cov­ erage in other words where we want to start scanning and where we want to finish. This can be done from gantry positioning lights and table position initially to obtain our pilot or scanogram, which is then used to setup the main scan from the console. We then decide how we want to collect our data by selecting para­ meters such as the slice thickness, whether we are going to collect a series of axial or spiral acquisitions which will control z axis resolution. The fan width in xy axis would be set to cover the body area we are interested in and then the matrix which together would control xy re­ solution. We also need to consider the signal, noise and dose by selecting appropriate exposure factors and acquisition times. The next few sections will explore these considerations alongside a few others to help you understand what has to be taken into account to optimise our CT scans.

SPATIAL RESOLUTION (SR) There are a number of factors that control spatial resolution but to begin with we are going to focus on the main ones related to the physical size of the voxel, which are; ◾ Slice thickness (z axis resolution) ◾ Matrix Size (xy axis resolution) Although SR is physically selected by the slice thickness in the z axis and the matrix size in the xy axis, there are some others with less ob­ vious influences on SR based on: ◾ Number of projections ◾ Filter convolution

195

Computed Tomography

Slice thickness (controls z axis SR) If slice thickness is reduced then it is clearly going to enable more detail from the finer slices and therefore increase in SR. z axis SR is directly proportional to slice thickness and hence voxel depth. There are some issues with using finer slices as they cover less of the body in the z axis. So in order to cover the same overall distance we will need more slices. This means more exposures, taking longer to acquire and slower overall scan times. More data also takes up more digital space and takes longer to process/reconstruct.

Matrix size (controls xy axis SR) The SR is physically set in the xy axis by selecting the matrix and SFOV. The bigger the matrix size, then more areas are created, which will go on to form finer pixels in the reconstructed slices. If we consider a SFOV of 240 mm then a 1024 matrix would give us 1,048,576 pixels each with a physical size of 0.23 mm2, whereas a 512 matrix would only give us 262,144 pixels each with a physical size of 0.47 mm2.

Number of projections Anything that increases the number of projections enables more data to be taken into account for scan reconstructions and helps provide more precise attenuation values within the voxel/pixel, improving overall image fidelity and hence SR (Figure 10.25). One way to increase or decrease the number of projections is by controlling scan rotation speed. If we assume the tube is pulsed at the same frequency of 1000 times per second then if it takes 2 seconds for a full rotation we can capture 2000 projections, whereas if it takes 1 second for a full rotation we will only acquire 1000 projections. Some scanners do not necessarily collect data from a full 360 de­ gree rotation arc. If the scan protocol you are using normally collects data from say a 180 or 270 degree arc that can be increased to 270 or 360 degree (respectively) which is an alternate way of collecting more projections. We can also use equipment design features or oversampling techni­ ques, which may be selectable or integral and manufacturer dependent. 196

Spatial Resolution (SR)

16 back projections

32 back projections

Figure 10.25 Demonstrates the extra spatial resolution that 32 back projections have in comparison to 16 back projection for a simple shape.

Some examples of this might be Dynamic Focal Spot or Quarter Detector Shift technology. These are essentially different methods of increasing the number of projections used to reconstruct slices produ­ cing more data and subsequently higher fidelity reconstructions, thereby increasing SR. You need to appreciate they exist and their purpose but detailed discussion of them is beyond the scope of this book. Time (sometimes referred to as temporal resolution) is one of the main tradeoffs with increasing the number of projections. Longer ac­ quisition times as a result of slower scan rotation times increases the possibility of movement unsharp-ness. More data also takes up more digital space and takes longer to process/reconstruct. Typically around 2000 projections might be used for high resolution scans. It could be 1000 projections for a standard scan and as few as 500 for a low re­ solution fast scan, that might be used to control movement for trauma patient.

197

Computed Tomography

High resolution filter convolution Is essentially a complex mathematical filter that can involve a number of steps. On a simplistic level the data is first interrogated for any stray signals which are minimised ‘cleaned’ and then a low pass filter is ap­ plied to reduce noise as we do not want to boost noise when we apply the high pass filter in the final stage. As such they work better with higher S/N ratios. The high pass filter itself is discussed in more detail in chapter 9 but essentially if two pixels next to each other have similar values they are not boosted by this form of convolution filter but if a neighbouring pixel is slightly higher it will be boosted even further. This type of filter can be over applied and boost subtle densities too much if not used ap­ propriately, creating an over grainy end result. Usually selected by including high resolution scanning algorithm as part of the protocol or selected from the control menu.

CONTRAST RESOLUTION (CR) Is mainly associated with the level and fidelity of the signal within the voxel, so anything that increases this will generally increase CR. The main contributors to CR are: ◾ Slice thickness ◾ Matrix size ◾ mAs

Slice thickness and matrix size For now it is enough to appreciate that CR increases with slice thickness and a coarser matrix (larger voxels) if there is no change in exposure, as there will be more quanta and hence higher signal per voxel. As a caveat though larger slices are more likely to suffer from what is known as partial volume effect, as larger volumes are more likely to include dif­ ferent types of tissue (partial volume effect is discussed later in this chapter). In addition the larger voxels, whilst producing relatively high signals also directly reduce SR which will fall proportionately with larger voxels. 198

The Link between Slice Thickness, Dose, Photon Density

mAs Alternatively we can increase the mAs to increase quanta and keep the same slice thickness and matrix to maintain SR. The downside is that as mAs increases so does patient dose. All manufacturers have a form of dose/mAs modulation where the mAs is constantly adjusted by software to enable the lowest dose and maintain the best signal. When the tube is at 0° the x-ray beam gen­ erally has less tissue (AP orientation) than when it is at 90° (lateral orientation), the software will reduce the mAs when at 0° and increase when at 90°. The software also varies mAs along the z axis, so throughout the scan it constantly monitors the signal at the detectors varying mAs in real time, to account for body habitus. So whilst the exposure varies it is also optimized for each projection without com­ promising image quality. The next section will explore these considerations and the inter­ relationship between dose, slice thickness, matrix, SFOV, S/N ratio, SR and CR in a little more detail.

THE LINK BETWEEN SLICE THICKNESS, DOSE, PHOTON DENSITY AND SIGNAL VALUE If we assume the photon density is uniform across the width of the X-ray beam then if we half the beam coverage by collimating from 10 mm to 5 mm then we effectively absorb half the quanta with the collimators, so only half the quanta now reaches the detectors (Figure 10.26). This means that the signal produced by the detectors is also half what it would have been. So to restore the signal to its previous value, we would need to double the photon density by doubling the mAs (Figure 10.27). There is another effect of thinner slices as we will require double the number of slices to cover same area. So when we consider changing from 10 mm to 5 mm slice thickness, if we want to maintain the original signal value and cover the same distance through the body it will ac­ tually require 4 times the dose. This is simply because the 5 mm slices 199

Computed Tomography

Signal Collimated to 10mm Noise

Total 20 mm section of tissue

Collimated to 5mm

Collimators

Amount of quanta contributing to signal

Signal Noise

Figure 10.26 By collimating from 10 mm to 5 mm with the same photon density (constant mAs) we effectively half the quanta to the detector and therefore half the signal.

need twice as much mAs to generate the same signal and we also need two 5 mm slices to replace one 10 mm slice.

The principles still apply when we consider multi-slice scanners Slice thicknesses are managed slightly differently, so that instead of using collimators over the tube and detectors to control slice thickness, the tube can be collimated to the whole width of the detector array. To select our slice thickness we actually select the number of small de­ tectors which collectively make up our reconstructed slice width. The individual detectors may only generate relatively small signals but we have 4 sets of signal for one exposure. If we want to reconstruct 200

The Link between Slice Thickness, Dose, Photon Density

Collimated to 5 mm

Signal Noise

Photon density of tissue is doubled by doubling the mAs in the volume below

Area of quanta contributing to signal is the same but as photon density is doubled the signal is also doubled

Signal Collimated to 5 mm Noise

Figure 10.27 The collimation is the same and therefore the area of quanta contributing to signal is the same but as photon density is doubled, the signal is also doubled.

thicker slices we can actually add these small signals from individual detectors together to generate a higher signal. Our reconstructed slice is actually made from 4 smaller slices or spirals of raw data matching the detector configuration. The example in Figure 10.28 is actually quite crude by modern standards as it essentially represents a 4 slice scanner. The reality (as mentioned earlier in this chapter) is typically a 40 mm wide detector array with 64 rows of 0.625 mm individual detectors. There are scanners with 128 rows of 0.625 mm detector rows giving a z axis coverage of 80 mm per rotation and even a 640 slice scanner, al­ though this is actually generated from 320 × 0.5 mm detector rows with total z axis coverage of 160 mm. The number of slices is often quoted to help categorise the type of scanner but this is really about how we collect the signals and the number 201

Computed Tomography

Collimated to 20 mm overall but using 4 x 5 mm detector channels simultaneously

Signal 1

Signal 2

Noise

Noise

Detector 1

Detector 2

Amount of quanta contributing to signal

Signal 3

Signal 4

Noise

Noise

Detector 3

Detector 4

Figure 10.28 We select 20 mm overall collimation spread over 4 x 5 mm detector channels providing 4 signals simultaneously.

