There is an increased desire for miniature ultrasound probes with small apertures to provide volumetric images at high frame rates for in-body applications. Satisfying these increased requirements makes simultaneous achievement of a good lateral resolution a challenge. As micro-beamforming is often employed to reduce data rate and cable count to acceptable levels, receive processing methods that try to improve spatial resolution will have to compensate the introduced reduction in focusing. Existing beamformers do not realize sufficient improvement and/or have a computational cost that prohibits their use. Here we propose the use of adaptive beamforming by deep learning (ABLE) in combination with training targets generated by a large aperture array, which inherently has better lateral resolution. In addition, we modify ABLE to extend its receptive field across multiple voxels. We illustrate that this method improves lateral resolution both quantitatively and qualitatively, such that image quality is improved compared with that achieved by existing delay-and-sum, coherence factor, filtered-delay-multiplication-and-sum and Eigen-based minimum variance beamformers. We found that only in silica data are required to train the network, making the method easily implementable in practice.

An ultrasound scan generates a huge amount of data. To form an image this data has to be transferred to the imaging system. This is an issue for applications where the data transfer capacity is limited such as hand-held systems, wireless probes and miniaturized array probes. Two-stage beamforming methods can be used to significantly reduce the data transfer requirements. In the first stage, which is applied in-probe, the amount of data is reduced from channel to scanline data. In the imaging system the data is then beamformed to obtain images that are synthetically focused over the entire image. Currently two approaches exist for the second stage. The first approach is a time-of-flight approach called synthetic aperture sequential beamforming (SASB) that has been developed for both linear and phased arrays. SASB does however introduce artefacts in the image that can be reduced by tapering the first stage scan lines at the cost of lateral resolution. The second approach is based on the wave equation, but a computationally efficient method for phased arrays that is producing sector scan data is lacking. Here we propose an algorithm that uses the fast Hankel transform to obtain a fast algorithm. The imaging performance of this method is evaluated with simulations and experiments. Compared with PSASB, which is an adaption of SASB for phased arrays, our method requires a similar amount of operations to construct the entire image and there is no trade-off between resolution and artefacts. These results show the advantage of using the wave equation instead of a time-of-flight approach.

This paper presents a front-end application-specified integrated circuit (ASIC) integrated with a 2-D PZT matrix transducer that enables in-probe digitization with acceptable power dissipation for the next-generation endoscopic and catheter-based 3-D ultrasound imaging systems. To achieve power-efficient massively parallel analog-to-digital conversion (ADC) in a 2-D array, a 10-bit 30 MS/s beamforming ADC that merges the subarray beamforming and digitization functions in the charge domain is proposed. It eliminates the need for costly intermediate buffers, thus significantly reducing both power consumption and silicon area. Self-calibrated charge references are implemented in each subarray to further optimize the system-level power efficiency. High-speed datalinks are employed in combination with the subarray beamforming scheme to realize a 36-fold channel-count reduction and an aggregate output data rate of 6 Gb/s for a prototype receive array of 24 x 6 elements. The ASIC achieves a record power efficiency of 0.91 mW/element during receive. Its functionality has been demonstrated in both electrical and acoustic imaging experiments.

Ultrasound front-end receive designs for miniature, wireless, and/or matrix transducers can be simplified considerably by direct-element summation in receive. In this paper we develop a dual-stage beamforming technique that is able to produce a high-quality image from scanlines that are produced with focused transmit, and simple summation in receive (no delays). We call this non-delayed sequential beamforming (NDSB). In the first stage, low-resolution RF scanlines are formed by simple summation of element signals from a running sub-aperture. In the second stage, delay-and-sum beamforming is performed in which the delays are calculated considering the transmit focal points as virtual sources emitting spherical waves, and the sub-apertures as large unfocused receive elements. The NDSB method is validated with simulations in Field II. For experimental validation, RF channel data were acquired with a commercial research scanner using a 5 MHz linear array, and were subsequently processed offline. For NDSB, good average lateral resolution (0.99 mm) and low grating lobe levels (<-40 dB) were achieved by choosing the transmit as 0.75 and the transmit focus at 15 mm. NDSB was compared with conventional dynamic receive focusing (DRF) and synthetic aperture sequential beamforming (SASB) with their own respective optimal settings. The full width at half maximum of the NDSB point spread function was on average 20% smaller than that of DRF except for at depths <30 mm and 10% larger than SASB considering all the depths. NDSB showed only a minor degradation in contrast-to-noise ratio and contrast ratio compared to DRF and SASB when measured on an anechoic cyst embedded in a tissue-mimicking phantom. In conclusion, using simple receive electronics front-end, NDSB can attain an image quality better than DRF and slightly inferior to SASB.

This paper presents a power-and area-efficient front-end application-specific integrated circuit (ASIC) that is directly integrated with an array of 32 × 32 piezoelectric transducer elements to enable next-generation miniature ultrasound probes for real-time 3-D transesophageal echocardiography. The 6.1 × 6.1 mm² ASIC, implemented in a low-voltage 0.18-μm CMOS process, effectively reduces the number of receive (RX) cables required in the probe's narrow shaft by ninefold with the aid of 96 delay-and-sum beamformers, each of which locally combines the signals received by a sub-array of 3 × 3 elements. These beamformers are based on pipeline-operated analog sample-and-hold stages and employ a mismatch-scrambling technique to prevent the ripple signal associated with the mismatch between these stages from limiting the dynamic range. In addition, an ultralow-power low-noise amplifier architecture is proposed to increase the power efficiency of the RX circuitry. The ASIC has a compact element matched layout and consumes only 0.27 mW/channel while receiving, which is lower than the state-of-the-art circuit. Its functionality has been successfully demonstrated in 3-D imaging experiments.

Microbubbles (MBs) have been shown to create transient or lethal pores in cell membranes under the influence of ultrasound, known as ultrasound-mediated sonoporation. Several studies have reported enhanced drug delivery or local cell death induced by MBs that are either targeted to a specific biomarker (targeted microbubbles, tMBs) or that are not targeted (non-targeted microbubbles, ntMBs). However, both the exact mechanism and the optimal acoustic settings for sonoporation are still unknown. In this study we used real-time uptake patterns of propidium iodide, a fluorescent cell impermeable model drug, as a measure for sonoporation. Combined with high-speed optical recordings of MB displacement and ultra-high-speed recordings of MB oscillation, we aimed to identify differences in MB behavior responsible for either viable sonoporation or cell death. We compared ntMBs and tMBs with identical shell compositions exposed to long acoustic pulses (500–50,000 cycles) at various pressures (150–500 kPa). Propidium iodide uptake highly correlated with cell viability; when the fluorescence intensity still increased 120 s after opening of the pore, this resulted in cell death. Higher acoustic pressures and longer cycles resulted in more displacing MBs and enhanced sonoporation. Non-displacing MBs were found to be the main contributor to cell death, while displacement of tMBs enhanced reversible sonoporation and preserved cell viability. Consequently, each therapeutic application requires different settings: non-displacing ntMBs or tMBs are advantageous for therapies requiring cell death, especially at 500 kPa and 50,000 cycles, whereas short acoustic pulses causing limited displacement should be used for drug delivery.