Nuclei with an odd number of protons and neutrons (such as hydrogen) have a magnetic moment causing the nucleus to act like a small magnet in the presence of an external magnetic field. The magnetic field of the nucleus aligns in the direction of the external magnetic field. Magnetic resonance imaging uses radio frequency (RF) pulses in a magnetic field in order to alter the spin of protons in the tissue. Coils detect the change in net magnetization, which after mathematical reconstruction provides spatial and compositional information of the tissue being imaged. Because clinical MRI usually detects magnetization of hydrogen, compositional information is limited to molecules containing hydrogen, such as water, body fat, and cholesterol. In a MRI scanner, proton spins in the body align in the direction of the external magnetic field. When an RF pulse is applied, the proton spins change, altering the magnetization. The time it takes for the spin to regain its alignment with the external magnetic field after the RF pulse is turned off depends on the molecule (size and structure) and its surroundings. By altering the sequence of the RF pulses and the gradient of the magnetic field, the location and type of tissue being imaged can be controlled. The signal received in an MR image reflects intrinsic factors of the tissue, either spin density or relaxation properties of the nuclei. Spin-lattice relaxation time (T1) is the time it takes a tissue to regain longitudinal magnetization after a 90° RF pulse makes the spins perpendicular to the external magnetic field. T1 is a measure of energy transfer to the surroundings (lattice) as the proton recovers its normal spin. T1 relaxation times generally are between 300-2000 msec. Spin-spin relaxation time (T2) is a measure of how long the proton spins remain in phase after an RF pulse. Interaction with other molecules (e.g. diffusion) affects the T2 relaxation time. As natural motion of the proton increases, such as in liquids, T2 increases. Water, therefore, has a long T2, and appears white in T2-weighted images. T2 relaxation times are shorter than T1 and can range from 30-150 msec. Inhomogeneities in the magnetic field can also affect T2. A static internal field (caused by large, slow- moving proteins or rigid trabeculae for example), may additionally alter the local magnetic environment and affect T2. T2* combines the effects of molecular interactions (T2) and these field inhomogeneities. In addition to relaxation times, more complicated measures may also be obtained from the MRI signal, such as phase analysis, relaxation time distribution, and chemical composition. MR images also can reflect the behavior of water or fat alone. Figure 2 shows a radiograph of a proximal femur and a comparative fat suppressed MR image.
The MR image clearly depicts the presence of a fracture. Bone tissue has low water content, extremely short T2 and thus relatively low MR signal, and therefore appears black in most MR images. The bone marrow in trabecular bone, however, has sufficient water and fat content to provide MR signal. The trabecular bone network may alter the properties of the marrow by creating magnetic inhomogeneities at the bone-marrow interface. Trabecular structure can be imaged by relaxometry, which measures the change in marrow properties due to trabecular structure, or by direct visualization of the black trabecular network. The effect of the trabecular network on marrow magnetic properties is prominent in T2* images. The inhomogeneities at the bone-marrow interface are dependent on the density of the trabecular structure, the size of the trabeculae and trabecular spaces, and the field strength. In general a denser network results in shorter T2* relaxation times due to more bone-marrow interfaces and increased inhomogeneities. The sequence and timing of RF pulses determines the image contrast. Common sequences in bone imaging include the spin-echo and gradient-echo sequences. An “echo” reverses the spin, which refocuses the magnetization and in effect cancels out external magnetic field inhomogeneities, which are intrinsic in the magnet of the scanner. In a spin-echo sequence a 90° pulse is followed by a 180° RF pulse, which produces the echo. In gradient echo sequences, the magnetic field is reversed to create the echo. The echo time (TE) is the time between the original RF pulse and the peak echo signal. The type of sequence affects the appearance of the trabecular structure. In both spin and gradient echo sequences the dimensions of the trabeculae may be amplified due to differences in magnetic susceptibility (the amount which a material becomes magnetized in a magnetic field) between the marrow and bone. The amount of distortion artifact is dependent on TE with longer TEs resulting in more distortion. In addition, gradient-echo sequences produce more susceptibility artifacts than spin-echo sequences. Representative images of the distal radius are shown in Figure 3. Spin-echo sequences, however, require a considerably longer scan time and require in-vitro samples or smaller fields of view (such as the finger and wrist) because of signal-to-noise and total imaging time considerations. Therefore, in vivo imaging of trabecular bone typically is performed using gradient-echo sequences with TEs as short as possible. Alternatively a fast large angle spin echo (FLASE) sequence can be used which uses an initial RF pulse greater than 90°. The following 180° pulse then partially restores the longitudinal magnetization and reduces the time to repeat (TR), making the spin-echo faster.
Figure 2 – (A) Radiograph and (B) fat suppressed MR image illustrating proximal femur fracture.
The typical maximum resolution of a 1.5T scanner is 78-200^m in-plane and 400-1000 ^m out-of-plane (slice thickness). Trabeculae are the same dimensions as the in-plane resolution, resulting in partial volume effects, in which the depiction of a trabecula in the image is a projection or average of multiple trabeculae. As a result the trabecular measures obtained from MRI are different than those obtained with histo- morphometry or microCT at higher resolutions (20 ^m). The magnetic field strength of the scanner affects the resolution and acquisition time of the scan. A 1.5T magnet is the standard scanner used clinically and can provide a maximum resolution of approximately 150x150x250 ^m. With high- resolution MRI requiring a stronger magnetic field strength (79.4 T) and a small-bore (limited to in vitro scans), resolutions can be improved to 50x50x100 ^m. Nuclear magnetic resonance imaging has even a smaller field of view (2-12 mm) but can obtain isotropic resolutions as high as 10 ^m. NMR imaging can additionally determine chemical shift making it possible to establish distribution of a given chemical. Generally, higher magnetic field strength improves signal-to- noise ratio, scan time, and image quality, but often with limited field of view and other factors such as tissue susceptibility to consider.
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Figure 3 – Axial images of the calcaneus using (A) spin-echo and (B) gradient echo sequences.