of detectors. We still tend to reconstruct to 5 or 10 mm slices and would not generally view 0.625 mm reconstructions in clinical practice. If we consider 20 cm z axis coverage of the brain for example it would result in 320 slices, simply too many for routine examinations. So on a 64 slice scanner we would typically collimate to the full 40 mm available and use all our 64 × 0.625 detectors but we actually group them together into channels. Typically we might use 8 channels (each containing 8 individual detectors) producing 8 × 5 mm slices at the same time. There are actually many more benefits and ways we can use our multi-slice detectors, we are not just restricted to grouping them into 202

Temporal Resolution and Control of Movement

channels. The following list gives an idea of the benefits of these systems and how we might use them; ◾ Very fine narrow slices (as fine as the detector) for smaller body areas but have to be aware of low S/N ratios, so this is more applicable where we have high subject contrast. ◾ Multiple axial slices or spirals obtained in a single rotation, in our example we produced 8 × 5 mm slices simultaneously. ◾ Fused axial slices into thicker slices, provides better CR, as signals from individual detectors can be added together boosting S/N ratios. ◾ Examination time reduced as collimation can be as wide as the detector bank per rotation of the tube increasing z axis coverage. ◾ Reduced artefacts as there is less chance of partial voluming, in particular, due to the finer slices. ◾ Isotropic voxels are possible with 64 slice (and above) enabling 3D and mult-planner reconstructions of the same resolution in any direction

TEMPORAL RESOLUTION AND CONTROL OF MOVEMENT Temporal resolution (TR) is essentially the time it takes to acquire data. We need to factor in the amount of movement likely to be encountered and will need greater temporal resolution for areas of the body with greater movement. The scanning technique itself can actually create movement. Any spiral scan for example, is effectively acquired whilst the patient is moving, as not only is the tube moving but so is the table. So for any spiral scan, all CT scanners have built in algorithms to account for movement of the table, the rotation speed and so on. Some movement may be due to the patient’s condition. We can control much of this movement by carefully selecting certain aspects of the scan protocol. So for a trauma scan we may consider a quick scan, whilst this will manage patient movement better by using shorter scan times, it usually means less data is collected per rotation and therefore reducing other aspects of image quality, so its use has to be carefully considered. 203

Computed Tomography

There are also some special software packages designed to help manage movement in different ways. We cannot discuss them all in detail but two packages are mentioned here to serve as examples. Cardiac gating software allows us to link the collection of our pro­ jections based on an electrocardiogram (ECG) signal. We choose a point in the heart rhythm usually between the ‘R’ waves when the heart is relatively still and only acquire data during this phase of the heartbeat (Figure 10.29). Effectively dealing with a moving heart. Scan phases when there is relatively little heart movement R T P Q S Figure 10.29 Linking scan acquisitions to certain phases of the heart beat.

Another example is something called Bolus tracking which can have a number of applications but we will use CT angiography as our example. We can select an area over an artery by using a region of interest. We then monitor this area in real time looking for an increase in Hounsfield values as contrast starts to travel round the body and through the artery we are interested in. Once it reaches a predetermined level it will trigger the scan to start at the optimum time.

ARTEFACTS AND OTHER SCAN CONSIDERATIONS It is not possible to cover all scanning limitations but there are a few you should be aware of; ◾ Beam hardening, ◾ Partial voluming, ◾ Cone beam 204

Artefacts and Other Scan Considerations

Beam hardening As the beam passes through the patient its properties alter. As it passes the ‘softer’ x-rays are effectively filtered by the tissues the beam passes through, whilst the ‘harder’ more powerful beam energies tend to pass through to greater depths. This means the beam is absorbed by differing amounts as it travels through the tissues. Contrast resolution and photo­ electric absorption are effectively greater on the entrance side to the exit side for similar tissue densities. The result is something known as beam hardening. The effect can lead to something we call ‘cupping’ where the centre of an object has artificially high attenuation values.

Partial volume artefact All voxels irrespective of their size are recorded as an average of the tissues they contain. This is not a problem if there is little tissue var­ iation within the voxel but if a voxel contains both high and low at­ tenuation tissues such as bone and air, it will be averaged and as such does not directly represent either tissue. ◾ It can be caused with tissues that only partly fill the X-ray beam hitting a particular detector. ◾ It can also be caused if the tissue is scanned during some of the projections but not in others. This can occur towards the outside of the fan beam We can reduce it using finer voxels/acquisitions by selecting thinner slices. It is much less of an issue using multi-slice and filtered inter­ polation algorithms. As we can see where the borders of different tissue types are more precisely and how the tissue is changing through the volume (Figure 10.30).

Cone beam artefact This is of greater concern the wider we make our fan beam in the z dimension. The shape of the beam changes from a relatively flat fan shape at say 5 mm and becomes progressively cone shaped the wider we make our detector array. It is similar to partial volume effect in that and the closer the object is to the periphery of our SOV and the more off centre it is the worse the artefact might be. It is more of an issue with wider detector arrays so gets progressively harder to manage as z axis width increases which is currently as wide as 205

Computed Tomography

Streak artefact caused by Streak artefact caused by Streak artefact reduced by partial volume effect and partial volume effect and thinner slices and beam hardening on a phantom beam hardening on a patient increasing mAs Figure 10.30 The effects of partial voluming and beam hardening.

16 cm on some scanners. It is usually managed using special manu­ facturer based algorithms that correct for the effect. It can also be managed by reducing fan beam thickness but then that defeats the point of having a wide detector array able to image large volumes of tissue in a single rotation (Figure 10.31).

X-ray tube at 0 degrees

This tissue is only being picked up by the detectors when the tube is in certain positons X-ray tube at 180 degrees

Figure 10.31 The causes of cone beam artefact.

206

MCQs

MCQs 1. Hounsfield scale a. Is based on air as the reference value of 0 b. Is based on water as the reference value of 0 c. Is based on bone as the reference value of 0 d. Is based on metal as the reference value of 0. 2. xy a. b. c. d.

resolution is adjusted by: Increasing the matrix Reducing the slice thickness Increasing the slice thickness Using a fast scan.

3. A 1K matrix will have approximately: a. 2000 pixels b. 4000 pixels c. 1,000,000 pixels d. 4,000,000 pixels. 4. If the scan field of view is 320 mm and we use a 1024 × 1024 matrix, what is the size of the pixel? a. 0.23 mm2 b. 0.31 mm2 c. 0.47 mm2 d. 0.66 mm2. 5. z axis resolution is adjusted by: a. Matrix b. Scan field of view c. mAs d. Slice thickness. 6. Spatial resolution is higher (finer) if we: a. Increase mAs b. Increase slice thickness c. Reduce the number of projections d. Increasing the matrix. 207

Computed Tomography

7. Contrast resolution is increased: a. Increasing the mAs b. Reducing slice thickness c. Using a fast scan d. Increasing the matrix. 8. Temporal resolution is better: a. Increase the number of projections b. Using a fast scan c. Increasing the matrix d. Reducing slice thickness. 9. Increasing the number of projections: a. Increases the S/N ratio b. Reduces the S/N ratio c. Provides better temporal resolution d. Increases slice thickness. 10. Partial volume effect can be reduced by: a. Reducing the mAs b. Using a fast scan c. Increasing slice thickness d. Reducing slice thickness.

208

CHAPTER 11 RADIATION PROTECTION AND SAFETY INTRODUCTION The aim of this chapter is to give the practitioner an understanding of the basic principles of radiation protection and safety. It is essential that any practitioner operating within an imaging department and using ionising radiation has a sound basis for their knowledge. You need to comprehend and be able to explain the factors affecting the protection of staff, patients and the general public. The primary purpose of radiation protection when using medical radiation is to reduce the associated risks to staff, patients and visitors to the imaging department to a minimal acceptable level. It should be remembered that the majority of anyone’s radiation dose will originate from naturally occurring radiations and there is little or nothing that can be done to reduce that dose. However, man-made radiation constitutes approximately 14% of an in­ dividual’s total radiation dose, 85–90% of which (i.e. 12% of the total ra­ diation dose) arises from the use of ionising radiation in medical and dental practice and here you can have an influence. The purpose of diagnostic imaging is to produce images of diagnostic quality at the lowest radiation dose. Therefore, the ultimate con­ sideration is to provide a diagnostic image and manage the patient effectively as a result of the test. Learning objectives The student should be able to: ◾ Understand and explain the basic principles of radiation protection for staff, patients and the general public. ◾ Quote the relevant legislation and apply the legislation to practice. 209

Radiation Protection and Safety

LEGISLATION There are two main pieces of legislation in Europe which determine protection of staff, patients and the general public. These are: 1. Ionising Radiations Regulation (IRR17) 2. The Ionising Radiation (Medical Exposure) Regulations 2017 (IR (ME)R 2017)

IRR17 The protection of staff is largely controlled through the application of IRR17. The legislation makes stringent requirements of employers to protect not only their staff, but also patients and members of the general public (a category which includes everyone who might be found in a hospital who could not be classified as either a patient or an employee). This document, among many other things estab­ lishes: ◾ Key roles in the legislation; radiation protection advisor ◾ A qualified expert in occupational radiation protection. ◾ Employer is legally obliged to consult an RPA for advice on fulfilling the requirements of IRR17. ◾ May be an individual or a body (organisation). ◾ Must be appointed in writing radiation protection supervisor. ◾ Responsible for ensuring local compliance with the requirements of IRR17 i.e. supervising work is carried out in accordance with the local rules. ◾ Must have completed an accredited training course. Medical physics expert (MPE) This is a person who holds a science degree or its equivalent and who is experienced in the application of physics to the diagnostic and therapeutic uses of ionising radiation’. ◾ Designation of classified persons. Defined as those employees who are likely to receive an effective dose in excess of 6 mSv per year or an equivalent dose which is likely to exceed three-tenths of any relevant dose limit. ◾ Notification and investigation of overexposure.

