Black-Blood Contrast in Cardiovascular MRI

Abstract

MRI is a versatile technique that offers many different options for tissue contrast, including suppressing the blood signal, so-called black-blood contrast. This contrast mechanism is extremely useful to visualize the vessel wall with high conspicuity or for characterization of tissue adjacent to the blood pool. In this review we cover the physics of black-blood contrast and different techniques to achieve blood suppression, from methods intrinsic to the imaging readout to magnetization preparation pulses that can be combined with arbitrary readouts, including flow-dependent and flow-independent techniques. We emphasize the technical challenges of black-blood contrast that can depend on flow and motion conditions, additional contrast weighting mechanisms (T₁ , T₂ , etc.), magnetic properties of the tissue, and spatial coverage. Finally, we describe specific implementations of black-blood contrast for different vascular beds. LEVEL OF EVIDENCE: 5 TECHNICAL EFFICACY STAGE: 5.

Keywords: black-blood contrast; blood signal suppression; vessel wall imaging.

FIGURE 1

Black‐blood contrast can be achieved with the spin echo technique due to through‐plane flow between the slice‐selective 90° excitation and 180° refocusing pulses (a). The slice profile shows through‐plane flow for three velocities (v) at the time t = TE/2: v = 0, v = z/TE, and v = 2z/TE, where z is the slice thickness and TE the echo time. In the first case, the tissue experience both 90° and 180° pulses, yielding the maximum signal at the echo time. In the second case, the tissue has partially left the slice and only half of the excited tissue experience the refocusing pulse, yielding intermediate signal intensity at the echo time. In the last case, the excited tissue has entirely left the slice before the refocusing pulse, resulting in no signal at the echo time. For fast spin echo techniques, where a train of refocusing pulses are applied for each excitation pulse, the blood suppression performance is given by the through‐plane flow at the effective echo time (TE_eff) which is typically around the middle of the echo train (b). ES = echo spacing.

FIGURE 2

The evolution of the 0th and 1st gradient moment (m₀(t) and m₁(t), respectively) during a spin echo pulse sequence between the excitation (RF₉₀) and echo at the end of the second gradient (G_x) (a). Notably, m₁(t) increase quadratically during the gradients and is nonzero at the echo time, unlike m₀(t), which increase linearly. The corresponding phase for two spins with different position (x1 and x2) and velocities (v1 and v2) (b). The position‐dependent phase terms are nulled for both spins at the echo time while the velocity‐dependent phase is different, proportional to the differences in velocities and lead to dephasing if the spins are within the same voxel at the echo time.

FIGURE 3

Dephasing of transverse magnetization due to velocity differences (turbulence) within a voxel. The left column illustrates three spins within a voxel for different flow conditions, no flow (top), constant velocity (second), mild turbulence (third), and strong turbulence (fourth), while there is an active bipolar gradient in the flow direction (x). The second column shows the resulting phase for the spins (individual spins in black, vector sum in gray) in the voxel where no or constant flow lead to the same phase for all spins, and subsequently maximum signal magnitude shown in the third column. Turbulent flow leads to a phase distribution across the voxel that results in a reduced magnitude, and complete signal suppression in the case of strong turbulence.

FIGURE 4

Dephasing of transverse magnetization during a fast spin echo sequence for spins moving through a vessel with constant velocity (v). If the actual refocusing pulses are lower than 180°, stimulated echo (STE) pathways will be created at each refocusing pulse. In this example, a portion of the M_xy magnetization at the second RF pulse are tipped back into the M_z direction and are not affected by the following readout gradient (G_x), unlike the portion of M_xy following the spin echo pathway that remains in the transverse plane. The third RF pulse acts as an excitation pulse on a portion of the STE pathway and because the phase is stored, the following G_x yields a stimulated echo coinciding with the second echo of the spin echo pathway (SE₂). However, differences in 1st gradient moment can lead to dephasing due to the motion‐induced phase difference between the SE₂ and STE pathways. The illustration of phase shows the evolution of phase during the application of the RF pulses and gradients, where dashed straight arrows show the starting phase, the dashed curved arrows the change in phase and direction, and the solid straight arrows the final phase at each timepoint for the two spins following the spin echo (green arrows) and STE (blue arrows) pathway, respectively. Note the amount of spins that follow either spin or stimulated echo pathways will depend on the effective flip angle of the refocusing pulses.

