Answers – 86 References – 87
3.1 Background
3.1.4 Overview of Brain Imaging Modalities
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systolic blood pressure, and aortic atherosclerosis are other risk factors for white matter hyperintensities [23–25]. These changes in white matter can be seen in periventricular and/
or in deep white matter locations. Despite relatively similar appearance on MRI (T2 and FLAIR sequences; see below for description), the nature of these lesions is somewhat hetero- geneous in terms of histological nature including white mat- ter infarction, gliosis, or plaques of demyelination [26, 27].
The appearance of periventricular white matter hyperintensi- ties in terms of the irregularity of lesion edges is thought to be related to their histological nature, i.e., lesions with irregular edges tend to be ischemic in nature, while those with smooth edges are more related to gliosis and demyelination [28].
The above changes in brain structure likely contribute to modification in brain networks in old age. Detailed discus- sion of these changes is beyond the scope of this chapter.
Briefly, there are two main lines of research into cognitive network activity and connectivity with aging: task-related fMRI and resting-state fMRI. Task-related fMRI involves obtaining brain activation pattern as it is related to a cogni- tive task (i.e., event-related, like working memory and epi- sodic memory tasks). Several studies identified modification in fMRI BOLD (described below) signal during cognitive tasks in older compared to younger adults. Examples of these modifications include loss of cortical specificity (dedifferen- tiation), recruitment of wider network to achieve the same task (indicating compensation), and more involvement of frontal rather than posterior networks when performing the cognitive task. This sizeable literature resulted in several models to explain cognitive aging including scaffolding the- ory of aging and cognition (STAG) [29], system vulnerability view [30], and brain maintenance hypothesis [31].
Resting-state fMRI (rsfMRI) connectivity studies have the advantage of not being task-performance dependent, which makes them more feasible and less subject to vari- ability. Converging evidence suggests reduction in connec- tivity in large-scale networks including DMN, SN, and CEN networks with aging. This change correlates with cognitive changes with aging and is likely the result of changes in brain structure described above (please see [32] for review).
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The quality of CT images including spatial resolution, noise, and contrast resolution is closely related to the absorbed x-ray dose. In turn, CT dose is governed by the amount of x-ray delivered (measured in mill amperage, mA), quality of x-ray beam delivered (measured in peak kilovoltage, kVp), and the length of time the patient is exposed to the x-ray (measured in seconds) [33]. The amount of radiation given to a patient is measured and tightly regulated to ensure that overexposure leading to tissue damage is avoided. A measure of the quantity of deposited radiation energy including x-ray, gamma ray, beta, or alpha particles is called the absorbed dose. The unit of absorbed dose is gray (Gy) or in metric, rad (radiation absorbed dose). 1 Gy = 100 rad. In CT, dose expo- sure is estimated for each patient examination and related to several factors such as dose distribution of the beam intensity along the patient path, thickness of the beam collimation, the number of detectors, and the pitch, which is the distance the table travels per tube rotation [34]. For simplicity, a calcu- lated absorbed dose reference level, referred to as the volume computed tomography dose index (CTDIvol), is provided on the scanner by the manufacturer for each scan when the appropriate imaging parameters are set. .Table 3.3 out- lines suggested imaging parameters for a routine CT brain scan applicable to geriatric psychiatry clinical indications.
Note that most clinical CT scans of the brain do not require exogenous contrast enhancement, except in cases related to cerebrovascular disease or neuro-oncology. As such the use of contrast media is not included and will not be further dis- cussed.
Teaching Point
In general for all forms of radiation, absorbed dose reduction can be achieved by minimizing the distance from the radioactive source, decreasing time exposed to the source, and increasing collimation thickness.
Magnetic Resonance Imaging (MRI)
Since the production of the first MR image of a live human nearly 40 years ago, MRI has rapidly advanced at a pace much faster than other imaging modalities [35]. The first human scan in July 1977, for instance, was an image of the chest acquired over 4 hours in a scanner twice the diameter and length of current clinical scanners (7http://www.smithsonian- mag.com/). Today, a similar image can be acquired in less than 5 minutes with ten times better spatial resolutions. Over the past 40 years, the strength of the magnetic field has gone from 0.5 Tesla in the mid-1980s to 7 Tesla. Tesla (T) is a measure of the magnetic flux density. One Tesla is equivalent to 10,000 gauss (a measure of magnetic flux). For comparison, the earth’s magnetic flux density is around 5 × 10−5 T or 0.5 gauss, and a refrigerator magnetic field strength is 0.005 T. Modern clinical MRI scanners have field strength of 1.5 T. MRI scanners used in clinical research can range from 3 T to 7 T.
