Additively manufactured porous scaffolds by design for treatment of bone defects

Toosi, Shirin; Javid-Naderi, Mohammad Javad; Tamayol, Ali; Ebrahimzadeh, Mohammad Hossein; Yaghoubian, Sima; Mousavi Shaegh, Seyed Ali

doi:10.3389/fbioe.2023.1252636

REVIEW article

Front. Bioeng. Biotechnol., 19 January 2024

Sec. Tissue Engineering and Regenerative Medicine

Volume 11 - 2023 | https://doi.org/10.3389/fbioe.2023.1252636

Additively manufactured porous scaffolds by design for treatment of bone defects

Shirin Toosi¹

Mohammad Javad Javid-Naderi²

Ali Tamayol³

Mohammad Hossein Ebrahimzadeh⁴

Sima Yaghoubian⁴

Seyed Ali Mousavi Shaegh^4,5,6*

¹Stem Cell and Regenerative Medicine Center, Mashhad University of Medical Science, Mashhad, Iran
²Department of Medical Biotechnology and Nanotechnology, Faculty of Medicine, Mashhad University of Medical Science, Mashhad, Iran
³Department of Biomedical Engineering, University of Connecticut Health Center, Farmington, CT, United States
⁴Orthopedic Research Center, Ghaem Hospital, Mashhad University of Medical Sciences, Mashhad, Iran
⁵Laboratory for Microfluidics and Medical Microsystems, BuAli Research Institute, Mashhad University of Medical Science, Mashhad, Iran
⁶Clinical Research Unit, Ghaem Hospital, Mashhad University of Medical Science, Mashhad, Iran

There has been increasing attention to produce porous scaffolds that mimic human bone properties for enhancement of tissue ingrowth, regeneration, and integration. Additive manufacturing (AM) technologies, i.e., three dimensional (3D) printing, have played a substantial role in engineering porous scaffolds for clinical applications owing to their high level of design and fabrication flexibility. To this end, this review article attempts to provide a detailed overview on the main design considerations of porous scaffolds such as permeability, adhesion, vascularisation, and interfacial features and their interplay to affect bone regeneration and osseointegration. Physiology of bone regeneration was initially explained that was followed by analysing the impacts of porosity, pore size, permeability and surface chemistry of porous scaffolds on bone regeneration in defects. Importantly, major 3D printing methods employed for fabrication of porous bone substitutes were also discussed. Advancements of MA technologies have allowed for the production of bone scaffolds with complex geometries in polymers, composites and metals with well-tailored architectural, mechanical, and mass transport features. In this way, a particular attention was devoted to reviewing 3D printed scaffolds with triply periodic minimal surface (TPMS) geometries that mimic the hierarchical structure of human bones. In overall, this review enlighten a design pathway to produce patient-specific 3D-printed bone substitutions with high regeneration and osseointegration capacity for repairing large bone defects.

1 Introduction

For the restoration of complex, critically sized difficult-to-heal, or non-healing bone damages and defects, there is no ideal solution. The current clinical approach is to employ bone autografts, allografts, or bone fillers at a defect site; however, the efficacy of such methods for bone regeneration depends on the defect size, its anatomical position, bone nature, and patient’s underlying conditions (Buttery and Bishop, 2005). These strategies are typically inadequate to regenerate all bone defects, and carry the risk of adverse immune response and diseases transmission. Tissue engineering (TE) methods have been developed to overcome these limitations of bone reconstruction through development of implantable porous scaffolds (Puppi et al., 2010; Henkel et al., 2013; Roseti et al., 2017). Scaffolds should be able to recreate the structure of the lost tissue, facilitate the regeneration, offer mechanical properties supporting tissue function, and eventually integrate with the native tissue (Hollister and Murphy, 2011; Ghelich et al., 2022). Bone tissue engineering scaffolds have been fabricated using different methods including gas foaming (Liu and Ma, 2004; Salerno et al., 2012; Bak et al., 2014; Poursamar et al., 2016), laser sintering (Williams et al., 2005; Duan et al., 2010; Maskery et al., 2018), electrospinning (Liao et al., 2008; Jang et al., 2009; Prabhakaran et al., 2009; Di Martino et al., 2011; Yu et al., 2016; Lin et al., 2020), and recently, additive manufacturing (AM) (Valainis et al., 2012; Mota et al., 2015; Bobbert et al., 2017; Bobbert et al., 2017; Düregger et al., 2018; Zadpoor, 2019a; Zadpoor, 2019b; Soro et al., 2019; Yuan et al., 2019). Among them, AM tools including 3D printing have been explored to fabricate scaffolds with controlled architectural, topological, biological, mechanical, and mass transport features (Bobbert et al., 2017; Faramarzi et al., 2018; Ostrovidov et al., 2019; Maydanshahi et al., 2021) to mimic the features of native bones.

