Medical Imaging

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This section covers the production of X-rays, stationary and rotating anode tubes, the control and indicating equipment of an X-ray machine, filters and grids, and the biological effects of X-rays.

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Nature and Production of X-rays • X-rays are electromagnetic radiation of wavelength about 0.01-10 nm and energy of roughly 100 eV to 100 keV, discovered by Röntgen in 1895.

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They are ionising, travel in straight lines at the speed of light, are undeflected by electric and magnetic fields, penetrate matter, cause fluorescence, blacken photographic film and produce biological effects.

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The useful relation is E (keV) = 1.24 / λ (nm). • Production requires three things: a source of electrons (a heated tungsten filament, by thermionic emission), a means of accelerating them (a high potential difference, the kVp, between cathode and anode) and a target to stop them (the anode).

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All of this takes place in an evacuated glass or metal envelope. • Two mechanisms of X-ray production: • Bremsstrahlung ('braking radiation') — an electron passing close to a tungsten nucleus is decelerated and deflected, and the lost kinetic energy appears as an X-ray photon.

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Because the deceleration varies continuously, this gives a continuous spectrum from zero up to a maximum photon energy numerically equal to the applied kVp (an electron that loses all its energy in one interaction).

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Bremsstrahlung accounts for the great majority of the beam in diagnostic radiology. • Characteristic radiation — an incident electron ejects an inner-shell (K-shell) electron of the target atom; an outer-shell electron falls into the vacancy and the energy difference is emitted as a photon of discrete energy characteristic of the target element (for tungsten the K-shell lines are at about 59 and 67 keV).

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It therefore appears as spikes superimposed on the continuous spectrum, and only above the threshold kVp (about 70 kVp for tungsten). • Efficiency: only about 1 % of the electron energy becomes X-rays; about 99 % becomes heat — which governs the entire mechanical design of the tube.

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Efficiency is proportional to Z × kV, which is why a high-atomic-number target (tungsten, Z = 74) is used.

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The X-ray Tube • Cathode: a tungsten filament (usually two, for a fine and a broad focus) in a focusing cup of molybdenum or nickel, negatively charged so as to condense the electron stream onto the focal spot.

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Filament current controls the temperature and hence, by thermionic emission, the tube current (mA), which is the number of electrons and therefore the quantity of X-rays. • Anode (target): tungsten or a tungsten-rhenium alloy on a molybdenum or graphite backing.

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Tungsten is chosen for its high atomic number (74), very high melting point (3,370 °C) and good thermal conductivity.

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In mammography a molybdenum or rhodium target is used instead, to give low-energy characteristic radiation suited to soft-tissue contrast. • The line-focus principle: the target face is angled (typically 6-17°) to the electron beam, so that the actual focal spot bombarded by electrons is large — spreading the heat — while the effective (apparent) focal spot projected towards the patient is small, giving good geometric resolution.

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Effective focal spot = actual focal spot × sin θ.

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The penalty is the anode heel effect: intensity is lower on the anode side of the field because those photons must traverse more target material, a difference of up to 45 % that is exploited by placing the thicker part of the patient towards the cathode. • Other components: the evacuated envelope (a vacuum is essential so that electrons are not scattered and the filament does not oxidise), the tube housing with lead shielding to absorb leakage radiation, the oil that provides insulation and heat transfer, the window and the collimator.

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Stationary anode tube Rotating anode tube Construction A fixed tungsten button embedded in a copper block A tungsten-faced disc of 7-15 cm diameter on a molybdenum stem, driven by an induction motor whose rotor is inside the vacuum and stator outside Stationary anode tube Rotating anode tube Heat handling All the heat is delivered to one small spot; copper conducts it away The focal track sweeps a large circular area, spreading the heat over a track hundreds of times larger Speed — 3,000-3,600 rpm (standard) or 9,000-10,000 rpm (high speed) Tube rating Low — limited mA and short exposures Much higher — permits the high mA and short exposure times needed to freeze motion Focal spot Relatively large Can be small while still handling high loads, giving better resolution Cost and complexity Simple, cheap, robust Complex, expensive, bearing wear is the usual failure Typical use Dental units, portable and low-output machines All general radiography, fluoroscopy, angiography and CT • Heat units: for single-phase equipment HU = kVp × mA × s, multiplied by 1.35 for three-phase six-pulse and 1.41 for three-phase twelve-pulse or high-frequency generators.

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Anode cooling charts and tube rating charts must not be exceeded. • Tube failure arises from filament evaporation and burnout, cracking or pitting of the anode surface, tungsten deposition on the glass envelope causing arcing, bearing failure, and loss of vacuum.

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X-ray Control and Indicating Equipment • The generator supplies the tube: a high-tension (step-up) transformer producing the kilovoltage, a filament (step-down) transformer supplying a few volts at high current, rectifiers to make the supply unidirectional (the tube itself is a self-rectifier but this is inefficient), and the timer.

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Generator types in ascending order of quality are single-phase half-wave and full-wave, three-phase six- and twelve-pulse, and high-frequency (constant potential) — the last being standard today because it gives an almost ripple-free output, higher mean photon energy, shorter exposure times and lower patient dose. • The three operator controls and what each does: kVp sets the accelerating potential and therefore the energy, penetration and beam quality; raising it increases both the quantity (roughly as kVp²) and the penetration, and reduces subject contrast. mA sets the tube current and therefore the number of photons.

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Time (s) sets the duration; mA × s = mAs governs the total quantity of radiation, the film density and, directly and proportionally, the patient dose.

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The classic exam rule is the 15 % rule — a 15 % increase in kVp approximately doubles the exposure, equivalent to doubling the mAs. • Automatic exposure control (AEC, phototimer) uses ionisation chambers or photomultiplier detectors behind the patient to terminate the exposure when the required radiation has reached the receptor, giving consistent density regardless of patient size.

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A backup timer limits the maximum exposure. • Indicating and safety devices: kV and mA meters, exposure and ready indicators, the line voltage compensator, space charge compensation, the collimator with its light beam diaphragm (which must coincide with the radiation field within 2 % of the SID), filtration indicators, exposure switches that require the operator to be behind the barrier, interlocks, and the dose-area product meter. • Fluoroscopy adds an image intensifier (input phosphor of caesium iodide → photocathode → electrostatic focusing and acceleration to about 25-30 kV → output phosphor), giving brightness gain from flux gain and minification gain, with automatic brightness control and a 5-minute cumulative timer; flat-panel detectors have now largely replaced the intensifier.

