Pulse Sequence Overview
A magnetization-prepared 3D bSSFP sequence with segmented k-space acquisition is used to produce FIA angiograms. The sequence starts with a magnetization-preparation module that has optional inversion-recovery and T2-preparation sections. When necessary, a nonselective inversion pulse is applied to reduce the signal from long-T1 fluids such as synovial fluid or edema. For improving the T2-dependent (blood/muscle, arterial/venous) contrast, a segmented adiabatic B1-insensitive rotation (BIR-4) pulse was used for T2-preparation (20). The BIR-4 pulse offers immunity to main field and radio-frequency (RF) excitation field inhomogeneities. A diagram of the sequence showing all the available modules is displayed in Fig. 1.
Figure 1. Pulse sequence diagram. Optional inversion-recovery and T2-preparation sections form the magnetization-preparation module. The bSSFP acquisition starts immediately following a ramped series of RF excitations.
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A limited number of phase encodes can be acquired with the desired contrast because of the transient nature of magnetization preparation. To capture this contrast effectively, centric phase-encode ordering [a square spiral for 3DFT (21)] is employed and k-space is segmented into several interleaves (16). After a certain number of phase encodes are acquired, the magnetization is allowed to recover to equilibrium and the preparation is repeated prior to the acquisition of the next interleaf. A linear ramp catalyzation is used, following magnetization preparation, to dampen transient signal oscillations (22).
Because frequent repetition of magnetization preparation reduces the scan efficiency, multiple phase encodes are acquired after a single preparation. This leads to a transient acquisition, where a mixture of prepared and steady-state contrast is captured. Figure 2 shows the transient bSSFP signal simulated for different preparation schemes: no preparation, only inversion recovery, only T2-preparation, and both inversion recovery and T2-preparation. The following approximate relaxation parameters were chosen from literature (13, 23–25): T1/T2 = 1,000/220 msec for arterial blood, 1,000/120 msec for venous blood [assuming 70% oxygen saturation in peripheral venous blood (26)], 870/50 msec for muscle and 4,000/2,000 msec for synovial fluid. The sequence parameters were α = 60°, TR/TE = 4.6/2.3 msec, TI (inversion time) = 2 sec, T2-preparation time = 80 msec, and a 10-excitation linear ramp catalyzation. The flip angle and T2-preparation time were chosen to optimize the initial blood/muscle contrast while maintaining as low a specific absorption rate (SAR) as possible. The inversion time was adjusted to produce low signal from synovial fluid during a 4–5 sec acquisition window.
Figure 2. The transient bSSFP signal is shown immediately after various types of magnetization preparation. The inversion-recovery sequence reduces the synovial-fluid signal for a finite time window at the expense of reduced blood signal. T2-preparation improves the initial blood/muscle and arterial/venous contrast.
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PS-SSFP places the fat and water resonance peaks in the centers of adjacent passbands, which have a π-radian phase difference. To achieve this placement at 1.5 T, a TR of about 4.6 msec is needed. PS-SSFP provides uniform fat suppression, but it can underestimate the water signal within a voxel containing a mixture of fat and water. These partial volume effects can be mitigated with sufficient spatial resolution. However, the resultant artifacts can be significant when the achievable spatial resolution is limited by SNR and scan time considerations. Because of the large volumetric coverage requirements, the resolution needs to be lowered to maintain an adequate SNR for lower leg and foot angiograms. Hence, partial volume artifacts with PS-SSFP reduce the quality of these angiograms.
On the other hand, fat-suppressing ATR-SSFP uses two consecutive repetition times (TR1 and TR2) to create a stopband centered at the fat resonance (18). Because ATR-SSFP employs a different phase-cycling scheme than bSSFP, its response is equivalent to a frequency-shifted bSSFP response at on-resonance. This slightly reduces the tissue signal for higher T1/T2 ratios, and the magnitude of the signal change increases with T1/T2. Therefore, the ATR signal exhibits slightly more T2-dependent characteristics than the bSSFP signal, both in the steady state and throughout the progression from the initial prepared state to the steady state.
Arterial blood, venous blood, and muscle signals were simulated for the previously listed relaxation parameters assuming α = 60°, TR1/TR2/TE = 3.45/1.15/1.725 msec, T2-preparation time = 80 msec, and a 10-excitation linear ramp catalyzation. The ATR-SSFP sequence, with 0–90−180–270 phase-cycling, yields higher signal and improved T2-contrast at the start of the acquisition following T2-preparation. The initial arterial blood/muscle contrast is approximately 3.9 for ATR-SSFP and 3.3 for bSSFP. This difference in signal behavior can lead to improved background suppression with ATR-SSFP. On the other hand, the arterial/venous contrast following T2-preparation is roughly 1.6 for both ATR-SSFP and bSSFP.