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Legislation ◾ Dose-equivalent limits for radiation workers and it is the

enforcement of such limits that have eliminated the illnesses and premature deaths seen in early radiation workers who were exposed to very large doses of ionising radiation, largely through ignorance. ◾ The application of the dose equivalent limits for different categories of workers and also the ALARP principle for the protection of non-employees. ALARP means as low as reasonably practicable (social and economic factors being taken into consideration) and refers to the radiation doses administered by the workers/employees. Applying the ALARP principle is one of your main considerations when undertaking an X-ray. This must be applied not only to the pa­ tient but to visitors, other staff and, of course, yourself. Largely this is done by ensuring the justificative process and ensuring that all X-ray room doors are closed. All non-essential staff are excluded during X-ray examinations.

IR(ME)R 2017 These regulations are intended to protect patients, comforters and carers against the dangers of ionising radiation while undergoing med­ ical exposures. These include diagnostic procedures such as X-rays, computed tomography (CT) scans and nuclear medicine examinations, as well as treatment such as radiotherapy, irrespective of where they are undertaken, e.g. hospital department, dental practice, chiropractic clinic. Medical exposures undertaken for research purposes are also covered by these regulations. The regulations define the duty-holders who carry responsibility under IR(ME)R 2017. These are: ◾ The employer who is (normally the NHS Trust in the hospitalbased environment) is the body responsible for putting all the necessary procedures and protocols in place to ensure that the IR (ME)R 2017 regulations can be fully applied ◾ The referrer, a registered health care professional (e.g. a medical or dental practitioner, radiographer, chiropodist, etc.) who is entitled, in accordance with the employer’s procedures, to refer patients for medical exposures 211

Radiation Protection and Safety ◾ The practitioner, a registered health care professional who is

entitled in accordance with the employer’s procedures and whose primary responsibility is justification of the individual medical exposure (e.g. the radiographer or radiologist) ◾ The operator, a person who is entitled in accordance with the employer’s procedures to undertake the medical exposure (e.g. an assistant practitioner or a radiographer) The 2018 regulations also introduced a medical physics expert (MPE) who is an expert in the use of ionising radiation for therapy & diagnosis (patients). They must be on the MPE register and involved in: ◾ All medical practices involving ionising radiation ◾ Optimisation ◾ Dosimetry and quality assurance ◾ Commissioning new facilities ◾ Training ◾ Advising the employer on compliance with IRMER Some of the main points to arise from the implementation of IR(ME)R 2017 are:

Justification ◾ No person shall carry out a medical exposure unless it has been

justified by the practitioner who has been provided with adequate information from the referrer.

Diagnostic reference levels (DRL) ◾ The employer must set DRLs and provide guidance and

procedures on how they are to be used. A diagnostic reference level is set for each standard radiological investigation. They should also be set for interventional procedures, nuclear medicine investigations and radiotherapy planning procedures.

Optimisation ◾ The practitioner and the operator, to the extent of their

respective involvement in a medical exposure, shall ensure that doses arising from the exposure are kept as low as reasonably practicable consistent with the intended purpose, of the exposure.

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Legislation ◾ The operator shall select equipment and methods to ensure that

for each medical exposure, the dose of ionising radiation to the individual undergoing the exposure is as low as reasonably practicable and consistent with the intended diagnostic or therapeutic purpose and in doing so shall pay special attention to: ◾ Careful/precise technique to minimise repeat examinations ◾ Quality assurance of equipment ◾ Optimisation of exposure factors to provide a diagnostic image within the DRLs set for each procedure ◾ Clinical audit of procedures and exposures

Training ◾ No practitioner or operator shall carry out a medical exposure

without having been adequately trained. This is enforced by the Care Quality Commission (CQC) It is impossible in a pocket-book such as this to provide a full coverage of all of the ionising radiation regulations. It is therefore recommended that you use published guidance, such as appropriate legislative in­ formation on government links.

Local rules and systems of work Drawn up in accordance with Regulation 17 of the Ionising Radiations Regulations 2017 (IRR17). They state all users must read them and sign a statement that they have read them, understood them and will comply with the local rules. There are certain responsibilities under the act, as outlined below.

Responsibilities ◾ Employer ◾ Compliance with regulations ◾ Radiation protection supervisor who is responsible for

ensuring compliance with the local rules ◾ Identification of radiation protection advisor (RPA) ◾ All employees ◾ Ensure their own and colleagues’ safety ◾ Follow employers safety instructions and use relevant monitoring

and relevant protection devices 213

Radiation Protection and Safety

There must also be designated areas: ◾ X-ray room is a controlled area. ◾ Warning signs must be present and lit when an exposure is taking place. ◾ Entry/exit doors must be closed during exposure and staff must not enter.

Radiation protection of staff Staff must ensure they are not accidentally exposed to radiation and have a legal requirement to protect themselves from risk. This can be achieved in a number of ways: ◾ X-ray and gamma camera rooms are designed to minimise the dose to staff. ◾ Only staff necessary should be present when an exposure is made. ◾ Shielding ◾ Stand behind protective screen when an exposure is made. ◾ Wear personal protective equipment (PPE) when undertaking fluoroscopy or exposures with mobile equipment. ◾ Radioactive sources must be shielded. ◾ Distance ◾ Use the inverse square law to stand far enough away from the X-ray or radiation source. There is a 2 m controlled area from the image intensifier in which the staff must wear appropriate PPE. ◾ Time ◾ Staff should minimise the amount of time spent in fluoroscopy or close to a radioactive source. Staff may also be monitored to ensure they do not exceed the legal dose limits for workers.

Radiation risk assessment ◾ The Employer must perform a “suitable and sufficient” assessment

of the risk to any employee or other person before commencing a new work activity involving ionising radiation. ◾ The risk assessment allows identification of the control measures needed to restrict exposure, to prevent over exposure and protects staff and the public! 214

Physical, Chemical and Biological Effects of Ionising Radiation ◾ If a risk from a particular radiation accident is identified, the

employer must take all reasonably practicable steps to: ◾ Prevent any such accident occurring ◾ Limit the consequences should it occur ◾ Provide employees with the information and adequate training

and equipment necessary to restrict their exposure The Health and Safety Executive have a ’5 step process to risk as­ sessment: 1. Identify the hazards 2. Decide who might be harmed & how 3. Evaluate the risks and decide on control measures 4. Record your findings & implement 5. Periodically review your assessment & update if necessary

DOSE LIMITS ◾ Adult employee per calendar year: ◾ Whole body: 20 mSv ◾ Lens of the eye: 20 mSv ◾ Skin: 500 mSv ◾ Extremities: 500 mSv ◾ Trainees under 18 years: 6 mSv ◾ Others: 1 mSv ◾ Comforters and carers (five years): 5 mSv ◾ Dose to fetus for the remainer of the pregnancy: 1 mSv ◾ Dose to abdomen of woman of reproductive capacity in three-

month period: 1.3 mSv.

PHYSICAL, CHEMICAL AND BIOLOGICAL EFFECTS OF IONISING RADIATION We now know that as X- and gamma rays penetrate matter they have the ability to ionise it (which is why we call them ‘ionising radiations’). 215

Radiation Protection and Safety

When energy is absorbed from the radiation by matter, it has the ability to split electrons from their associated atoms or molecules, leading to the production of a negative ion (the dissociated electron) and a positive ion (the remainder of the atom or molecule, now minus one of its negatively charged electrons). This initial process can subsequently give rise to a four-stage process: ◾ Stage 1. Energy from the photon of radiation is absorbed by atoms or molecules of the material giving rise to the physical process of ionisation. ◾ Stage 2. The electron(s) ejected from the atom(s) may have sufficient energy to ionise other atoms in which case secondary ionisation has occurred and highly reactive free radicals are produced. ◾ Stage 3. The free radicals can then go on to interact with other cellular chemicals to produce chemical changes. ◾ Stage 4. The chemical changes can cause the cell function to change, which is itself a biological change within tissue. The biological effects of ionising radiation can be grouped under two headings:

CANCER EFFECTS (STOCHASTIC) The likelihood of this effect occurring is governed by the laws of probability and the chance of it occurring is therefore directly related to the dose of ionising radiation received. The guiding principles arising from this probability-related effect are two fold: 1. There is no ‘threshold’ limit below which a stochastic effect cannot occur. However, the greater the dose received, the greater the risk of the effect occurring. It therefore follows that if we calculate dose equivalent limits for the radiations at our disposal, we can use them to calculate the statistical risk that an individual runs by exceeding that limit. 2. While the chance of a cancer effect occurring may be probabilityrelated, the severity of any resultant effect will not be. We know from statistical records that populations which are exposed to ionising radiations (over and above natural background radiation which 216

