Echocardiography allows a comprehensive analysis of the heart in an objective and feasible manner, and it is the modality of choice for routine diagnostics given its low cost and high accessibility.[1] It has become the working horse of cardiology, with widespread use and a pivotal role in routine practice. Various echocardiographic parameters based on measurements of heart geometry as well as of blood flow or tissue velocity are regularly used for the clinical assessment of the heart, but none allows a comprehensive description of all aspects of cardiac function. In particular, the relation between myocardial wall stress and deformation is difficult to assess, as forces and pressures are not accessible by noninvasive imaging.
Recent advances in ultrasound instrumentation and computer technology enabled the practical implementation of high-frame-rate cardiac ultrasound imaging[2] with frame rates up to 5000 Hz, which is close to one hundred times the frame rate of conventional echocardiography. The acquisition of such high-frame-rate data introduces a new era, as it offers the possibility to visualize fast myocardial tissue movements that may reveal information on the relationship of stress and strain and allows a better quantification of myocardial deformation as well as complex cardiac blood flow.[3] This review will give insights into this developing technology and will provide an overview of potential clinical applications.
Two-dimensional (2D) echocardiography still represents the imaging modality of choice for the noninvasive evaluation of myocardial function. Novel quantitative techniques in echocardiographic deformation imaging, such as tissue doppler imaging (TDI) and speckle-tracking echocardiography (STE), have been shown to be more sensitive in detecting subtle changes in myocardial deformation than conventional echocardiography.[3,4] Most of the current echocardiographic imaging systems, however, operate at approximately 60 frames per second for 2D imaging and even considerably less for 3D,[5] which allows a good visualization of heart motion and morphology and the assessment of certain functional aspects during the cardiac cycle. However, these frame rates are insufficient to assess rapid and short-lived myocardial events, for example, during isovolumetric contraction/relaxation periods.[2] Therefore, while some aspects of regional myocardial deformation can be analyzed to some extent with the conventional approach, higher temporal resolution could provide additional information on pathophysiological mechanisms.[6–8]
In conventional imaging, focused ultrasound beams are sent sequentially from different directions into the tissue, and their echoes are used to reconstruct a respective line of the image sector (Figure 1A). Given a velocity of sound in soft tissue of about 1,530 m/s, and the typical propagation distance of 15 cm (round-trip distance 30 cm), one single ultrasound pulse takes approximately 200 µs before it returns to the transducer, so that 5000 pulses can be sent and received per second. Thus, an image composed of 180 lines requires 36 ms, which allows a frame rate of about 28 frames per second. Because the number of pulses per second is limited by the image depth, frame rate can only be increased if the density of scan lines or the image depth are reduced.[2] This trade-off between frame rate, depth, and number of lines often presents a challenge in the use of 2D and, in particular, 3D echocardiographic imaging, since both good temporal and spatial resolution are desired.
Conventional versus high-frame-rate ultrasound. (A) Traditional echocardiography makes use of focused transmit beams. (A–C) High frame rate imaging either transmits several focused beams in parallel (two in the example in (B)); defocused or diverging waves (C). (From Voigt14 with permission.)
A commonly used technique to speed up the acquisition process without compromising the density of scan lines or the image depth is multiline acquisition (MLA).[9–11] In this approach, the sent beams are less focused and allow the reconstruction of two or more scan lines simultaneously through differential focusing on receive (Figure 1B). In this way, the currently typical rate of 60 frames per second of modern ultrasound systems is reached.[12] Furthermore, by sending a “plane wave” or a completely unfocused beam (“divergent wave”), it is possible to reconstruct an entire image from a single ultrasound pulse, thus reaching 5000 frames per second[8,13,14] (Figure 1C). Because the transmitted sound energy per image is much lower, this technique comes at the cost of lower signal-to-noise ratio, but several such single-pulse, diverging wave images can be summed up to obtain images at high frame rate in reasonable quality.[12,15–17]
Dedicated research systems and clinical ultrasound machines adapted for research purposes already allow the clinical application of such high-frame-rate ultrasound imaging in cardiology, so that blood and tissue motion, as well as tissue properties, are already being explored using this new technology.