The ability of MRI to produce high spatial resolution and good soft tissue contrast makes MRI appealing for
investigating subtle anatomical (and to a larger extent, func- tional) abnormalities. An MR image of the brain is created when the head is exposed to the electromagnetic field pro- duced by a strong MRI magnetic field. The human body is made up of large pools of hydrogen atoms (H+, a component of water or H2O), each consisting of one subatomic particle in the nucleus, called a proton. Hydrogen protons rotate freely around their own axes, in a manner similar to a toy- spinning top after it has been hit. When hydrogen atoms in the body are exposed to the strong magnetic field of an MRI, the protons are “magnetized,” meaning the initial random nondirectional motion of their collective spins is now aligned with the direction of the MRI’s magnetic field. In the case of a
.Table 3.3 Standard clinical CT brain imaging parametersa Param-
eter Description Set value
Mode Helical scanners can be operated in sequential or non-spiral mode, often referred to as axial mode or step-and- shoot. Helical offers faster imaging and less motion artifacts
Helical
kVp Peak kilovoltage; the maximum voltage applied to the tube
120
mA Tube current or the number of electrons produced by the tube per second. The lower the tube current the lower the dose
300
FOV Vendor preset field of view (FOV);
small, medium, or large body. For head, small body is used to maintain isotropic imaging and minimize radiation exposure
SFOV
Rotation time
The lower the rotation time, the lower the radiation dose but the higher the image noise
0.5 seconds
Pitch Applies to helical CT; the distance the table travels in mm in one 3600 tube rotation. The larger the pitch, the lower the patient dose
0.531:1
Speed (mm/
second)
The table feed speed; how fast the table translates through the gantry per tube rotation in mm per second
10.62
CTDIvol (mGy)
CT dose index; the diagnostic reference level set by the American College of Radiology for brain imaging is 75 mGy
54.15
Image recon- struction
One with standard algorithm and second one with bone algorithm;
thickness and interval = 5 mm
2
aThis is based on typical imaging parameters used on a 32-multislice GE Healthcare CT scanner. Variations between vendors might exist. For more information, other vendors suggested CT brain imaging parameters (see AAPM Adult Routine Head CT Protocols Version 2.0 March 1, 2016) (7https://www.aapm.org/pubs/CTProtocols)
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patient lying on their back for a brain scan, majority of their hydrogen protons will line up in the direction of the head or feet when placed inside the magnetic field of the MRI scan- ner. The protons aligned in either the head or feet will cancel out, leaving a few protons that are not perfectly aligned to the MRI field. If the MRI radiofrequency pulse is applied to the remaining protons, it will cause the protons to spin at a certain frequency and in a given direction. When the radiofrequency pulse is removed, the hydrogen protons will return back to their natural orientation, releasing energy that is detected as MRI signals. This process is called relaxation.
Thus, the MRI contrast in an image depends on the amount of relaxation of the hydrogen protons.
Hydrogen proton relaxation is measured using two MRI values, T1 and T2. T1 and T2 images are acquired using T1-weighted and T2-weighted pulse sequences, respectively (see suggested optimal imaging parameters in .Table 3.4).
Each imaging sequence is generated by manipulating two dis- tinct timing parameters measured in milliseconds (ms), echo time (TE) and the repetition time (TR). Echo time refers to the time duration after application of a radiofrequency pulse and before readout or sampling of the MRI signal following
proton relaxation. Repetition time is the time period between successive radiofrequency pulses. T1-weighted images are created by combination of a relatively short TE and TR times, producing brain images where tissues with bound hydrogen protons (such as fat) give off high signal and appear bright, while fluids containing large pools of free hydrogen protons such as cerebrospinal fluid (CSF) give off low signal and appear dark or black. This is because the free pool of hydro- gen proton in CSF has not fully relaxed back to their natural state giving off only a small fraction of their maximal signal, before signal readout by the computer. The large difference between T1 values of CSF, gray matter, and white matter produces high tissue contrast necessary for evaluating areas of subtle changes in gray matter structure and volume. On the other hand, T2-weighted imaging is the reverse of T1, a combination of long TR and long TE. Fluids such as CSF are the brightest in T2 images, because their long T2 values are ideal for longer TE and TR times, ensuring that a larger fraction of their maximum signal is readout by the computer.
