#### **Table 2.** Studies selected in the systematic search according to information related to bone properties.

**Table 3.** Studies selected in the systematic search according to information related to additively manufactured (AMd) porous Ti structures' mechanical, geometrical, and dimensional properties.

Through the systematic research performed in this study and by reviewing the medical device regulations from the Food and Drug Administration (FDA), a Quality Target Product Profile has been established including proposed values for the Ideal Quality Attributes for mechanical and dimensional properties of porous bone implants. These target properties are aimed for porous metal implant structures designed for load bearing implant applications. Following this, the results of the selected studies were compared and discussed, from a biomimetic point of view, with the characteristics of natural human bone to identify implants with properties similar or superior to human bone and current medical standards.

#### *3.1. QbD Step 1: Ideal Quality Target Product Profile*

The Quality Target Product Profile (QTPP) is critical for formulating the ideal features of a product considering both performance and safety. To direct the product development process, it is vital to understand user needs. Using QTPP, design failures can be identified early in the product development process to reduce costs and time. According to Martinez-Marquez et al. [36], the quality of bone implants should be defined from three perspectives: product-based, manufacturing-based, and user-based. In the context of fully porous Ti microstructures, the dimensions of quality considered most relevant are performance, features, reliability, conformance, durability, and perceived quality. The requirements relating to bone implants corresponding to these quality dimensions have been identified through studies of existing scientific research results.

#### *3.2. Bone Implant Requirements*

An additively manufactured bone implant must act as a stable scaffold that is biocompatible without causing inflammation or leaching material toxins into surrounding tissue. It requires a suitable surface that promotes cell adhesion and differentiation as well as provides a constant flow of cell nutrients and metabolic waste. This allows for bone tissue formation [81]. The scaffold material must have mechanical properties matching those of the surrounding tissues to avoid stress shielding and mechanical failure [81,82]. The bone pore size, geometry, interconnectivity, and porosity are microscopic features that make for the foundation of bone regeneration, cell growth, osteoconduction, and cell proliferation [74]. Designing implants with adequate pore dimensions allows for a constant flow of cell nutrients and waste. It also allows for sufficient connections to establish between the local bone area and the scaffold [74]. If the right attributes are chosen for the implant microstructure, it can mimic human bone's natural characteristics, which is the end goal of biomimetic implant design [79,80]. Vasireddi and Basu [140] completed a list of general requirements for 3D printed implants:

In similar research, Jabir et al. [141] described the main fundamental requirements for implants as biocompatibility, good manufacturability, geometric precision, appropriate design, biomechanical stability, resistance to implant wear, corrosion and aseptic loosening, bioactivity, and osteoconduction. Since the environment in the human body is highly corrosive and biomaterials are usually bioactive, the implant will interact with its environment after implantation [140]. The implant must therefore be designed with both useful functions and biological safety, providing utmost biocompatibility. The implant design must have a high degree of reproducibility, which will ensure faster and cheaper manufacturing as well as predictability and reliability.

Furthermore, the implant must be durable and of initial strength for safe handling during sterilisation, transport, and surgery, and to survive physical forces in vivo after implantation [142]. The implant will be subjected to constant load in the body, from walking and further strenuous

movements [127]. Its mechanical strength is vital for its viability, where the implant must last the entirety of its expected lifetime without defects or failure [127].

Taking all the above into consideration in conjunction with the ideal eight quality dimensions of AMd bone implants proposed by Martinez-Marquez at al. [74], we proposed seven ideal quality dimensions of porous internal architecture of Ti bone implants. These quality dimensions are based on three quality perspectives, namely product, manufacturing, and user-based, as presented in Table 4.

**Table 4.** The ideal quality dimensions of porous internal architecture of Ti bone implants.

#### *3.3. QbD Step 1.1: Ideal Quality Attributes*

The Ideal Quality Attributes (IQA) are the tissue or biological construct characteristics that must be mimicked to imitate the desired tissue biological architecture and functions. The IQA can be dimensional, physicochemical, mechanical, biological, or functional. However, if any technological and regulatory limitations exist it is important to consider that these IQA serve just as 'ideal models' to imitate even if it is not possible to achieve them. Therefore, the IQA can provide an ideal goal for the product development process in any tissue engineering project.

By studying the properties of natural human bone, we can find the different IQAs for bone implants [79,80]. As mentioned in Section 3.1, the implant must be of adequate strength while the elastic modulus must be similar to that of surrounding tissue to avoid damage [81]. The structure must be porous, using unit cells that allow for fluid movement and bone ingrowth, and the right porosity will reduce the implant stiffness. Interconnectivity completes the flow within the structure and facilitates bone ingrowth [82]. A combination of cancellous and cortical bone properties applied on bone implants can allow the implant to function like a natural bone, yet stronger without impairing surrounding native tissue. Another specific property of bone is its ability to heal itself when fractured [90,143]. The process involves cell migration, differentiation, and cellular proliferation [125]. The ability of natural bone to selfheal must be considered when designing a Ti implant, since the structure will not be able to do the same.

#### 3.3.1. Structure and Composition of Bone

To fully understand the microstructure requirements of fully porous Ti implants, one must first understand the properties of human bone. Bone is essentially an open-cell composite material of fibrous protein, collagen, and calcium phosphate crystals, with an intricate vascular system forming various structures and systems in a five-level hierarchically organised structure [84,144]. According to Rho, et al. [145], these five hierarchical levels are dimensional scales ranging from the macro to the sub-nano levels (Figure 3).

**Figure 3.** Hierarchical structural organization of bone: (**a**) cortical and cancellous bone; (**b**) osteons with Haversian systems; (**c**) lamellae; (**d**) collagen fiber assemblies of collagen fibrils; (**e**) bone mineral crystals, collagen molecules, and non-collagenous proteins. Image reproduced from Ref [81].

Bone's macrostructure is composed of cancellous and cortical bone, which are two different regions of bone with different density. Cortical bone forms the outside layer of the bone providing a strong, compact structure, leaving only 3–10% of the volume for its biological elements, such as blood vessels, osteocytes, erosion cavities, and canaliculi [81]. Cancellous bone forms the inside of the bone and is spacious and highly porous [81]. The pores are filled with bone marrow and the spacious architecture allows space for metabolic waste and nutrients to flow. Cancellous bone has an active metabolism and regenerates quicker compared with cortical bone [146]. By changing its density, cancellous bone can reorganise its structure depending on the stress direction [107]. These features cause cancellous bone mechanical properties to vary from bone to bone and change longitudinally [144].

The microstructure of bone ranges from 10 to 500 µm and it contains three major cavities. These are Haversian canals, osteocyte, lacunae, and canaliculi [88]. Cortical bone microstructure is composed of cylindrical structures called osteons with diameters ranging between 70 to 140 µm [81,87,92]. Along osteons' central axis are pores called haversian systems, with diameters ranging between 20 to 50 µm, containing blood vessels and nerves [92]. Cancellous bone micro-architecture is composed of irregular units called trabeculae which create its porous structure. The pore geometry and pore size of cancellous bone porous structure is critical for cell distribution and cell migration [81]. Cancellous bone is naturally stochastic, with random pore distribution of pores of different size.

The pores in cancellous bone are ellipsoidal in the natural direction of loading and are usually 300–600 µm wide [147]. The interconnectivity between the pores of cancellous bone is essential for nutrient and waste diffusion.

At the micro level are also located the three types of differentiated bone cells osteoclasts, osteoblasts, and osteocytes, as shown in Figure 4. Osteoclasts and osteoblasts are vital for the functions of developing and healing bone tissue [86]. Osteoclasts (Figure 4a) are the main cells responsible for resorption of old bone tissue. Osteoblasts (Figure 4b) are bone cells responsible for bone formation, remodelling, fracture healing (for which they are critical), and bone development [148]. Osteocytes (Figure 4c) are osteoblasts cells present inside mature bone and serve as mechanosensory cells to control the activity of osteoclasts and osteoblasts [149].

Bone's sub-microstructure, from 1 to 10 µm, is composed of lamellae which in cortical bone compromise the concentric layers of osteons and in cancellous bone lamellae forms the trabeculae volume [150,151]. The sub-nanostructure of bone, below a few hundred nanometres, is composed molecular constituent elements, such as collagen, non-collagenous organic proteins, and mineral. From a few hundred nanometres to 1 µm is bone's nanostructure comprising fibrillar collagen with embedded hydroxyapatite nanocrystals [152].