High-frame-rate imaging enables a new method for estimating the elastic properties of the myocardium (i.e., stiffness): shear wave imaging (SWI)[18–21] can measure and visualize certain vibrations in the myocardium (shear waves, SW) that propagate at a velocity that is directly related to its stiffness. In a large, isotropic, elastic medium, the velocity of shear waves (
Although these assumptions do not completely hold for myocardium (which is neither large nor isotropic), this approach allows for the first time a noninvasive, destruction-free, direct estimation of tissue stiffness
Shear wave imaging with intrinsic excitation. (A) Physiologic events, such as aortic or mitral valve closure can generate a shear wave that propagates along the myocardium. The propagation of the wave is then measured by high frame rate imaging. (B) Anatomic M-mode, derived from a straight line drawn along the anteroseptal wall (dotted line) from the base towards the apex in a parasternal long axis echo image. (C) The colors in the M-mode display code the radial acceleration of the myocardium. Note the pronounced waves (light blue) propagating from the anteroseptal base towards the apex immediately after aortic valve closure (AVC) with a velocity of 3.6 m/s. AVC = aortic valve closure; MVC = mitral valve closure; LV = left ventricle. (Adapted from Santos et al.[34] with permission.)
So far, researchers have mainly investigated natural SWs in parasternal, long axis images of the interventricular septum because of the advantageous anatomical position of this cardiac structure.[13,34] Note that 2D imaging of SWs measures only the projection of the SW velocity vector on the image plane. A recent study on three healthy volunteers (HV) demonstrated the feasibility of estimating myocardial SW velocities in 3D along different directions.[35] An example of a mechanical wave in the heart visualized with fast 3D imaging technology is given in Figure 3.[36]
3D visualization of mechanical wave propagation along the left ventricle at atrial systole. The normalized tissue acceleration is mapped in 3D over time. The white arrow indicates the maximum acceleration. (From Salles et al.[40] with permission.)
Shear wave imaging with external excitation. (A) External excitation by a strong focused ultrasound impulse induces a translational wave (shear wave) that propagates along the myocardium and is then measured by high frame rate imaging. (B) The shear wave propagation can be visualized on B-mode acquisition with high-frame-rate imaging; example of pushing locations shown by 3 red dots. (C) The propagation velocity of a shear wave is directly related to the stiffness of the tissue, which varies within a cardiac cycle depending on the pushing time. LV = left ventricle. (Adapted from Pernot et al.[38] with permission.)
Repeated application of ultrasonic push pulses in Langendorff-perfused, isolated rats with generation of artificial SWs through the entire cardiac cycle, showed dynamic stiffness variations during systole and diastole (Figure 4). Moreover, myocardial stiffness measured during systole was found to be strongly dependent on the inotropic state modulated with calcium or beta-adrenergic agonists, indicating that peak systolic myocardial stiffness has the potential to be used as an index of myocardial contractility.[30] The same author showed that end-diastolic stiffness measured by SWI using artificial SWs could also be used to determine the passive diastolic properties of the myocardium in an open-chest animal study.[43] Myocardial stiffness was also quantified in explanted porcine heart grafts (
Pernot et al. evaluated the potential of SWI using acoustic radiation force to quantify passive diastolic myocardial stiffness in an ovine model of ischemic cardiomyopathy and showed that SWI can differentiate between a stiff infarcted wall and a softer wall containing stunned myocardium.[43] Wouters et al. reported that the distinction of a scarred, nonviable septum and a thin, dysfunctional, remodeled septum in heart failure patients with left bundle branch block (LBBB) is possible using natural SWI after MVC.[48]
Myocardial stiffness in healthy volunteers of different ages and patients with HCM. Comparison of myocardial stiffness in healthy volunteers (HV, green bar) and heart failure patients with hypertrophic cardiomyopathy (HCM, red bar). Values denote the myocardial stiffness in kPa. Based on the ROC curve analysis, the optimal cut-off value of myocardial stiffness for detection of HCM-HFpEF was 8 kPa. (From Villemain et al.[44] with permission.)
