# Implementation of Constrained Swept Synthetic Aperture Using a Mechanical Fixture
## 1 Introduction  
Ultrasound imaging is widely used for both screening and diagnosis in soft tissues due to advantages in cost, safety and portability over competing imaging technologies [1; 2; 3], particularly with the advent of portable devices [4]. Image contrast is based on differences in the acoustic impedance of tissues, providing good visualization of cysts, tumors, vessels and similar structures. For example, ultrasound has long been the primary screening tool for hepatocellular carcinoma (HCC). However, target detectability in images is based on the combination of contrast and resolution [5]--the latter is a particular weakness of ultrasound imaging. In HCC screening, a recent meta-analysis found that only 45% of early-stage HCCs were detected by ultrasound alone [6], bringing into question its utility without added blood tests or other imaging modalities. While axial resolution is related to the length of the transmitted imaging pulse (often approximately a wavelength, \(\lambda\)), lateral resolution is based on the wavelength and the f-number, \(F=z/D\) for depth \(z\) and imaging aperture size \(D\). This leads to an anisotropic resolution that varies drastically through depth. In low-contrast lesions, a loss of resolution blurs lesion boundaries and fills off-axis information from surrounding tissues into the hypoechoic spaces. In abdominal or obstetric imaging, where imaging depths can exceed 10-20 cm, particularly in obese patients, the lateral resolution can be worsened to several millimeters. This makes the characterization of very small suspected HCC lesions (\(<\)1 cm) impossible until further growth has occurred [7], thus, lessening the advantages of early detection. Current LI-RADS protocols suggest the alternatives of contrast-enhanced ultrasound or CT/MRI for the diagnosis and staging of HCC [8; 9] with standard B-mode imaging unable to provide clear tumor size measurements. While the imaging frequency (and wavelength) of ultrasound is limited by the required penetration, with higher attenuation at higher frequencies, the aperture size is fundamentally only limited by the available acoustic window. A standard curvilinear imaging array used in abdominal applications has a total aperture size around 5-6 cm, although these elements point radially outward to increase field of view and result in an even smaller coherent active aperture size. My colleagues and I have recently demonstrated the value of increasing the aperture size to improve the lateral resolution and detectability at depth in simulation [10], ex vivo experiments [11] and in vivo imaging [12] even in the presence of image-degrading clutter effects. Multi-transducer experiments have demonstrated improvements in resolution that scale with the effective aperture extent [13; 14; 15] and provided enhanced visualization of in vivo liver and kidney structures. Despite the promise of large-aperture imaging strategies, there are practical barriers to the implementation of such systems. The cost, weight and electronics complexity of arrays scales with the number of array elements required, particularly when considering 2-D array systems [16]. Point-of-care systems are unlikely to implement large apertures as they are designed to be low cost and highly portable, often imposing data bandwidth constraints on the system as well. Colleagues and I have previously introduced a reduced complexity method for large aperture imaging termed \"swept synthetic aperture\" (SSA) imaging [17; 12] that may be more widely applicable, using the motion of the ultrasound array to form an effective coherent array that is larger than the physical footprint of the device. This is similar to a monostatic imaging approach [18] with the advantages of improved signal-to-noise ratio and signal redundancy provided by an imaging array.