T2 images are ideal for evaluating white matter lesions and intracerebral vascular changes such as hemorrhage, inflam- mation, and edema.
Teaching Point
The large difference between T1 and T2 values of CSF, gray matter, and white matter produces the high tissue contrast necessary for evaluating changes in brain struc- ture and volume.
Other combinations of TR and TE are possible, each producing varying tissue contrasts. For example, a short TE and long TR will produce a proton-density weighted image.
Advanced imaging sequences such as arterial spin labeling (ASL), used to directly measure cerebral blood, or blood- oxygen- level-dependent (BOLD) MRI, for relating changes in deoxyhemoglobin to brain function, are advanced variations of T1/proton-density-weighted and T2-weighted sequences, respectively. ASL and BOLD are increasingly gaining clini- cal relevance especially for mapping of functional brain net- works in populations with major neurocognitive disorders [36] and localizing eloquent regions in epilepsy patients [37]. ASL is a noninvasive MRI-based perfusion imaging technique where the inflow of magnetized blood spins from the neck is measured downstream in the brain, a short time after arrival. The cerebral blood flow (CBF) images from ASL scans are similar to CBF images from PET measurements of radiolabeled 15O-water tracer [38]. Regional CBF measure- ment with ASL also correlates well with regional cerebral metabolic rate measurement using 18F-FDG PET in normal healthy adults [39–41] and Alzheimer disease patients [42, 43]. These observations suggest that these noninvasive and radiation-free MRI techniques are ideal for detecting and long-term monitoring of functional changes in the brain.
In the interim, clinically available T1- and T2-weighted imaging sequences can provide added quantitative
.Table 3.4 Standard T1 and T2 brain imaging sequences (at 3 T)
Sequence Parameters
T1-weighted TR/TE/T1 = 2000/2.96/900 milliseconds Flip angle = 90
Slices/slab = 176 FOV = 256 × 256 Acceleration factor = 2 Voxel size = 1.0 mm3 isotropic Bandwidth = 240 Hz/Px Three-dimensional acquisition
Total acquisition time = 4 minutes 8 seconds T2-weighted
FLAIR (fluid-attenu- ated inversion recovery)
TR/TE/T1 = 5000/395/1800 milliseconds Slices/slab = 160
FOV 256 × 256 Acceleration factor = 2 Voxel = 1.0 mm3 isotropic Bandwidth = 781 Hz/Px Three-dimensional acquisition
Total acquisition time: 5 minutes, 52 seconds Note: 3 T scanner is twice the field strength of 1.5 T. Expect increased signal-noise ratio, temporal resolution (faster imaging), spatial resolution, T1 relaxation times, and specific absorption rates (amount of radiofrequency energy deposited per unit mass of tissue) compared to 1.5 T. Sequence parameters are optimized for a Siemens 3 T Verio MRI scanner. Imaging parameter names may vary with vendor
T1 = inversion time, time between a 180° and a 90° pulse for spin echo-based sequences
FOV = field of view in X (longitudinal) and Y (transverse) plane Acceleration factor increases acquisition speed by a set factor value
Bandwidth is the frequency range allowed in the MRI signal in hertz per pixel (Hz/Px)
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information beyond qualitative visual inspections currently practiced. Statistical modeling techniques can be applied to T1 and T2 images to reveal subtle regional differences over time within an individual or groups of patients or between patients and controls [44]. Voxel-based morphometry (VBM) and cortical thickness mapping (CTM) are auto- mated techniques for segmenting T1 or T2 images and statis- tical mapping of local changes in gray matter or white matter concentration, cortical shape, and cortical thickness. Clinical T1 scans can be readily analyzed using open-source and free- for-use software such as Statistical Parametric toolbox (SPM, University College, London, UK, 7www.fil.ion.ucl.ac.uk/
spm/) and FMRIB Software Library (Oxford University, Oxford, UK, 7http://fsl.fmrib.ox.ac.uk) for VBM analysis or Free Surfer (Harvard University, Boston, USA, 7https://
surfer.nmr.mgh.harvard.edu) for CTM. Several VBM and CTM studies in large cohorts of individuals over the human life span have revealed temporal patterns of age-related brain atrophy [45, 46] and distributed patterns of atrophy related to neurodegenerative disease [47, 48]. It is evident that VBM and CTM could also serve as sensitive diagnostic biomarkers for detecting, characterizing, and quantifying neurobehav- ioral disease progression.