In
A recent study also suggested a potential role for SWI in diagnosing graft dysfunction, one of the most common complications after heart transplant (HTx).[55] Histological changes accompanying allograft dysfunction include myocyte hypertrophy and diffuse myocardial fibrosis, which can in turn lead to diastolic dysfunction due to increased myocardial stiffness.[56] Clinical evaluation of graft dysfunction therefore aims at tissue characterization and assessment of diastolic function. Given that allograft dysfunction may lead to a stiffer myocardium, SWI might aid in the early detection of this condition. A study by Petrescu et al. showed that in 52 HTx recipients SW velocity at MVC correlated better with native myocardial T1 values (
Shear wave propagation velocities in patients after HTX. CMR T1 mapping of the mid left ventricular segment of the anteroseptal wall and acceleration M-mode map of the anteroseptal wall in three heart transplant recipients (HTx1-3). Note that shear wave velocities at mitral valve closure (MVC) increase with increasing native T1 value. The highlighted region of the ECG indicates the time interval covered by the M-mode map. HTx = heart transplant recipient; MVC = mitral valve closure. (From Petrescu et al.[32] with permission.)
The quantitative estimation of myocardial stiffness by SWI without the need for invasive procedures, cardiac magnetic resonance, or histology has been shown to be feasible for characterizing diffuse myocardial fibrosis, ischemic scar, and other pathologic conditions in addition to the noninvasive assessment of myocardial systolic and diastolic function. Further clinical validation is needed, however, before this promising technique can be applied in the clinical routine. Nevertheless, current preliminary results are promising and suggest that SWI will complement our echocardiographic toolbox in the coming years.
Global and segmental measurements of myocardial strain provide additional information on cardiac mechanics beyond conventional parameters such as ejection fraction and visual wall motion analysis. Strain can be obtained by tissue Doppler or speckle tracking echocardiography (STE). [39] While STE is more feasible in the clinical routine than tissue Doppler, it suffers from the low frame rate of the underlying grey scale images (40–80 fps). This results in under-sampling of fast and short-lived events, e.g., in the isovolumic phases of the cardiac cycle and the in accuracy of temporal derivatives of motion and strain, i.e., velocity and strain rate.[57,58] Recently, dedicated high-frame-rate STE algorithms have been developed in order to combine the high temporal resolution of ultrafast ultrasound imaging with the comfort of 2D speckle tracking algorithms.[57,59–61] Joos et al. applied STE on high-frame-rate B-mode images at 500 fps that were obtained by special coherent compounding methods based on motion compensation. A global longitudinal strain of the left ventricle was obtained from STE in 10 subjects and compared to the results provided by a clinical scanner. Interestingly, group means were not statistically different (
These results show that dedicated high-frame-rate STE algorithms allow speckle tracking at frame rates that so far only tissue Doppler could provide. Studies need to show how clinical decision-making can benefit from such temporally resolved myocardial deformation information.[2,59,60]
Coronary ultrafast Doppler angiography (CUDA) was developed as a specific application for imaging coronary arteries enabling the noninvasive imaging of the coronary vasculature and the assessment of coronary flow reserve. The technique involves acquiring ultrasound images of the beating heart at high frame rates and using specific spatiotemporal clutter filters[66] in order to distinguish the myocardium with high echogenicity and relatively slow motion from blood vessels with low echogenicity and faster motion.[67] Thus, CUDA is able to visualize the coronary vasculature down to 100 µm with high sensitivity and without the use of contrast agents. It can also quantify intramyocardial blood flow changes in a wide range of flow conditions (1 mm/s to 1 m/s). The first experiments in the beating heart of large animal models that showed that CUDA accurately quantifies the change of flow rate in coronary arterioles (with diameters ranging from 500 to 100 µm) during hyperemia were performed by Maresca et al.[66] The authors compared the measurements to the gold standard by placing a proximal flow meter probe on the epicardial artery. Quantification of the coronary flow changes during hyperemia was in good agreement with gold standard measurements (r2 = 0.89), as well as the assessment of coronary flow reserve (2.35 ± 0.65 vs. 2.28 ± 0.84;
2D ultrafast color Doppler imaging of intramural coronary vasculature. Flow in epicardial and intramyocardial vessels an open-chest swine experiment. Vessels below 100 µm remain below resolution. The scale bar represents 3 mm. A and B are longitudinal images, C and D are short axis images. Panels A and C: in a systole, venous blood flow moves upwards from the endocardium to the epicardium (red) and is collected in the epicardial veins. Panels B and D: in diastole, arterial blood from the epicardial vessels flows downwards (blue) in the myocardium. (From Maresca et al.[68] with permission.)