Since ATR-SSFP reduces the fat signal at the time of data acquisition, the corresponding partial volume effects are reduced. It is important to note that the level of suppression in the ATR stopband is a function of off-resonance. At the edges of the stopband where large field inhomogeneity is experienced, the signal reaches up to ∼50% of the passband signal, yielding ineffective fat suppression. Methods with improved stopbands are needed when the fat signal cannot be sufficiently reduced through the whole imaging volume. Although ATR-SSFP achieves adequate fat suppression in the lower leg, high field inhomogeneity is observed in regions such as the foot because of their irregular shape. In this case, an improved fat suppression method that relies on two ATR-SSFP acquisitions can be used instead (19).
In regular ATR-SSFP, the phase of the RF excitation prior to the TR2-interval is selected to place the central null of the stopband at the fat-resonance (−220 Hz at 1.5 T). This selection divides the stopband into two equal-width segments, where one segment is in-phase and the other is out-of-phase with the on-resonant water signal. By decreasing the RF phase, we can shift the central null toward higher frequencies and extend the width of the in-phase segment. Alternatively, the out-of-phase segment can be extended by increasing the RF phase.
Fat suppression can be achieved by summing two separate acquisitions where the whole stopband is respectively in-phase and out-of-phase with the water signal (19). In one variation of this method, RF phases of 45° (0–45−180–225) and 135° (0–135−180–315) are used to create a stopband as wide as the regular ATR-SSFP stopband, assuming TR1/TR2 = 3. Because the magnitude profiles of the two acquisitions are very similar around the fat resonance, the resulting stopband suppression is significantly improved. In another variation, RF phases of 45° and −45° are used to generate in-phase and out-of-phase acquisitions for all frequencies in the spectrum except for the ones within the passband. This choice of RF phases increases the width of the stopband three times (for TR1/TR2 = 3).
Figure 3 compares the simulated magnetization profiles for the aforementioned fat suppression methods. The following parameters were assumed: α = 60°, TR/TE = 4.6/2.3 msec (bSSFP) and TR1/TR2/TE = 3.45/1.15/1.725 msec (ATR-SSFP and the double-acquisition fat suppression method), and T1/T2 = 260/80 msec for fat (25). PS-SSFP uses the bSSFP profile, which has high passband signal at the fat resonance (−220 Hz at 1.5 T). ATR-SSFP reduces this signal by creating a stopband; however, the amount of suppression is limited with field inhomogeneity. The ATR-based double-acquisition method can yield substantially reduced remnant signal in the stopband. Alternatively, the stopband can be widened to further improve the immunity to field inhomogeneity at the expense of a lower level of suppression and ∼11% reduction in the passband signal.
Figure 3. The transverse magnetization profiles were simulated for the a: bSSFP, b: ATR-SSFP sequences, and c: the ATR-based double-acquisition fat suppression method. ATR-SSFP creates a stopband around the fat resonance; however, the residual stopband signal with increasing field inhomogeneity limits the level of achievable suppression. On the other hand, the double-acquisition method robustly suppresses the fat signal. Furthermore, this method can be modified to widen the stopband and provide increased immunity to field inhomogeneity.
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Lower leg and foot angiograms were produced with the 3DFT bSSFP FIA sequence using a single-channel transmit-receive linear extremity coil (26 cm in length, 18 cm in diameter). The experiments were performed on a 1.5 T GE Signa Excite scanner with CV/i gradients. The fat signal was suppressed with PS-SSFP, ATR-SSFP, and the double-acquisition ATR method. The acquisition parameters were α = 60°, TR/TE = 4.6/2.3 msec for PS-SSFP, and TR1/TR2/TE = 3.45/1.15/1.7 msec for ATR-SSFP, 1 mm isotropic resolution, 25.6 cm × 12.8 cm × 12.8 cm FOV, ±125 kHz bandwidth, 256 × 128 × 128 encoding matrix, T2-preparation time = 80 msec, a 10-excitation catalyzation, 4 interleaves, at 19 sec data acquisition window for each interleaf, and a 4 sec recovery time. The acquisition time for a single dataset was 1 min 28 sec. The irregular shape of the foot led to high field inhomogeneity. Therefore, the double-acquisition method was modified to widen the stopband as previously mentioned. All datasets were zero-padded to twice the initial matrix size prior to a maximum-intensity projection (MIP) to improve the visualization of the vasculature.
This study was approved by our institutional review board. Four volunteers were scanned to produce FIA angiograms of the lower leg and foot. Written, informed consent was obtained from all subjects. To quantitatively compare the level of background suppression and the reliability of vessel contrast, the arterial/venous and arterial blood/muscle contrast as well as scan-time-normalized contrast-to-noise-ratio (CNR) were measured. In the lower leg, the measurements were performed in the popliteal, peroneal, and posterior tibial arteries. The arterial blood/muscle CNR was computed with the mean arterial signal. In the foot, the lateral plantar artery and dorsal metatarsal arteries were used to measure the arterial blood/muscle CNR. Homogeneous regions of arterial as well as neighboring venous and muscle signal were selected on the source images for the measurements. The noise was then estimated by computing the standard deviation in the muscle region. The same tissue regions were used when comparing different techniques. The measurements from all subjects were averaged.