Contraindications for Structural Imaging
MRI and CT are complimentary structural imaging tools. In certain cases one modality could be preferred over the other.
Contraindications and advantages for each imaging tools are outlined in .Table 3.5. In general, both CT and MRI are poor at handling metallic objects embedded within the skull, jaw, or brain. Surgical aneurysm clips, orthodontic implants, and dental crowns create signal voids in MRI and streak- ing artifacts from x-ray beam hardening and scatter in CT images. Known allergic reactions to MRI contrast media for evaluation of vascular lesions, neoplasms, stroke, and inflam- mation may permit the use of a complimentary non-contrast CT scan. Evaluation of bone-related neoplastic processes or evaluation of cortical ossicles is better obtained from CT scans compared to MRI.
Functional Brain Imaging: SPECT and PET The primary role of functional imaging in neuroradiology has long been recognized as a means of detecting subtle disease- related brain changes. Patterns of functional brain abnormalities serve as biomarkers for detecting and charac- terizing brain disorders and often precede deficits in cogni- tive function, which in turn precede structural abnormalities [49]. The fundamental understanding that brain function is related to local changes in cerebral blood flow (CBF) or glu- cose metabolism (CMRglc) is central to the use of functional imaging techniques such as PET and SPECT in evaluation of functional deficits. Earlier measurements of brain function in humans used invasive serial sampling of arterial/venous blood during continuous inhalation of inert gas mixtures to measure changes in global CBF [50] or linked indirect changes in blood volume from cortical surface pulsations measures to task-evoked brain responses [51]. The advent of
nuclear engineering following World War II led to discov- eries of safe radioisotopes capable of capturing in vivo local changes in CBF or CMRglc, revolutionizing the field of neu- roimaging.
PET and SPECT imaging are in vivo molecular imag- ing methods used in mapping the distribution of trace amounts (nano- to picomoles) of radioactive compounds injected into the body. Radiotracers are radioactive isotopes attached to targeted analogs of biologically active molecules found within the body such as a sugars, proteins, enzymes, or hormones. Using highly sensitive scintillation detec- tors, the minute amount of radioactive atoms bound to the radiotracers within the body are detected each time an atom disintegrates (decays), from an unstable radioactive atom to its stable chemical form. For instance, fluorine-18 (18F), the most common PET tracer, is a radioactive fluorine isotope that decays after a short period of time to a stable 18oxygen (18O) molecule. Depending on the atomic composition of the radioactive isotope, each radioactive decay results in release of either gamma photons, beta particles, or alpha
.Table 3.5 Comparison of structural brain imaging modalities
CT MRI
Radiation burden
1 scan (~2.2 mSv/year)a is less than the average annual radiation dose from natural back- ground radiation (3.1 mSv/year)a
None
Tissue contrast
Good soft tissue contrast
Excellent soft tissue contrast
Scan time Faster, order of seconds to less than 1 minute;
less prone to patient motion
Slow; few minutes long for each sequence; prone to patient motion Claustro-
phobia
Relatively larger bore;
easier for
claustrophobic- prone patients
Smaller enclosed bore; sedation may be offered
Trauma- friendly
Yes In certain cases;
transfer from emergency room to MRI-safe monitoring devices and intravenous lines are required
Metallic implant contraindi- cation
Safe to scan; metallic streaking artifacts might exist
Conditional-safe;
thorough screening is required before admitting patient to MRI suite
aNCRP Report No. 160, Ionizing Radiation Exposure of the Population of the United States. Available at 7www.ncrpon- line.org, Accessed November 19, 2016
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particles. The amount of time it takes for a radioisotope to decay is measured by the length of time it takes a quantity of the radioisotope to decrease by half, known as the half- life. A quantity of radiation emitted by a radioisotope mea- sured in SI unit is the becquerel (Bq), or non-SI metric unit is the curie (Ci), named after the 1903 Nobel Prize in Physics winners, Henri Becquerel, Pierre Curie, and Marie Curie.