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Filters and Grids • Filtration removes the low-energy (soft) photons that would be absorbed in the patient's skin and contribute dose but not image.

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Inherent filtration (the tube window, oil and housing, about 0.5-1.0 mm Al equivalent) plus added filtration (aluminium sheets) must total at least about 2.5 mm Al equivalent above 70 kVp.

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The effect is beam hardening — the mean energy of the beam rises while the total intensity falls.

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A compound (K-edge) filter such as erbium or the molybdenum filter of mammography selectively removes a band of energies; a wedge or compensating filter evens out exposure across a body part of varying thickness. • Attenuation obeys I = I₀ e−μx, where μ is the linear attenuation coefficient.

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The half-value layer (HVL) — the thickness of a stated material reducing the intensity by half, HVL = 0.693/μ — is the standard practical measure of beam quality.

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The interactions that matter diagnostically are the photoelectric effect (dominant at low energies and in high-Z materials, varying roughly as Z³/E³, and the source of most subject contrast and of patient dose) and Compton scattering (dominant above about 30 keV in soft tissue, nearly independent of Z, and the source of the scattered radiation that degrades contrast and endangers staff). • Grids are placed between patient and image receptor to absorb scattered radiation, which otherwise adds a general fog and destroys contrast.

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A grid is an array of thin lead strips separated by a radiolucent interspace (aluminium, fibre or carbon). • Grid parameters: grid ratio = height of the lead strips ÷ width of the interspace (commonly 5:1 to 16:1 — a higher ratio removes more scatter but demands more accurate alignment and more dose); grid frequency (lines per cm); focused, parallel or crossed construction; and stationary or moving (the Potter-Bucky diaphragm, which blurs the grid lines out of the image).

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Selectivity and contrast improvement factor describe performance, and the Bucky factor is the increase in exposure needed when the grid is used. • Grid cut-off — loss of primary radiation at the edges or across the image — occurs with off-level, off-centre, off-focus (wrong SID) or upside-down focused grid errors.

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A grid is indicated when the body part exceeds about 10 cm or above about 70 kVp; alternatives are the air gap technique and tight collimation, which is always the first and best means of reducing scatter.

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Biological Effects of X-rays • X-rays are ionising and act by direct damage to DNA and, more commonly, indirectly through radiolysis of water producing free radicals.

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The critical target is DNA, and double-strand breaks are the lesions most likely to be lethal or mutagenic.

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The law of Bergonié and Tribondeau states that cells are most radiosensitive when they are rapidly dividing, undifferentiated and have a long mitotic future — hence the high sensitivity of bone marrow, gonads, lymphocytes, intestinal crypt cells and the developing fetus, and the relative resistance of nerve and muscle. • Deterministic (tissue reaction) effects have a threshold dose, and their severity increases with dose: skin erythema, epilation, cataract, sterility, and the acute radiation syndromes.

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They are the concern of fluoroscopy and interventional radiology. • Stochastic effects have no threshold; the probability (not the severity) increases with dose.

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These are cancer induction and heritable genetic effects, and they are the reason for the linear-no-threshold model used in protection. • Units: exposure in coulomb per kilogram (old unit roentgen); absorbed dose in gray (Gy) = 1 J/kg (old unit rad, 1 Gy = 100 rad); equivalent dose in sievert (Sv) = absorbed dose × radiation weighting factor (1 for X-rays, photons and beta, 20 for alpha); and effective dose in sievert, weighting each organ by its tissue weighting factor to give whole-body risk. • Radiation protection rests on three ICRP principles — justification, optimisation (ALARA — as low as reasonably achievable) and dose limitation — and in practice on time, distance and shielding, distance being the most powerful because intensity obeys the inverse square law, I ∝ 1/d².

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Practical measures: collimation to the area of interest, correct kVp and mAs, filtration, gonad and thyroid shielding, lead aprons (0.25-0.5 mm Pb) and lead glass, protective barriers and warning signs, personal dosimeters (film badge, TLD, OSL), the 10-day rule and pregnancy enquiry, and never holding the patient or cassette by hand.

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This section introduces computed tomography, its basic principles and mathematics, the generations of CT scanners, the system components, reconstruction methods and image quality.

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Introduction and Basic Principles • Computed tomography produces cross-sectional (tomographic) images by measuring the transmission of a rotating X-ray beam through the patient from many directions and reconstructing the distribution of attenuation coefficients mathematically.

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It was invented by Godfrey Hounsfield with theoretical work by Allan Cormack (Nobel Prize 1979), the first clinical scanner being installed in 1971. • Why it was needed: a plain radiograph superimposes all the structures along the beam path into one image and has poor low-contrast (soft-tissue) resolution.

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CT removes superimposition and, by measuring transmission precisely, distinguishes tissues differing in attenuation by less than 1 %. • Principle: each measurement (a ray sum) gives ln(I₀/I) = Σ μᵢ·Δx, that is the line integral of the attenuation coefficient along that path.

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A complete set of such measurements at one angle is a projection (view); hundreds of projections around 360° provide enough equations to solve for the attenuation coefficient of every element (voxel) in the slice. • The Hounsfield unit expresses the result on a standard scale:

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HU = 1000 × (μtissue − μwater) / μwater.

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Hence water = 0 by definition, air = −1000, fat ≈ −100 to −50, water-equivalent soft tissue ≈ 20-60, clotted blood ≈ 60-80, and dense cortical bone ≈ +1000 or more. • Windowing: the reconstructed data span about 2,000-4,000 HU, but a display can show only about 256 grey levels and the eye far fewer, so the operator selects a window width (the range of HU displayed, controlling contrast) and a window level or centre (the midpoint, controlling brightness) — a narrow window for brain, a wide one for lung and bone.

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Nothing is lost: the raw data remain and can be re-windowed.

