Sankara Nethralaya's Atlas of Imaging in Ophthalmology Ambika Selvakumar, Veena Noronha, Padmaja Minakshi Sundaram
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1Physical Principles of Computed Tomography and Magnetic Resonance Imaging
SECTION OUTLINE_________
  • 1. Physical Principles of Computed Tomography and Magnetic Resonance Imaging2

Physical Principles of Computed Tomography and Magnetic Resonance ImagingChapter 1

 
BASICS OF COMPUTED TOMOGRAPHY (CT)
The first CT scanner was developed by Sir Godfrey Hounsfield in 1972; since then this modality has become an important tool in diagnostic radiology. From the first scanner to the present day multislice helical scanner, CT technology has revolutionized the world of imaging and enhanced patient management.
 
BASIC PHYSICS
Computed tomography uses X-rays to obtain cross-sectional, two-dimensional images of the body. The cross- sectional image is produced by 360° rotation of the X-ray tube around the patient. The transmitted radiation is measured by the detectors located inside the gantry like a ring around the patient. The final image is generated from these measurements. The gantry of the CT machine houses the X-ray tube and the detectors (Fig. 1.1).
 
TYPES OF SCANNING TECHNIQUES
 
Axial (Sequential) Scanning
In sequential scanning, single slice is obtained with single 360° rotation of the tube (Fig. 1.2A). The disadvantage is that the time taken for an individual study is long, hence prone to motion artifacts and quality of reformations is suboptimal.
 
Helical (Spiral) Scanning
With the advent of slip ring technology, the continuous rotation of the X-ray tube around the patient is made possible during continuous patient table movement. This led to the development of helical scanning (Fig. 1.2B). The transmitted radiation thus takes the form of a helix or spiral around the patient acquiring volume of data. This allows larger anatomical coverage of the body part to be imaged, thereby reducing the possibility of artifacts caused by patient movement especially pediatric and critically ill patients. Faster scanning also increases patient throughput.
Multislice or multidetector machines utilize the principles of the helical scanner but incorporate multiple rows of detector rings. They can therefore, acquire multiple slices per tube rotation, thereby increasing the anatomical coverage in a shorter time (Fig. 1.3).
 
CT TERMINOLOGIES
 
Pixel and Voxel
Every CT image is made of a square of matrix called the picture element (pixel) and an object volume called voxel (Fig. 1.4). The obtained CT image is subdivided into a matrix of up to 512 × 512 or 1024 × 1024 elements. The pixel width is determined by the field of view (FOV) and matrix size, i.e. FOV/matrix. The voxel volume = pixel area × slice thickness.
Higher the pixel, better is the image resolution.
 
HOUNSFIELD UNIT OR CT NUMBER
Each voxel is traversed during the scan by numerous X-ray photons and the intensity of the transmitted radiation is measured by the detectors. From these intensity readings, the density or attenuation value, viz Hounsfield unit or CT number is calculated and assigned to every tissue.4
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Fig. 1.1: Cross-sectional view of the gantry showing the orientation of the X-ray tube and detectors in a fourth generation CT scanner
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Fig. 1.2A: Sequential scan—single cross-sectional slice of the patient in a single rotation
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Fig. 1.2B: Helical scan—rotation of the tube around the patient with continuous table movement
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Fig. 1.3: Multislice imaging—generation of six slices per rotation of the tube in a multidetector scanner
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Fig. 1.4: Pixel represents the matrix and voxel represents the slice thickness
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Fig. 1.5: Scale representing the range of Hounsfield numbers of the tissues seen in the body
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Each pixel is assigned a numerical value (CT number), based on the attenuation of X-rays by the tissue. This number is compared to the attenuation value of water and displayed on a scale of arbitrary units named Hounsfield units (HU) after Sir Godfrey Hounsfield.
This scale assigns water as an attenuation value (HU) of zero. Each number represents a shade of gray with +1000 (white) and −1000 (black) at either end of the spectrum (Fig. 1.5).
CT number of various tissues in the body is as follows:
  • Brain gray matter → 35 – 40 HU
  • Brain white matter → 30 – 35 HU
  • Blood → Flowing blood – 40 HU
Acute hematoma → 70 to 90 HU (density depends on the Hb concentration and coagulation profile)
  • Calcification → + 80 and above
  • Fat → − 10 to − 100
  • CSF → 0 to 10 HU
  • Bone → + 800 to 1000 (depends on the type of bone).
 