Cancer Effects (Stochastic)

everyone receives to a lesser or greater extent, depending largely on where they live) have a greater risk of developing cancer. However, if the disease is contracted, its severity will not be related to the radiation dose received. There is a much reduced risk of developing a cancer following a low-dose chest X-ray than there would be following a high-dose examination, such as a CT scan of the chest. However, if a cancer did result from either examination, its rate and extent of progression would be related to the genetic make up of the individual, not the dose of radiation which may have caused it. There are two types of cancer effects resulting from exposure to ionising radiations: The risk of developing a cancer from a medical irradiation is related to the dose of radiation received as we have said previously. However, the general risk is very small across all examinations. It is clear from research data that some tissues are more sensitive and therefore at greater risk of developing a malignancy than others. This can be helpful in that most cancer cells develop more quickly than their surrounding tissues and therefore may respond to the use of ionising radiation as a treatment to destroy or shrink the tumour. The downside though is that some naturally occurring tissues, such as stem cells and fetal cells develop extremely rapidly and are therefore at greater risk also. It is for this reason that we have radiation protection regulations which protect vulnerable groups such as pregnant women and those who think they may be pregnant (see Special protection measures for women of reproductive capacity), and we aim to minimise the risk of cancer development to all by applying the ALARP principle every time we undertake an X-ray examination. Radiation induced damage to the genetic material carried in an individual’s germ cells carried in the ova or spermatozoa may cause genetic effects in children. Radiation-induced damage to the genetic material carried in an in­ dividual’s germ cells carried in the ova or spermatozoa. Ionising radiation can cause biological damage to genes which may cause faulty combina­ tions of chromosomes to occur. Such damage can only show itself in future generations and may or may not become part of the broader human gene pool depending on whether the resulting individuals themselves, go on to reproduce. 217

Radiation Protection and Safety

The aim in any case is to make every effort to reduce the radiation dose received to the gonads to an absolute minimum – using sound radiographic techniques and appropriate protective devices – as the effects are unpredictable and potentially far-reaching.

Deterministic (tissue reactions) effects 1. In most cases the severity will increase with increasing dose, e.g. radiation burns erythema will increase with increasing dose (i.e. the effect is proportional to the dose). 2. There is a threshold below which the effect will not occur. It is reasonable to say that heritable effects will always occur if a threshold dose is achieved (unlike stochastic effects which may occur whatever the dose received). Deterministic effects are associated with what we would consider to be in diagnostic imaging terms, very large radiation doses of an order of magnitude only experienced as a result of nuclear or other radiation-related accidents. They may occur in radiotherapy. Examples of deterministic effects might be: ◾ Radiation burns (also called ‘erythema’ or skin reddening) ◾ Radiation-induced cataracts and sterility Erythema will occur from 1 to 24 hours after a radiation dose of 2 Sieverts and cataracts will occur after 2–10 Sieverts have been received, but may take many years to develop. To apply some perspective for diagnostic radiographers, 1 Sievert is equivalent to around 50,000 chest X-rays or 100 whole-body CT scans. In very basic terms, the aim of the radiographer when undertaking a radiographic investigation is, first and foremost, to secure an optimum di­ agnostic image while at the same time, minimising the risk of a cancer event and subjecting the patient to no risk at all of a heritable event occurring.

STANDARD OPERATING PROCEDURE FOR WOMEN OF REPRODUCTIVE CAPACITY There should be a standard operating procedure (SOP) which sets out the duties and responsibilities under the IR(ME)R 2018 regulations for 218

Standard Operating Procedure for Women of Reproductive Capacity

practitioners and operators. This identifies the procedure for imaging investigations involving the use of ionising radiation on women of childbearing age from referrers. Any irradiation of a fetus should always be avoided whenever possible and alternative imaging techniques/ diagnostic procedures should be considered before a decision is taken to expose a woman of childbearing age to ionising radiation. Where pregnancy is not necessarily suspected by the woman herself, checks must be made to ensure that there is no possibility of pregnancy before the Operator proceeds with the examination, particularly if the abdomen or pelvic area is to be irradiated. All women between the ages of 12 – 55 years must be consulted about the possibility of pregnancy prior to undergoing imaging investigations using ionising radiation. It may be necessary to extend this age range if there is a possibility of pregnancy either earlier in the age range or extended beyond it. IR(ME)R includes the requirement to make enquiries of individuals of childbearing potential, this should accurately reflect the diversity of the gender spectrum in the population. If in any doubt ask the patient if there is a possibility they could be pregnant.

Pregnancy enquiry procedure The patient should be asked the date of their last menstrual period (LMP). If possibility of pregnancy exists, order a pregnancy test (e.g. if patient is to receive substantial pelvic irradiation) and obtain results before the pelvis is irradiated. The practitioner (in consultation with referrer) makes a determination whether or not to do the X-ray ex­ amination based on the clinical need of the examination. This process has replaced the ‘28 day rule’ and ‘10 day rule’ in most departments, however an explanation of these rules may be useful. For high dose examinations, e.g. CT scans on the abdomen the 10 day rule is applied. It is unlikely that a female patient will become pregnant in the first 10 days of her menstrual cycle. There should be a clearly defined protocol stating which examinations are classed as ‘high dose’ where the 10 day rule is used. Other examinations which are relevant should apply the 28 day rule. Table 10.1 illustrates a typical pregnancy flowchart; taken from Clark’s Pocket Handbook for Radiographers 2nd edition (Figure 11.1).

219

Radiation Protection and Safety Pregnancy Ask patient: ‘Are you or might you be pregnant?’

Answer: Yes

Answer: Unsure

Answer: No, not pregnant

For low-dose procedures, e.g. plain radiography of abdomen, spine or extremities

For high-dose procedures, e.g. CT of abdomen or pelvis, barium enema

Record LMP. Proceed with examination providing LMP was within previous 28 days”

Patient has missed a period/their period overdue, i.e. more than 28 days” ago

Patient has not missed a period, i.e. period less than 28 days” ago

Re-book examination within first 10 days of onset of menstural cycle

Record LMP. Proceed with examination providing LMP was within previous 28 days”

Review justification for examination with referring clinician

If overriding clinical reasons for examination exist, then proceed using dose reduction strategies

If patient subsequently is found to be pregnant, then review justification for procedure with referring clinician

Figure 11.1 A typical ‘pregnancy rule’ flowchart. CT, computed tomography; LMP, last menstrual period.

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Standard Operating Procedure for Women of Reproductive Capacity

Minimising radiation risks in pregnancy If the examination does proceed the relatively small radiation risk to the patient/fetus will be outweighed by the benefit of the diagnosis and subsequent treatment of potentially life-threatening or serious condi­ tions. These could present a much greater risk to both parties if left undiagnosed. To minimise the risks when examining pregnant women, the radiographer should adopt the following strategies: ◾ Use of the highest imaging speed system available, e.g. 800 speed or equivalent settings for computed radiography/direct digital radiography ◾ Limiting collimation to area of interest ◾ Use of shielding (can the uterus be shielded without significant loss of diagnostic information?) ◾ Use of the minimum number of exposures to establish a diagnosis ◾ Use of projections that give the lowest doses

Inadvertent fetal exposures These can occasionally happen in the imaging department and may be an oversight when the patient was not asked if they could be pregnant or when the operator followed the correct procedures and was assured that the patient was not pregnant. All accidental or unintended ex­ posures to the fetus must be investigated by the RPA usually assisted by the MPE. They also need to be reported to the Care Quality Commission (CQC) to determine what lessons can be learnt to prevent the incident happening again. The patient should be counselled and a duty culture of candour applied (regulation 20). According to ICRP 84, termination of pregnancy at fetal doses of less than 100 mGy is not justified based upon radiation risk. At fetal doses between 100 and 500 mGy, the decision should be based upon the individual circumstances.

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Radiation Protection and Safety

REPORTING OF RADIATION INCIDENTS The Ionising Radiation (Medical Exposure) Regulations 2017 and the Ionising Radiation (Medical Exposure) Regulations (Northern Ireland) 2018) are designed to protect people while undergoing examinations and treatment using ionising radiation. When there is an accidental or unintended exposure to ionising radiation, and the IR(ME)R employer knows or thinks that it is significant, they must investigate the incident and report it (under Regulation 8(4)).