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Generations of CT Scanners Generation Geometry and motion Remarks First Pencil beam, single detector; translate-rotate The original Hounsfield design; about 4-5 minutes per slice; head only Second Narrow fan beam with a small detector array (about 3-30); translate-rotate Several projections per translation, reducing scan time to about 20 s per slice Third Wide fan beam covering the whole patient with a large curved detector array; both tube and detectors rotate together No translation; about 1 s or less per rotation; the dominant modern geometry; liable to ring artefacts if a detector mis-calibrates Fourth A complete stationary ring of detectors (4,000+); only the tube rotates Immune to ring artefacts and self-calibrating, but expensive, with more scattered radiation; largely abandoned Fifth (electron beam, EBCT) No moving parts — an electron beam is swept electromagnetically over stationary tungsten target rings Scan times of about 50 ms; developed for cardiac imaging Spiral / helical Continuous tube rotation with simultaneous continuous table movement, made possible by the slip ring Acquires a volume rather than slices; allows a single breath-hold, overlapping reconstruction and true 3-D and multiplanar reformatting Multi-detector (MDCT) Multiple detector rows (4, 16, 64, 128, 320) acquiring many slices per rotation Faster, thinner slices, isotropic voxels; enables CT angiography and cardiac CT with ECG gating; dual-source and dual-energy systems add material discrimination • Pitch, the key parameter of helical scanning, is table movement per rotation ÷ nominal beam collimation.

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Pitch = 1 means contiguous coverage; pitch > 1 (extended) covers more quickly with less dose but with some loss of resolution; pitch < 1 (overlapping) gives better quality at higher dose.

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System Components Component Function Gantry Houses the tube, detectors, collimators, filters and generator; can be tilted; the aperture is typically 70 cm X-ray tube High-output rotating anode with a very large heat capacity (several MHU), operating at 80-140 kVp continuously during the scan Collimators and filters Pre-patient collimation sets the slice thickness and limits dose; post-patient collimation rejects scatter; a bow-tie filter equalises the beam across the elliptical patient Detectors Solid-state scintillators (cadmium tungstate, gadolinium oxysulphide, rare-earth ceramics) coupled to photodiodes — now universal, with high efficiency (> 90 %) and fast response; earlier systems used xenon gas ionisation chambers Data acquisition system (DAS) Amplifies, integrates and digitises the detector signals; logarithmic conversion produces the ray sums Slip ring Transfers power and data to the continuously rotating gantry without cables — the enabling technology for helical scanning Patient table (couch) Low-attenuation carbon-fibre top with accurate, reproducible indexing Computer and reconstruction hardware Performs the reconstruction (formerly array processors, now GPUs) and post-processing Operator console and viewing workstation Protocol selection, windowing, MPR, MIP and 3-D rendering; archiving to PACS (4.5) Reconstruction • Simple back-projection smears each projection back across the image plane; summing all projections concentrates value where the object lay, but produces a characteristic 1/r blurring (star artefact). • Filtered back-projection (FBP) corrects this by applying a high-pass convolution filter (kernel) to each projection before back-projecting, and was for decades the standard method: fast, predictable and linear.

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The choice of kernel trades spatial resolution against noise — a sharp/bone kernel gives high detail and high noise, a smooth/soft-tissue kernel the reverse. • Iterative reconstruction begins with an estimate, forward-projects it, compares with the measured data and corrects, repeating until convergence (ART, SIRT, and the modern statistical and model-based algorithms).

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It models noise and system geometry and therefore achieves the same image quality at substantially lower dose, at a much higher computational cost — it is now standard on modern scanners.

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Deep-learning reconstruction is the current development. • Fourier-based methods rest on the central slice (projection-slice) theorem: the one-dimensional Fourier transform of a projection equals a radial line through the two-dimensional Fourier transform of the object.

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Image Quality, Artefacts and Dose • Spatial resolution (about 0.5-0.7 mm) is set by the focal spot, detector size, sampling, slice thickness, reconstruction kernel and matrix (usually 512 × 512).

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Contrast (low-contrast) resolution is CT's great strength and is limited by noise, which falls with the square root of the dose: noise ∝ 1/√(mAs), so halving noise requires four times the dose.

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Temporal resolution matters in cardiac work. • Artefacts: beam hardening (cupping and dark streaks between dense structures, since the beam is polychromatic), metal artefacts (severe streaking from prostheses and dental work), motion artefacts, partial volume averaging (a voxel containing two tissues is assigned their average HU), ring artefacts (a mis-calibrated detector in a third-generation scanner), streak and aliasing from undersampling, and photon starvation through the shoulders and hips. • Dose:

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CT contributes a disproportionate share of medical radiation exposure.

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It is expressed as CTDIvol (mGy) and dose-length product DLP (mGy·cm), from which effective dose is estimated by a conversion factor.

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Reduction strategies are automatic tube current modulation, iterative reconstruction, lower kVp in children and slim adults, restriction of the scan range, avoidance of unnecessary multiphase scans, and above all justification of the examination. • Contrast media: iodinated agents (ionic or, preferably, non-ionic low-osmolar) given intravenously, orally or intra-arterially raise attenuation in vessels and perfused tissue.

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Risks are allergic-type reactions, contrast-induced nephropathy and extravasation, so renal function and allergy history must be checked.

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This section covers the fundamental physics of nuclear magnetic resonance, the principles and parameters of MRI, image formation by gradients and pulse sequences, contrast-enhanced MRI, clinical applications and safety.

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Fundamental Concepts • MRI images the body using a strong magnetic field and radiofrequency waves, with no ionising radiation.

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It relies on nuclei with an odd number of protons or neutrons, which possess spin and therefore a magnetic moment.

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In practice the hydrogen nucleus (a single proton) is used, because hydrogen is by far the most abundant element in the body (in water and fat) and has the largest gyromagnetic ratio of the biological nuclei. • In the main field B₀ the proton magnetic moments align either parallel (low energy) or antiparallel (high energy) to the field, with a very slight excess in the parallel state that produces a measurable net magnetisation vector (M₀) along the z axis.

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They also precess about the field direction. • The Larmor equation is the central formula of MRI: ω₀ = γB₀, or f₀ = (γ/2π)·B₀, where γ/2π for hydrogen is 42.58 MHz per tesla.

