Porosity of 3D biomaterial scaffolds and osteogenesis

Abstract

Porosity and pore size of biomaterial scaffolds play a critical role in bone formation in vitro and in vivo. This review explores the state of knowledge regarding the relationship between porosity and pore size of biomaterials used for bone regeneration. The effect of these morphological features on osteogenesis in vitro and in vivo, as well as relationships to mechanical properties of the scaffolds, are addressed. In vitro, lower porosity stimulates osteogenesis by suppressing cell proliferation and forcing cell aggregation. In contrast, in vivo, higher porosity and pore size result in greater bone ingrowth, a conclusion that is supported by the absence of reports that show enhanced osteogenic outcomes for scaffolds with low void volumes. However, this trend results in diminished mechanical properties, thereby setting an upper functional limit for pore size and porosity. Thus, a balance must be reached depending on the repair, rate of remodeling and rate of degradation of the scaffold material. Based on early studies, the minimum requirement for pore size is considered to be ∼100 μm due to cell size, migration requirements and transport. However, pore sizes >300 μm are recommended, due to enhanced new bone formation and the formation of capillaries. Because of vasculariziation, pore size has been shown to affect the progression of osteogenesis. Small pores favored hypoxic conditions and induced osteochondral formation before osteogenesis, while large pores, that are well-vascularized, lead to direct osteogenesis (without preceding cartilage formation). Gradients in pore sizes are recommended for future studies focused on the formation of multiple tissues and tissue interfaces. New fabrication techniques, such as solid-free form fabrication, can potentially be used to generate scaffolds with morphological and mechanical properties more selectively designed to meet the specificity of bone-repair needs.

Introduction

A key component in tissue engineering for bone regeneration is the scaffold that serves as a template for cell interactions and the formation of bone-extracellular matrix to provide structural support to the newly formed tissue. Scaffolds for bone regeneration should meet certain criteria to serve this function, including mechanical properties similar to those of the bone repair site, biocompatibility and biodegradability at a rate commensurate with remodeling. Scaffolds serve primarily as osteoconductive moieties, since new bone is deposited by creeping substitution from adjacent living bone [1]. In addition to osteoconductivity, scaffolds can serve as delivery vehicles for cytokines such as bone morphogenetic proteins (BMPs), insulin-like growth factors (IGFs) and transforming growth factors (TGFs) that transform recruited precursor cells from the host into bone matrix producing cells [1], thus providing osteoinduction. Finally, osteogenesis occurs by seeding the scaffolds before implantation with cells that will establish new centers for bone formation [1], such as osteoblasts and mesenchymal cells that have the potential to commit to an osteoblastic lineage. Genetically transduced cells that express osteoinductive factors can also be used. Combining scaffolds, cytokines and cells to generate ex vivo tissue-engineered constructs is hypothesized to provide more effective bone regeneration in vivo in comparison to biomaterial matrices alone. In addition, improved bone-like tissue growth in vitro offers new options to study disease progression.

Scaffolds for osteogenesis should mimic bone morphology, structure and function in order to optimize integration into surrounding tissue. Bone is a structure composed of hydroxyapatite (Ca₁₀(PO₄)₆(OH)₂) crystals deposited within an organic matrix (∼95% is type I collagen) [2]. The morphology is composed of trabecular bone which creates a porous environment with 50–90% porosity (typical apparent density values for femoral cortical bone 1.85±0.06 g/cm³) [3] (for relation between porosity and apparent density refer to Methods to measure porosity and pore size section) and pore sizes at the order of 1 mm in diameter [4], with cortical bone surrounding it. Cortical bone has a solid structure with a series of voids, for example haversian canals, with a cross-sectional area of 2500–12,000 μm² that results in 3–12% porosity [5] (typical apparent density values for proximal tibial trabecular bone 0.30±0.10 g/cm³ [3]). The degree of mineralization varies within different bone tissues: for example, in trabecular bone from the calcaneus was measured at 1.135±0.147 g/cm³, while in trabecular bone from the iliac crest it was measured 1.098±0.077 g/cm³ [6]. Four cell types are present in bone tissue: osteoblasts, osteoclasts, osteocytes and bone lining cells [2]. Bone is at a constant state of remodeling with osteoblasts producing and mineralizing new bone matrix, osteocytes maintaining the matrix and ostoclasts resorbing the matrix [2]. Bone lining cells are inactive cells that are believed to be precursors for osteoblasts [2]. Various hormones, such as parathyroid hormone (PTH) and 1α, 25(OH)₂ vitamin D₃, and cytokines, such as IGFs, platelet-derived growth factor (PDGF), fibroblast growth factors (FGFs), vascular endothelial growth factors (VEGFs), TGFs and BMPs are sequestered in bone matrix and regulate bone metabolism, function and regeneration [7].