Window Level (WL) and Window Width (WW)
The term ‘window level’ represents the central Hounsfield unit of all the numbers within the window width. The window width covers the HU of all the tissues of interest and these are displayed as various shades of gray. Tissues with CT numbers outside this range are displayed as either black or white. Both the WL and WW can be set independently on the computer console and their respective settings affect the final displayed image (Figs 1.6A and B).
 
Slice Thickness
It is the collimation of the X-ray beam as it emerges from the X-ray tube. The slice thickness can be varied depending on the anatomical region to be covered by varying the beam collimation. For example, orbit scanning is done using 2 to 3 mm slice thickness, posterior fossa using 4 to 5 mm slice thickness and supratentorial brain parenchyma using 10 mm slice thickness.
 
Pitch
Pitch is the terminology used in helical scanning and denotes the distance traveled by the table (in millimeters) during one complete rotation of the X-ray tube, divided by the slice thickness (millimeters). Increasing the pitch by increasing the table speed, reduces dose and scanning time, but at the cost of decreased image resolution (Figs 1.7A and B).
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IMAGE POST-PROCESSING
Post-processing the acquired volumetric data during spiral CT is done in ways appropriate to the clinical situation such as:
  • Multiplanar reformatting (MPR): After obtaining the serial axial volumetric data, the computer can reconstruct the data in sagittal and coronal planes. With the current multislice CT scanner, it is possible to obtain isotropic sagittal and coronal reconstructions thereby reducing patient radiation dose. These are useful in pediatric patients and trauma patients who cannot be positioned for direct coronal (Fig. 1.8).
  • Three-dimensional imaging: The acquired data can also be projected as a three-dimensional model to display spatial information or surface characteristics (volume and surface rendering). This is useful in pediatric craniofacial anomalies and post-trauma for maxillofacial injuries to guide the surgeon in treatment planning (Fig. 1.9).
  • CT angiography (CTA): This involves injection of 100 to 120 ml of contrast media rapidly using a pressure injector at a predetermined rate of injection. Serial axial images are obtained. These images are then used for reconstruction of the data using maximum intensity projection to get a display of the vascular tree. By altering the time of image acquisition and contrast injection, we can obtain only the arterial or venous phases (Figs 1.10A and B).
 
CT Contrast Media
These are iodine containing compounds. Iodine absorbs X-rays within the CT range (120 KVp) since iodine has an atomic number of 53 and atomic weight of 127.
There are two types of contrast agents used:
  1. Ionic contrast: These are sodium or methylglucamine combined with tri-iodinated benzene ring to form soluble salts. These are hyperosmolar and hence are likely to cause severe contrast reactions. These are contraindicated intrathecally.
  2. Nonionic contrast: These are near iso-osmolar and hence tend to produce fewer side effects and considered relatively safe for patients.
Absolute contraindication for contrast:
  1. Previous contrast sensitivity
  2. Abnormal renal parameters
Patients with diabetes and multiple myeloma are more likely to develop altered renal function post-IV contrast injection.6
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Figs 1.6A and B: (A) Soft tissue window settings of an axial CT scan of the brain (WW = 100, WL = 30) and (B) Bone window setting of an axial CT scan of the brain (WW = 2000, WL = 220)
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Figs 1.7A and B: (A) A low pitch—tight helic and (B) A pitch of more than one—loose helic—shorter scan time at the cost of image resolution
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Fig. 1.8: Coronal reformation of the face showing fractures involving the lateral wall of the left maxillary sinus, zygoma, lateral wall of the left orbit and frontal bone
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Fig. 1.9: Three-dimensional volume rendered CT image of the skull
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Figs 1.10A and B: CT angiogram of the brain: (A) Showing axial MIP image and (B) Volume rendering showing the intracranial circulation
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Patients with myasthenia gravis, sickle cell anemia and pheochromocytoma are at risk of developing contrast-induced symptoms.
 