Due to the staggering number of patients suffering from bone fracture and loss, the materials and fabrication processes used for bone scaffolds have been critically reviewed elsewhere and will not be discussed in detail here. One area that has not been properly reviewed is the importance of the pore geometrical features on the bone tissue regeneration. In the following sections, initially, a brief overview of the physiology of bone regeneration is provided to provide a better understanding of scaffold design considerations in practice. Then, effects of design parameters on various features of scaffold with a particular emphasis on structures having triply periodic minimal surface (TPMS) geometries are highlighted. TPMS structures provide a higher surface-area-to-volume ratio in comparison with conventional lattice structures that promote cell adhesion, cell migration and cell proliferation. In overall, the challenges limiting the fabrication of scaffolds by design and the opportunities for overcoming these barriers are discussed.

2 Bone regeneration physiology and pathophysiology of large bone defects

Bone is considered as a rigid organ in body which support and guard some other organs and facilitate mobility of live body (Martin et al., 2008; Wang and Yeung, 2017). Bone is a porous composite material, which could be divided into categories of (i) compact and (ii) cancellous bones. Compact section is a hard outer-shell of bone (compact bone), which possesses lower porosity, while cancellous section has a highly porous structure inside a bone (cancellous bone), which is less dense than the outer surface. Porosity of compact bone is within the range of 5%–10% with apparent density of 1.5–1.8 g/cm³; that is the reason why it is called “compact” bone. Cancellous bones have 30%–95% porosity with pores sizes range from 200 μm to 1,000 µm. A desirable porous bone substitute is the one that mimics human bone properties with hierarchical architecture which enable tissue ingrowth and movement of bodily fluids through itself that are required for cell proliferation (Bobbert et al., 2017). Bone is mostly made of hard apatite minerals along with soft collagen protein networks (Currey, 1969). Such composite construction generates the stiffness and the suitable function of bone tissue. For instance, ear bone content of over 80% mineral allows its vibration to transmit sound, however, it is unable to resorb energy (Wang and Yeung, 2017). On the other hand, deer antlers consist of less dense mineral content for absorbing high energy levels (Wang and Yeung, 2017).

Bone formation is continued dynamically through two different processes, known as modelling and remodelling (Wang and Yeung, 2017) that also contribute to bone fracture recovery. Bone modelling progression starts with formation of a new bone with no prior bone resorption, however, during the course of bone remodelling, bone formation occurs in following with bone resorption (Wang and Yeung, 2017). Bone modelling starts at the early ages, changes the shape and size of body bones as it grows to. Bone modelling stops once body reaches its adult age (Kimmel, 1993). In contrast, remodelling is a lifelong procedure, which starts at early life and maintains bone health for proper functionality by constantly substituting impaired bone with new bone (Kimmel, 1993). In contrary to other tissues, bone healing enables body to repair a damaged bone and fully restore it to its previous composition, construction, and functionality (Einhorn and Gerstenfeld, 2015). Bone repair can be defined into direct, i.e., primary bone healing, as well as indirect or secondary bone healing procedures. Direct bone healing mainly starts when small and narrow gaps, usually less than 0.1 mm fractures, happen and the fraction site is rigidly stabilized. As direct bone healing progresses, bone gap is filled continuously through ossification and following Haversian remodelling (DeLacure, 1994). Indirect bone healing happens once the fracture edges are smaller than twice the injured bone diameter. It includes several actions, such as the formation of blood clot, inflammatory response, and formation of fibro-cartilage callus at the site of injury; as well as intramembranous and endochondral ossification, and bone remodelling. The bone fracture repair mechanism initiates with anabolism, increasing bone, differentiation of recruiting stem cells, and retardation with chondrocyte apoptosis (Lee et al., 1998; Einhorn and Gerstenfeld, 2015). Some event such as high-energy trauma, disease, revision and secondary surgeries, developmental deformities, and tumour resections can deteriorates bone healing and create large segmental bone defects (Gugala and Gogolewski, 2002; Wildemann et al., 2007; Reichert et al., 2009). These large bone losses can affect blood circulation and tissue differentiation that finally can lead to bone fracture, that may result in non-union without interventions (Claes et al., 2003). In addition, defect size is not the only parameter that determines a critical bone defect (Lindsey et al., 2006), but defect length is also an important factor that should be considered for bone healing (Khan et al., 2005; Lindsey et al., 2006). Regarding the impressive improvement in the field of bone healing, still non-properly healed fractures or bone defects can extremely affect the quality of patients’ lives because of treatment costs and prolong period of healing (Figure 1).