#### 3.3.2. Bone Mechanical Properties

Most bones in the body are load bearing and require high mechanical strength. Bone tissue is anisotropic and stronger in compression than in tension. The mechanical properties must be measured in two orthogonal directions: longitudinal, which is the natural loading direction, and transverse [86]. The mechanical strength of bone is complex to measure as it varies with age, health, activity, and position in the body [86]. Bone becomes stiffer and less ductile with age and its ability to heal decreases [155]. It is also likely to weaken with immobilisation, such as for a person with movement disabilities or limited physical activity.

The two different types of bone, cortical and cancellous, have completely different mechanical properties; therefore to specify the properties of bone as one material, both bone types must be considered. Cortical bone is highly dense an act like a shell that provides the greatest stiffness and resistance to bending. In contrast, the mechanical properties of cancellous bone are determined by its apparent density and trabecular architecture. The trabecular structure of cancellous bone is arranged accordingly to the stress distribution of load, as shown in Figure 5. As a result, the least material is used in the most strategic locations to carry the greatest loads with the least strain [156].

**Figure 4.** Coloured scanning electron micrographs of bone cells: (**a**) Activated osteoclast and resorption pit by kind permission of Timothy Arnett Ref [153]; (**b**) Osteoblast growing on a bone scaffold made of calcium oxide and silicon dioxide with added strontium and zinc by kind permission of Guocheng Wang from [154]; (**c**) Osteocytes embedded in the bone matrix with long cytoplasmic extensions reaching into the bone tissue, by kind permission of Kevin Mackenzie. Here, the minerals in the bone have been removed by embedding in resin and etching with perchloric acid. This reveals the spaces in the bone and the shape of the osteocyte cells.

**Figure 5.** Cortical and cancellous bone; image adapted from [157,158] by kind permission of Alan Boyde.

The mechanical properties of cortical and cancellous bone are difficult to measure, and tend to vary depending on bone orientation, test methods, mathematical formulas, and assumptions [146,151,159]. According to Chen and Thouas [57], the elastic modulus of cortical bone is approximately 11–21 GPa in the longitudinal direction and 5–13 GPa in the transverse direction. Another study suggested

18–22 GPa [147], whereas Wang et al. [81] suggested a range of 3–30 GPa. Cortical bone has a porosity of less than 10% [85]. The elastic modulus of cancellous bone is estimated to be 0.02–6 GPA and it has a high porosity of 50–90% [85,86].

To achieve adequate strength in bone implant design, one must understand the strength requirement of bone. Especially cortical bone must be considered, because the strength of such tissue is the minimum strength required by the implant. The compressive yield point of bone represents the threshold from where the structure accumulates irreversible deformation. Unlike metals such as steel, the yield point cannot be clearly distinguished, and it is rather associated with a continuous transition zone [160]. Furthermore, it has been proven that the compressive yield strength of bone varies depending on the anatomic site [161]. The challenges in determining the yield strength of cortical bone results in varying values in the literature. Researchers have estimated cortical bone to have a compressive yield strength of 133.6 ± 34.1 MPa [162]. The same characteristic was estimated to be 108–117 MPa by Yeni and Fyhrie [163]. Further studies have tested compressive yield strain using uniaxial compression and achieved 141.0 ± 5.0 [164], 111.0 ± 18.6 [165], 112.5 ± 9.5 [166], and 115.1 ± 16.4 MPa [167].

If the loading surpasses the yield point for bone, it will eventually reach the ultimate point. This point represents the ultimate compressive strength the bone can withstand until irreversible strains and damage occur. Past this point, macrocracks are formed and fracture occurs [160]. Unlike yield strength, the ultimate compressive strength of cortical bone can be exactly determined, using a stress– strain experiment [160]. However, due to bone properties differentiating, values vary in the literature. For example, Wang et al. [84] suggested that cortical bone has an ultimate compressive strength of 180–200 MPa, whereas Calori et al. [8] suggested a wider range of 130–290 MPa and Henkel et al. suggested 100–230 MPa [168]. Tables 5 and 6 summarise the mechanical and dimensional properties of natural bone considered in this study.

**Table 6.** Summary of mechanical properties of natural human bone.

*3.4. Comparison of Properties of Porous Ti Sca*ff*olds Fabricated by AM and Ideal Quality Attributes*

#### 3.4.1. Unit Cell Geometry

Metamaterials can be rationally designed by changing their geometry at the microscale of the constituting unit cells of the porous structure. In this systematic search, a total of 169 porous scaffolds were identified as a rationally designed and fabricated for bone implant applications. It was found that there are three preferred strategies for fabricating bio-metamaterials: beam-based, sheet-based, and including irregular porous structures [172].

According to Figure 6, the preferred design approach was beam-based, which represent 74.6% of the total scaffolds produced in the selected studies. The beam-based bio-metamaterials' micro-architecture is composed of a lattice structure created using unit cells based on platonic solids, Archimedean solids, prisms and anti-prisms, and Archimedean duals [173–177] to mimic bone porous macro structure and mechanical properties such as modulus of elasticity. Nevertheless, the biological performance of bio-metamaterials created with beam-based geometries is limited by their inaccurate description of complex natural shapes due to their straight edges and sharp turns [109].

**Figure 6.** Distribution of experimental studies on beam-based, triply periodic minimal surfaces (TPMS), and random based geometries for bone implant applications.

The unit cells of sheet-based geometries, on the other hand, are based on triply periodic minimal surfaces (TPMS), which are present in different organisms and cellular structures [178]. Therefore, it is no surprising that the second most used design strategy identified in this systematic search was sheet-based representing 20.1% of the total scaffolds produced, from which all used a TPMS as unit cell. Bio-metamaterials based on TPMS can mimic the various properties of bone to an unprecedented level of multi-physics detail in terms of mechanical properties and transport properties [79,172]. Moreover, the bone-mimicking mean surface curvature of zero of TPMS eliminates the effect of stress concentrators at nodal points [120].

In the case of irregular porous structures, are created in a random way to generate irregular porous structures that mimic trabecular bone geometry and mechanical properties [179]. These irregular structures are generated using the Voronoi and Delaunay tessellation methods. Irregular structures have been found to further enhance scaffold's permeability and bone ingrowth compared with porous structures designed with regular unit cells [180]. According to our results only 5.3% of the studies used a randomised design approach to create irregular porous structures to mimic trabecular bone. This result was surprising considering that irregular porous structures were the first additively manufactured coatings used for orthopaedic implants in the medical industry. However, unlike random porous scaffolds, the great advantage of using regular repeating arrays of unit cells made of beam or sheet-based geometries is that they allow the creation of metamaterials with properties that can accurately be predicted [181]. This explain why these design strategies are preferred in research.

From all the different possible unit cells that can be used to produce metamaterials, a total of 17 types of unit cell were used by the selected studies, as presented in Figure 7. According to Figure 7, the beam-based diamond unit cell (59 studies) was the most used, followed by the cubic (18 studies), and the gyroid TPMS (17 studies). These results correlate with the opinion of different experts who have stated that the diamond unit cell is the most studied for the development of metamaterials due to its biomimetic mechanical properties [121,182]. The high mechanical properties and self-supporting properties of the diamond unit cell are due to its unique geometrical arrangement, where one node is tetrahedrally surrounded by four other nodes coming from the crystal structure of the diamond

crystal [113], as shown in Figure 8a. Moreover, this arrangement gives 48 symmetry elements to the diamond structure, making this unit cell invariant to different symmetry operations such as translations, reflections, rotations, and inversion [183]. Similarly, extensive research for bone regeneration has been performed to study the cubic unit cell (Figure 8b). The research interest in the cubic unit cell is because it is based on one of the simplest and easy–to–manufacture platonic solids thanks to its struts at an angle of 90◦ [62]. In the case of porous metamaterials based on the gyroid TPMS (Figure 8c), they have been found to exhibit similar topology to human trabecular bone, and also superior mechanical properties compared with metamaterials based on other types of TPMS [120]. For example, according to Yang et al. [184] metamaterials based on the gyroid TPMS have a more homogeneous stress distribution, which can provide equal mechanical stimulation to bone cells.

**Figure 7.** Comparison of the number of experimental studies of each unit used for bone regeneration identified in the reviewed articles.

**Figure 8.** Some examples of unit cells with their corresponding meta-biomaterial scaffold below: (**a**) diamond beam-based unit cell; (**b**) cubic beam-based unit cell; (**c**) Gyroid triply periodic minimal surface based (TPMS-based) unit cell; (**d**) Diamond TPMS-based unit cell; and (**e**) Voronoi (top) and Delaunay (bottom) irregular porous structures.