ADVANTAGES AND CLINICAL USE OF COMPUTED TOMOGRAPHY
  • CT is readily available in most hospitals and cost- effective
  • It is a rapid imaging modality with excellent image resolution hence useful in trauma, pediatric and un-cooperative patients
  • Patients in whom MRI is contraindicated.
 
DISADVANTAGES
  • Radiation: The effective doses from diagnostic CT procedures are typically estimated to be in the range of 2 to 5 mSv.
 
BIBLIOGRAPHY
  1. Bushberg JT, Seibert JA, Leidholdt EM, Boone JM. The essential physics of medical imaging, 2nd edn. Philadelphia, Pa: Lippincott Williams and Wilkins,  2001.
  1. Barrie Grossman C. Magnetic resonance imaging and computed tomography of the head and spine. Williams and Wilkins,  1996.
  1. Huda W, Chamberlain CC, Rosenbaum AE, Garrisi W. Radiation doses to infants and adults undergoing head CT examinations. Med Phys 2001;28:393–9.
  1. Mahadevappa Mahesh. Search for Isotropic Resolution in CT from Conventional through Multiple-Row Detector. Radiographics 2002;22:949–62.
  1. Seeram E. Computed tomography: Physical principles, clinical applications, and quality control. Philadelphia, Pa: Saunders,  2001.
  1. Zatz L. General overview of computed tomography instrumentation. In Potts D (Ed). Radiology of the skull and brain: Technical aspects of computed tomography. St Louis, Mo: Mosby  1981;pp.4025–57.9
 
BASICS OF MAGNETIC RESONANCE IMAGING
Magnetic resonance imaging (MRI) is based on the principles of nuclear magnetic resonance (NMR).
 
PHYSICAL PRINCIPLES OF MR IMAGING
Magnetic resonance imaging is based on the absorption and emission of energy in the radiofrequency range of the electromagnetic spectrum. The human body is primarily made of fat and water which has many hydrogen atoms (almost 63%). The hydrogen atom (1H) consists of a single positively charged proton which spins around its axis. These charged particles create an electromagnetic field, similar to that of a bar magnet (Fig. 1.11).
The proton possesses a property called spin which has a small magnetic field. These spinning particles have a net magnetic moment which has both magnitude and direction. In the absence of an external magnetic field, these protons are randomly oriented.
When placed in a magnetic field of strength B, the protons align themselves parallel or antiparallel to the external magnetic field. There is a low energy state where the poles are aligned N-S-N-S and a high energy state N-N-S-S. These particles can undergo a transition between the two energy states by the absorption of a photon. A particle in the lower energy state absorbs a photon and ends up in the upper energy state. The energy of this photon must exactly match the energy difference between the two states.
Application of a radiofrequency (RF) pulse of appropriate duration and amplitude excites these protons from the lower energy state to the higher energy state.
The MRI signal results from the energy difference of the spins emitted during transition from the higher energy state to the lower energy state. The signal is thus proportional to the population difference between the states (Figs 1.12A and B).
When the RF pulse is applied, the protons are tipped into the horizontal or X-Y plane by an angle termed as the flip angle depending on the type of RF pulse. The rate at which the protons precess is termed as frequency and the angular position of the precessing spin is called the phase of the spin.
The frequency of precession (f) is called the Larmor frequency and is characteristic of the specific nucleus and strength of the external magnetic field and is expressed as:
f = γB
Where f = mHz/sec, B is expressed in Tesla and γ is the gyromagnetic ratio of the specific nucleus and expressed as mHz/T. Hydrogen has the highest gyromagnetic ratio and is the most abundant body element, hence is the natural choice for H signal.
 
Radiofrequency Field
Every nucleus in the body precesses at its own Larmor frequency and will produce an MR signal only when the RF energy is delivered at the correct frequency. The excitation RF pulses are delivered by coils that produce an RF field perpendicular to the external magnetic field. The RF is absorbed by the nuclei and the magnetic moment is tipped away from the Z axis, i.e. axis of the external magnetic field depending on the duration and amplitude of the RF pulse.
 