FIGURE 1

FIGURE 1. Schematics for bone healing process with implantation of scaffold in bone fracture.

3 Role of porous scaffolds as bone substitute in bone tissue engineering

In addition to the properties of bulk materials used for scaffolds preparation, the architectural features of scaffolds affect their function and characteristics. For example, porosity, pore size distribution, and their interconnectivity affect the transport properties and the mechanical properties (Figures 2, 3). In the following section, we will discuses the effects of porosity on various characteristics of scaffolds.

FIGURE 2

FIGURE 2. Design considerations for a bone scaffold. Three parameters including scaffold permeability, mechanical strength of the scaffold and the interfacial adhesion at the interface will affect bone ingrowth, osteoinduction, osteointegration, bone resorption.

FIGURE 3

FIGURE 3. Schematic illustration of the fabrication of 3D porous scaffolds: properties and application.

3.1 Effect of porosity, pore size and porous scaffold on bone regeneration

Size and geometry of the scaffold’s pores, as well as interconnectivity of porous structure highly affect cellular penetration and distribution, their proliferation and differentiation, and formation of blood vessels.

Osteoblasts size ranges within 10–50 μm (Sugawara et al., 2005), however, larger pores (100–200 μm) are more suitable for their function to regenerate mineralised bone post implantation (Toosi et al., 2016; Abbasi et al., 2020). In this way, macrophages are allowed to infiltrate and eliminate bacteria. In addition, infiltration of other cells involved in colonisation, migration and vascularisation is supported (Iviglia et al., 2019).

However, smaller pore size (<100 μm) may induce the creation of non-mineralised osteoid or fibrous tissue (Liu et al., 2018; Iviglia et al., 2019). Previous studies reported that significant bone formation was observed in scaffolds with 800-μm pore sizes. Fibroblasts preferred to fill smaller pores while larger pores were filled bone cells revealing a 800 μm-scaffold could be more suitable for homing and ingrowth of bone cells (Roosa et al., 2010). In a study by Karageorgiou and Kaplan (2005), scaffolds with pore sizes greater than 300 μm were found to be more appropriate to repair large bone defects as the formation of new bone and capillaries were enhanced (Hollister and Murphy, 2011). Excellent osteoinductions were reported for scaffolds having pore sizes ranges from 500 μm to 1,200 μm (Hutmacher, 2000; Van Bael et al., 2012).

Similar observation has been made in the case of hydrogel-based scaffolds. Further, pore size significantly affects individual cell’s response including its attachment, growth as well as proliferation (Al-Munajjed et al., 2008). Highly porous scaffolds are easier for cells to penetrate since material degradation through expression of matrix metalloproteinase (MMPs) is not needed to create space for cell migration. This can potentially reduce the level of inflammation as the over expression of MMPs can induce inflammatory responses (Rosso et al., 2005; Anderson et al., 2011). However, for scaffolds that are expected to support bone regeneration throughout their volume, pore size distribution and their interconnectivity are critical features. In another study, it was shown that bone cells grow faster and differentiate in scaffolds with pores within the range of 100 μm–325 µm (Abbasi et al., 2020). Another important study was performed on scaffolds with small (90–120 µm) and large (350 µm) pore diameters as implanted in rats. Scaffolds with small pores showed chondrogenesis before osteogenesis, while in those with larger pores, direct bone formation was observed due to enhanced vascularization through the pores. The vascularization observed in the larger pores facilitated mass transport through the pores for sufficient oxygen and nutrient delivery as required for direct osteogenesis. Cheng et al. (2016) employed magnesium scaffolds in pore sizes of 250 μm and 400 μm. They observed that formation of mature bone was more in the larger pores owing to improved vascularisation. In this way, sufficient level of oxygen and nutrients could be delivered to maintain osteoblastic activity that result in upregulation of osteopontin (OPN) and collagen type I with direct impact on augmentation of bone mass (Cheng et al., 2016).