Significant accidental or unintended exposures (SAUE) These are categorised by the CQC as: ◾ Accidental exposure: an individual has received an exposure in error when no exposure of any kind was intended, e.g. incorrect patient. ◾ Unintended exposure: although the exposure of an individual was intended, the exposure they received was significantly greater or different to that intended. eg, in the dose received, the modality or technique carried out, anatomy, radiopharmaceutical or timing of exposure. These can happen for many reasons including procedural, systematic or human error. Unintended exposures can also include exposures to individuals re­ sulting from an equipment malfunction. There must be a record of the investigations into the incident and what they found. Records in accordance with your local procedures and with Regulation 8(3) must be kept. This must be done regardless of whether an incident needs to be notified to the appropriate enforcing authority or not. For SAUE incidents, the employer must send a report on the outcome of the investigation to the appropriate enforcing authority. The report should include: ◾ What happened ◾ An estimate of the dose(s) received by the exposed individual(s) 222

Reporting of Radiation Incidents ◾ A detailed account of the root causes and contributory factors ◾ Whether any similar previous incidents have occurred where

◾ ◾ ◾ ◾

individuals might have been over or under exposed, or if there are any trends that show a possible systematic failure Whether local duty of candour requirements have been met Whether local procedure, required under Regulation 8(1), schedule 2(l), has been applied Any learning from the investigation and how this has been shared The corrective measures adopted and/or remedial actions implemented to reduce the likelihood or prevent this type of incident from happening again.

Under-exposures Regulation 8(4)(b) requires employers to make notifications of radio­ therapeutic exposures that are significantly lower than intended. This includes molecular radiotherapy, brachytherapy and intraoperative therapy. You do not need to make a notification of exposures lower than intended for non-radiotherapeutic modalities.

Over-exposure An overexposure may be defined as an exposure in excess of a relevant dose limit (employee or member of the public). This is an exposure much greater than intended, therefore is a dose to a patient that exceeds the guideline multiplying factors in Health and Safety Executive (HSE) publication PM77. The notification guidelines are broadly representative of patient exposure, i.e. effective dose or mean glandular dose. Suitable mea­ surements for determining these quantities are: ◾ Dose–area product (DAP) metre ◾ Duration of exposure ◾ Product of tube current and time (mAs) ◾ Volume of tissue irradiated ◾ Activity administered

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Radiation Protection and Safety

MCQs 1. Deterministic effects of radiation include: a. Cataracts b. Sterility c. Erythema d. Leukaemia. 2. Stochastic effects of radiation include: a. Carcinogenesis b. Heridity c. Erythema d. Leukaemia. 3. Manmade radiations account for what percentage of background dose: a. 3% b. 10% c. 14% d. 20%. 4. The person responsible for justification of an exposure to X-rays is the: a. The employer b. The operator c. The practitioner d. The referrer. 5. The operator is responsible for: a. Setting DRL’s for each examination b. Writing the local rules c. Optimising the radiation dose to the patient for each examination (ALARA) d. Training referrers to fill in the request form correctly.

224

MCQs

6. The controlled area extends to which distance from the X-ray source (tube) a. 1 metres b. 2 metres c. 3 metres d. 4 metres. 7. Designated workers are defined as those employees who are likely to receive: a. 20 mSv per year b. 1 mSv per year c. 5 mSv per year d. 6 mSv per year. 8. Cancer effect of ionising radiation have the following characteristics: a. No threshold dose below which it cannot occur b. A threshold dose below which it cannot occur c. The same risk regardless of the radiation dose d. The effect is proportional to the dose. 9. What does SAUE stand for in relation to radiation incidents: a. Sustained accidental or unintended exposures b. Significant accidental or unsuitable exposures c. Sustained accidental or unsuitable exposures d. Significant accidental or unintended exposures (correct one). 10. What is the aim of the X-ray examination? a. Optimum images with no risk b. Optimum images with minimum risk c. Optimum images regardless of the radiation dose d. Optimum images with risk of heritable effects.

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CHAPTER 12 BENEFIT-RISK INTRODUCTION The aim of this chapter is to give the practitioner an understanding of the basic principles of benefit-risk. It is essential that any practitioner using ionising radiation has the knowledge and skills to enter into a dialogue with patients concerning benefit-risk. You need to comprehend and be able to explain the benefits of examinations using X-rays and the factors affecting radiation dose Learning objectives The student should be able to: ◾ Understand and explain the basic principles of benefit-risk ◾ Have a dialogue with patients and staff on the benefit-risk of examinations using X- and gamma-radiations

BENEFIT-RISK ANALYSIS Benefit-risk analysis is one of the best tools for managing risk on a daily basis in all walks of life (you probably use it for yourself dozens of times a day), but its value as a key weapon in the armory of the radiographer has to be considered differently as you will be utilising it in your professional capacity on a daily basis, not for your own benefit, but for that of others. While benefit-risk analysis can, at least in theory, be applied for every patient examination we undertake, the same cannot be said when we 227

Benefit-Risk

consider the radiation dose that workers such as radiographers and other health professionals receive. There is no benefit for these people from the dose of radiation they could potentially receive (if you discount earning an income for now) and therefore radiation protection has to be considered differently for this group of people. Risk analysis is still a very big part of the process of protecting radiation workers but in this case, has to be considered as a non-beneficial risk. Under the Management of Health and Safety at Work Regulations 1999 (MHSWR), ‘all employers must undertake a risk assessment and this should consider the risks from radiation in the workplace. Under IRR17 employers must undertake a radiation risk assessment before they start any new activity with ionising radiation, giving particular consideration to exposure levels and accident situations.’ Once the work commences two regulations also come into force: ◾ Regulation 3, which requires the recording of the assessment (if there are 5 or more employees) and the maintenance of the risk assessment to keep it up to date where there has been a significant change in the matters to which it relates. ◾ Regulation 5 of MHSWR also requires employers to make arrangements for effective planning, organisation, control, monitoring and review of preventative and protective measures, including those for restricting exposures to ionising radiations.’ We now have a clear understanding that the use of X-rays carries with it, on the one hand, associated dangers or risks, yet on the other, distinct benefits for mankind (can you imagine a twenty-first century hospital providing its services without using X-rays or radioisotopes for that matter in either a diagnostic or therapeutic capacity?). It therefore follows that it is necessary to control some of the factors associated with the use of X-rays. Radiation legislation in the UK requires that no one should be irradiated intentionally unless there is a valid clinical indication. In making this judgement, the clinician must determine that the benefit to the patient in having the examination will outweigh the risk. This is the process of justification and if the result of having the X-ray examination will change the clinical management of the patient then that examination can be said to be justified. You do have to be aware, however, that requesting an X-ray to exclude an injury or a disease process can be perfectly justifiable also, 228

Benefits of Examinations Using X-rays and Gamma Photons

providing it is not possible to be reasonably sure of the outcome by other, less risky means. An example of ‘less risky’ in this case might be a thorough clinical examination. So justification is the first step in a radiation protection strategy because the best way to reduce the radiation dose to the patient is not to undertake the examination in the first place, if this is considered to be an appropriate course of action. The justification of an X-ray examination is a two-stage process and should be the responsibility of both the requesting clinician along with the radiographer responsible for undertaking the request. Determining whether an examination is justified will vary depending on factors such as the age of the individual, the pregnancy status or the availability of other diagnostic procedures.

BENEFITS OF EXAMINATIONS USING X-RAYS AND GAMMA PHOTONS The benefits of having a X‐ray examination, CT or study using radioactive products are associated with managing the treatment and/or diagnosis of the patient. These may include: ◾ Saving the person’s life by providing the correct diagnosis which may not be able to be made without the use of X-rays, e.g. chest X-ray examination to demonstrate extent of pathology or nuclear medicine bone study to see if a primary tumour has metastised. ◾ Giving the patient the correct treatment as a result of the correct diagnosis. ◾ Eliminating disease/disorders which affect the management of the patient, e.g. determining if a patient has a fracture and how best to manage the patient’s fracture. ◾ Managing the treatment of a patient by imaging the response to treatment, e.g. images to determine the effect of radiotherapy. ◾ Making a diagnosis with an examination which has less morbidity and mortality than an alternative test, e.g. computed tomography (CT), rather than invasive surgery. 229

Benefit-Risk

RISKS FROM X- AND γ-RADIATION There is no safe radiation dose limit and all exposures to ionising radiations carry some risk. The purpose of a benefit-risk discussion should therefore justify the examination to the patient, discuss the need for the examination and quantify the risk. The Health Protection Agency produces an excellent leaflet, called ‘X-ray Safety Leaflet’ which outlines the common imaging procedures and levels of risk for common X-ray and isotope procedures. To quote them: You will be glad to know that the radiation doses used for X-ray examinations or isotope scans are many thousands of times too low to produce immediate harmful effects, such as skin burns or radiation sickness. The only effect on the patient that is known to be possible at these low doses is a very slight increase in the chance of cancer occurring many years or even decades after the exposure.

It also gives approximate estimates of the chance or risk that a particular examination or scan might result in a radiation-induced cancer later in the lifetime of the patient. There are a number of ways of describing the risk. These include: ◾ Equivalent background dose, expressed in equivalent period of natural background radiation, e.g. a few days to several years. ◾ Statistical risk, expressed in numbers, e.g. risk of cancer is 1 in 1,000,000. ◾ Comparisons to general risks of cancer, i.e. the population have a 1 in 2 chance of getting cancer. ◾ Comparison to everyday activities: ◾ For example, airline flights are very safe with the risk of a crash being well below 1 in 1,000,000. ◾ A chest X-ray exposes you to the same risk as a 4-hour flight. ◾ Smoking or drinking alcohol. ◾ Driving or undertaking dangerous sports, such as skydiving. ◾ Lost life expectancy, given in days.

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Risks from X- and γ-Radiation Table 12.1 Table of typical radiation dose and risk for common examinations.