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Thus at 1.5 T the resonance frequency is about 63.9 MHz and at 3 T about 127.7 MHz — in the radiofrequency band. • Excitation: an RF pulse at exactly the Larmor frequency is applied, which is the resonance condition.

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It tips the net magnetisation away from the z axis (a 90° pulse into the transverse plane, a 180° pulse inverting it) and brings the spins into phase coherence.

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The rotating transverse magnetisation induces a voltage in a receiver coil — the MR signal. • Relaxation returns the system to equilibrium by two independent processes:

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T1 (longitudinal, spin-lattice) relaxation — the recovery of magnetisation along z as energy is given to the surrounding lattice, an exponential process whose time constant T1 is the time to recover 63 % of longitudinal magnetisation; and T2 (transverse, spin-spin) relaxation — the loss of phase coherence in the transverse plane through interaction between neighbouring spins, with T2 the time for the transverse signal to fall to 37 % of its initial value.

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T2 is always shorter than or equal to T1, and T2* is the still faster decay seen in practice, because static field inhomogeneity adds to true T2 dephasing; the raw signal after a single pulse is therefore the free induction decay.

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Tissue T1 at 1.5 T T2 Appearance Fat Short (~250 ms) Short-intermediate Bright on T1 Water / CSF Long (~4,000 ms) Long (~2,000 ms) Dark on T1, very bright on T2 White matter ~780 ms ~90 ms Brighter than grey on T1 Grey matter ~920 ms ~100 ms Brighter than white on T2 Cortical bone, air, calcification — Very short Signal void (black) on all sequences Principles and Parameters • TR (repetition time) is the interval between successive excitation pulses and controls the degree of T1 weighting;

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TE (echo time) is the interval from the excitation pulse to the sampling of the echo and controls T2 weighting. • The three image weightings, which must be known cold:

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T1-weighted = short TR (< 700 ms) and short TE (< 30 ms) — good anatomical detail, fat bright and fluid dark;

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T2-weighted = long TR (> 2,000 ms) and long TE (> 80 ms) — fluid and most pathology bright, the sequence for detecting disease; proton density-weighted = long TR and short TE — signal proportional to the number of protons, with little T1 or T2 influence. • Other parameters: flip angle, number of excitations/averages (NEX), field of view, matrix size, slice thickness and gap, and bandwidth.

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The governing compromise is that signal-to-noise ratio improves with voxel volume, NEX and field strength but worsens as resolution and speed increase; scan time for a basic spin echo is TR × number of phase-encoding steps × NEX. • Pulse sequences: spin echo (a 90° pulse followed by a 180° refocusing pulse that cancels static inhomogeneity and so measures true T2 — the reference sequence); fast/turbo spin echo (an echo train, much faster); gradient echo (a reversing gradient instead of a 180° pulse — fast, low flip angle, but sensitive to inhomogeneity, so it measures T2* and is used for susceptibility and blood-product imaging); inversion recovery variants STIR (short TI, suppresses fat) and FLAIR (long TI, suppresses CSF and so reveals periventricular lesions); echo planar imaging (a whole image after one excitation — the basis of diffusion and functional MRI); and diffusion-weighted imaging, whose restricted diffusion signal detects acute stroke within minutes.

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Image Formation • Since every proton in the body would resonate at the same frequency in a uniform field, spatial information is encoded by deliberately making the field non-uniform with gradient coils, in three orthogonal directions (Gx, Gy, Gz). • Slice selection: a gradient along the slice axis makes the Larmor frequency vary with position, so an RF pulse of a defined narrow bandwidth excites only one slice; the slice thickness is set by the RF bandwidth and the gradient steepness. • Phase encoding: a gradient applied briefly after excitation leaves the spins in each row with a different phase; the sequence is repeated with a different gradient amplitude for each phase-encoding step, which is why scan time is proportional to their number. • Frequency encoding (readout): a gradient applied during signal sampling makes the frequency vary along the remaining axis. • k-space and reconstruction: the signals fill a matrix of spatial-frequency data called k-space, and the image is obtained by a two-dimensional inverse Fourier transform.

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The centre of k-space carries contrast and signal, the periphery carries fine detail and edges — a favourite examination point. • Hardware: the main magnet (superconducting niobium-titanium coils in liquid helium for 1.5 T and 3 T systems — the standard; permanent and resistive magnets for low field), shim coils for homogeneity, gradient coils (whose rapid switching against the main field produces the loud knocking noise), RF transmit and receive coils (body, head, surface and phased-array coils), the RF and gradient amplifiers, the Faraday cage shielding the room, and the computer.

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Contrast-Enhanced MRI • MRI contrast agents do not themselves produce signal; they alter the relaxation times of nearby water protons. • Gadolinium chelates are the standard agents.

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Free gadolinium is toxic, so it is chelated (DTPA, DOTA and similar).

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Gadolinium is paramagnetic, with seven unpaired electrons, and predominantly shortens T1, so enhancing tissue appears bright on T1-weighted images.

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Enhancement indicates increased vascularity, increased permeability or breakdown of the blood-brain barrier — hence its value in tumours, infection, inflammation, demyelination and MR angiography. • Safety: gadolinium agents are generally well tolerated, but nephrogenic systemic fibrosis may occur in patients with severe renal impairment, particularly with the older linear agents, so renal function must be checked and macrocyclic agents preferred; gadolinium deposition in brain and bone has been described and is under continuing review. • Other agents: superparamagnetic iron oxide (SPIO/USPIO), which predominantly shortens T2/T2* and causes signal loss, used for liver and lymph nodes; manganese; and hepatobiliary agents.

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Non-contrast techniques also produce angiographic and functional images — time-of-flight and phase-contrast MRA, arterial spin labelling for perfusion, and the BOLD (blood oxygen level dependent) effect of functional MRI, which exploits the paramagnetism of deoxyhaemoglobin.

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Clinical Applications and Safety • Strengths: outstanding soft-tissue contrast, no ionising radiation, direct imaging in any plane, and a wide range of functional and physiological contrasts.