In the case of random or stochastic structures, Kou et al. [185] suggested that scaffolds based on this structures are more realistic; that is, they look more like natural bone with random non-uniform pore distribution and pore size [185]. Such structures are believed to provide benefits such as improved mechanical properties, including strength, fluid dynamics, surface area, and surface-to-weight ratio. They combine advantages of small and large pores without necessarily decreasing the mechanical strength or reducing the bone in-growth to levels that are inappropriate in application [186]. Figure 8 presents the five most representative unit cells identified in this study.

#### 3.4.2. Porosity

It is known that the degree of micro-porosity in bone implants directly affects their biological and mechanical properties. The porosity of natural bone is crucial for vascularisation, diffusion of cell nutrients and metabolic waste, and cell migration [187], and in a similar way it is important for metal bone implants. Moreover, several studies have considered porosity as the main parameter affecting stiffness and strength of porous biomaterials. Increased porosity reduces the strength of the implant [187,188]. As a result, porous metallic biomaterials are used as coatings in many medical applications, but more recently porous biomaterials have been proposed for load-bearing applications beyond surface coatings [138,142]. Ti and Ti alloys are commonly used for load-bearing implant applications due to their relatively low elastic modulus fatigue resistance, high strength to weight ratio, and corrosion resistance [82,94]. However, bulk Ti and Ti alloys do not completely match all the mechanical properties of natural bone such as modulus of elasticity. Therefore, it is a need of the hour to accomplish specific mechanical properties for Ti or Ti-based alloys by controlling the porosity and pore characteristics for customised implants [131]. However, the ideal porosity for medical implants seems controversial in the literature [181].

In this systematic review a total of 49 articles out of 50 recorded porosity of various degrees, as shown in Figure 9. For example, Stamp et al. [186] recommend using a porosity above 65% in medical implants whereas Ghanaati et al. [189] found that vascularisation increased in vivo when reducing the porosity from 80 to 40%. Sarhadi et al. [190] and Schiefer et al. [191] have recommended using a porosity of approximately 50%. According to Will et al. [192], the porosity that best promotes vascularization in porous scaffolds is 40–60%. Pattanayak et al. [193] manufactured porous Ti implants and reported an increase in compressive strength from 35 MPa to 120 MPa when reducing the porosity from 75 to 55%. Murr et al. reduced the porosity from 88 to 59% with an increase in stiffness from 0.58 GPa to 1.03 GPa [131]. As mentioned, natural cancellous bone has a porosity of 50–90% [85,86].

Zou et al. [102] designed three implants with similar pore size and shape. By reducing the porosity from 72 to 53% they achieved a compressive strength of 200 MPa and Young's modulus of 4.3 GPa, instead of a compressive strength of 60 MPa and Young's modulus of 2 GPa. Even though both implants achieved a stiffness close to that of human bone, only the implant with 53% porosity achieved a compressive strength greater than human bone. Hence, the implant with 72% porosity would not

qualify as a load-bearing bone implant. These results, along with previous research by Pattanyak et al. and Murr et al., confirm the influence of porosity on mechanical properties [193]. They also show that the implant porosity may be adjusted within limits to increase strength and adjust bone stiffness, which varies for each patient's characteristics.

From a medical regulatory perspective, the U.S Food and Drug Administration (FDA) only approves implant porosities of 30–70% for porous coatings on solid Ti implants [194]. The range is relatively large but can be used to one's advantage since both the elastic modulus and strength of the implant can be adjusted by adjusting the porosity. Implants with porosity outside of this range do not comply with FDA regulations and cannot enter the market. It is known that medical regulations, especially for implantable medical devices, are based on strong scientific evidence. Therefore, it is vital to adhere to these regulations when designing bone implants.

Taking into consideration the medical regulations for porous implants, we selected a porosity range of 30–70% as the IQA for porous metal implants to identify studies in the systematic search that fabricated Ti scaffolds with porosity values within this range. From the 49 articles that recorded porosity of various degrees, a total of 167 results were extracted and compared with the selected IQA porosity range. The results of this comparison are presented in Figure 9. According to our results, 56.6% of the porous scaffolds studied in the 49 selected articles had porosity values within the acceptable porosity range (30–70%) required to satisfy medical regulations such as the FDA. By contrast, a total of 26 studies explored the properties of porous scaffolds with porosity values above the acceptable porosity range representing 37.3% of the total results extracted in this systematic literature review. There were several reasons for these studies to explore porosity levels higher than 70%. For example, Zhang et al. [100] and Amin Yavari et al. [106] fabricated different porous structures with various porosities to explore their mechanical properties and deformation mechanisms. Moreover, porous Ti scaffolds with high levels of porosity can serve as storage for mesenchymal stem cells to facilitate bone tissue regrowth, and also to improve cell oxygenation and nutrition [100]. On the other hand, Ti porous scaffolds with porosity levels lower than 30% can provide similar mechanical properties to cortical bone [98,100].

#### 3.4.3. Macropore Size

Since macro pore size is directly related to the strength, porosity, and stiffness of the implant, it is an important property for implant design [195]. Pore size has a profound effect on the behaviour of osteogenic cells even in an organ culture system [196]. The implant's macro porosity determines whether bone cells can successfully penetrate and grow within the structure, and many studies have discussed the influence of pore size on the biological properties of implants [197]. Furthermore, several studies have shown that a minimum pore size of 100 µm is required for vascularization and bone ingrowth, but pores larger than 100 µm increase bone in-growth by allowing improved vascularization and oxygenation [86,91,193]. A minimum macropore size limit of 100 µm is supported by further research as vascular penetration has been found to be restricted in smaller pore interconnections [85,168,188,192].

Studies have found that pores greater than 300 µm are required for vascularisation and bone ingrowth [86,168]. Tang et al. [188] found that 200–350 µm is the optimal macropore size, and various studies have found that bone ingrowth is less likely to occur beyond 400 µm [90,195,198]. However, research that used pore sizes of 300, 600, and 900 µm in porous Ti scaffolds found that those with macropores sizes of 600 and 900 µm had much higher bone ingrowth compared with the scaffolds with 300 µm pores [101]. Bose et al. [85] suggested that all macropore sizes between 100 and 600 µm are osteoconductive. Fukuda et al. [199] experienced greater results in 500 and 600 µm pores compared with 900 and 1200 µm pores. Xue et al.'s [98] results showed that macropore sizes in the range of 100–600 µm possess the optimum ability for cell growth into the pore structure of porous titanium.

According to the FDA regulations, macropore sizes of 100–1000 um are approved for coatings for Ti implants [194]. Large macropores have a smaller surface area than do small pores, decreasing the cell attachment on the implant [86]. However, large macropores increase scaffold vascularisation, which is vital for supplying oxygen and nutrients to the tissue as well as osteoblast proliferation and migration [80], but they decrease the mechanical strength of the material. The limit of how much the macropore size can be increased while maintaining sufficient mechanical strength depends on both the material and the processing conditions. Therefore, regulatory guidelines for surface coating may be misleading for fully macroporous implants, since the strength-to-weight ratio differs between a porous and solid structure. Since the porosity decreases the strength of the implant, and large pore sizes decrease the strength of the internal architecture, large pore sizes must be avoided to increase the structure's strength. A more defined pore size range is therefore sought [86].

In this study, the macroporosity used in different studies was explored. It was found that no consensus currently exists on what upper limit to macropore size that is ideal, but somewhat of a consensus on the lower limit exists (100 µm). According to FDA regulations, porous implants should have macropore sizes between 100 and 1000 µm [194]. It has further been found that macropores start to lose their osteogenic functionality when larger than 500–600 µm [8,16,85,98,199]. Considering that a fully porous structure is weaker than a solid structure, and that high strength is vital for implants, it can be assumed that there is no need to design a structure with pores larger than what is needed to cater for all functions within the implant.

These findings made us choose a macropore size range of 100–600 µm as the IQA for porous metal implants to identify the studies in the systematic search that fabricated Ti scaffolds with pore size values within this range. From the 42 articles that recorded pore size of various degrees, a total of 144 results were extracted and compared with the selected IQA pore size range, as seen in Figure 10. According to our results, 51.4% of the results of all studies had a macropore size within 100–600 µm. It was further noted that 86.8% of the experimental results of all studies had a macropore size within the FDA recommended range of 100–1000 µm. From these results, we could infer that most of the research studies identified through the systematic search somewhat considered the macroporosity range required to satisfy medical regulations.