In another study, Lim et al. (2010) testified that 200 μm–350 μm was optimum size for osteoblast proliferation whereas cell attachment was not affected at pores with larger sizes (500 μm). Cell aggregation and proliferation could be more controlled in smaller pores (Chen et al., 2018), however, such scaffolds may stimulate endothelial cell proliferation due to exogenous hypoxic condition (Bianco et al., 2017). In addition, proinflammatory higher levels of cytokines including tumour necrosis factor α and interleukin 6, 10, 12, and 13 could be produced in pores with larger size that can activate bone regeneration responses (Mukherjee et al., 2019).

Micropores, however in contrast to macropores, promote protein adhesion and cell attachment over the scaffolds in vitro (Wang et al., 2016; Diaz-Rodriguez et al., 2018; Sokolova and Epple, 2021). O’Brien et al. (2005) found that pores with size of 95 μm could provide the best environment for initial cell adhesion at in vitro conditions (Mukherjee et al., 2019). In another study it was reported that scaffolds with pore size of 100 μm–325 μm was optimal for bone engineering in vitro (Mukherjee et al., 2019). Some previous studies claimed that although pores larger than 50 μm (i.e., >50 μm, macropores) provide favourable effects to enhance osteogenic quality; cell infiltration is limited in small pore size at in vitro conditions.

Owing to these facts, it could be concluded that scaffolds designed to have a gradient in pore size and porosity might provide an optimal solution for bone regeneration. Gradient PCL scaffolds could enhance the osteogenic differentiation of human mesenchymal stem cells (MSCs) at in vitro through increased level of calcium content and ALP activity as a consequence of improved supply of oxygen and nutrients in larger pores (Di Luca et al., 2016). Effect of gradient porosity on cell-seeding was evaluated by Sobral et al. (2011) on 3D poly (ε-caprolactone) scaffolds with pore size of 100–700–100 μm and 700–100–700 μm. In static conditions, the gradient porosity showed higher seeding efficiency, as increased from 35% for uniform porosity to about 70% in the gradient pore sizes (Sobral et al., 2011). In addition, for degradable scaffolds, pore size and porosity regulate the degradation rate. As for PLA scaffolds, Xu et al. (2014) reported that square shape pores had higher degradability and scaffold weight loss.

Porosity is a morphological property of a porous structure and is independent of the structure material that highly affects the biological response of scaffolds (Karageorgiou and Kaplan, 2005; Lv et al., 2021). Such interconnectivity is one of the most essential requirements for tissue ingrowth (Lee et al., 2019; Ferrández-Montero et al., 2020; Dong et al., 2021; Jamee et al., 2021; Jiang et al., 2021; Kumar et al., 2021).

Porosity (P), defined as void space percentage of a solid structure, is determined by Eq. 1 as shown in below (Léon and León, 1998; Yuan et al., 2019):

P = (1 - \frac{P s t r u c t r e}{P m a t e r i a l}) \times 100 % (1)

where P_material shows the density of the bulk material and P_structure is the density of the porous structure (Karageorgiou and Kaplan, 2005).

3.2 Effect of porosity and pore size on permeability of porous structures

The microstructure of TE scaffolds is generally characterized by porosity, pore size, interconnectivity and tortuosity. However, these parameters are not sufficient to predict the success of a porous scaffold. On the other hand, permeability is an important parameter in the assessment of biological performance, including mass transport parameter and can be considered as an independent design parameter (Kemppainen and Hollister, 2010; Pennella et al., 2013; Lipowiecki et al., 2014; Ali and Sen, 2017; Montazerian et al., 2017; Rahbari et al., 2017; Daish et al., 2019; Lv et al., 2022a). Therefore, the mass transport through porous bone substitutes that is mainly measured by permeability, should be well designed to allow for sufficient oxygenation and delivery of nutrients to residing cells (Karande et al., 2004; Hollister, 2005; Dias et al., 2012; Truscello et al., 2012; Bobbert et al., 2017).