Radiograph Abdomen AP Chest AP Chest LAT Chest PA Cervical spine AP Cervical spine LAT Knee AP Knee LAT Lumbar spine AP Lumbar spine LAT Pelvis AP Shoulder AP Skull AP/PA Skull LAT Thoracic spine AP Thoracic spine LAT

ESD per radiograph (mGy) 4 0.2 0.5 0.15

0.3 0.3 5.7 10 4 0.5 1.8 1.1 3.5 7

No. of rooms 167 53 47 285

DAP per radiograph (Gy cm2) 2.5 0.15 0.1 0.15 0.15

40 32 192 185

1.5 2.5

204 34 21 21 104 104

2.2

1.0 1.5

AP, antero-posterior; DAP, dose-area product; ESD, entrance skin dose; LAT, lateral; PA, postero-anterior.

Table 12.1 demonstrates the recommended national reference doses for individual radiographs on adult patients. (Taken from table 28 from Health Protection Agency (2012). These can be used to support the benefit-risk discussion The purpose of managing radiation dose in diagnostic procedures using X-ray or gamma radiation is to avoid deterministic health effects and to reduce the probability of stochastic health effects of ionising radiation.

231

Benefit-Risk

POSSIBLE EFFECTS OF IONISING RADIATIONS ON HUMAN CELLS If the DNA within cell(s) is damaged there are three possible outcomes: 1. The cell(s) die. The death of a few cells in the millions within the human body has no significant effect. 2. Significant numbers are damaged to observe a clinical effect, either immediately (erythema) or delayed (cataracts). 3. The damage is incorrectly repaired leading to mutation of the DNA. These cell(s) may subsequently die or may lead to radiationinduced malignancy. Practitioners must be educated in the benefits and risk and the radiation dose given from X-ray procedures. Whichever one you decide to use, make sure you have the correct information at hand and always discuss benefit versus risk. You can also use the principles of justification and optimisation to inform the patient that X-rays are not undertaken without a valid clinical reason. Any examinations will optimise the exposure and use the lowest dose compatible with making a diagnosis (ALARP). Don’t forget the patient always has the right to decide not to have the examination. Table 12.2 is a list of common examination’s, the effective dose and additional lifetime risk of a fatal carcinoma. Table 12.2 Table of common X-ray examinations with the effective dose in mSv and the additional lifetime risk of a fatal carcinoma.

Examination

Typical effective dose (mSv)

Risk*

Hand/foot Chest Mammography Abdomen Lumbar spine CT head Barium enema CT body

0.01 0.02 0.06 0.7 1.3 2 7.2 9

1 in a few million 1 in 1 000 000 1 in 300 000 1 in 30 000 1 in 15 000 1 in 10 000 1 in 2800 1 in 2200

*Additional lifetime risk of fatal cancer. CT, computed tomography.

232

MCQs

MCQs 1. Discussions with patients should always include: a. The risks b. The benefits c. Benefit-risk d. Statistical analysis of the risk. 2. Risk of X-rays include: a. Saving the person’s life b. Giving the patient the correct treatment as a result of the correct diagnosis c. Making a diagnosis with an examination which has less morbidity and mortality than an alternative test d. Hair loss. 3. The risk of developing a carcinoma in the general population is: a. 1 in 2 b. 1 in 3 c. 1 in 4 d. 1 in 5. 4. The everyday risk associated with a chest X-ray is approximately flying for: a. 1 hour b. 2 hours c. 3 hours d. 4 hours. 5. An a. b. c. d.

X-ray described as minimal risk is: Less than 1 in 1,000,000 1,000,000 to 100,000 100,000 to 10,000 10,000 to 1000.

6. The following are ways to minimise the risk from X-rays: a. Justification b. Optimisation 233

Benefit-Risk

c. Observing DRLs d. All of the above. 7. The purpose of managing radiation dose is: a. To avoid deterministic effects b. To avoid stochastic effects c. To avoid deterministic effects and minimise stochastic effects d. To increase the probability of stochastic effects. 8. Who is jointly responsible for justifying an exposure to X-rays? a. Practitioner and referrer b. Referrer and employer c. Patient and employer d. Patient and referrer. 9. Which of the following statements is true? a. The patient must have an X-ray if the referrer sends them to the imaging department b. The patient has the right to refuse an X-ray c. Practitioners must X-ray patients if the referrer tells them to do it d. Requesting physicians have no obligation to inform the patient of the risk from an X-ray. 10. For an X-ray to be justified it must: a. Clearly demonstrate the suspected injury or pathology b. Have a risk of 1,000,000 to 100,000 of resulting in a radiation-induced cancer c. Change the management of the patient d. Have an exposure within the DRL.

234

ANSWERS TO MCQ’S CHAPTER 1 1. 2. 3. 4. 5. 6. 7. 8. 9. 10.

c d a c c d d c c d

CHAPTER 2 1. 2. 3. 4. 5. 6. 7. 8. 9. 10.

b c d b b a b c c b

CHAPTER 3 1. b 2. d 235

Answers to MCQ’s

3. 4. 5. 6. 7. 8. 9. 10.

a a c b a c d c

CHAPTER 4 1. 2. 3. 4. 5. 6. 7. 8. 9. 10.

d a b b c c a a d a

CHAPTER 5 1. 2. 3. 4. 5. 6. 7. 8. 9. 10.

236

b c d a c a c b a d

Answers to MCQ’s

CHAPTER 6 1. 2. 3. 4. 5. 6. 7. 8. 9. 10.

a a b d c c b d a c

CHAPTER 7 1. 2. 3. 4. 5. 6. 7. 8. 9. 10.

a c d d a c c b a c

CHAPTER 8 1. 2. 3. 4. 5. 6.

c b d c a b 237

Answers to MCQ’s

7. 8. 9. 10.

a c c d

CHAPTER 9 1. 2. 3. 4. 5. 6. 7. 8. 9. 10.

b b d b d c c a a c

CHAPTER 10 1. 2. 3. 4. 5. 6. 7. 8. 9. 10.

238

b a c c d d a b a d

Answers to MCQ’s

CHAPTER 11 1. 2. 3. 4. 5. 6. 7. 8. 9. 10.

d c c c c b d a d b

CHAPTER 12 1. 2. 3. 4. 5. 6. 7. 8. 9. 10.

c d a d b d c a b c

239

CHAPTER FORMULAS CHAPTER 1 Magnification =

Area of unsharpness =

FRD FOD Focus × ORD FOD

CHAPTER 2 mAs × kVp4 mAs × kVp4 = grid factor × FRD2 grid factor × FRD2 10,000 100 a = r sin

1

1 d2

Size of image A B C FRD = = or = a b c FOD Size of the patient

241

Chapter Formulas

CHAPTER 3 (mass number )12 C (atomic number )6 1 Newton = 1 kg × m/s2 1 joule = Newton × metres

1 A = 1 coulomb of charge flowing/s 1 becqueral = 1 disintigration per second n

p+

+

p

n+

+

=

+

+

CHAPTER 5

=

+

Attenuation = Absorption + Scatter Z3 1 E3 242

Chapter Formulas

Z3 E3 1 E electron density

CHAPTER 7 Ug =

Focal spot size × ORD FOD

CHAPTER 10 HU = 1000 × water

water air

243

INDEX A absorbed dose (of radiation) 141–142; definition in grays 146; gray (Gy) as unit of 141, 146; lumbar spine 144–145 active matrix array (AMA) 108–109 alpha particle emission (α - emission) 34–35; daughter nucleotide after 35; penetrating power of 38; quarks and 34, 35 American Association of Physicists in Medicine (AAPM) 150 ampere (A) (unit of electrical current) 44 analogue image intensifier (II) 118 analysis of variance (ANOVA) 24 anode assembly (part of X-ray tube) 52–54 as low as reasonably practicable (ALARP) 2, 125, 146, 147, 150, 211, 217, 232 atomic balance 32–33 atomic number 28 atomic structure 27–28; electrons in 28, 29–30; neurons in 28; protons in 28; quarks in 28; shells of atoms 29, 30; sub-atomic particles 28, 29 atomic symbols 32 attenuation (of an X-ray beam) 74–77, 176, 177; Compton scatter process 78, 83–86; in diagnostic radiography 77–86; elastic scatter process 78; equals absorption plus scatter 74–75, 77–78; linear attenuation coefficient 75–76; mass attenuation coefficient 76–77; pair production process 78; photoelectric absorption process 78–83

automatic exposure control (AEC) circuits 93–96; good collimation essential for use 95 automatic exposure devices (AEDs) 140 B basic iterative back projection 186 becquerel (SI unit of radioactivity) 34 benefit-risk analysis 117–229; applied to radiation exposures 227–229; as tool for managing daily risks 227–228; for X-ray procedures 230–231 beta particle emission (β - emission) 34, 35–36; beta decay processes 35, 36; beta minus and beta plus types 35–36; imbalance between neutrons and protons 35; penetrating power of 38 binding energy of an electron 31; determination of 31; nucleus as strong attractive force for electrons 31 Bolus tracking 204 Bremsstrahlung curve 65–66; size of 65 Bremsstrahlung X-ray production 58, 62–64; beam outputs 63–64; factors influencing intensity and quality of X-ray beam in 65–70; heterogenous beam from 63; occurrence of 62 bucky assemblies/antiscatter grids 7–8; grid frequency 8; specifications of different grids based on 7–8; used to reduce noise and improve contrast 7–8 C cancer effects (stochastic) from ionising radiation 216–218; 245