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Weaknesses: long scan times and motion sensitivity, high cost, noise and claustrophobia, poor depiction of cortical bone and lung, and the extensive contraindications below. • Applications: neuroimaging (tumours, stroke with DWI, multiple sclerosis with FLAIR, epilepsy, dementia, spine and cord); musculoskeletal (menisci, ligaments, cartilage, marrow, soft-tissue tumours); cardiac (function, viability with late gadolinium enhancement, congenital disease); abdominal and pelvic (liver lesions, MRCP for the biliary tree, prostate, gynaecological and obstetric imaging); breast MRI;

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MR angiography; and the advanced techniques fMRI, spectroscopy, diffusion tensor tractography and perfusion imaging. • Safety — the dominant hazard is the static field, which is always on.

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Ferromagnetic objects become projectiles (the missile effect), and rigorous screening and zoning of the MR environment are mandatory.

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Contraindications and cautions: cardiac pacemakers and ICDs (unless MR-conditional), cochlear implants, some aneurysm clips, neurostimulators, metallic intraocular foreign bodies, and the first trimester of pregnancy as a relative caution.

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RF heating, limited by the specific absorption rate (SAR, W/kg); acoustic noise (over 100 dB, requiring ear protection); peripheral nerve stimulation from rapidly switched gradients; and quench — the sudden boiling-off of liquid helium, which requires an effective quench pipe and oxygen monitoring to prevent asphyxiation.

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This section covers the physics of ultrasound, the construction and properties of transducers, the ultrasonic beam, the modes of ultrasound imaging, Doppler ultrasound, clinical applications and biological effects.

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Physics of Ultrasound • Ultrasound is a longitudinal mechanical (pressure) wave of frequency above 20 kHz, the limit of human hearing; diagnostic imaging uses 2-15 MHz.

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Being mechanical, it requires a medium and cannot travel through a vacuum — the reason coupling gel is essential and why air is such an effective barrier. • Basic relations: c = f λ.

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The speed of sound depends on the medium, not on the frequency, and machines assume an average soft-tissue value of 1,540 m/s (air ~330, fat ~1,450, blood ~1,570, muscle ~1,580, bone ~3,500-4,080 m/s).

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Hence at 1 MHz the wavelength in soft tissue is 1.54 mm, and at 5 MHz about 0.3 mm. • The pulse-echo principle: the transducer emits a short pulse and listens for returning echoes; from the go-and-return time t the machine computes depth d = c·t / 2.

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In soft tissue, 13 μs of round-trip time corresponds to 1 cm of depth. • Acoustic impedance Z = ρ·c (density × speed), measured in rayls, is the property that governs reflection.

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At a boundary the intensity reflection coefficient R = [(Z₂ − Z₁)/(Z₂ + Z₁)]².

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A large impedance mismatch reflects almost all the energy: at a soft tissue-air interface about 99.9 % is reflected, and at soft tissue-bone about 50 % — which is why gel must exclude air, why bone and gas cast acoustic shadows, and why lung and bowel are so difficult to image. • Interactions: specular reflection at large smooth interfaces (angle-dependent), scattering from structures smaller than the wavelength (which produces the speckle texture of parenchyma and is angle-independent), refraction at oblique interfaces of differing speed (Snell's law, causing edge artefacts), diffraction, and absorption of energy as heat. • Attenuation is the total loss of intensity with depth, expressed in decibels, and in soft tissue is approximately 0.5 dB per cm per MHz — so it increases with frequency, which is the origin of the single most important trade-off in ultrasound.

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Attenuation is compensated by time-gain compensation (TGC, depth-gain compensation), which amplifies later-arriving echoes more.

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Transducers • The transducer both transmits and receives, exploiting the piezoelectric effect — the property of certain crystals of developing a voltage when deformed (reception) and of deforming when a voltage is applied (transmission, the reverse piezoelectric effect).

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The material is lead zirconate titanate (PZT), a synthetic ceramic, or newer single-crystal and composite materials; quartz is the natural example.

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Heating above the Curie temperature destroys the polarisation permanently, which is why probes must never be autoclaved. • Construction: the PZT element, whose thickness determines the resonant frequency (thickness = half a wavelength); the backing (damping) block, usually tungsten in epoxy, which shortens the pulse by absorbing backward energy — improving axial resolution and broadening the bandwidth at the cost of sensitivity; the matching layer of intermediate impedance (ideally √(Z₁Z₂), a quarter-wavelength thick) which improves transmission of energy into tissue; electrodes; acoustic insulation; and the housing. • Transducer types: linear array (rectangular image, high frequency, for superficial structures, vessels and small parts), curvilinear (convex) (wide sector at depth, for abdomen and obstetrics), phased array (small footprint with electronic steering, for cardiac imaging between the ribs), endocavitary (transvaginal, transrectal — high frequency close to the target), transoesophageal, intravascular, and mechanical or matrix arrays for 3-D/4-D imaging. • Resolution: axial (along the beam) resolution = half the spatial pulse length, so it improves with higher frequency and shorter pulses (better damping) and is always better than lateral resolution; lateral resolution equals the beam width and is best at the focus; elevational (slice-thickness) resolution depends on the lens; and temporal resolution is the frame rate, which falls as depth, sector width and line density increase.

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The Ultrasonic Beam • A beam from a disc of diameter D has two regions.

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The near field (Fresnel zone) is roughly cylindrical and converges to a natural focus at length L = D²/(4λ) = D²f/(4c); the far field (Fraunhofer zone) beyond it diverges at an angle given by sin θ = 1.22 λ/D.

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Hence a larger aperture and a higher frequency give a longer near field and less divergence — better lateral resolution at depth. • Focusing may be fixed (an acoustic lens or a curved element) or, in arrays, electronic — by firing the elements with small time delays, which also allows multiple transmit focal zones (at the cost of frame rate) and dynamic receive focusing throughout the depth.

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Beam steering in a phased array is achieved by the same delay principle. • Artefacts arise directly from the assumptions the machine makes (constant speed of 1,540 m/s, straight-line propagation, echoes from the main beam only, attenuation uniform): acoustic shadowing (behind stone, bone or gas), posterior acoustic enhancement (behind a fluid-filled structure such as a cyst — the classic sign of a cyst), reverberation and its special case comet tail/ring-down, mirror image, side lobe and grating lobe artefacts, refraction (edge shadowing), speed displacement, and anisotropy in tendons.

