## 1 Introduction  
Tissue engineering scaffolds should provide adequate three-dimensional (3D) microenvironments for cell adhesion, proliferation, differentiation, and natural extracellular matrix (ECM) deposition [1]. The physical and chemical properties of these scaffolds are key factors in tissue regeneration [2, 3]. With the increasing understanding of the intrinsic interactions between cells and their microenvironment, the preparation of scaffolds mimicking the compositional and structural aspects of ECMs has received growing attention [4, 5]. Since naturally occurring ECMs are mainly composed of nano-scale collagen fibers, to date, several fabricating methods (e.g., phase separation, self-assembly, and electrospinning) have been applied to prepare biomimetic scaffolds with nanofibrous structures [6, 7, 8]. Of these methods, electrospinning is the only technique that allows the continuous production of fibers at micro or nanoscales [9]. Electrospinning has attracted great attention because of its material versatility, straightforward nature, and cost-effectiveness [10, 11]. Although electrospinningning-manufactured nanofibers possess a large specific surface area, high porosity, and spatial interconnectivity, electrospun nanofibrous mats still lack macroporous structures [12]. As a result, the intrinsically small interfiber pore size of these materials hinders cell penetration across the layers, which is necessary in forming integrated tissues. In addition, the small interfiber pore size of electrospun nanofibrous mats constrains infiltration into the cell itself, hindering its integration with the host tissue upon in vivo implantation [13]. Sacrificial microbeads, particulate leaching, freeze drying, and cryogenic electrospinning have been used in an attempt to increase the pore size of these products [14]. However, nanofibrous mats with large pore sizes (ca. 100-200 um) and highly interconnected pore structures for the transport of cells and metabolites have not been well achieved yet [15; 16]. Three-dimensional (3D) porous scaffolds play an important role in tissue engineering, as they can provide vital frameworks for cell proliferation, migration, and differentiation into specific tissues while secreting ECM components [17]. Scaffolds with complex, and often anisotropic, porous microstructures have been prepared via freeze-drying cycles by controlling the freezing temperature and rate [18]. These porous microstructures provide nutrients, transfer oxygen, and remove waste. In addition, pores of varying sizes are critical for tissue repair, since neovascularization, fibroblast ingrowth, skin regeneration, and bone regeneration require different pore sizes [19]. However, the inherent lack of mechanical strength associated with porosity is a common problem with these porous scaffolds and limits further applications of these materials. Synthetic polymers such as polycaprolactone, polylactic acid, and poly(lactic-co-glycolic acid) are frequently used to fabricate fibers or porous scaffolds in tissue engineering, and previous studies already showed their compatibility to the cells. By in vivo studies, these biomaterials showed too slow degradation rates [20], which are not suitable for some of the applications, such as soft tissue and skin regeneration. Both chitosan and collagen are natural derived materials that have been widely used in tissue engineering because of their excellent biocompatibility, biodegradability, biological activity, and low immunogenicity [21; 22]. Chitosan, obtained by the deacetylation of chitin, is one of the most abundant polysaccharides, while collagen is a significant component of ECMs. In addition, chitosan and collagen are usually combined to mimic the composition of natural ECMs and improve their physical properties [23; 24]. Collagen used in tissue engineering is mainly obtained from animal tissues and considered protein of xenogeneic origin, and its purity and quality are not always consistent [25].