FIGURE 5

Sequence diagram for ECG‐triggered double‐inversion recovery in short axis view of the heart. A nonselective inversion pulse (NS INV) is applied after the R‐wave which inverts the longitudinal magnetization (M_z) globally. This is immediately followed by a selective inversion pulse (S INV) which reinverts the area of the imaging slice (with some margin in the slice direction). The inversion delay is timed such that the inverted blood reaches the M_z zero‐crossing during the acquisition of the center of k‐space (k₀), during which time it should also have replaced the reinverted blood that flows out of the imaging plane. This effectively nulls signal from blood flowing into the slice, while static tissue experiences both inversion and reinversion pulses and have full M_z.

FIGURE 6

Conventional bright‐blood late gadolinium enhancement (LGE) (top row) and blood signal suppressed LGE (bottom row) in three patients with subendocardial and papillary muscle scar that can be clearly visualized using the dark blood technique. Images courtesy of Mr. Robert Holtackers, Maastricht University, Maastricht, The Netherlands.

FIGURE 7

Images of pulmonary veins in a patient with congenital heart disease, where flow‐induced signal loss impede visualization using bright‐blood 3D bSSFP (a), while black‐blood 3D VFA‐FSE allow clearer depiction (red arrows) (b). Coronal view of superior caval vein (SCV), superior cavopulmonary anastomosis, and proximal branch pulmonary arteries in a 4‐year‐old female who underwent a superior cavopulmonary anastomosis (c,d). Three‐dimensional bSSFP demonstrates poor visualization of the SCV and proximal pulmonary arteries due to dephasing in the vessels (c), while 3D VFA‐FSE improves visualization of the SCV and branch pulmonary arteries (d). RPA = right pulmonary artery; LPA = left pulmonary artery.

FIGURE 8

Black‐blood images from motion‐corrected volume (left), acquired from a fetus with hypoplastic left heart syndrome with total anomalous venous at 32 weeks gestation, displayed in coronal (Cor), sagittal (Sag), and transverse (Tra) planes. The 3D volume of fetal thorax was reconstructed retrospectively from multiple orthogonal input stacks of 2D images acquired using 2D FSE. A 3D segmentation of the fetal heart and vascular anatomy generated from this volume is shown on the right (posterior projection). * = ascending vein; IV = innominate vein; C = pulmonary venous confluence; DAo = descending aorta; SCV = superior caval vein; LV = left ventricle; RV = right ventricle. Images courtesy of Dr. David Lloyd, King's College London, London, UK.

FIGURE 9

Forty‐seven‐year old male who presented with left‐sided transient ischemic attack. Pre‐ (a) and postcontrast (b) transverse source images of 3D VFA‐FSE show eccentric, enhancing atherosclerotic plaque in the carotid siphon (orange arrows). A second, smaller eccentric plaque further distally in the left intracranial carotid artery is seen in figure (d,e) highlighted by green arrows. In (c f) the corresponding postcontrast coronal reformations are shown. Images were acquired at 7.0T.

FIGURE 10

Sixty‐three‐year old male patient with bilateral carotid artery aneurysms. Maximum intensity projection of contrast‐enhanced MR angiography shows cervical and cranial vasculature in the coronal plane (a, center), and in double oblique projections of the right (a, left panel) and left (a, right panel) carotid arteries. White arrows denote the aneurysms. Contrast‐enhanced black‐blood 3D VFA‐FSE transverse source image (b,d) and coronal reformation (c,e) of the right and left carotid arteries show the enlarged vascular lumen and thin, heterogeneously enhancing vessel wall (b,d). Note excellent black‐blood contrast despite the presence of contrast agent.

FIGURE 11

Susceptibility‐weighted imaging (SWI) from four subjects (a,b), (c,d), (e), and (f,g). To characterize vessel wall components of the carotid arteries, the magnitude images (a,c) are used for anatomical orientation and indicate areas of altered susceptibility at the vessel walls (yellow and green arrows). The phase image of the first patient shows a positive susceptibility that is indicative of intraplaque hemorrhage (b, yellow arrow), while the second patient has a negative susceptibility that suggests calcification (d, green arrow). Ferumoxytol‐enhanced brain SWI enable angiography and venography in a single scan (e). Before ferumoxytol injection SWI images yield black‐blood contrast for deoxygenated venous blood that has a relatively high susceptibility compared to surrounding tissue (red arrow) unlike arterial blood (light blue arrow) (f). Administration of ferumoxytol increases susceptibility in both venous and arterial blood that enables black‐blood venograms and angiograms (g). Images courtesy of Drs. Chaoyue Wang and Qi Yang, Capital Medical University, Beijing, China, and Dr. Mark Haacke, Wayne State University, Detroit, Michigan, USA.