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Modes of Ultrasound Imaging Mode Description and use A-mode (amplitude) A one-dimensional plot of echo amplitude against depth — the original mode; now used mainly in ophthalmic biometry (axial length for IOL power) and echoencephalography B-mode (brightness) Echo amplitude displayed as brightness of a dot; sweeping the beam builds the familiar two-dimensional grey-scale image — the basis of all modern imaging M-mode (motion, TM-mode) A single line of B-mode plotted against time, giving very high temporal resolution — used in echocardiography for valve motion, chamber dimensions and fetal heart activity Real-time B-mode Rapid repeated B-mode scanning producing a moving image; the standard clinical mode 3-D and 4-D Volume acquisition and rendering;

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4-D is real-time 3-D, widely used in obstetrics Doppler modes Continuous wave, pulsed wave, colour flow and power Doppler — see below Harmonic imaging Receives at twice the transmitted frequency, exploiting non-linear propagation; reduces artefact and improves resolution in difficult patients Elastography Measures tissue stiffness, used for liver fibrosis and for characterising breast and thyroid lesions Contrast-enhanced ultrasound Gas-filled microbubbles that resonate in the beam, used for perfusion and lesion characterisation Doppler Ultrasound • The Doppler effect is the change in observed frequency when source and observer move relative to one another.

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In ultrasound, sound reflected from moving red blood cells returns at a shifted frequency, and the Doppler shift is given by Δf = 2 f₀ v cos θ / c, where f₀ is the transmitted frequency, v the velocity of blood, θ the angle between the beam and the direction of flow, and c the speed of sound.

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The factor 2 arises because the cells act both as moving receiver and moving source. • The angle is everything: because of the cos θ term, the signal is maximal at 0° (flow along the beam) and zero at 90° (flow perpendicular to the beam) — the commonest practical error in vascular scanning.

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For quantitative velocity measurement the angle of insonation should be kept below 60°, since beyond that small angle errors produce large velocity errors.

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Conveniently, the shift for diagnostic frequencies and physiological velocities falls in the audible range, so the signal can be heard. • Modes: continuous wave (CW) — separate transmitting and receiving crystals, measuring any velocity without aliasing but with no depth resolution (range ambiguity); pulsed wave (PW) — one crystal with range gating, giving depth-specific measurement but subject to aliasing; colour flow Doppler — mean velocities colour-coded and superimposed on the B-mode image, conventionally red towards and blue away from the transducer (BART: blue away, red towards); and power (energy) Doppler — displaying the amplitude of the Doppler signal rather than its frequency, which is more sensitive to slow flow and almost angle-independent, but gives no direction or velocity information. • Aliasing occurs in pulsed Doppler when the Doppler shift exceeds half the pulse repetition frequency — the Nyquist limit; the display wraps around and high velocities appear reversed.

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Remedies are to raise the PRF/scale, shift the baseline, use a lower transmit frequency, increase the angle, select a shallower sample volume, or switch to continuous wave. • Indices and uses: peak systolic and end-diastolic velocity, the resistive index (PSV − EDV)/PSV and the pulsatility index.

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Applications include carotid and peripheral arterial stenosis, deep vein thrombosis, transcranial Doppler, echocardiography with valve gradients from the simplified Bernoulli equation ΔP = 4v², renal and hepatic vessels, testicular and ovarian torsion, and obstetric umbilical and middle cerebral artery studies.

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Clinical Applications and Biological Effects • Applications: obstetrics (dating, growth, anomaly scanning, liquor volume, placental localisation, biophysical profile and fetal Doppler); abdomen (liver, gallbladder and biliary tree, pancreas, kidneys, spleen, aorta, free fluid and the FAST scan in trauma); pelvis; small parts (thyroid, breast, testis, salivary glands); musculoskeletal; vascular; echocardiography; paediatric neurosonography through the fontanelle; and interventional guidance for vascular access, biopsy, drainage and regional anaesthesia. • Advantages: no ionising radiation, real-time and dynamic, portable and inexpensive, no known harmful effects at diagnostic levels, and excellent for fluid, vessels and guidance.

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Limitations: highly operator-dependent, blocked by bone and gas, limited penetration in obese patients, a restricted field of view, and poor reproducibility between examiners. • Biological effects are of two kinds.

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Thermal — absorbed acoustic energy is converted to heat, which matters most at bone-soft tissue interfaces and in Doppler modes, where intensities are far higher than in B-mode; this is indicated by the thermal index (TI), subdivided into TIS (soft tissue), TIB (bone) and TIC (cranial bone).

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Mechanical (non-thermal) — cavitation, the formation, oscillation and possible violent collapse of gas bubbles, which is stable at low intensities and inertial (transient) at high, together with radiation force and microstreaming; this is indicated by the mechanical index (MI) = peak negative pressure ÷ √frequency. • Safety practice: diagnostic ultrasound has an excellent safety record and no confirmed harmful effect at diagnostic exposure levels, but it is not assumed to be without risk.

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The governing principle is ALARA: use the lowest output and shortest exposure time consistent with a diagnostic result, be especially careful with Doppler in the first trimester and over the fetal eye and brain, monitor the displayed TI and MI, and avoid non-medical use such as keepsake fetal imaging.

1

This section introduces digital imaging, computed radiography and direct digital radiography, and the picture archiving and communication system with the DICOM standard.

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Introduction to Digital Imaging • A digital image is a numerical array: the field is divided into pixels (picture elements; the corresponding tissue volume is a voxel), each holding a number representing brightness.

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The matrix size (512 × 512, 1024 × 1024, 2048 × 2048 or more) and the field of view together set the pixel size = FOV ÷ matrix, which governs spatial resolution; the bit depth (8, 10, 12, 16 bits) sets the number of grey levels available, 2n — 12 bits giving 4,096 levels. • Conversion from the analogue detector signal involves sampling (in space and time) and quantisation (into discrete levels), performed by an analogue-to-digital converter; the Nyquist criterion requires sampling at at least twice the highest spatial frequency present, otherwise aliasing occurs. • Advantages of digital over film: post-processing (windowing, edge enhancement, magnification, measurement, inversion); a wide dynamic range so that exposure errors can largely be corrected without a repeat; immediate availability; no chemical processing, darkroom or film store; simultaneous multi-site viewing; archiving and transmission (teleradiology); integration with reporting and the electronic record; and generally lower patient dose.