By contrast, a total of 68 results out 144 showed pore size values above the acceptable macropore size range representing 48.6% of the total results extracted in this systematic literature review. There were several reasons for these studies to explore pore sizes above 600 µm. For example, the FDA approves macropore sizes between 100 and 1000 µm [194]. Hara et al. [96] tested four porous structures with different macropore size to explore their mechanical properties. Taniguchi tested 300, 600, and 900 µm pore sizes and found that the structures with 600 and 900 µm pore size exhibited higher bone ingrowth. Large macropores increase scaffold vascularisation, which is vital for supplying oxygen and nutrients to the tissue as well as osteoblast proliferation and migration [80]; however, larger pores also have a

smaller surface area compared with small pores, decreasing the cell attachment on the implant [86]. Larger macropores also result in higher porosity, and this reduces the strength of the implant [187,188].

The ideal macropore size for bone implants is controversial and undefined, and according to Otsuki et al., a reason for the varied pore size data may be that the interconnectivity of pores was not considered [198]. Macropore size regulations have been developed for the first generation of porous implants, which use a single-scaled porous network with repeated, equally sized, and shaped pores. However, bone grows in a naturally random structure with pores of different sizes, shapes, and directions similar to the structure of a sponge [130,155]. The use of a multiscale porous scaffold that combines smaller and larger pore sizes within the same structure is a recent strategy to optimise the internal architecture of implants [142,188]. This method combines advantages of both small and large macropores without decreasing the strength or reducing the bone in-growth to levels that are inappropriate in application [188]. According to our results, a total of 8 out of 50 studies used some sort of multiscale pore approach, but it was observed that the design method varied, and that the researchers failed to provide the percentage of the total structure that used each pore size. It was further observed that a multiscale porous structure occurred in some implants where the manufacturing of a single-scaled structure resulted in varying pore sizes due to unprecise manufacturing tolerances.

#### 3.4.4. Pore Inter Connectivity

According to our results, 46% of the collected studies registered pore interconnectivity. Interestingly, all of these studies designed their porous scaffolds with an interconnectivity of 100%. The pores in a porous bone implant must be interconnected to ensure movement and the supply of necessary nutrients through ingrowth of tissue and bone [200]. Interconnected pores tend to facilitate the flow of fluids and biological cells through the structure which is essential for bone tissue formation [185]. According to Nyberg et al. [201] the integration of artificial material tissue with native tissue can be improved by interconnected pores. Tang et al. [188] suggested that an increased pore interconnectivity increases the number and size of blood vessels formed in scaffolds. The interconnectivity is also a critical factor for ensuring that all cells within the structure are within a 200 µm range from a blood supply to provide transfer of nutrients and oxygen [202]. According to the FDA's recommendations for porous metal coatings, pores in such structures must be interconnected [194]. Although this requirement is for surface coatings, it also indicates the importance of an interconnected porosity for fully porous implants. In the systematic review, it appears as though a vast majority of studies had indicated the importance of an interconnected porosity. Therefore, to guarantee all processes and fluid movements necessary for tissue and bone ingrowth, the selected IQA for pore interconnectivity would ideally be 100%.

#### 3.4.5. Elastic Modulus

It was observed in this systematic review that an elastic modulus is a property commonly reported in AM porous scaffolds studies (by 89% of all studies). A controlled modulus of elasticity has proved to be critical in prostheses and scaffolds to avoid stress shielding [81,82]. Stress shielding occurs when there is a stiffness mismatch between the implant and surrounding bone, and it can cause inflammation and the need for revision surgery [197]. Ti and common implant Ti alloys have an elastic modulus of roughly 100–120 GPa [81,84,138,197]. A reduced modulus is necessary to avoid stress shielding and can be achieved by designing implants with a porous structure [90].

Defining an ideal specific modulus of elasticity for porous bone implants is not practical because Since the mechanical properties of human bone, especially the elastic modulus, change drastically with factors, such as age, physical activity, and health. For example, femoral bone specimens from patients aged 3, 5, and 35 years had an elastic modulus of 7, 12.8, and 16.7 GPa, respectively, indicating a dramatic change with age [90]. As previously shown in Table 6, the elastic modulus of human bone varies in the literature. Chen and Thouas [57] estimated the elastic modulus of cortical bone to be approximately 11–21 GPa in the longitudinal direction, whereas Lee et al. [147] suggested 18–22 GPa. Wang et al. [81] suggested a wider range of 3–30 GPa. These findings indicate that the stiffness of an implant may need to be adjusted specifically for the person it is intended for, and that the target value for the elastic modulus may be specific to each patient. Therefore, it is more practical to think that for the design of porous scaffolds, an ideal target area exists for the modulus of elasticity. Based on this, the IQA for elastic modulus is proposed to be 3–30 GPa for fully porous Ti implants.

Figures 11 and 12 show the elastic modulus that was reported in the reviewed articles and these values were compared with the proposed IQA. From the extracted of elastic modulus results, 55.5% reached the target area of 3–30 GPa. These implants achieved an elastic modulus within the range of natural bone and would therefore eliminate risk of stress shielding. By contrast, 40% of the results exhibited an elastic modulus below 3 GPa, and only 3.6% of the results reported an elastic modulus higher than 30 GPa. These results clearly demonstrated that most studies are aiming towards a modulus of elasticity closer to the bone modulus.

**Figure 11.** Elastic modulus of porous metamaterials based on triply periodic minimal surfaces (TPMS) compared with the Ideal Quality Attribute (IQA) target zone.

The elastic modulus of metals such as titanium and its alloys naturally have a much higher elastic modulus compared with bone [81]. However, research shows that the elastic modulus of metals can be readily adjusted by modifying their porosity. Porous metals with a low modulus of elasticity correspond to high levels of porosity. For example, Wang et al. [84] explored five types of porous

structures using the same material (TiNbZr) and pore size (550 µm) but ranging porosity. His results revealed that four of the implants with porosities ranging from 42% to 69% all achieved an elastic modulus within the approved range of 3–30 GPa; however, the implant with the highest porosity (74%) achieved the lowest elastic modulus of 1.6 GPa. Similarly, Li et al. [203] used a porosity of 91% resulting in a low elastic modulus of 0.8 GPa. Furthermore, Chen et al. [204] received an elastic modulus of 44.4 GPa for a porous titanium structure using 30% porosity but by increasing the porosity to 40% the elastic modulus was reduced to 24.7 GPa.

Using the data obtained through the systematic literature search we calculated two multiple linear regressions to predict the modulus of elasticity of beam and TPMS-based AMd Ti scaffolds. The regression model used the independent variables of pore size, relative density (porosity), and the interaction of pore size–porosity. According to our results, a regression equation was found for beam-based AMd Ti scaffolds (F(3,75) = 54.139, *p* < 0.0001), with an R<sup>2</sup> adj of 0.671, as shown in Figure 13a. The residuals of the multiple linear regression are randomly scattered around the centre line of zero with no obvious pattern. The predicted compressive yield strength of beam-based scaffolds is equal to 27.738 – 0.078 (pore size) – 27.417 (porosity) + 0.0689 (pore size\*porosity), where pore size is coded or measured in µm, and relative density expressed as porosity as a percentage. The beam-based scaffolds' modulus of elasticity decreased 0.078 MPa for each µm, 27.417 MPa per 1% of porosity increment, and increased 0.0689 MPa for the interaction pore size\*porosity. Both pore size (*p* < 0.0005) and porosity (*p* < 0.0001) were significant predictors of beam-based scaffolds' modulus of elasticity, including the interaction between pore size and porosity (*p* < 0.0001).

**Figure 13.** Overall predicted model of elastic modulus actual versus an elastic modulus predicted (**a**) beam-based scaffolds, and (**b**) TPMS-based scaffolds.

In the case of the multiple linear regression of TPMS based AMd Ti scaffolds a significant regression equation was also found (F(3,28) = 4.897, *p* < 0.0073), with an R<sup>2</sup> adj of 0.273, as shown in Figure 13b. The residuals of the multiple linear regression are randomly scattered around the centre line of zero with no obvious pattern. The scaffolds' predicted modulus of elasticity is equal to 0.008 − 0.002 (pore size)–2.342 (porosity) + 0.002 (pore size\*porosity). The TPMS based scaffolds' modulus of elasticity decreased 0.002 MPa for each µm, 2.342 MPa per 1% of porosity increment, and 0.002 MPa for the interaction pore size\*porosity. Pore size was a significant predictor of TPMS-based scaffolds' modulus of elasticity with a *p*-values < 0.0563. However, porosity was not a significant predictor with *p*-values < 0.553 and 0.843, respectively. Moreover, no interaction between pore size and porosity was found regarding to modulus of the elasticity.