Permeability (k) is a proportionally constant between the average velocity of liquid passing through a porous structure at an applied pressure gradient and is defined by Darcy’s law as presented in Eq. 2 (Zhianmanesh et al., 2019):

Ū = - \frac{k}{µ} \nabla p (2)

where Ū shows the average fluid velocity, $\nabla$ P is the applied pressure gradient, and µ is the dynamic viscosity. Several studies have characterized the permeability of porous scaffolds for biomedical applications (Tamayol and Bahrami, 2009). In this way, the constructs with minimal surfaces have received special attention. They have specific geometrical properties that make them appealing for bone tissue regeneration. For these surfaces, the mean curvature is zero which resembles the mean curvature of trabecular bone (Bobbert et al., 2017).

Scaffold pore size is a vital parameter in TE since it promotes cell adhesion, proliferation and differentiation. In this way, modulation of pore size distribution would change the permeability of the TE scaffolds. Al-Munajjed et al. (2008) showed for hyaluronic-collagen scaffolds that the permeability and porosity of scaffolds were increased as pore size was enlarged. Larger pores create less resistance for fluid to pass through the scaffold and provide higher Darcy’s constant (Al-Munajjed et al., 2008).

As for scaffolds with higher permeability, cell suspension experiences less resistance once permeate through the scaffold. This leads to faster stream, which give cells shorter time period to attach to a solid surface. Thus, seeding could be more productive for structures with smaller pores, i.e., lower permeability values (Van Bael et al., 2012).

3.3 Effect of porosity on interfacial adhesion and vascularization

In general, pore structure is a significant consideration for TE constructs. Pores must be highly interconnected to allow for cellular, migration and proliferation and diffusion of required substances. Specific surface area per unit mass is an important design parameter that affects the interfacial cell adhesion of a scaffold. For the scaffolds with small pores, formation of cellular capsules around the edges of pores can limit the delivery of nutrients and oxygen (Mostafavi et al., 2021; Lv et al., 2022b; Lv et al., 2022c). On the contrary, too large pores reduce the surface area and limit the cellular adhesion (Murphy et al., 2010).

In a study, Torres-Sanchez et al. (2017) found that small pores enhanced cell attachment and showed higher cell growth rate until the third day of cell culture owing to the larger surface area. However, larger pores supported the cell proliferation and had larger cell growth rate after the third day (Yuan et al., 2019). Due to the importance of surface area on cell adhesion, the upper and lower values of pore size is a major design consideration for collagen scaffold. O’Brien et al. (2005) tested this hypothesis that the level cell attachment is modulated by the average pore size. It was observed that cell binding and activity could be altered significantly affected by the type of cell, as well as composition and pore size of a scaffold. It could be expected that TE of each construct should require an appropriate pore size [O’Brien et al. (2005)].

The second parameter that affects the interfacial adhesion is the surface structure of scaffold substrate. Surface structure of an implant plays a crucial role in biocompatibility, bioactivity and osseointegration of a scaffold. After implantation, surface of a scaffold directly starts to interact with surrounding bio fluids and tissues. Adsorption of proteins is affected by the scaffolds surface texture and chemistry. Scaffold surface chemistry can affected by the roughness of the surface, response to wettability, and its mechanical attributes (O’Brien et al., 2005; Dave and Gomes, 2019). In this way, cellular-related activities including binding, proliferation and differentiation are highly regulated by the composition and type of absorbed proteins on the surface of the scaffold (Civantos et al., 2017).

4 Additively manufactured scaffolds

3D printing, also known as additive manufacturing (AM), has shown great potential for bone tissue engineering by enabling the fabrication of customized porous scaffolds that mimic the structural properties of natural bone extracellular matrix (Tables 1, 2) (Praveena et al., 2022). The key 3D printing technologies explored for bone scaffold fabrication can be categorized as follows.