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Disadvantages: high capital cost, dependence on networks and IT, the risk of dose creep (since over-exposure no longer darkens the image and may go unnoticed), and the need for quality control and data security. • The exposure indicator is therefore an essential feature of digital radiography: since the displayed image looks correct over a wide range of exposures, the system reports a numerical index (S value, EI, or the standardised EI and deviation index DI) that tells the radiographer whether the exposure was appropriate.

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Computed Radiography (CR) • CR is cassette-based indirect digital radiography and was the first widely adopted digital system, because it could be used with existing X-ray equipment simply by replacing the film cassette. • Mechanism: the cassette contains a photostimulable phosphor imaging plate, usually barium fluorohalide doped with europium (BaFBr:Eu).

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X-ray exposure raises electrons into metastable traps (F-centres), forming a latent image that persists for hours.

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In the reader, a red laser scans the plate, releasing the trapped electrons; they return to the ground state emitting blue light — photostimulated luminescence, proportional to the original exposure.

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A photomultiplier tube detects this light, an ADC digitises it, and the plate is then flooded with bright white light to erase it for reuse (several thousand cycles). • Characteristics: advantages — retrofits to existing equipment, portable and usable for bedside and theatre work, cassettes of standard sizes, lower cost of entry.

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Disadvantages — a separate reading step taking around 30-60 s, cassette handling, somewhat lower spatial resolution and detective quantum efficiency than DR, and plate wear and dust artefacts.

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Direct Digital Radiography (DR) • DR uses a flat-panel detector permanently in the table or wall stand (or as a portable panel) that converts X-rays into a digital signal and delivers the image to the console within a few seconds, with no separate reading step. • Indirect conversion DR: a scintillator — caesium iodide (structured, columnar, which limits light spread) or gadolinium oxysulphide — converts X-rays to light, which is then detected by a matrix of amorphous silicon photodiodes with thin-film transistors.

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Two conversion steps, but high absorption efficiency. • Direct conversion DR: a photoconductor, amorphous selenium, converts X-rays directly into electrical charge, which is collected by a TFT array under an applied electric field.

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Because no light is produced there is no lateral light spread, giving the highest spatial resolution — hence its use in mammography. • CCD and CMOS systems with lens or fibre-optic coupling are used in dental and small-field devices. • CR versus DR:

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DR offers higher detective quantum efficiency and therefore lower dose, much faster workflow and higher throughput, and better image quality, at a higher capital cost and with less flexibility for awkward projections.

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DQE — the efficiency with which a detector converts incident X-ray quanta into image information — is the key figure of merit, alongside the modulation transfer function (MTF) for resolution. • Related digital modalities: digital fluoroscopy with flat panels and digital subtraction angiography (DSA), in which a pre-contrast mask image is subtracted from the contrast-filled image to display the vessels alone; digital mammography and tomosynthesis; and dual-energy subtraction.

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PACS and DICOM • PACS (Picture Archiving and Communication System) is the integrated system that acquires, stores, distributes and displays medical images electronically, replacing film entirely in a filmless department. • Components: image acquisition devices (modalities) — CR/DR, CT, MRI, ultrasound, nuclear medicine; a secure network of adequate bandwidth; the PACS server and archive, usually in tiers of short-term (online) storage, near-line and long-term (offline or cloud) archive, with redundancy and disaster recovery; display workstations with calibrated, high-resolution diagnostic monitors (calibrated to the DICOM grey-scale display function); and the software for worklists, viewing, post-processing and reporting. • Integration:

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PACS works with the RIS (Radiology Information System), which handles scheduling, patient demographics, worklists and reports, and with the HIS (Hospital Information System) and electronic health record.

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Non-image data are exchanged using the HL7 standard, and the IHE initiative defines profiles for how these systems should work together.

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A modality worklist pushes the correct patient data to the scanner, greatly reducing identification errors. • DICOM (Digital Imaging and Communications in Medicine) is the universal standard for the format, storage and exchange of medical images.

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A DICOM file contains both the pixel data and a header of metadata — patient identification, study, series and instance UIDs, modality, acquisition parameters, orientation and window settings — which is why a DICOM image is self-describing and why de-identification (anonymisation) is required before research or teaching use.

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DICOM also defines network services such as C-STORE, C-FIND, C-MOVE and modality worklist.

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Its existence is what makes vendor-independent interoperability possible. • Benefits: no lost films, simultaneous access from many locations, instant retrieval of prior studies for comparison, teleradiology and remote reporting (of great practical importance in Nepal's geography), reduced cost of film and storage, and audit and workflow data.

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Challenges: capital cost and the need for reliable power and network infrastructure, data security, privacy and access control, long-term storage growth, backup and disaster recovery, monitor quality control, migration between vendors, and training.

1

This section covers the types of radioactivity, radionuclides used in medicine, the gamma camera, single photon emission computed tomography and positron emission tomography.

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Types of Radioactivity • Radioactivity is the spontaneous transformation of an unstable nucleus with the emission of radiation.

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It is random and independent of temperature, pressure and chemical state, and obeys N = N₀ e−λt, where λ is the decay constant; the physical half-life T½ = 0.693/λ.

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Activity is measured in becquerel (Bq) = 1 disintegration per second (the old unit being the curie, 1 Ci = 3.7 × 10¹⁰ Bq). • Three half-lives must be distinguished: the physical half-life (decay), the biological half-life (excretion from the body) and the effective half-life, given by 1/Teff = 1/Tphys + 1/Tbiol, so the effective half-life is always the shortest of the three.

