## 1 Introduction  
In recent years, several efforts were made to study new mechanical models to analyze biological soft tissues [1]. These materials display nonlinear behavior, where the mechanical response depends on time and the previous load state. The diversity of biological soft tissues and the complexity of their structure, made up of several layers and material properties, make mechanical model development difficult. The need for in silico mechanical models for simulating surgical procedures boosted the development of biological mechanical properties and the study of multiple biological materials, such as muscles [2; 3; 4; 5], tendons [6], brain [7; 8], skin [9; 10], arteries and veins [11; 12; 13], and eye tissues [14; 15]. In all cases, material properties are obtained from adjusting experimental data, where time dependence is evaluated by tests run at different strain rates [16]. Of all human tissues, the cornea is the only avascular tissue, and several investigations were carried out due to the number of clinical applications that include refractive surgery corrections, intracorneal rings, evaluation of keratoconus progression analyses, and cross-linking treatment optimization [17]. The in vivo corneal biomechanics study is currently based on biomechanical parameters evaluated from the results of noncontact tonometers [18]. The human cornea protects the eye from external agents, and it is a unique tissue due to its transparency and strength properties. The cornea has an aspherical shape and is usually represented by means of a biconical surface [19]. It is composed of several layers, where the stroma is the main structural layer with near 90% of total thickness [20]. The stroma is based on 200-400 lamellar layers with collagen fibers embedded in an extracellular matrix, these fibers run parallel in each lamella, but at different angles between them. Lamellar thickness varies across corneal thickness. Lamellae are twice as thick in the posterior zone compared to anterior ones [21], and shear stiffness is in 3-fold greater in the anterior zone due to the percentage of interweaving among lamellae. Both orientation and distribution were quantified by X-ray techniques [22]. According to the obtained results, collagen fibers are distributed mainly in the nasal-temporal and superior-inferior directions in the central cornea area and circumferentially in the limbus [23]. [PERSON] and [PERSON] [24] estimated that two thirds of fibrils lie in the sectors contained at \(\pm 45^{\circ}\) around the horizontal and vertical meridians. This distribution can be observed in the central 7 mm zone. According to secondary harmonic generation images, a clear 3D arrangement distribution appears and allows analyzing the mechanical influence of collagen fibers with thickness [25]. Collagen net damage causes loss of stiffness and a change in corneal curvature, which results in the corneal bulging and its local thinning, whose origin is natural or physical. This disease is known as keratoconus. Recent studies [26] established that the deterioration of the transversal bonds between fibers could be the reason for this weakening due to loss of fiber orientation in lamellae. When the origin is due to corrective vision surgery, it is called postsurgical ectasia, and, despite advances, the incidence of the unpredictability of surgical outcomes remains close to 2.8% [27]. The improvement of the surgery success percentage is associated with the development of patient-specific models, for which considerable efforts were made [28]. The development of patient-specific models starts with the topographic corneal measurement and mechanical modelling improvement [21], which strongly depends on the employed material model and materials' properties. The material properties of the cornea can be evaluated by mechanical tests [17].