#### 3.4.6. Compressive Yield Strength

For an adequate functioning of any load-bearing implant, it is vital that its design withstand the required forces and loading cycles. Mechanical strength is one of the implant's most crucial features for avoiding implant failure. To withstand the loads of daily activities, load-bearing implants must have at least the same yield strength as the bone that they replace [81]. The yield point of bone represents the threshold from where the structure accumulates irreversible deformation. Unlike bulk metals such as steel, the yield point of bones cannot be clearly distinguished; it is rather associated with a continuous transition zone [160]. Strain beyond the yield point will deform the structure beyond its point of resilience causing material damage, usually occurring as micro-cracks [205]. Bone tissue has evolved to mainly support compressive stress [206,207]. Bone is 30% weaker under tensile stress, and 65% weaker under shear stress [208]. Therefore, load-bearing implant scaffolds require a high compressive strength to prevent fractures and improve functional stability [209]. The compressive yield strength of cortical bone varies in the literature. As previously shown in Table 6, the compressive yield strength of cortical bone varies approximately between 90 and 170 MPa. To replace like with like, using a biomimetic approach for comparison purposes, a minimum and a maximum compressive yield strength of 90 MPa and 170 MPa were selected as the IQA for fully porous Ti implants.

The systematic search identified that 37 out of 50 studies recorded compressive yield strength, from which a total of 133 experimental results were extracted and compared, as shown in Figures 14 and 15. Figure 14 presents the results of studies using TPMS structures and Figure 15 presents the results of studies using porous beam-based metamaterials. Both comparisons show high numbers of studies resulting in a compressive yield strength below 90 MPa. A total of 55.7% of all results had a compressive yield strength below the defined IQA target and 25% of the studies achieved a compressive yield strength within the bone region. On the other hand, only 19% of the extracted experimental results had a strength above the bone region. Such implants would have strengths similar to or higher than cortical bone and are expected to not experience permanent deformation caused by the expected bone compressive loading conditions in the human body.

**Figure 14.** Compressive yield strength of porous metamaterials based on triply periodic minimal surfaces (TPMS) compared with the Ideal Quality Attribute (IQA) target zone.

*Materials* **2020**, *13*, 4794

**Figure 15.** Compressive yield strength of porous beam-based metamaterials compared with the Ideal Quality Attribute (IQA) target zone.

Decreased strength of a porous implant can result from high porosity and large pore sizes, [81,188]. For example, Zhang et al. [113] fabricated porous scaffolds based on the TPMS diamond unit cell with a wide range of compressive yield strengths from 36 MPa to 140 MPa just by varying the scaffolds' porosity and maintaining the pore size constant. The type of unit cell used to design porous scaffolds can also drastically change it mechanical properties. For example, Zhao et al. [116] fabricated four porous scaffolds with the same pore size and similar porosities using two different unit cells (tetrahedron and octahedron). However, the scaffolds based on the octahedron unit cell registered almost double the compressive strength compared with those based on the tetrahedron unit cell [116]. The compressive yield strength of porous metals can also be enhanced by gradually changing the porosity level along the radial direction of the scaffold. This was demonstrated by Zhang et al. [100], who reported functionally graded porous scaffolds based on the diamond unit cell with superior comprehensive mechanical properties to the biomaterials with uniform porous structures.

Using the data obtained through the systematic literature search, we calculated two multiple linear regressions to predict compressive yield strength based on pore size, porosity, and the interaction of size–porosity for beam-based and TPMS-based AMd Ti scaffolds, respectively.

According to our results, a regression equation was found for beam-based AMd Ti scaffolds (F(3,75) = 31.452, *p* < 0.0001), with an R<sup>2</sup> adj of 0.539, as shown in Figure 16a. The residuals of the multiple linear regression are randomly scattered around the centre line of zero, with no obvious pattern. The scaffolds' predicted compressive yield strength is equal to 380.557 − 0.075(pore size)–350.828 (porosity) + 0.557 (pore size\*porosity), where pore size is coded or measured in µm, and porosity is measured as a percentage. The beam-based scaffolds' compressive yield strength decreased 0.075 MPa for each µm, 350.828 MPa per 1% of porosity increment, and increased 0.557 MPa for the interaction of pore size\*porosity. Both pore size (*p* < 0.0378), porosity (*p* < 0.001), and the interaction between pore size and porosity (*p* < 0.0048) were significant predictors of beam-based scaffolds' compressive yield strength. The interaction between pore size and porosity was found to be significant with a *p*-value < 0.0048.

b b

**Figure 16.** Overall predicted model of compressive yield strength Actual versus compressive yield strength Predicted of (**a**) beam-based scaffolds, and (**b**) TPMS-based scaffolds.

Regarding the multiple linear regression of TPMS-based AMd Ti scaffolds, a significant regression equation was also found (F(3,27) = 65.547, *p* < 0.0001), with an R <sup>2</sup> adj of 0.872, as shown in Figure 16b. The residuals of the multiple linear regression are randomly scattered around the centre line of zero, with no obvious pattern. The scaffolds' predicted compressive yield strength is equal to 524.780 − 0.008(pore size)–625.266 (porosity) + 0.183 (pore size\*porosity). The TPMS-based scaffolds' compressive yield strength decreased 0.008 MPa for each µm, 625.266 MPa per each 1% of porosity increment, and increased 0.183 MPa for the interaction of pore size\*porosity. Porosity was a significant predictor of TPMS based scaffolds' compressive yield strength with *p*-values < 0.0001. However, pore size and the interaction between pore size and porosity were non-significant, with *p*-values < 0.7018 and 0.3260, respectively.

#### 3.4.7. Ultimate Compressive Strength

If loading surpasses the yield point of bone, it will eventually reach the ultimate point. This point represents the maximum compressive strength that a material can withstand without irreversible strains and damage occurring. Past this point, macrocracks are formed and fracture occurs [160]. Bone implants in load-bearing applications must withstand high stress within the body, to a degree where no permanent deformation occurs during the load that the implant is expected to be exposed to. Hence, controlled ultimate compressive strength is a crucial property to study in bone implant research. Natural bone is estimated to have an ultimate compressive strength of 180–200 MPa [84]. However, results vary in the literature. For example, Calori et al. [8] suggested a more widespread range of 130–290 MPa, whereas Henkel et al. suggested 100–230 MPa [168]. To replace like with like using a biometric approach, bone implants should have an ultimate compressive strength similar to that of bone [8]. Taking into consideration the compressive yield strength suggested previously and the three results presented in Table 5, the proposed IQA region for the ultimate compressive strength is between 180 MPa and 290 MPa.

In this systematic search a total of 60 experimental results of ultimate compressive strength from 19 different studies were extracted. Figures 17 and 18 show the ultimate compressive strength of the different studies compared with the defined IQA target of between 180 MPa and 290MPa. Figure 17 corresponds to experimental results of porous scaffolds composed of TPMS unit cells compared with the IQA target. According to Figure 17, only one study with three different experimental results measured the ultimate strength of porous scaffolds based on TPMS unit cells. In this study by Yanez et al. [120], three ultimate compressive strengths of 17, 47.5, and 83.5 MPa were achieved using the gyroid unit cell. Dramatic improvement in the ultimate compressive strength of the scaffold was achieved. This improvement in mechanical properties was possible by slightly changing the gyroid unit cell into an elongated gyroid. However, none of the experimental results obtained by Yanez

et al. [120] were able to reach the minimum IQA ultimate compressive strength proposed in this study. The low strength of Yanez at al.'s [120] samples can be attributed to their high porosity values which ranged between 75% and 90%. This increased the stress concentration, reducing the ultimate compressive strength.

**Figure 17.** Ultimate compressive strength of porous metamaterials based on triply periodic minimal surfaces (TPMS) compared with the Ideal Quality Attribute (IQA) target zone.

**Figure 18.** Ultimate compressive strength of porous beam-based metamaterials compared with the Ideal Quality Attribute (IQA) target zone.

Figure 18 is composed of 19 different studies on beam-based unit cells where 57 experimental results are compared with the IQA target. According to Figure 18, 10 experimental results reached higher ultimate compressive strengths than bone. The highest ultimate compressive strength (830 MPa) recorded was achieved with a scaffold based on the diamond unit cell by Zhang et al. [100]. From all results recorded, a total of 16.7% achieved an ultimate compressive strength above the proposed IQA ultimate compressive strength; 10% had similar ultimate compressive strengths to bone; and 73.3% had lower ultimate compressive strengths than bone. The scaffolds that did not fulfil the required IQA ultimate compressive strength would risk fracturing due to macrocracks occurring during high loads. Porosity within a structure has been proven to decrease the strength of a structure [81,188], which may explain the high number of scaffolds with low ultimate strength.