1 Ci = 3.7 × 1010 Bq and 1 Bq = 1 disintegration per second.
When a radiotracer is injected, inhaled, or ingested into the body, the tracer interacts physiologically with the targeted body tissue, decays, and emits gamma photons detected by scintillation detectors. The location of distributed radiotrac- ers in the body can be tracked by the path the photon traveled after exiting the body to strike the detector crystals, known as the line of response (LOR). In SPECT, a single gamma photon is released after tissue absorption and each radiation decay. The LORs are shaped by collimators mounted over the large flat scintillation detectors to allow or block photons, depending on the LOR angle. Photons are counted if they pass between septal walls of the collimator and fall within a set energy window, typically around the radioisotope’s peak energy. The sensitivity (the amount of photons detected) and spatial resolution (the full-width-half- maximum (FWHM) of the line spread function) are largely influenced by the shape, width, and length of the collimator. For brain imaging, fan-beam or cone-beam collimators are preferred to parallel- hole collimators due to their ability to converge and magnify objects at a distance away from the detector [52]. Parallel hole is more readily available and can be used if placed in close proximity to the head. SPECT cameras with more than two detectors (dual-head systems) are necessary for brain imaging, to improve sensitivity and minimize motion from long scan times.
In contrast to SPECT, PET tracer LORs are defined by the coincidence detection of two photons arriving at the same time in almost opposite directions, close to 1800 apart. In PET imaging, a positively charged beta particle, called a pos- itron, is initially emitted following radioactive decay, exists briefly, and is quickly annihilated to two gamma photons.
The gamma photons exit the body at opposite directions and are counted if they interact with the detector crystals at the same time. As such, PET systems are rings of detector blocks mounted on a gantry similar to CT systems, but unlike CT, they are stationary. Because of the coincidence detection nature of PET, physical collimation is not required, giving PET a much higher sensitivity profile compared to SPECT (see .Fig. 3.2). However, electronic collimation is used in PET to improve signal-to-noise ratio, limited by noncol- linearity of annihilated photons and detection of random coincidences. Other contrasting features of PET and SPECT are listed in .Table 3.6.
Because of the added costs for production of SPECT tracers and more so for PET, anatomical imaging using CT and MRI techniques is more prevalent in clinical neuroim- aging and often used as the first line of investigation before functional imaging. The introduction of hybrid PET/CT and recently PET/MRI has seen a resurgence in demand for PET
imaging. Ten years ago, 18F-2-fluoro-2-deoxy-D-glucose (18F-FDG) was the sole PET tracer approved for clinical diagnostic brain imaging in the United States and Europe.
Recently, the US Food and Drug Administration (FDA) approved the use of three PET ß-amyloid tracers for demen- tia imaging [53] and one new SPECT dopamine transporter (DaT) tracer for parkinsonism syndromes [54]. In addition, a number of promising brain radiotracers with high clinical application have been approved for use in clinical trials or human research studies in the United States, Canada, and Europe (see .Table 3.7).
Nonetheless, 18F-FDG still remains the most widely used tracer for PET imaging since it was first synthesized as an analog to Louis Sokoloff and colleagues’ 14C-2-deoxyglucose [55]. The normal regional distribution of 18F-FDG in the brain is well established. 18F-FDG is an analog of glucose where one of the normal hydroxyl groups in position 2 of the glucose carbon cycle has been replaced with 18F. When injected into the bloodstream, 18F-FDG initially behaves sim- ilar to glucose and is transported into brain cells and phos- phorylated, but then cannot undergo further glycolysis. This
mM
μM
Sensitivity (molar conentration) nM
0.5
Spatial resolution (mm)
1 4 10
CT
MRI/MRS
Metabolic
Molecular SPECT PET
Structural
.Fig. 3.2 Comparison of sensitivity and spatial resolution limits across neuroimaging modalities
.Table 3.6 Key differences between PET and SPECT
Characteristics PET SPECT
Radiation decay Two positrons Single gamma photon Half-life Short (minutes to
hours)
Longer (hours to days)
Energy (kiloelec- tron volts, kev)
~511 93–364,
99mTC = 140 Production Cyclotron- produced Generator-
produced Neuroimaging in Clinical Geriatric Psychiatry