## 1 Introduction  
The number of cardiovascular disease (CVD) deaths has increased by 41% from 1990 to 2013 [1]. A better understanding of CVD pathogenesis, aetiology, and haemodynamic behaviours may help to improve cardiovascular interventions and therapeutic procedures. Various in vivo measurements, in vitro measurements, and numerical modelling have provided a better insight into the vascular system and haemodynamics [2; 3; 4; 5]. However, in vivo measurements may not provide high precision. For instance, transcranial Doppler sonography is a common non-invasive measurement method but it has limited spatial resolution [6]. In addition, the induced signal from slow-moving organs is not detectable in ultrasound techniques [7]. Computational fluid dynamics (CFD) have been extensively used for haemodynamic modelling [8; 9; 10]. However, most of these studies need experimental validation to be reliable for clinical use and can lack complex fluid-wall interaction. Therefore, in vitro modelling can be used to investigate the haemodynamics and optimise catheter-based vascular interventions. Particle image velocimetry (PIV) of forced flow in silicone arterial phantoms has been frequently used for in vitro modelling of haemodynamics. Several experimental modelling investigations have been performed in rigid silicone phantoms [11; 12; 13; 14]. However, some studies that utilised fluid-structure interaction (FSI) reported that rigid wall assumptions can lead to overestimation of wall shear stress (WSS) in comparison with a compliant phantom [15; 16; 17; 18]. In addition, wall stiffness is a measure of arterial health, and the pulse waveform differs in stiff and compliant arteries [19]. Failing to mimic arterial compliance reduces the precision of the model [20]. Compliance is the ability of the artery to dilate and recoil due to the changes in transmural pressure. Compliance is an important factor in pulsatile flow haemodynamics. Hence, matching compliance across the in vitro phantom and in vivo artery is important to obtain meaningful haemodynamics during experimentation. Since wall thickness is inversely proportional to compliance, the fabrication of the phantom must yield accurate and precise wall thickness to enable successful experimental results. Although the wall thickness varies in different arterial regions, significant thickness variation in compliant phantoms can cause unphysiological compliance distribution and unrealistic flow fields [21]. Such problems reduce the clinical value of experimental outcomes. Different manufacturing methods have been used for the construction of compliant phantoms [22; 23; 24; 25; 26]. Dip coating of a rapid prototyped mould has been a common casting method for compliant phantom construction [27; 28]. In this method, the mould is dipped and coated several times in a silicone bath until the desired thickness is achieved. However, it is difficult to ensure consistent wall thickness. [PERSON] et al. [29] fabricated a compliant model of the aortic arch using dip coating. A wall thickness of 6.35 mm (1/4\") was developed. This is considerably larger than the wall thickness required for in vivo compliance matching [24; 30]. [PERSON] and [PERSON] [31] constructed different compliant phantoms using dip spin coating. The 3D printed mould was spun on a two-axis mechanism subsequent to each coating step. They determined that horizontal dipping produced more uniform thickness compared with vertical dipping. Their results showed 0.60 mm and 0.58 mm of vertical and horizontal thickness bias respectively, for different diameter locations. [PERSON] et al. [32] fabricated a compliant phantom of the aortic arch including the superior branches for in vitro PIV analysis. A hollow patient specific mould was designed, and 3D printed. The inner surface of the mould was coated with RT620 elastomer and rotated to ensure even distribution of the mixture. A 1.5 mm wall thickness was reported with a \(\pm\)0.5 mm tolerance.