TABLE 1

TABLE 1. Summary of the composition, printing technique, pore size, and biological effect of different 3D porous composite scaffolds.

TABLE 2

TABLE 2. Materials employed for 3D printing of tissue engineered bone substitute.

4.1 Extrusion-based techniques

Fused deposition modeling (FDM) is a widely used methodology due to its simplicity and cost-effectiveness. It employs a thermoplastic polymer filament that is heated and then extruded layer by layer through a nozzle, allowing for the creation of a desired scaffold structure. FDM allows for the incorporation of ceramic particles into polymer to improve bone bioactivity. However, high temperatures limit direct printing of cells or bioactive factors. Overall, FDM is ideal for rapidly fabricating customized biopolymer bone scaffolds (Lee et al., 2019; Jiang and Ning, 2020; Winarso et al., 2022; Zhang et al., 2023). Numerous 3D porous scaffolds were created using FDM method (Lee et al., 2019; Ferrández-Montero et al., 2020; Jiang and Ning, 2020; Dong et al., 2021; Jiang et al., 2021).

4.2 Inkjet printing techniques

Inkjet bioprinting utilizes thermal or piezoelectric mechanisms to eject bioink droplets containing cells, growth factors, and other components onto a platform in order to create tissues. This methodology facilitates the generation of bioactive bone scaffolds with exceptional accuracy. However, there are limitations regarding the consistency of bioinks that can be printed (Jamee et al., 2021; Kumar et al., 2021; Parodi et al., 2023).

Generally, inkjet bioprinting offers the advantages of high output, cost-effectiveness, easy implementation, and compatibility with low viscosity biomaterials (Yenilmez et al., 2019; Dell et al., 2022). Consequently, inkjet bioprinting is extensively used in preclinical research and clinical applications (Lv et al., 2018; Li et al., 2020).

4.3 Laser-based techniques

4.3.1 Selective laser sintering (SLS)

SLS is a manufacturing technique that uses a laser to fuse layers of a powdered material based on a three-dimensional (3D) model. The process involves heating and fusing a thin layer of powder using a laser beam that follows a predetermined scanning path. This process is repeated layer by layer until a 3D porous scaffold is formed. The SLS technique has proven to be effective in producing a variety of scaffolds suitable for bone tissue engineering, with an optimal pore structure and improved mechanical properties. This technique has been successfully used to create porous scaffolds using both polymers and metals (Sun et al., 2016; Shuai et al., 2018).

4.3.2 Stereolithography (SLA)

SLA is a 3D printing technique that uses a UV laser to selectively cure and solidify liquid photopolymer resins in layer-by-layer fashion to build a 3D object. This approach possesses a remarkable level of precision and has the capability to generate complex internal configurations. Nonetheless, the range of available materials is limited, and there is a potential hazard associated with the use of toxic resins (Shirvan et al., 2021; Raguraman and Rajan, 2023). Previous researches have reported successful fabrication of 3D scaffolds for bone repair using SLA technique (Elomaa et al., 2020; Ronca et al., 2021).

4.4 Low-temperature printing techniques

Low-temperature deposition manufacturing (LDM) utilizes a process of extrusion to avoid the risks of temperature-induced damage to cells and proteins, thereby ensuring their intact wellbeing. Freeze-drying removes solvents from printed parts. LDM can create bone scaffolds with nano-scale pores for cell infiltration. However, harsh solvents are involved, and weaker structures are produced. LDM uniquely facilitates room-temperature printing of hierarchically porous bioactive bone scaffolds (Zerankeshi et al., 2022). The applications of LDM to fabricate 3D porous polymer-metal composite bone scaffolds have been reported (Ma et al., 2020; Long et al., 2021; Ali et al., 2022).