Free-induction Decay
When the RF pulse is switched off, then magnetic momentum of the nuclei begins to return to its original position, thereby transferring the absorbed energy and inducing alternating current signal in the receiver coil. This is termed as free-induction decay (FID). As this occurs immediately after the RF pulse, this signal is not used for image data. The magnetization is manipulated to generate a useful signal termed as echo, which produces the image.
 
T1 and T2 Relaxation
When the RF pulse is switched off, two processes take place simultaneously:
  1. Recovery of the net magnetic moment in the Z axis— termed as longitudinal or T1 relaxation. T1 is the time required for the build-up of 63 percent of the original magnetization along the Z axis (Figs 1.13A to C).
  2. Loss of phase coherence in the X-Y plane or transverse plane, termed as T2 relaxation.
The nuclei while returning to the ground state dissipate their excess energy to their surroundings, which is called the lattice. This process is named as spin-lattice relaxation (Fig. 1.14). Smaller molecules reorient more rapidly than larger molecules. The medium-size molecules, such as lipids relax faster as their frequency of rotation is closer to the Larmor frequency than that associated with pure water or larger molecules, such as proteins. Thus, T1 relaxation times depend on magnetic field strength because the latter affects the Larmor frequency. Thus, water has long T1s.10
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Fig. 1.11: Every spinning particle possesses a magnetic moment (µ) and creates a magnetic field similar to a bar magnet
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Figs 1.12A and B: (A) Showing protons outside a magnetic field and (B) Showing exited protons in a magnetic field moving from a lower energy level to a higher energy level with two distinct energy levels. The population difference is directly proportional to the magnetic field strength
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The transverse magnetization occurs due to magnetic field generated by the surrounding electrons allow the precessing nuclei to experience different field strength. Loss of transverse magnetization (phase coherence) occurs as the magnetic moments get out of phase as a result of their mutual interaction. Anything that changes the magnetic field strength also changes the precessional frequency and causes a loss of phase coherence (dephasing) and shrinking of the transverse magnetization. This is called T2 relaxation or spin-spin relaxation (Fig. 1.15). It denotes the loss of phase coherence caused by interactions between neighboring magnetic moments. T2 is the time required to reduce the transverse magnetization to 37 percent of its original value.
In biological tissues, the main contribution to T2 relaxation is from the relatively static magnetic field from neighboring protons. Large molecules, which tend to reorient more slowly than small molecules, promote T2 relaxation and have shorter T2 times. Free water has a longer T2 than water associated with macromolecules. The T2 is relatively independent on the field strength.
 
Repetition Time
The time between two RF excitation pulses is called the repetition time (TR). The TR can be chosen from a certain minimum value, depending on the imaging technique and the MR system, to very long times.
Longer values of TR allow more T1 relaxation to occur, and this property can be exploited to manipulate the contrast between tissues with different T1s or the signal-to-noise ratio in an image.
 
Echo Time
The time from the center of the RF excitation pulse to the center of the echo is the echo time (TE). The amplitude of the transverse magnetization at the echo peak depends on TE and T2 of the tissue. As TE is prolonged, the transverse magnetization becomes weaker. Adjusting TE influences the contrast between tissues that have different T2s.
 
Slice Orientation
The orientation of a slice, i.e. axial, coronal or sagittal depends on which of the three magnetic field gradients is activated during the RF pulse. An RF pulse in the presence of the z-gradient creates a transverse slice. The x- and y-gradients select slices in the sagittal and coronal orientations, respectively. Oblique slices are created by activating two or more gradients during an RF pulse.
 
Slice Position
Slices are located where the Larmor frequency matches the frequency of the RF pulse. The slice-selection gradient lowers the Larmor frequency on one side of the center of the magnet and raises it on the other side. Slice position is controlled by changing the frequency of the RF pulse because changing the amplitude of the slice-selection gradient would inadvertently alter the thickness of the slice.
 