Index deterministic (tissue reactions) effects 217; from medical irradiation 217–218; as probability-related effect 216–217; types of 217–218 cardiac gating software 203–204 Care Quality Commission (CQC) 213, 221, 222 cathode assembly (filament) (part of X-ray tube) 54–55 characteristic X-ray production 58–62; excitation in 58; homogeneous beam from 62; identification of materials in 60; ionisation in 58–60 charged coupled device (CCD) technology 105–107, 111–112; advantages of 111–112; compared to TFT technology 111; design considerations 112; linked to image intensifier (II) 118; optical fibre coupling 112–113, 120; optically coupled by mirror and high-quality lens 114–115; slot scan techniques 112–114; solid state X-ray image intensifier (SSXII) and 120 chi-squared test 24 Clark’s Pocket Handbook for Radiographers, 219 compounds 33 Compton scatter 77; as attenuation process 78, 83–86; coefficients 85–86; process that leads to production of 83–84 computed radiography (CR) 2, 9; amount of attenuation based on Hounsfield number 176–177; compared to IDR 108; criticisms of 105; ideal exposure for 133; as indirect system 101; necessity of exposure indicators in 150–151; reader 102–103; scanning detector 103; time to read imaging plates 105; typical resolutions 103–105; 246

uses imaging plates to produce digital radiographic images 101, 102 computed tomography (CT): benefits of compared to risks from 229–231; contrast resolution (CR) 195–198; principles of 176 computed tomography (CT) detectors 89 computed tomography (CT) scan 175; beam hardening limitation 204–205; cone beam artifact limitation 204, 206; description of radiation output from 146; partial volume artifact limitation 204, 205–206; process 194–195; spatial resolution (SR) 195–198 conduction (as form of heat transfer) 40 contrast resolution (CR) 198–199; mAs factor 198–199; matrix size factor 198; slice thickness factor 198, 199–200 convection (as form of heat transfer) 40 coulombs per kilogram of air (C/kg) (unit of exposure for air) 140 Currie (old unit of radioactivity) 34 D “Dated radiography detectors — a technical overview” (Lanca & Silva) 120 demagnification 115 deoxyribonucleic acid (DNA) 231–232; radiation effects on 231–232 detective quantum efficiency (DQE) 9, 90–91, 132, 152; signal-to-noise ratio (SNR) 91, 92, 161–162 detector dose indicators (DDI) 148 detector exposure index 150; 152

Index diagnostic imaging: factors affecting quality of image and/or radiation dose 2 diagnostic radiography: attenuation in 77–86; Compton scatter process 78, 83–86; photoelectric absorption process 78–83 diagnostic reference levels (DRLs) 146–147, 212, 213 diamentor (DAP) reading 11 digital fluoroscopic systems 118–120; fluoroscopic flat panel detectors 119–120; solid-state Xray image intensifier (SSXII) 120 digital images: brightness and contrast in viewing 133–134; checks for optimal exposure 133–134; lookup table (LUT) 133; recorded as raw data 155–157; scatter effects on contrast 136 digital imaging and communications in medicine (DICOM) 171 digital radiography (DR) 2, 9; ideal exposure for 132; necessity of exposure indicators in 149–150; time to read imaging plates 105 digital X-ray detectors (photostimulable storage phosphor (PSP)) 9 direct digital radiography (DDR) 100, 116–118; TTF technology 116–117; works similar to ionisation chambers 116 direct ionisation (to detect and measure radiation) 139 directly scanned sequential slices 190–192 dose area product (DAP) metres 139 dose length product (DLP) 145

to measure 143; for X-ray examinations 145 electrical charge 43; as flow of electrons within a conductor 43 electrical circuit 44 electricity 43–44; alternating current 44; direct current 44; electrical charge 43; electrical circuit 44; as flow (current) of electrons around a circuit 44 electrocardiogram (ECG) 201 electromagnetic radiation (EMR) 44–45; measurement of 44 electromagnetic spectrum 46 electronically-balanced atom 32 electronically-imbalanced atom 32–33 electron multiplying CCD (EMCCD) 120 electrons: binding energy of 31; maximum number of in shells 30, 31; nucleus as strong attractive force for 31; orbitals of 29–30; as part of atomic structure 28; shells of related to nucleus 32; sub-shells and 30 elements 33 energy 39; kinetic energy (KE) 39; potential energy (PE) 39 equivalent dose (of radiation) 141; 142–143; sievert (Sv) as unit of 142 excitation 58 exposure calculations, mathematics of 16; absorbed dose gray 142; absorbed dose (gray Gy) vs. equivalent dose (sievert Sv) 142, 143 exposure creep phenomenon 151 exposure indicators (EI) 150–152

E effective dose (of radiation) 143–144; tissue weighing factor

F fetal exposures 221 film-screen radiology 150

247

Index filtered back projection (FBP) 187–189 filtered interpolation 193–194 flat panel detectors (FPD) 122, 140 fluoroscopic flat panel detectors: initial problems with 121; recent development improvements 120–121 fluoroscopy units 139 focus to object distance (FOD) 4 focus to receptor distance (FRD) 4 Food and Drug Administration (FDA) 189 force 39 G gamma camera 96, 139 gamma ray emission (γ - emission) 34, 36–37; penetrating power of 38; as by product of either α or β decay 36; X-rays and 37, 38 Gaussian kernel 169 Geiger-Muller tube 139 gold leaf electroscope (quartz fibre dosimeter) 139 graphs 19 H Health and Safety Executive (HSE) 223 Health Protection Agency 230, 231 heat 39–40; conduction as form of transfer of 40; convection as form of transfer of 40; radiation as form of transfer of 40–41; X-ray tube interactions producing 57–58 HL-7 (health level 7) standard 171 hospital information system (HIS) 171 Hounsfield scale 177–178; tissue types and 177–178 I ICRP publication 84 termination of pregnancy 221 248

IEC exposure index 149, 151 IEEE.802.3 (ethernet) standard 171 IEEE.802.5 (token ring) standard 171–173 image display 155; manipulation of 155, 159–162; standards 168–170 image interpolation 158–159 image manipulation 155, 159–162; tools 160–161 image production: distortion aspect of 129–130; evaluation of exposure indicator (EI) 133, 134; formation of penumbra (unsharpness) 127–128; fourstage process of 97–98; geometry of 4–8, 127–136; ideal set-up 2, 127–128; magnification aspect of 4–5, 129–130; principles of 1–2; resolution/definition aspect of 4; skill of practitioner in 11; unsharpness aspect of 4, 5–8; variations within human body 135; visual display units (VDU) 10 image production pathway 155 image quality: characteristics for evaluating 126–127; definition of optimum and poor 126; matrix size effects 158; movement unsharpness 132; signal-to-signal noise ratio as assessment of 130–131; subjective 126; unsharpness 131; unsharpness measurement 133 image reconstruction processes: edge enhancement by high-pass spatial filtering 133, 160, 169–170; issues in 187; noise reduction by background subtraction 160, 162–164; noise reduction by low-pass spatial filtering 160, 167–168; simple edge enhancement 174; spatial domain filtering for smoothing

Index and sharpening 164–166; windowing 160, 161; xy dimension raw data creation 179–189; z dimension raw data creation 179, 190–194; zooming and enlarging 160, 161–162 image storage 155 image transmission 155 imaging plates (photostimulable phosphors (PSP)) 101; construction 101–102 imaging system: resolution of 131; spatial resolution of 131 indices 18; multiplication of 18 indirect digital radiography (IDR) 101, 105–112; compared to CR technology 108; systems used in 105–106; workings of 106–107 inherent filtration 68 International Commission for Radiation Protection (ICRP) 150 International Electrotechnical Commission (IEC) 150 international system of units 17; derived SI units 17; SI base units 17; units used in radiography 17 inverse square law for radiation 23, 24 ionisation chambers 91–92, 149; automatic exposure control (AEC) circuits as 93–96; constituents of 91–92; phasing out of and replaced by virtual chambers 95; workings of 91–93 ionisation of air 140; exposure as traditional measure of 140 ionisation radiation process 9–10, 45, 58–60; biological effects of 216–218; effects of on human cells 231–232; four stages in 216 Ionising Radiation (Medical Exposure) Regulations (Northern Ireland) 222 The Ionising Radiation (Medical Exposure) Regulations 2017 (IR (ME)R 2017) 210, 211–211, 222