3D ultrafast Doppler coronary angiography. (1) Coronary blood flow at baseline (left) and during reactive hyperemia (right) after transient occlusion of the LAD artery. (2) Flow velocity mapping using 3D ultrafast Doppler coronary angiography at early diastole. (From Correia et al.[70] with permission.)
CUDA provides the detection of coronary flow in epicardial and intramural arteries with characterization of coronary flow reserve and intramural perfusion without the use of contrast agents or exposure to ionizing radiation. This technique could be used, for example, in patients with microvascular angina, but it will need further clinical validation and study.
The recent development of high-frame-rate ultrasound made a new approach to blood flow assessment possible. High-frame-rate imaging-based blood speckle tracking (BST) is a novel blood flow visualization that relies on tracking native blood on high-frame-rate images without the need for contrast agent injections. It can be used for the evaluation of blood flow energetics by describing the blood vortex inside the ventricles[65,72,73] or to visualize and quantify flow in congenital defects, valve disease, or pulmonary hypertension (Figure 9). Vortex visualization in fetal cardiac ventricles could also contribute to a better evaluation of cardiac function before birth and might improve prenatal detection of coarctation of the aorta. Furthermore, BST allows the visualization of altered flow patterns in ventricles with reduced function and by using high-frame-rate flow data in the calculation of intraventricular pressure gradients.[74]
Cardiac blood speckle tracking imaging in a neonate with double-outlet right ventricle. The figure displays frames from the parasternal long-axis view at four different moments in the cardiac cycle. Ao, aorta; LV, left ventricle; RV, right ventricle. (From Nyrnes et al.[78] with permission.)
BST is so far the first commercially available application of high-frame-rate imaging in cardiology and has mainly been used to image cardiac blood flow patterns in neonates[72] and pediatric patients.[65] The feasibility of this method was also investigated in the adult heart (Figure 10),[36,70,75,76] but image depth is limited to approximately 8 cm (Nyrnes). An increase in depth would require a probe with a lower frequency, leading to reduced spatial resolution and, thus, providing insufficient tracking results.
3D blood speckle tracking imaging. The figure shows a 3D representation of the flow path lines in three parts of the cardiac cycle corresponding to diastole, diastasis, and systole by using blood speckle tracking. Complex swirling flow patterns can be displayed throughout the cardiac cycle. Ao = aorta; LA = left atrium; LV = left ventricle. (From Wigen et al.[77] with permission.)
Intracardiac blood flow analysis may contribute to a deeper understanding of complex flow physiology. However, further studies are needed to evaluate the potential of this novel technique for improving diagnosis, prognosis, and treatment of cardiac disease.
SWI can determine the fiber direction in skeletal muscle[77] and myocardium[78,79] as tissue stiffness and—with this—SW speed is higher in the direction of the fibers than across. Elastic tensor imaging (ETI)[78] has been proposed to quantify this anisotropy. SWs are induced by acoustic radiation force, and ultrafast ultrasound imaging is used to track the SW propagation in different directions and in different layers of the tissue (Figure 11).[71,80] This method has been tested
Myocardial fiber orientation assessed by Elastic Tensor Imaging. (A) The probe is placed at the surface of the myocardium and rotated around a single axis z, depth of the myocardial wall. (B) Shear waves were generated by using acoustic radiation force and were measured for different depths within the myocardium (in percentage) for each angle of rotation. (From Ngo et al.[82], modified after Lee et al.[73] with permission.)
The link between myocardial structure and cardiac function may have important potential clinical applications; however, it remains vastly unexplored because of technical challenges.