The goal of designing a porous structure with enough porosity and pore size without diminishing strength is a difficult task, and as researchers aim to create highly porous structures with low elastic modulus, many structures experienced low ultimate strength. Attar et al. [111] manufactured a porous titanium structure by SLM with rectangular pores and 17% porosity and achieved an ultimate compressive strength of 747 MPa. Using the same material, manufacturing method, and unit cell, a different structure with 37% porosity achieved an ultimate strength of 235 MPa. In a similar manner, Chen et al. [204] designed three structures of the same material and manufacturing method. Using porosities of 30%, 40%, and 50%, the ultimate compressive strengths recorded were 524, 301.7, and 120.3 MPa, respectively, indicating that increased porosity reduces the strength of the structure.

#### 3.4.8. Fatigue Strength

During normal daily activities, load-bearing implants experience just a fraction of the material's ultimate stress [209–211]. However, after years of use, the high cyclic loading to which load-bearing implants are subjected eventually leads to the accumulation of small stresses, causing progressive and localised material damage that results in implant failure [212]. For instance, one of the most critical mechanical properties for load-bearing implants is fatigue strength. However, fatigue strength is the most difficult mechanical property to determine [213].

The required fatigue resistance of a load bearing implant and its components mainly depends on their cyclic loading conditions and the required life span. For example, it is estimated that lower limb prostheses are subjected to up to 2 million gait cycles per year [214], and in the case of orthodontic prostheses these can reach up to 300,000 loading cycles per year [213]. Therefore, a large variety of medical standards exist for testing fatigue strength. Some of these fatigue tests differ depending on the type of load applied such as tension–tension, compression–compression, and tension–compression. In the case of load-bearing bones, their loading conditions in real-life activities are complex [135]. However, bone is mainly loaded in compression [206,207]. Therefore, to test the fatigue life of metamaterials for bone implant applications, compression fatigue tests are preferred due to the simplicity of the test setups [215].

Regarding the number of cycles that load bearing implants and their components need to have tested for fatigue strength, all the different medical standards agree that such products need to have a fatigue life within the high-cycle fatigue region (N > 10<sup>4</sup> cycles). For example, the ASTM standard F2777 – 16 recommends testing tibial inserts' endurance and deformation under high flexion with a minimum number of cycles of 2.2 <sup>×</sup> <sup>10</sup><sup>5</sup> , and in the case of dental implants they are typically tested up to 5 million cycles [213,216]. Nevertheless, for a component of a load-bearing implant to have at least 25 years of life span [217], the highest number of cycles that must be tested is 10<sup>7</sup> cycles [213,216]. Taking into consideration current medical standards for load-bearing implants, the high-cycle fatigue region between 10<sup>4</sup> cycles and 10<sup>7</sup> cycles was selected as the IQA fatigue life for porous titanium metamaterials for bone regeneration.

Using the selected high-cycle fatigue region, this systematic literature search identified a total of 13 different studies on fatigue resistance, among which 11 studies performed compression–compression fatigue tests. Then, for comparison purposes, a total of 51 experimental results were extracted and compared. Moreover, to facilitate the comparison of the results of the studies, they were classified according to the type of unit cell used to produce porous structures as beam and TPMS-based as resented in Figures 19–21. According to our results, the TPMS porous structures that withstood the highest stresses at the high-cycle fatigue region were achieved by Bobbert et al. [79]. The primitive TPMS structure presented the highest stress within the high-cycle fatigue region, with 232 MPa at <sup>3</sup> <sup>×</sup> <sup>10</sup><sup>4</sup> cycles, as shown in Figure 19. The TPMS porous structures that were able to withstand the second and third highest stresses within the high-cycle fatigue region were the I-WP and diamond structures, with 227 MPa at 3 <sup>×</sup> <sup>10</sup><sup>5</sup> cycles and 204 MPa at 3 <sup>×</sup> <sup>10</sup><sup>6</sup> cycles as shown in Figures 19 and 20. Remarkably, the primitive TPMS structure was the only one to pass the 10<sup>7</sup> threshold with 80 MPa at <sup>3</sup> <sup>×</sup> <sup>10</sup><sup>7</sup> cycles, as presented in Figure 20.

**Figure 19.** Comparison of the results of different studies on the fatigue strength of porous metamaterials based on triply periodic minimal surfaces (TPMS) within the high-cycle fatigue region between 10<sup>4</sup> and 10<sup>5</sup> cycles.

**Figure 20.** Comparison of the results of different studies on the fatigue strength of porous metamaterials based on triply periodic minimal surfaces (TPMS) within the high-cycle fatigue region between 10<sup>6</sup> and 10<sup>7</sup> cycles.

**Figure 21.** Comparison of the results of different studies on the fatigue strength of porous beam-based metamaterials within the high-cycle fatigue region between 10<sup>6</sup> and 10<sup>7</sup> cycles.

In the case of beam-based metamaterials the study performed by Zhao et al. [116] achieved the highest fatigue strength with 130 MPa at 10<sup>6</sup> cycles using a lattice structure based on the octahedron unit cell, as shown in Figure 21. In second and third place with the highest fatigue strength are the tetrahedron and the cubic porous lattice structures with 90 MPa at 1 <sup>×</sup> <sup>10</sup><sup>6</sup> cycles and 90 MPa at 1 <sup>×</sup> <sup>10</sup><sup>6</sup> cycles fabricated by Amin Yavari et al. [106] and Zhao et al. [116], as presented in Figure 21.

According to the results of this systematic search it could be seen that TPMS structures provide superior fatigue strength to porous bio-metamaterials compared with beam-based unit cells. Moreover, in terms of fatigue resistance, it was identified that the primitive, I-WP, and diamond TPMS provided the best performance, whereas the octahedron, cubic, and tetrahedron are the best-performing lattice unit cells. However, it is crucial to note that several factors can affect the fatigue life of additively manufactured Ti metamaterials. Some of these factors are residual stresses and stress concentrators caused by high surface roughness and manufacturing defects [218]. Moreover, the fatigue strength of bulk materials significantly degrades when they are in porous form or when voids and pores are developed during fabrication [219]. In the case of additively manufactured components, it has been proven that their fatigue strength is extremely sensitive to localised and nonuniform heat and uncontrolled cooling cycles during fabrication [220,221].

#### *3.5. Discussion and Summary of the Findings*

Table 7 presents a summary of proposed IQA target values for porous Ti and Ti alloy bone implants aimed at load-bearing applications. These values are based on scientifically supported values found in human bone research, federal regulations (FDA), and research articles on porous Ti implants manufactured using AM between 2000 and 2020. These properties are part of the first step of the QbD framework of "Define the Quality Target Product Profile." The IQAs are necessary for a systematic and qualitative design approach. These properties will provide benchmark guidance to facilitate future research on porous bone implant design.

**Table 7.** Summary of Ideal Quality Attributes proposed in this study.

An IQA target zone for the porosity of porous Ti implants has been proposed as 30–70% based on results found in research articles as well as in current FDA regulations. We found that 56.6% of all studies in this review achieved a porosity within this range. AM was found to produce porous structures with highly controlled porosity. Numerous studies as well as FDA regulations have discussed the importance of using porous structures for tissue ingrowth, which have numerous advantages to non-porous implants. Whereas porous coatings are used in some instances, it was observed in this study that many researchers believe in using a fully porous structure to achieve a biomimetic structure mimicking natural bone. All data used extracted in this systematic literature review came from studies that used scaffolds with constant porosity. However whether a repeated lattice structure throughout the entire structure is sufficient or a biomimetic "sponge-like" structure with irregular, elongated pores should be used has not yet been confirmed in research; nevertheless implants using a biomimetic design approach replicating natural bone received attention in recent research [86]. Since bone has a random and stochastic structure with pores of different sizes, a multiscale porous structure with pores of different sizes deserves further research. Multiscale porous scaffold structures have been proven to perform better than one-dimensional structures [142]. Nonetheless, to properly compare implants with a multiscale structure, pore distribution and size must be recorded.