INSTRUMENTATION
The key components of an MR system are the magnet, the gradient, the radiofrequency subsystem, and the computer.
 
The Magnet
The magnet is the main component of the MR system. There are three types of magnets in common use for MRI—permanent magnets, resistive electromagnets, and superconducting electromagnets. The higher the field strength better is the signal-to-noise ratio. The strength of the magnetic field is measured in Gauss (G) or Tesla (T) units (10,000 G = 1 T). Diagnostic MR systems usually employ magnets with operating field strengths ranging from 0.02 to 3T. Research systems operate above 3T.
 
Superconducting Magnets
These are most commonly used magnets and operate at field strength above 0.5T. Some metals (e.g. Hg) and alloys (e.g. niobium/titanium, Nb/Ti; niobium/tin, Nb3Sn; and vanadium/gallium, V3Ga) lose their electrical resistance at very low temperatures and become superconductors. The superconductor most widely used in the construction of clinical magnets is Nb/Ti. This alloy becomes superconducting at 10° Kelvin (K) in the absence of an external magnetic field. This temperature is provided by a bath of liquid helium (4° K) (Fig. 1.16).
 
Resistive Magnets
A resistive magnet is an electromagnet in which the magnetic field is generated by the passage of electrical current through a wire. The disadvantage is their high power consumption limiting field strength.
 
Permanent Magnets
It uses a horse-shoe magnet. An advantage of these low-field permanent magnet systems is that their C-shaped design is patient friendly and therefore useful in claustrophobic patients.12
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Figs 1.13A to C: (A) Alignment of the proton along the direction of the external magnetic field (B0) in the z-axis, (B) After applying the RF pulse of an appropriate frequency, the magnetization (M0)/protons are tipped away from its equilibrium in the x-y plane and (C) If a longer pulse lasting twice as long the magnetization is inverted
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Fig. 1.14: Spin-lattice relaxation time
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Fig. 1.15: Spin-spin relaxation—T2 relaxation—loss of magnetization in the x’-y’ plane is faster than the loss of magnetization in the z-direction due to loss of phase coherence of the microscopic components
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Fig. 1.16: Schematic representation of the superconducting MR systems. The bore is surrounded by the coils of the wire through which electric current is passed and cooled by liquid helium to achieve magnetization and desired field strength. The current is disconnected once magnetized
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Their field strength is limited to 0.5T (Fig. 1.17).
 
Magnetic Field Gradients
Magnetic field gradients are activated as pulses for a short duration at timed intervals. It is a magnetic field that increases in strength along a particular direction, e.g. x-, y-, and z-gradients, according to the direction of change of the magnetic field strength. The strength of a gradient refers to the rate at which its magnetic field changes with distance.
 
RF System
The excitation of the nuclei is done with a short duration RF pulse close to or at the Larmor frequency of the nuclei. The desired frequency is produced by a frequency synthesizer.
The receiver detects signals in the high- and very-high-frequency (HF and VHF) range. The magnetic resonance signals are typically a few µV in amplitude.
 
Transmitter and Receiver Coils
The body part to be examined is placed inside a coil. Separate coils can be used for transmitting and receiving or a single coil can be used for both excitation and detection (transceiver coil).
A coil is a winding of low-resistance wire, usually made of copper. Volume coils are used for large body parts. Surface coils are coils used to study small region such as orbit. The advantage of surface coils is that their signal-to-noise ratio is better as the part is close to the coil. Surface coils can receive a good signal from the tissues within the depth of half of its diameter.
 
COMMONLY USED PULSE SEQUENCES
 
Spin-Echo Pulse Sequence
In a spin-echo pulse sequence two RF pulses, i.e. 90° and 180° are applied spaced by a time interval of TE/2. After the nuclei are excited by a 90° pulse, the spins dephase in the x’-y’ plane, this is followed by a refocusing 180° pulse. The faster spins lie behind the slower ones, but at time TE/2 they make up, thus producing an echo. The 180° pulse results in reversal of the phase of each spin. The position of the spins has not changed, so they will continue to rotate in the same direction. However, the 180° pulse causes the spins to return towards their starting point (alignment), rather than rotating further away from it. This 90°-180° pulse sequence is called spin echo (SE) sequence (Fig. 1.18).
By altering the echo delay time (TE), and the sequence repetition time (TR), the SE sequence can be used to obtain T1, T2, or proton density images.
The spin echo sequence has been largely replaced by faster sequences such as fast spin echo and fast GRE (Gradient recalled echo).
 
