## 1 Introduction  
Over the last decades, volumetric cardiac ultrasound imaging has gained momentum as the modality of choice to assess cardiac morphology and visualize global heart motion [1]. Recently, 3D cardiac ultrasound has also been used to quantitatively assess regional cardiac dynamics and several commercial products for 3D speckle tracking are now readily available [2]. One of the challenges for 3D motion estimation remains the relatively low image quality of the volumetric ultrasound data set. Particularly, its relatively low spatiotemporal resolution is of concern. Indeed, state-of-the-art commercial systems make use of a combination of several techniques in order to improve frame rate while maintaining image quality, such as limiting the field of view (FOV) [3], decreasing the line density [4], using electrocardiography (ECG) gating [5], and/or applying parallel receive beamforming (i.e., multi-line acquisition, MLA) [6]. This typically results in a frame rate of up to \(\sim\)30 Hz when scanning with a moderate line density, when using a representative wide-angle field of view (i.e., \(-60^{\circ}\times\sim 60^{\circ}\)) and when gating is performed over four to six cardiac cycles. ECG gating not only lengthens the acquisition but can also induce artefacts due to arrhythmias, breathing, motion or incorrect gating. A frame rate of 30 Hz for a \(60^{\circ}\times 60^{\circ}\) opening angle is adequate to evaluate cardiac morphology and visualize global heart motion [1]. However, when trying to visualize the mechanical activation wave associated with the contraction of the heart, a frame rate around 500 Hz is needed [4], while assessing the mechanical properties of the heart by measuring the speed by which (acoustically induced) shear waves propagate through the myocardium requires even higher frame rates [7]. Based on the desired cardiac application, the temporal resolution of the imaging system should thus be adapted. Recently, two other solutions have been proposed to increase the temporal resolution of the 3D systems. On the one hand, Diverging Wave Imaging (DWI) has been introduced, which uses a sparse virtual array located behind the probe [8]. This allows volumetric imaging at very high rates since the line density is only restricted by the reconstruction time. However, the use of unfocused beams implies that the energy of the beam is spread over a wider area resulting in low pressures which do not allow for harmonic imaging. In addition, the overall signal-to-noise ratio drops, having a negative impact on motion estimators (as illustrated by the Cramer-Rao Lower Bound [9]). On the other hand, multi-line transmit beamforming (MLT) has been proposed to increase temporal resolution by simultaneously transmitting multiple focused beams allowing harmonic imaging [10]. While its frame rate gain is more limited than that of diverging wave imaging [11] and although potential artifacts such as cross-talk might be introduced due to the interaction between neighboring MLT beams, it preserves image quality [10] and is an attractive imaging approach. Indeed, it has recently been demonstrated that cross-talk artifacts can be suppressed by using proper apodization [11], by avoiding the main directions of the transducer [6], by frequency-coding the different MLT beams [12; 13], and/or by physically separating the beams in space [9]. Furthermore, MLT can easily be combined with MLA to further increase the temporal resolution. More precisely, the width of the transmit beams can be customized based on the desired amount of receive parallelization for each transmit beam. Based on Expression (10) of [14], at least 4 MLA (two in azimuth and two in elevation) can be reconstructed per focused beam in 3D for a typical cardiac 2D matrix array transducer. In order to allow a higher number of MLA, broader beams can be transmitted by reducing the transmit aperture. It should also be noted that safety issues due to the acoustic superposition of the MLT beams in the near field are of little concern.