4.5 Biological 3D printing

This technique focuses on printing cell-laden hydrogel-based bioinks to generate living bone tissue constructs. Key techniques include inkjet bioprinting, extrusion bioprinting, and laser-assisted bioprinting. These methods allow for the printing of bone constructs encapsulating living cells, growth factors, etc. They are essential for fabricating vascularized, functional bone grafts with a physiological cell distribution (Arastouei et al., 2021; Zhou et al., 2021; Maresca et al., 2023). Among the available techniques, FDM is the most widely used 3D printing technique for bone scaffolds due to its simplicity, low cost, and ability to process a range of biomaterials. However, extrusion-based techniques like FDM provide lower resolution compared to light-based methods like SLA or SLS. Inkjet bioprinting enables high precision cell printing but has limitations on the viscosity of printable. The optimal choice depends on factors such as desired resolution, mechanics, and incorporation of biological components. Each 3D printing approach has its own advantages and disadvantages that make it suitable for different bone tissue engineering applications (Ngo et al., 2018; Gharibshahian et al., 2023).

5 Triply periodic minimal surfaces (TPMS)

Triply periodic minimal surface (TPMS) structures have a higher surface-area-to-volume ratio in comparison with conventional lattice structures. Specifically, the TPMS sheet constructs exhibit significantly large surface areas. The advantage of high surface area for TPMS-based scaffolds is an enhancement in cell adhesion, migration and proliferation. Besides, these geometries have an infinitely continued surface with smooth joints that causes lower levels of stress concentration and enhances the mechanical strength of the scaffold. TPMS surfaces known to Schwartz works and defined as periodically infinite structures along with three independent axes that have zero mean curvature of the surface (Karcher, 1989; Yuan et al., 2019; Lv et al., 2022d).

Titanium alloys have been commonly employed as the most appropriate materials for biomedical devices. Such metal alloys have elastic modulus greater than bone tissues that this mismatch in mechanical properties may lead to stress shielding (Claes et al., 2003; Truscello et al., 2012; Liu et al., 2018; Kondiah et al., 2020; Samandari et al., 2022). In this case, stress is removed from bone and majority of exerted forces are bypassed through adjacent implanted scaffolds. Stress shielding results in failure of bone scaffold (Wang et al., 2019) due to the reduction of bone density. A healing bone usually remodels itself at the presence of mechanical loadings. This remodelling process helps bone to adjust its mechanical properties in response to loading. Reduction in load, due to the existence of the implanted scaffold, would cause a bone to become thinner and weaker because there are no stimuli to induce remodelling (Brien et al., 2005). To solve this problem, either a metallurgical method can be applied or porosity can be introduced into the metals (Wang et al., 2019; Yuan et al., 2019). Porosity in metallic constructs reduces the elastic modulus of metal materials which result in stress transfer between bone tissue and its adjacent construct, and allows the porous construct, i.e., scaffold, to get integrated into bone tissue, making long-term osteointegration achievable (Wang et al., 2019). Several studies have been performed to find correlations between design parameters and mechanical properties of porous scaffolds (Bobbert et al., 2017; Maskery et al., 2018; Al-ketan et al., 2019; Wang et al., 2019). Among different methods of fabricating porous structure, the additive manufacturing techniques are prominent because of their abilities in optimization of TPMS structures due to their dimensional and high level of design flexibility with periodic regular structures (Bobbert et al., 2017) (Figure 4). Li et al. (2019) used graded TPMS porous scaffolds and showed that these types of scaffolds are more suitable technique for implant fabrication.

FIGURE 4

FIGURE 4. Topological design based on six different types for TPMS porous biomaterial fabrication.

In the last decade, TPMS scaffolds have been fabricated from a variety of materials including metals, polymers (Karageorgiou and Kaplan, 2005; Karageorgiou and Kaplan, 2005; Faramarzi et al., 2018; Martinez-marquez et al., 2018; Cai et al., 2019; Efraim et al., 2019; Ostrovidov et al., 2019; Yan et al., 2019; Gerdes et al., 2020; Javid-Naderi et al., 2023), ceramics (Hulbert et al., 1970; Kokubo, 1996; Will et al., 2004; Zhang et al., 2016; Bobbert et al., 2017; Elsayed et al., 2017; Diaz-Rodriguez et al., 2018; Toosi et al., 2022), as well as hydrogels (Somo et al., 2015; Bianco et al., 2017; Mohammadi et al., 2018; Wong et al., 2019). 3D printed porous scaffolds from metals, polymers, and ceramics offer mechanical properties comparable to native bones. The structural and mass transport properties of these scaffolds, especially permeability, have a significant effect on bone formation and implant integration.