GRADIENT-ECHO IMAGING
Gradient-echo imaging is an imaging technique by which images can be acquired in much shorter times than conventional pulse sequences. The basic difference between spin-echo and gradient-echo imaging is that, gradient-echo uses gradient reversals to get an echo and spin-echo uses 180-degree rephrasing pulse and gradient- echo uses flip angle < 90 degrees (Fig. 1.19).
 
INVERSION RECOVERY IMAGING
The inversion recovery sequence uses a 180-degree inverting pulse, a 90-degree pulse, and a rephrasing 180-degree pulse. The inversion time (TI) is determined by the TR and T1 of the tissue needed to be suppressed (Fig. 1.20). Commonly used inversion recovery pulse sequences are:
  • Fluid attenuated inversion recovery (FLAIR), whereby the CSF bright signal is suppressed. It is now a routinely used sequence in brain imaging and especially to image periventricular plaques in multiple sclerosis
  • Short-tau inversion recovery (STIR) sequence, mainly used in imaging the optic nerves. It suppresses the orbital fat and highlights the lesions within the optic nerve mainly in optic neuritis.
 
MAGNETIC RESONANCE ANGIOGRAPHY (MRA)
Advantages of MRA vs catheter angiogram:
  • Noninvasive or minimally invasive
  • Three-dimensional information can be obtained
  • Can give surrounding soft tissue details.
Disadvantage: Flow dynamic information is lacking.
 
Techniques of Magnetic Resonance Angiography
The commonly used techniques in clinical practice are:
  • Time-of-flight (TOF) MR angiogram
  • Phase contrast (PC) MR angiogram
  • Contrast-enhanced MR angiogram (CE-MRA).
 
Time-of-flight Angiogram
This is the most widely used MR angiography technique for imaging the intracranial circulation.14
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Fig. 1.17: Schematic diagram of a permanent MR system showing the generation of the magnetic field in a vertical direction by magnetized ceramic blocks
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Fig. 1.18: Diagram of a spin-echo pulse sequence—spin-echo pulse sequence. The spin system is excited by a 90° pulse. After a time delay (τ), one or several 180° pulses follow. This leads to the formation of an echo. The time between the 90° pulse and the peak of the echo is called echo time (TE). TR is the repetition time between two complete pulse sequences
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Fig. 1.19: Formation of a gradient echo. Instead of the 180° pulse, a gradient pulse (–G) is used followed by a second gradient pulse of opposite polarity (+G). In gradient-echo sequence, the signal decay is determined by T2*, which is always less than T2
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Fig. 1.20: Pulse sequence diagram of an inversion recovery pulse sequence. The 180° inverting pulse is followed by a 90° pulse and 180° rephasing pulse
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It gives reliable vascular information without the need for intravenous contrast.
Basic principle involves suppression of the static background tissue and retaining the signal from the flowing blood. The saturation of the stationary tissue is done by using very short TR so that the stationary spins do not have enough time to regain their longitudinal magnetization. The flowing (unsaturated) spins which enter the slice are unaffected by the slice selective RF pulse, will be fully magnetized producing a bright signal (Fig. 1.21). The signal produced is directly proportional to the velocity of the flowing blood.
The flow saturation will occur when the spins in the imaging volume are not entirely replenished after each pulse.
The time-of-flight angiogram can be obtained using 2D or 3D sequences. In a 2D sequence, sequential thin sections are obtained whereas in 3D a slab of tissue is excited. Each of them has their advantages and disadvantages.
2D angiograms are used to evaluate slow flowing blood, but susceptible to turbulent flow. It has less spatial resolution.
3D angiograms have high spatial resolutions and less susceptible to turbulent flow.
 