Mechanical properties of bone depend on age; 3, 5, and 35-year-old femoral specimens had modulus of elasticity values of 7.0, 12.8, 16.7 GPa, respectively [8]. It is generally reported that, after maturation, the tensile strength and modulus of elasticity of femoral cortical bone decline by approximately 2% per decade [3]. Mean values for bone modulus of elasticity and ultimate strength are presented in Table 1. The complexity of architecture and the variability of properties of bone tissue (e.g. porosity, pore size, mechanical properties, mineralization or mineral density, cell type and cytokines gradient features), as well as differences in age, nutritional state, activity (mechanical loading) and disease status of individuals establish a major challenge in fabricating scaffolds and engineering bone tissues that will meet the needs of specific repair sites in specific patients.

Scaffold properties, depend primarily on the nature of the biomaterial and the fabrication process. The nature of the biomaterial has been the subject of extensive studies including different materials such as metals, ceramics, glass, chemically synthesized polymers, natural polymers and combinations of these materials to form composites. Properties and requirements for scaffolds in bone tissue engineering have been extensively reviewed and recent examples include aspects of degradation [9], [10], [11], [12], mechanical properties [9], [13], [14], [15], [16], [17], cytokine delivery [18], [19], [20], [21], [22], [23], [24], [25] and combinations of scaffolds and cells [23], [26], [27], [28], [29], [30].

Porosity is defined as the percentage of void space in a solid [31] and it is a morphological property independent of the material. Pores are necessary for bone tissue formation because they allow migration and proliferation of osteoblasts and mesenchymal cells, as well as vascularization [32]. In addition, a porous surface improves mechanical interlocking between the implant biomaterial and the surrounding natural bone, providing greater mechanical stability at this critical interface [33]. The most common techniques used to create porosity in a biomaterial are salt leaching, gas foaming, phase separation, freeze-drying and sintering depending on the material used to fabricate the scaffold. The minimum pore size required to regenerate mineralized bone is generally considered to be ∼100 μm after the study of Hulbert et al., where calcium aluminate cylindrical pellets with 46% porosity were implanted in dog femorals [34]. Large pores (100–150 and 150–200 μm) showed substantial bone ingrowth. Smaller pores (75–100 μm) resulted in ingrowth of unmineralized osteoid tissue. Smaller pores (10–44 and 44–75 μm) were penetrated only by fibrous tissue [34]. These results were correlated with normal haversian systems that reach an approximate diamter of 100–200 μm [34]. However, using laser perforation techniques and titanium plates, four different pore sizes (50, 75, 100 and 125 μm) were tested in rabbit femoral defects under non-load-bearing conditions [35]. Bone ingrowth was similar in all the pore sizes suggesting that 100 μm may not be the critical pore size for non-load-bearing conditions [35].

In the present review pore size and porosity for different biomaterials are reviewed in the context of mechanical properties and extent and type of bone formation in vitro and in vivo. Based on this assessment conclusions are drawn regarding the relationship between these morphological and functional features to provide guidance regarding design choices for scaffolds related to bone repair.