All images were acquired using a whole-body Philips 3T Achieva scanner (Philips Medical System, Best, The Netherlands) equipped with 80 mT/m gradients. RF was transmitted using the body coil and SENSE reception (31 (link)) was employed. A series of consecutive direct saturation and CEST scans were performed using the 8-element knee coil for both the glycogen phantom and in vivo human calf muscle. To minimize leg motion, foam padding was placed between the subject’s lower leg and the knee coil. In all cases, second order shims over the entire muscle on the imaging slice were optimized to minimize B0 field inhomogeneity. Notice that the width of Z-spectra depends on T2 and that the WASSR procedure provides an absolute field frequency map so that there is no need for higher order shimming for the CEST acquisition. Clinical imagers generally employ a prescan to center the bulk water signal of the object/subject, optimize the flip angle and shim the field. Note that no such “prescan” should be made between direct saturation and CEST scans to maintain the same field reference conditions. For both scans, saturation was accomplished using a rectangular RF pulse before the turbo spin echo (TSE) image acquisition, as previously described by Jones et al. (21 (link)).
The power level needed for each saturation experiment depended on the load and was optimized by measuring sets of Z-spectra under these different conditions. For WASSR, the power and pulse lengths were chosen as small as possible to have sufficient direct saturation, while minimizing any MT effects. For CEST, the maximum pulse length allowed for the body coil within the protected clinical software (500 ms) was used and the power was optimized for maximum effect at the phantom and muscle loads. WASSR was obtained at higher frequency resolution than CEST, but over a smaller frequency range as only the direct saturation region needs to be covered. The WASSR range was chosen sufficiently large to validate the simulated results, consequently leading to a larger number of frequencies needed in vivo than for the phantom.
Single-slice glycogen phantom imaging was performed using SENSE factor = 2, TSE factor [number of refocusing pulses] = 34 (two-shots TSE), TR = 3000 ms, TE = 11 ms, matrix = 128 × 122, FOV = 100 × 100 mm2, slice thickness = 5 mm, NSA = 1. Imaging parameters for human calf muscle experiments were identical to those in phantom experiments except for the following: FOV = 160 × 160 mm2. The saturation spectral parameters for WASSR and CEST are indicated inTable 1 .
The power level needed for each saturation experiment depended on the load and was optimized by measuring sets of Z-spectra under these different conditions. For WASSR, the power and pulse lengths were chosen as small as possible to have sufficient direct saturation, while minimizing any MT effects. For CEST, the maximum pulse length allowed for the body coil within the protected clinical software (500 ms) was used and the power was optimized for maximum effect at the phantom and muscle loads. WASSR was obtained at higher frequency resolution than CEST, but over a smaller frequency range as only the direct saturation region needs to be covered. The WASSR range was chosen sufficiently large to validate the simulated results, consequently leading to a larger number of frequencies needed in vivo than for the phantom.
Single-slice glycogen phantom imaging was performed using SENSE factor = 2, TSE factor [number of refocusing pulses] = 34 (two-shots TSE), TR = 3000 ms, TE = 11 ms, matrix = 128 × 122, FOV = 100 × 100 mm2, slice thickness = 5 mm, NSA = 1. Imaging parameters for human calf muscle experiments were identical to those in phantom experiments except for the following: FOV = 160 × 160 mm2. The saturation spectral parameters for WASSR and CEST are indicated in
