Читать книгу Magnetic Resonance Microscopy - Группа авторов - Страница 61

Shorter superconducting solenoid magnets have clear benefits for patient acceptance and ease of siting and use in an ED setting. In contrast to the situation for cryogenic equipment, the design of superconducting MRI magnet windings is little changed over the past few decades; see Lvovsky et al. for a thorough review [97]. In standard design optimizations, the position and number of turns of discrete winding groups are optimized on the cylinder. In addition to the primary field-producing windings, counter-wound shielding windings (usually larger-diameter coils at the bore end) attempt to reduce the stray field around the magnet. The optimization seeks a short magnet on a predetermined diameter cylindrical former that achieves the target homogeneity over the imaging region (typically defined as a given diameter spherical volume (DSV) with a minimum wire cost (length of wire). The optimization either assumes a small number of winding groups (e.g. six) [100] or uses linear programming and a sparsity-promoting norm to limit the winding groups [101]. The magnet homogeneity is a relative measure (ppm) and expressing the magnet length in unitless terms (as L/DSV, where DSV is the imaging “diameter spherical volume”) is also helpful. Xu et al. [102] showed that for a 1 ppm target homogeneity, the cost-optimized magnets followed a linear formula: L_optimal = 1.19 DSV + 0.77D, where L_optimal and D are the magnet windings length and diameter, and DSV is the diameter of the spherical homogeneity region. Reducing the magnet length further (<L_optimal) results in skyrocketing wire costs (which are proportional to conductor volume). A similar analysis was also applied to gradient coil lengths [103].