Ionising Radiations Regulation (IRR17) 11, 210–211 ions 33 isotopes 33 iterative filtered back projection (IFBP) 177, 187–189 J joule (J) (as unit of measure of work and heat) 39 K kerma dose (of radiation) 141 keV as X-ray tube voltage 66; effects of increasing 66–67 kinetic energy (KE) 39 L Laplacian filter 169–170 large field detectors (LFDs) 99–100 linear attenuation coefficient 75–76 linear energy transfer (LET) 148 line focus principle 19–21; relationship between real and apparent focus 20–21 logarithms (logs) 18–19; common type of 19; natural type of 19 look-up table (LUT) 10, 133; luminescence (to detect and measure radiation) 139 M mA as measure of current flowing across X-ray tube 66; beam intensity is proportional to 66; beam quality unaffected by changes to 66 Magnetic Resonance Imaging (MRI) 43 magnetism 42–43 magnification (of image production) 4–5; formula 5 Management of Health and Safety at Work Regulations 1999 (MHSWR) 228 249

Index Mann-Whitney U test 24 mass attenuation coefficient 76–77 mass number 28 measurement prefixes (powers) 18; division of 18; multiplication of 18 Medical Physics Expert (MPE) 210, 212, 221 Modern Direct Radiography (DR) systems 8; dose reading on 11 modulation transfer function (MTF) 132 multipoint interpolation 193–194 multislice CT detectors 99 multislice CT scanners 190, 199–212; benefits 202–203 N newton (N) (unit of measure of force) 39 nuclear magnetic resonance (NMR) 43 nuclear medicine: examinations 209; scintillation crystals/ photocathode multiplyers in 96 O object to receptor distance (ORD) 4 optically stimulated luminescent dosimeter (OSLD) 139, 145–146, 148–149; compared to TLDs 145, 146; sensitivity range 144–145 P penumbra (unsharpness), principle of 127–128; measurement of 129–130 periodic table 32 personal protective equipment (PPE) 214 photoelectric absorption 78–83; coefficients 80–83 photostimulated luminescence (PSL) 140 picture archive and communication systems (PACS) 171, 172 250

pipeline principle of image reconstruction 187–189 Positron Emission Tomography (PET) scanning 78 potential energy (PE) 39 power 39 pregnancy: minimising risks in 221; procedure for imaging women of childbearing age 219–221; termination of after fetal dose 221 pregnancy enquiry procedure 219–221; check last menstrual period (LMP) 219; protocol 220 publication PM77, Health and Safety Executive (HSE) 223 Q Quantum Detection Efficiency (QDE) 2 R radiation: benefit-risk analysis of workplace exposure 228–229; different types cause different biological damage 147–148; as form of heat transfer 40–41 radiation detection 139–140; linear energy transfer (LET) 147; methods 139–140; relative biological effectiveness (RBE) 147 radiation detectors: characteristics of 90; detective quantum efficiency (DQE) to compare 90–91, 92; large field detectors (LFDs) 99–100; scintillation crystals/photocathode multipliers 96; scintillation crystals/silicon diode multipliers 98–99; types 89 radiation dose 11, 139–153, 231; dose area product (DAP) reading 11; limits 215; recording of 11 radiation exposure (from diagnostic X-rays) 138; accidental exposure

Index 220; cancer risks 139; incident reporting 222–223; notification guidelines 223; sources of 139; underexposures 223; unintended exposure 222 radiation measurement 139–140; methods 139–140 radiation monitors: ionisation chambers 149; optically stimulated luminescent dosimeter (OSLD) 149–150; thermoluminescent dosimeter 149 radiation protection 209; aim of 209; justification step 229; legislation for patients’ protection 211–213; legislation for staff protection 210–211; legislation in UK 228; local rules and systems 213–214 radiation protection advisor (RPA) 210, 213, 221 radiation risk assessment 214–215 radioactive decay 34; Alpha particle emission (α - emission) type 34–35; Beta particle emission (β emission) 34, 35–36; Gamma ray emission (γ - emission) 34, 36–37 radioactive emissions, penetrating powers of 38 radioactivity 34–38; as number of disintegrations per second of a radionuclide 34; process 34 Radiography 122 radiology information system (RIS) 172 radionuclide imaging (RNI) 34; radionuclides examples 37–38 raw data attenuation values 179; creation of xy dimension 179–189; creation of z dimension 179, 189–194 raw data image matrix 156–158; display of 156–157; grid of

values collected on CR/DR detector plate 155–156 ray sum intensity profile 180–181, 185 Regulation 3, Management of Health and Safety at Work Regulations 1999 (MHSWR) 228 Regulation 5, Management of Health and Safety at Work Regulations 1999 (MHSWR) 228 Regulation 8(1), schedule 2(1), Ionising Radiations Regulations 2017 (IRR17) 223 Regulation 8(3), Ionising Radiations Regulations 2017 (IRR17) 222 Regulation 8(4), Ionising Radiations Regulations 2017 (IRR17) 222 Regulation 8(4)(b), Ionising Radiations Regulations 2017 (IRR17) 223 Regulation 17, Ionising Radiations Regulations 2017 (IRR17) 213–214 Regulations for practitioners and operators, The Ionising Radiation (Medical Exposure) Regulations 2018 (IR(ME)R 2018) 219 relative biological effectiveness (RBE) of radiation 148 resolution/definition (in image production) 4, 8; diagnostic quality measured by definition 8; measurement of 8 S safety 209; legislation 210–212; local rules and systems 213–214; radiation dose limits 215; radiation risk assessment 214–215; staff radiation protection 214 scan field of view (SFOV) 179, 180 scintillation (to detect and measure radiation) 139

251

Index scintillation crystals/photocathode multipliers 96 scintillation crystals/photocathode X-ray image intensifier 97 scintillation crystals/silicon photodiode multipliers 98–99 semi-conductor detection (for radiation) 139 significant accidental or unintended exposures (SAUE) 222–223; reports of 221–223 Silva, A 122 similar triangles 21–22 slot scan techniques 112–114, 120; dedicated digital mammography system 114–115; slot scan chest radiography 112–113 solid-state charged couple device (CCD)-based camera 97–98 solid-state X-ray image intensifier (SSXII) 120 sound 41–42 spatial filtering 160; high-pass 160, 169–170; low-pass 160, 167–168 spatial resolution (SR) 195–198; high resolution filter convolution factor 195, 197–198; matrix size factor (xy axis resolution) 195, 194; number of projections factor 195, 196–197; slice thickness factor (z axis resolution) 195 standard operating procedure (SOP) 219 statistics 23–24 T TCP/IP (transmission control protocol/internet protocol) 172 temporal resolution (TR) 203–204 thermoluminescent dosimeters (TLDs) 139, 145, 146, 148 thin-film transistor (TFT) technology 105–110; active matrix array (AMA) in 108–110; 252

compared to CCD technology 111; in direct digital radiography (DDR) 116–117; image interpolation in 157–158 t-test 24 U unsharpness (in image production) 4, 5–8; calculation of 6–7; factors in 5–6; minimising 6–7 V visual display units (VDU) 10; lookup table (LUT) for pixels 10 W waves 41; components of 41; wavelength 41 work 39 X X-ray beam 49; attenuation 74–75; characteristics 2–3; cone beam artefact limitation 204, 206; control of by collimation 3–4; factors in Bremsstrahlung production influencing 65–70; filtration impact on 68–70; hardening limitation 204–205; intensity of inversely proportional to filtration 79; intensity of is measure of number of X-ray photons 65; possibilities for photons leaving the tube 3; quality of proportional to filtration 70; quality of radiation measures overall energy of 65; radiation dose if energy deposited in patient’s tissue 74; understanding effect of filtration to calculate impact on patient 69 X-ray circuit 55–64; function of 55–56 X-ray detection process 96–97 X-ray detectors 9; computed

Index radiography (CR) type 9; digital detectors (photostimulable storage phosphor (PSP)) 9; digital radiography (DR) type 9 X-ray image intensifier 97 X-rays 49; benefits of 229; Bremsstrahlung interaction to produce 58, 62–64; characteristic radiation interaction to produce 58–62; components of 50; gamma emission (γ - emission) and 37; interactions in X-ray tube to produce 58–64; principles of interactions to matter 73–77; risks of 230; risks of compared to benefits of 228, 229–230 “X-ray Safety Leaflet” (Health Protection Agency) 230 X-ray tube 49, 50–55; anode assembly part of 52–54; apparent focus 20; Bremsstrahlung X-ray production in 58, 62–64; cathode assembly (filament) part of 54–55; characteristic X-ray production in 58–62; components of 50; filament circuit part of 54; focal area variable 19–20; focusing cup part pf 55; functions of 50–51; insert (envelope) part of 51–52, 53; interaction of high-energy electrons with matter in 56;

interactions between incoming electrons and outer-shell electrons in tungsten 56–57; interactions producing heat in 57–58; interactions producing X-rays in 58–64; mA measure of current flowing across 66; materials used in construction of 51 X-ray tube shield 51 xy dimension data: filtered back projection (FBP) 179, 187–189; image matrix 179–180; iterative filtered back projection (IFBP) 187–189; matrix size 195, 196; number of projections (samples) 179, 180–181; pixel size 179, 180; Scan Field of View (SOV) 179; simple back projection 179, 181–187 Z z dimension data 190–192; directly scanned sequential slices 190–192; multi-spiral scanning 192–193; spatial resolution (SR) 195; spiral acquisition 193 zooming as image reconstruction process 161–162; Geometric zoom (GZ) 161, 162; Interpolated zoom (IZ) 161–162; Reconstructed zoom (RZ) 161, 162

253