In the case of scaffolds' pore size, the IQA was proposed to be 100–600 µm based on numerous performed research trials. Whereas the FDA requirement is a pore size between 100–1000 µm, a more defined pore size range of 100–600 µm is supported by research due to numerous research trials having experienced reduced bone ingrowth in pores larger than 600 µm. Strut thickness is the thickness of the pore walls within a porous structure. It may have a substantial impact on implant mechanical properties [90] and can serve as a unit cell's characteristic to compare different mechanical properties in relation to it. However, it was found that struct thickness is not commonly reported in the literature.

Research has proven that the elastic modulus of bone changes drastically with age and that patient customisation is a necessity in specific cases. As a result, in this study, an IQA for elastic modulus was proposed to be 3–30 GPa based on values from research studies. According to data analysed in this study a porous structure has a significant influence on mechanical properties in Ti-based implants. Therefore, a porous structure can be altered to provide an elastic modulus comparable to that of natural bone and be adjusted to modify the elastic modulus according to the patient's age and health condition. This will further reduce the risk of stress shielding between the implant and surrounding tissue.

As reported by most studies it was found in the regression analysis that the elastic modulus of Ti and its alloys is directly influenced by the implant porosity, and that by increasing the porosity the elastic modulus is increased and vice versa. A total of 55.5% of the studies in this review recorded an elastic modulus within this range. Regarding the effect of unit cells on the modulus of elasticity of porous structures, it was found that the studies that used beam-based unit cells covered a wider range of modulus of elasticity values than did those than used TPMS-based unit cells (Figure 22a). Moreover, the studies that used beam-based unit cells also obtained modulus of elasticity values within the whole range of human bone Young's modulus, as shown in Figure 22a.

**Figure 22.** Comparison in logarithmic scale between mechanical properties of bone, beam-based, and triply periodic minimal surfaces (TPMS): (**a**) modulus of elasticity; (**b**) compressive yield strength; and (**c**) ultimate compressive strength.

Compressive yield and ultimate strength were studied. The compressive strength values for cortical bone were found to vary in the literature, which is why a range of values was selected to represent this bone characteristic. The IQA for compressive yield strength was proposed to be a minimum of 90 MPa and the IQA target for ultimate compressive strength a minimum of 180 MPa. A total of 44.1% of the studies achieved a compressive yield strength above 90 MPa, whereas only

26.7% achieved an ultimate compressive strength above 180 MPa. Both TPMS-based and beam-based scaffolds provided a wide range of compressive yield strengths covering the whole range of human bone compressive yield strength, as presented in Figure 22b. However, according to Figure 22c only beam-based scaffolds provided an adequate range of ultimate compressive yield strength values that include bone properties. The narrow range of ultimate compressive yield strength values of TPMS-based scaffolds can be explained by the lower number of studies that have addressed this mechanical with this type of unit cell. This indicates that a need exists for TPMS-based scaffold research to report the ultimate compressive yield strength.

Just as with the elastic modulus, the strength of implants was found to be directly affected by scaffold porosity. Increased porosity was found to rather dramatically decrease the strength of the implant as demonstrated by the regression model performed with the results of several different studies. Moreover, it was found that large pores are directly related to lowered strength and elastic modulus, and that several studies have struggled to produce highly porous structures with strength higher than or close to natural bone. Therefore, to ensure the adequate strength of implants, pores larger than required (>600 µm) and porosity values higher than necessary (>70%) are recommended to be avoided as these have been found to likely result in low-strength structures.

#### **4. Conclusions**

This systematic literature review presented an overview of the reported properties in research studies of fully porous Ti bone implants manufactured with AM received in the last two decades. The study focused on implants made of Ti and Ti alloy since they are considered ideal biomaterials for load-bearing applications. This study followed a QbD approach and includes the first step of the QbD system that defines the QTPP for properties relating to the porous internal architecture of fully porous Ti implants designed for load bearing applications. IQA, part of the QbD system, have been proposed supported by properties of natural human bone, governmental regulations, and scientific research relevant to bone implants. Unit cell geometry, porosity, elastic modulus, compressive yield strength, ultimate compressive yield strength, and compressive fatigue strength were systematically reviewed and benchmarked against the proposed IQA.

This study found that many implant geometrical, mechanical, and dimensional characteristics were directly related to each other. Scaffolds' pore size influences the porosity of the structure, and the porosity alters the elastic modulus as well as the strength of the implant. The unit cell geometry was also found to directly affect the Young's modulus and strength of porous scaffolds, and naturally would also impact the structures' interconnectivity. Moreover, by using ranges rather than set values where possible, such as for elastic modulus, porosity, and pore size, there is a flexibility to the design that allows the implant to be adjusted to its purpose and patient. Such design flexibility is necessary as bone properties vary with patients' age and anatomic site.

Despite the variety of scaffold characteristics reviewed in this study, future systematic literature searches should also focus on other properties, such as fluid dynamics, surface finish (topography), creep, and hardness, as well as surface coatings. Moreover, it should be considered that bone ingrowth may modify implants' mechanical properties; hence, they may be dependent on the level of bone ingrowth. Therefore, it is important to measure the changes in strength and stiffness in metallic bone scaffolds after bone ingrowth has occurred. Since implant location may determine the importance of each implant property, future studies would ideally find IQA for different implant locations within the human body. For example, mechanical features are important to study in load bearing implants, whereas fluid dynamics and vascularisation might be more important features for facial implants that require good aesthetic results and that are placed close to fragile tissue and nerves.

This study was possible due to the abundance of data available in research articles. However, to further develop an effective QbD engineering strategy for the development of specific bone implants it is important to identify the degree of importance of each bone scaffold characteristic according to the implant's future location within the human body. This can be identified by following

the second step in the QbD system, namely Critical Quality Attributes (CQA), as detailed described in previous work [222]. CQA are product characteristics that must fall within specific limits to comply with the quality standards defined in the QTPP. They can be identified through prior knowledge and experimental data from systematic research based on scientific and risk management rationale that considers regulatory and business requirements [223]. Each step of the QbD system must be further developed to facilitate the strategic, qualitative development of bone implants and increase the rate of successful research studies.

Overall, Ti and Ti alloy porous bone implants are well underway to achieving the ideal properties that will fully allow them to replace natural bone. With the help of the QbD system, consistent and qualitative design of medical implant devices can be achieved.

**Author Contributions:** Conceptualisation: D.M.-M., Y.D., and R.A.S.; Formal analysis: D.M.-M. and Y.D.; Funding acquisition: R.A.S.; Investigation: D.M.-M. and Y.D.; Methodology: D.M.-M., Y.D., S.S., and R.A.S.; Project administration: D.M.-M. and R.A.S.; Resources: D.M.-M. and Y.D.; Supervision: D.M.-M., S.S., and R.A.S.; Visualisation: D.M.-M. and Y.D.; Writing—original draft: D.M.-M. and Y.D.; Writing—review and editing: D.M.-M., S.S., and R.A.S. All authors have read and agreed to the published version of the manuscript.

**Funding:** This research received no external funding.

**Acknowledgments:** This work was supported by Griffith School of Engineering and Built Environment.

**Conflicts of Interest:** The authors declare no conflict of interest.

#### **References**

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### *Review* **Bone Morphogenetic Proteins, Carriers, and Animal Models in the Development of Novel Bone Regenerative Therapies**

**Nikola Stokovic 1 , Natalia Ivanjko 1 , Drazen Maticic 2 , Frank P. Luyten <sup>3</sup> and Slobodan Vukicevic 1, \***

**Abstract:** Bone morphogenetic proteins (BMPs) possess a unique ability to induce new bone formation. Numerous preclinical studies have been conducted to develop novel, BMP-based osteoinductive devices for the management of segmental bone defects and posterolateral spinal fusion (PLF). In these studies, BMPs were combined with a broad range of carriers (natural and synthetic polymers, inorganic materials, and their combinations) and tested in various models in mice, rats, rabbits, dogs, sheep, and non-human primates. In this review, we summarized bone regeneration strategies and animal models used for the initial, intermediate, and advanced evaluation of promising therapeutical solutions for new bone formation and repair. Moreover, in this review, we discuss basic aspects to be considered when planning animal experiments, including anatomical characteristics of the species used, appropriate BMP dosing, duration of the observation period, and sample size.

