## 1 Introduction  
Muscle hardness, defined as a mechanical property that represents the transverse muscle stiffness [1], exhibits an intrinsic viscoelastic tension [2]. At the molecular level, it has been shown that muscle fiber stiffness is proportional to the number of attached cross-bridges and changes linearly as the muscle fiber develops the force [3]. In the network of the tensional tissues (biotensorignity) [4], passive tension is independent of the central nervous system's contribution (non-neural), resulting from the intrinsic molecular interaction of the actomyosin filaments in sarcometric units and myofiblast cells of the skeletal muscle [2]. However, the neural component of the muscle tension, which is mostly defined as the active muscle tension, has a physiological implication to the muscle's tone at rest [5]. The active and passive tension, made up of the passive elements within and surrounding the muscle, together with the active elements of the muscle, need to act in accordance in order to maintain the dynamic behavior of the musculoskeletal stability [6]. When the joint angle changes, the amount of lengthening (elongation) of the muscle becomes important for understanding the intrinsic properties of the muscle. The lengthening of the skeletal muscle, accompanied with an angular motion, exhibits a resistance caused by the passive stiffness when its motor neuron is quiescent and myofibrils are not actively contracting [7]. This resistance becomes measurable and increases with increasing tension when the muscle goes beyond its slack length, which is usually assumed to be the length measured with the joint in middle-range position and when the net joint torque is minimized [8]. With aging, the intrinsic properties of the muscle and tendon also change and hence alters the muscle stiffness and musculoskeletal stability [6]. On the other hand, even though the muscle morphology remains similar for six decades, the aging factor could begin to affect the characteristics of skeletal muscle as early as at 30 years of age [9; 10]. To investigate the hardness of the muscular tissue, a non-invasive assessment of the localized muscle mechanical properties has been increasingly conducted using the ultrasound shear wave elastography (SWE) technology [11; 12]. The ultrasound probe generated a radiation force followed by tissue deformation, resulting in propagation of a transient shear wave. This propagation was captured by the ultra-fast ultrasound image to calculate the tissue hardness [13]. This technology makes it possible to assess the stiffness of both superficial and deep muscles [10]. The ultrasound SWE has opened the window towards a better understanding of the alterations in muscle hardness, and has thus inspired researchers to find out the effect of the stretching treatment on muscle stiffness under various conditions. However, there have been variations in the revealed results in shear modulus, such as 9.24 \(\pm\) 1.98 kPa in Gastrocnemius Medialis (GM) muscle with knee extended [11], and 16 \(\pm\) 5 kPa in lateral gastrocnemius during ankle plantar-flexion and 26 \(\pm\) 9 kPa during ankle dorsi-flexion while the muscle is isometrically contracting [12]. The application of the measurement technique is another important factor that influences muscle shear modulus, which is already sensitive to (i) specific tension, (ii) degree of anisotropy, (iii) pennation angle, and (iv) angle between muscle fibers and transducer direction [14]. The muscle shear modulus can be affected by the placement of the ultrasound probe, which merits further study to introduce a reliable ultrasonic assessment protocol. The muscle shear modulus was reported to be significantly different when an ultrasound probe was positioned at the proximal one third, the distal one third and in the middle, along with the contraction direction of the GM muscle [15]. The previous studies described the probe location with reference to the recommended placements of the electrodes of electromyographic (EMG) recording [16; 17].