Phase Contrast MR Angiogram
Moving spins undergo a phase shift in the presence of paired opposing gradients. This phenomenon is utilized in phase contrast MR angiography. The amount of phase shift increases with increasing flow velocity. When the flowing blood (moving spins) moves along the direction of the gradient field, it precesses faster as the field increase and undergoes a phase change. Thus, the motion is phase encoded giving it both direction and magnitude (Fig. 1.22).
The amount of phase shift is directly proportional to the flow velocity, gradient strength and time interval between the gradient applications. By choosing an appropriate velocity encoding value (VENC), fast or slow flowing blood can be imaged.
Phase contrast MR angiography can be acquired as both 2D and 3D sequences.
Advantages of phase contrast MRA: It gives:
  • Flow quantification
  • Flow direction
  • Excellent background suppression
  • Can be used for imaging areas of slow flow.
Disadvantage: Long scan time.
 
Contrast-enhanced MR Angiography (CE-MRA)
The limitations of TOF and PC angiograms such as flow saturation, flow-related artifacts, breathing and pulsation artifacts made depiction of blood vessels in the body especially the abdomen difficult.
By using intravenous contrast and rapid gradient imaging, it is now possible to obtain MR angiography images almost at par with conventional angiogram.
The technique involves capturing of high magnetization strength during the first pass of the vascular contrast, i.e. gadolinium by appropriate timing using 3D acquisition (Fig. 1.23).
Advantages:
  • Insensitive to saturation effects of the RF pulse as against TOF angiogram. Therefore, can cover vessels over larger field of view
  • Useful in large aneurysms where flow is complex.
 
MR CONTRAST
Most of the contrast agents in clinical use enhance tissue relaxation. Gadolinium is a rare earth element and toxic by itself, hence it is chelated with multidentate ligands for safety such as diethylenetriamine pentetate (DTPA)— tetra-azacyclododecane tetra-acetic acid (DOTA). It is a paramagnetic substance and shortens the T1 relaxation and hence makes the tissues with contrast appear bright.
 
Safety
  • These contrast agents are considered safe with rate of adverse reaction such as nausea and vomiting (1 to 2%) and hives (1%). Severe anaphylactoid reactions have been reported with an estimate rate at 1 in 200,000 and 1 in 400,000
  • These contrast agents can be safely used in children above 2 years
  • They should not be used in patients with compromised renal function. There have been cases reported of nephrogenic systemic fibrosis in patients with compromised renal function
  • Should not be used in pregnancy as its bioeffect on the fetus has not been established.
 
Newer Advanced MR Imaging Techniques
 
Diffusion Weighted Imaging
It is based on the principle of Brownian motion, which is dispersion or random translation of a molecule in a liquid due to thermal agitation.
Motion of molecules in biological tissues is complex. Neuronal tissue consists of tightly and coherently packed axons surrounded by glial cells.16
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Fig. 1.21: Schematic representation of time-of-flight angiogram
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Fig. 1.22: Schematic diagram of a phase contrast MR angiogram
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Fig. 1.23: Contrast-enhanced TRICKS (Time Resolved Imaging of Contrast Kinetics) angiogram image of the brain
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Fig. 1.24: Axial diffusion weighted image showing restricted diffusion in the right occipital lobe suggestive of acute right posterior cerebral artery territory infarct
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The movement of water molecules is hindered in a direction perpendicular to the orientation of the axonal fibers. Thus, motion of molecules in biological tissues is anisotropic. The cell membranes are thought to be responsible for anisotropic diffusion rather than myelin. The restricted diffusion appears as a bright signal on diffusion weighted images (Fig. 1.24).
Application:
  • Stroke
  • Multiple sclerosis
  • Tumors
  • Trauma
  • Abscess.
Lesions bright on diffusion images:
  • Acute infarct
  • Bacterial abscess
  • Acute demyelination
  • Epidermoid cyst
  • Tissues with high cellularity
  • Subacute hemorrhage.
 