AM has shown major potential for manufacturing complex structures such as TPMS with extremely extending surfaces. Researchers have demonstrated the feasibility of 3D printing TPMS structures for bone regeneration applications. In a study TPMS structures fabricated with 316L stainless steel by using SLM technology, and discovered their mechanical properties and energy absorption capacities (Zhang et al., 2018). In another one, with SLM printed TPMS specimens compare between finite element method and experimental data then examine elastic modulus, yielding strength, stress strain distribution and the failure occurrence mechanisms of Ostrovidov et al. (2019). In other study, Abueidda et al. (2017) used SLS technology to make different TPMS samples and established that different porosity can alter the mechanical properties of structures. Maskery found that the polymer 3D printed Schwarz primitive lattice displayed stretching and bulking, whereas the Gyroid and diamond deformed in a blending manner (Maskery et al., 2018). A list of scaffolds with TPMS structures is provided in Table 3.

TABLE 3

TABLE 3. Summary of the TPMS structure, substrate materials, printing technique, porosity and mechanical properties.

6 Conclusion and future directions

In summary, this paper provides a review on the effect of porosity, pore size, pore structure and interfacial adhesion on the exchange of nutrient, vascularization, and bone formation. Special attention is given to the AM porous structures, especially TPMS scaffolds. The main conclusions are as follows:

(1) Porous structures that facilitate cell differentiation; migration and formation of blood vessels are desirable for implant applications. Porous scaffolds with 200–350 µm pore size, which mimics the porosity of cancellous bone, facilitate bone ingrowth. It is shown that increasing pore size increases permeability and porosity. In addition, in vivo experiments suggest that larger pore size enhances vascularization, and higher porosity enhances osteogenesis and bone formation. The specific surface area is another factor that affects osteointegration. It is shown that smaller pore size for a scaffold provides a larger specific surface area that is an important factor for cell attachment. Thus, there is a compromise between the optimum pore size for the cell migration and the specific surface area.

(2) Use of porous metallic scaffolds reduces the effect of stress shielding, which results in stress transfer between bone tissue and scaffolds. This effect results in the integration of porous structure in bone tissue and in making long-term physiological fixation. Among different methods of fabricating porous structures, AM of the TPMS surfaces is the optimal method.

(3) Permeability, which can be considered as an independent design parameter, is an essential parameter in determining the mass transport properties of a scaffold for sufficient delivery of nutrients and proper oxygenation.

(4) Engineering scaffolds with multiscale porosity is challenging and 3D printing is a powerful tool for achieving that. However, advanced biomaterials inks and printing nozzles are needed to be developed to achieve a combination of micro to macropores. Another important consideration is the method for the implantation of scaffolds. The fixation of polymeric and hydrogel scaffolds in place is challenging and therefore new strategies are needed to facilitate the implantation of scaffolds. One emerging technique is in vivo 3D (bio) printing, where the scaffold is directly built inside the defect site (Samandari et al., 2022). Such technique eliminates the need to implantation and also facilitate the formation of scaffolds that seamlessly fit the defect site. This area is expected to advance the field of bioprinting for treating complex injuries.

(5) 3D-printed biodegradable metallic scaffolds are also emerging that can facilitate improved bore regeneration. In this way, the incorporation of other materials including polymers/hydrogels into the metallic scaffolds can create a composite structure that can provide proper mechanical strength with suitable microenvironment for cell seeding.

(6) Another critical design consideration is easy handing and manipulation during implantation and surgery for practical use. In this way, scaffold, i.e., bone substitute, should maintain its mechanical shape and integrity during implantation. In addition, any employed materials for scaffold production should enable the conduct of sterilization process prior implanting. Importantly, quality control protocols and regulatory considerations should be also applied to those patient-specific scaffolds and bone substitutes that are produced in factories or hospitals with advanced manufacturing capabilities.

Author contributions

All authors listed have made a substantial, direct, and intellectual contribution to the work and approved it for publication.

Acknowledgments

SAMS would like to appreciate funding from Iran’s National Elite Foundation to support writing this article. AT gratefully acknowledges the financial support of National Institutes of Health (NIH, R01-AR077132).

Conflict of interest

The authors declare that the research was conducted in the absence of any commercial or financial relationships that could be construed as a potential conflict of interest.

Publisher’s note

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References

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