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Type Nature Properties and medical relevance Alpha (α) A helium nucleus — 2 protons + 2 neutrons, charge +2 Very high ionising power but very low penetration (stopped by paper or skin); no use in imaging, dangerous if internalised; used in targeted alpha therapy (radium-223) Beta minus (β⁻) An electron emitted when a neutron converts to a proton Moderate ionisation, penetrates a few millimetres of tissue; not used for imaging but ideal for therapy — iodine-131, yttrium-90, strontium-89 Beta plus (β⁺, positron) A positive electron emitted when a proton converts to a neutron Travels a short distance, then annihilates with an electron, producing two 511 keV photons emitted at almost exactly 180° to each other — the basis of PET Gamma (γ) Electromagnetic radiation from an excited nucleus No charge or mass, highly penetrating, weakly ionising — the radiation used for imaging Electron capture / internal conversion A nucleus captures an orbital electron / transfers energy to an orbital electron Produce characteristic X-rays and Auger or conversion electrons; thallium-201 decays by electron capture Isomeric transition A metastable nucleus decays to its ground state by emitting a gamma ray only The ideal imaging decay — technetium-99m is the example Radionuclides in Medicine • Properties of the ideal imaging radionuclide: pure gamma emission with no particulate radiation (which would add dose without contributing to the image); a photon energy of about 100-200 keV — high enough to escape the body, low enough to be efficiently absorbed by the detector and easily collimated; a half-life comparable to the duration of the study (a few hours); ready availability at reasonable cost; and chemistry allowing it to be bound to a variety of pharmaceuticals. • Technetium-99m meets all of these and accounts for the great majority of nuclear medicine studies: it emits a single 140 keV gamma photon by isomeric transition, has a physical half-life of 6 hours, and is obtained on site from a molybdenum-99/technetium-99m generator (the 'moly cow'), eluted with saline, Mo-99 itself having a 66-hour half-life.

6

Labelled to different pharmaceuticals it images almost every system — MDP for bone, MAG3 and DTPA for kidneys, sestamibi for myocardium and parathyroid, MAA for lung perfusion, HIDA for the biliary tract, sulphur colloid for liver and sentinel node, pertechnetate for thyroid. • Other single-photon nuclides: iodine-131 (beta and gamma, 8 days — therapy for thyroid carcinoma and thyrotoxicosis, with imaging), iodine-123 (159 keV, 13 h — imaging only), thallium-201 (myocardial perfusion), gallium-67 (infection and tumour), indium-111 (labelled leukocytes), xenon-133 (ventilation). • Positron emitters are cyclotron-produced and short-lived: fluorine-18 (110 min), carbon-11 (20 min), nitrogen-13 (10 min), oxygen-15 (2 min) and the generator-produced rubidium-82 and gallium-68.

7

F-18 FDG (fluorodeoxyglucose) is the dominant PET agent: a glucose analogue taken up by cells in proportion to glycolytic activity and then trapped after phosphorylation, which is why it accumulates in tumours, infection, inflammation, brain and myocardium. • The fundamental difference from radiology:

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X-ray, CT, MRI and ultrasound are transmission or reflection techniques showing anatomy, whereas nuclear medicine is an emission technique in which the source is inside the patient and the image shows physiology and function — often revealing disease before any anatomical change, at the cost of relatively poor spatial resolution.

9

The Gamma Camera • The gamma (Anger) camera is the basic instrument of nuclear medicine, producing a two-dimensional planar image of the distribution of activity. • Components in order: the collimator — a lead plate perforated by thousands of holes that accepts only photons travelling in the required direction and rejects all others; it therefore determines the spatial resolution and sensitivity of the camera, and is the principal reason for its low efficiency (fewer than 1 in 10,000 photons is used).

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Types are parallel-hole (the standard), pinhole (magnifies small organs such as the thyroid), converging and diverging, in low-, medium- and high-energy designs. • Next the scintillation crystal — a large single crystal of sodium iodide doped with thallium, NaI(Tl), typically 30-50 cm across and 9.5 mm thick, which converts each gamma photon into a flash of light; then a light guide and an array of photomultiplier tubes (37-91) that convert the light to an amplified electrical pulse; then the position logic circuit, which computes the X and Y coordinates of each event from the relative pulse heights, and the pulse height analyser, which accepts only pulses within an energy window (typically 140 keV ± 10 %) and so rejects scattered photons; and finally the computer and display. • Performance: intrinsic spatial resolution of about 3-4 mm but system resolution of roughly 8-12 mm once the collimator is included, energy resolution about 10 %, and uniformity, linearity and sensitivity all subject to daily and weekly quality control. • Imaging modes: static (a single image), dynamic (a rapid series showing function over time, as in renography), gated (synchronised to the ECG, as in a MUGA scan of ventricular function) and whole body (a moving couch).

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SPECT and PET SPECT PET Radiation detected Single gamma photons from the nuclide (e.g.

12

140 keV from Tc-99m) Pairs of 511 keV annihilation photons emitted at 180° after a positron annihilates Nuclides Tc-99m, I-123, Tl-201, In-111 — many generator-produced and widely available F-18, C-11, N-13, O-15, Ga-68 — mostly cyclotron-produced, short-lived, needing nearby production Detection principle One or more gamma camera heads rotate around the patient; requires a lead collimator A stationary ring of detectors (LSO, LYSO, BGO crystals) records coincidence events within a few nanoseconds — electronic collimation, so no lead collimator is needed Sensitivity Lower — the collimator discards most photons Much higher (1-2 orders of magnitude) Spatial resolution About 8-12 mm About 4-6 mm Quantification Semi-quantitative Truly quantitative — the standardised uptake value (SUV) Cost and availability Relatively low; widely available High; requires a cyclotron or a regional supply network Typical uses Myocardial perfusion, bone, renal, brain perfusion, parathyroid, sentinel node Oncological staging and response with FDG, epilepsy focus localisation, dementia, cardiac viability • Reconstruction in both is tomographic — filtered back-projection or, more usually now, iterative (OSEM) reconstruction, with corrections for attenuation, scatter, randoms and dead time. • Hybrid imaging is the decisive modern development:

13

SPECT/CT and PET/CT combine the functional information of the emission scan with the anatomical localisation and attenuation map of the CT in a single session, and PET/MRI adds superior soft-tissue contrast with no CT dose.

14

The CT provides both precise anatomical correlation and the attenuation correction that makes quantification possible. • Radiation protection in nuclear medicine differs from radiology because the patient becomes the source: doses are handled in shielded syringes and vials behind lead-glass screens, staff use distance and time, patients (especially after I-131 therapy) may need isolation and instructions about contact with children and pregnant women, and radioactive waste must be stored for decay.

15

The ALARA principle and national regulations govern practice.