Functional Imaging
It is the demonstration of brain activation to a specific stimulus based on the functional anatomy of the brain, for example, the primary visual cortex is activated using a flicker display or alternating checkerboard pattern as a visual stimulus. Once the brain is activated using a stimulus, there is change in the blood flow to the particular region to supply the increased demand for oxygen and glucose. This increase in the oxygen, i.e. deoxyhemoglobin concentration causes local susceptibility effects which is used to receive the signals using appropriate pulse sequences. This is termed blood oxygen level dependent (BOLD) contrast imaging (Fig. 1.25).
 
MAGNETIC RESONANCE SPECTROSCOPY (MRS)
MR spectroscopy utilizes the differences in the resonance frequency of nuclei due to their different chemical bond. This is also termed as chemical shift imaging. The frequency difference varies with the magnetic field and is directly proportional to the external magnetic field. It is expressed in parts per million (ppm). The advantages of a higher field strength while performing spectroscopy is that it provides better signal-to-noise ratio and better separation of metabolite peaks.
1H (proton) spectroscopy is used for brain imaging as it is easy to perform and gives a better signal-to-noise ratio as compared to 23Na and 31P. Of all the atomic nuclei, 1H has the strongest response and found in all biochemicals. MR spectroscopy thus provides details of the brain chemistry (Figs 1.26A and B). The spectrum is read from right to left and the metabolites detected on brain spectroscopy are:
  • Lipid 0.9 –1.4 ppm
  • Lactate 1.3 ppm
  • N-acetyl asparatate (NAA) at 2 ppm
  • Creatine (Cr) 3.0 ppm
  • Choline (Cho) 3.2 ppm
  • Myo-inositol 3.5 ppm.
The TE affects the metabolites detected, thus short TE ~ 30 ms shows metabolites with short and long T2 relaxation times and with long TE ~ 270 ms only metabolites with long T2 relaxations times are detected, therefore the spectrum primarily consist of NAA, Cr and Cho. Another advantage of long TE ~ 144 ms is that the lactate peak at 1.3 ppm gets inverted.
Rather than absolute concentrations, one should rely on the various ratios to give a clinical diagnosis.
Ratio
Normal
Abnormal
NAA/Cr
2.0
< 1.6
NAA/Cho
1.6
< 1.2
Cho/Cr
1.2
> 1.5
Indications:
  • Radiation necrosis vs recurrence
  • Infections
  • Neurodegenerative disorders
  • Metabolic brain disorders
  • Stroke.
MR spectroscopy should be carefully interpreted and correlated with MR images to make a final diagnosis.
 
BIBLIOGRAPHY
  1. Barrie Grossman C. Magnetic resonance imaging and computed tomography of the head and spine. Williams and Wilkins,  1990.
  1. David D Stark, William G Bradley. Jr Magnetic resonance imaging, 3rd edn. CV Mosby. 
  1. Edelman R, Hesselink R, John, Zlatkin B Michael, Crues V John. Clinical magnetic resonance imaging, 3rd edn, Philadelphia: Saunders-Elsevier,  2006.
  1. Peter A Rink. Magnetic Resonance in Medicine. The Basic textbook of the European Magnetic Resonance Forum, 5th revised edn, 2003.
  1. Patric Hagmann, Lisa Jonasson, Philippe Maeder, Jean-Philippe Thiran, Van J Wedeen, Reto Meuli. Understanding Diffusion MR Imaging Techniques: From Scalar Diffusion-weighted Imaging to Diffusion Tensor Imaging and Beyond R. Radiographics 2001;21:767–79.
  1. Ross BD, Colletti P, Lin A. MR spectroscopy of the brain: Neurospectroscopy in Edelman. Hesselink, Zlatkin and Crues (Eds): Clinical Magnetic Resonance Imaging, 3rd edn, Philadelphia: Saunders-Elsevier,  2006; pp.1840–910.18
 
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Fig. 1.25: fMRI showing activation of the motor cortex after right finger-tapping
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Figs 1.26A and B: Multivoxel MR spectroscopy: (A) Showing the voxel placed in the normal parietal white matter with NAA color map and (B) Showing normal spectrum