**Keywords:** animal model; bone fracture; bone healing; posterolateral spinal fusion; regenerative medicine; bone morphogenetic proteins

#### **1. Introduction**

Bone tissue possesses unique regenerative properties, and bone fractures regularly heal under physiological conditions. However, large segmental bone defects resulting from severe trauma or extensive tumor resection cannot be restored by endogenous selfrepair mechanisms, decrease quality of life, and may sometimes lead to limb amputation. Indeed, the management of large segmental defects is one of the most challenging issues in orthopedic medicine, typically due to the biologically hampered microenvironment [1,2]. The standard of care for the healing of large bone defects requires the use of an autologous bone graft (ABG), which is usually harvested from the iliac crest. ABG is also used as a gold standard to achieve spinal fusions, including posterolateral spinal fusion (PLF). PLF is a commonly performed surgical procedure used for the treatment of degenerative diseases of the spine, including degenerative disc disease, spondylolisthesis, spinal instability, and symptomatic scoliosis [3–6].

However, ABG possesses several disadvantages, including a limited amount of bone that might be harvested, the potential transfer of contaminating agents, acute and chronic pain, skin scarring, and deformity at the donor site [4,7]. In addition, the use of ABG increases the blood loss, duration, and cost of surgical procedures. Therefore, there remains an imminent need for the development of novel bone regeneration strategies to enrich or replace ABG. Among these, osteoinductive devices are under investigation for clinical use in PLF and healing of large segmental long bone defects.

In the last few decades, numerous preclinical studies using animal models have been conducted to test novel bone bridging or fusion strategies [8,9]. The principles of

**Citation:** Stokovic, N.; Ivanjko, N.; Maticic, D.; Luyten, F.P.; Vukicevic, S. Bone Morphogenetic Proteins, Carriers, and Animal Models in the Development of Novel Bone Regenerative Therapies. *Materials* **2021**, *14*, 3513. https://doi.org/ 10.3390/ma14133513

Academic Editor: Antonio Boccaccio

Received: 18 May 2021 Accepted: 21 June 2021 Published: 24 June 2021

**Publisher's Note:** MDPI stays neutral with regard to jurisdictional claims in published maps and institutional affiliations.

**Copyright:** © 2021 by the authors. Licensee MDPI, Basel, Switzerland. This article is an open access article distributed under the terms and conditions of the Creative Commons Attribution (CC BY) license (https:// creativecommons.org/licenses/by/ 4.0/).

the rational use of animal models in the evaluation of novel bone regenerative therapies have been previously described [8]. Hence, we further investigated the use of animal models in the development of osteoinductive therapies of large segmental bone defects and PLF procedures, in particular the selection of a proper anatomical model, treatment dose, observation period, and sample size. We specifically analyzed published in vivo studies looking into the development of bone morphogenetic protein (BMP)-based bone inducing implants.

#### **2. Bone Regeneration by Bone Morphogenetic Protein Devices**

BMPs are well-known osteoinductive molecules, required and sufficient for ectopic bone induction, and powerful agents for the restoration of large orthotopic bone defects [10]. BMP2 is the most widely used osteoinductive BMP in preclinical testing, and it is a part of an osteoinductive device (Infuse™, Medtronic, Dublin, Ireland), currently approved for anterior lumbar interbody fusion (ALIF), acute tibial fractures, and maxillofacial reconstructions [11–15]. However, BMPs have been used off- label in various spinal indications, including cervical spine fusion, posterior lumbar interbody fusion (PLIF), transforaminal lumbar interbody fusion (TLIF), posterolateral spinal fusion (PLF), and thoracolumbar fusions [16,17]. Reported side effects in patients included implant displacement, infection, swelling of the adjacent tissue and dysphagia, formations of seroma, radiculitis and nerve root compression, ectopic bone formation, osteolysis, and retrograde ejaculation [11,12,16–21]. These side effects eventually resulted from the use of supraphysiological doses as registered BMP2-based devices contain 4–12 mg recombinant protein, while the human body contains only a total of 2 mg of BMPs [22].

Other commonly used osteoinductive BMPs are BMP7, which is no longer in clinical use, and more recently, BMP6 [23]. We demonstrated that BMP6 appears to be superior to BMP2 and BMP7 in promoting osteoblast differentiation in vitro and inducing bone formation in vivo [23,24]. The superiority of BMP6 may arise from its resistance to noggin inhibition and affinity across the BMP type I receptors. Therefore, BMP6-based devices are expected to be more efficacious at lower doses compared to BMP2 and BMP7.

BMPs require a carrier/delivery system that will sustain the BMP concentration and allow prolonged BMP release [25–28]. Moreover, the ideal BMP carrier should be biocompatible, enable vascular and cellular infiltration, resist compression, and define the contours of the resulting bone [25,26,29]. BMP carriers can be divided into four major groups: natural polymers, synthetic polymers, inorganic materials, and combinations between these groups [25,26,30].

Natural polymers include collagen, hyaluronic acid, gelatin, fibrin, chitosan, alginate, and silk and have been extensively evaluated in preclinical studies [6,31–38]. The advantages of natural polymers are biocompatibility, biodegradability, and resorbability in the physiological environment [25–27]. The most commonly used is bovine tendon collagen which delivers BMP2 in the clinically approved Infuse™ device. However, collagen has significant disadvantages, including a low affinity for BMPs, immunogenicity due to its animal origin, and weak biomechanical properties resulting in compression by surrounding tissues [25–27].

Biocompatible and biodegradable synthetic polymers, such as polylactic acid (PLA), polyglycolic acid (PGA), poly(D, L-lactide-co-glycolide) (PLGA), polyethylene glycol (PEG), poly-E-caprolactone (PCL), and polypropylene fumarate (PPF), as well as their block polymers have been evaluated as potential BMP carriers to overcome the disadvantages of natural polymers, including immunogenicity and disease transmission risk [39–42]. They are also moldable into highly porous three-dimensional scaffolds, linearly oriented scaffolds, fibers, sheets, blocks, or microspheres [26]. Apart from these advantages, synthetic polymers decrease local pH as a result of acidic breakdown byproducts, have poor clearance, cause bulk degradation, and cause chronic inflammation associated with high-molecular weight polymers, resulting in substantial disadvantages [26]. They have also been tested with other potentially osteogenic molecules, such as PGE2 and PGE4 prostaglandin receptor analogs [43], and materials such as calcium silicate (CaSi) and dicalcium phosphate dihydrate (DCPD) [44–47].

Inorganic materials as potential BMP carriers include calcium phosphate (CaP) ceramics, calcium phosphate and calcium sulfate cement, and bioglass [2,5,29,32,38,42,48–72]. The most commonly used inorganic preclinical materials are CaP ceramics, further subdivided into hydroxyapatite (HA), tricalcium phosphate (TCP), and biphasic calcium phosphate (BCP) containing both HA and TCP at various ratios. We have recently shown that the chemical composition of ceramics does not affect the amount of newly formed bone induced by the osteoinductive device [73,74]. However, HA and TCP significantly differ in resorbability (HA is very stable, while TCP is more resorbable), which would eventually result in different residual ceramic volumes. The resorbability might be adjusted by varying HA/TCP ratios in BCP ceramics [75]. Moreover, CaP ceramics might be formulated into particles or blocks in a broad range of sizes and geometrical shapes while porosity, pore size, and interconnectivity are adjusted during the sintering process [73,75,76]. We demonstrated that particle size affects the volume of newly formed bone; smaller particles (74–420 µm) combined with rhBMP6 resulted in higher bone volume than larger particles (1000–4000 µm) [73]. Another important determinant of ceramics is the pore size since pores from 300 to 400 µm promoted the formation of the largest bone volume [51].

The fourth group of BMP carriers are composites of the aforementioned materials which have been introduced to overcome the encountered limitations of a single component. The most typical combinations are composites containing either natural or synthetic polymers with CaP ceramics [39,77–85]. In these combinations, ceramics increase the biomechanical properties of the implants and are used to address compressibility issues. Less frequent, natural, and synthetic polymers might be combined.

We have recently developed an autologous bone graft substitute (ABGS) comprised of BMP6 delivered within an autologous blood coagulum to which a compression-resistant matrix, such as allograft or synthetic ceramics, can be added [22,73,74,76,86–92]. Moreover, the volume of newly induced bone increased with the elevation of the CRM amount, which might be attributed to the enlargement in an overall surface area [73].

#### **3. Animal Models**

Animal models are routinely used in the development of novel bone regenerative therapies [8]. Models might be categorized according to the species (mouse, rat, rabbit, sheep, non-human primate) and tested indication (ectopic model, critical-size defect, PLF). In this review, we suggested classification based on the stage of preclinical development, namely as initial, intermediate, and advanced testing of osteoinductive devices (Figure 1). Initial testing includes rodent ectopic and rodent critical-size defect models for rapid comparison of different osteoinductive responses. Intermediate evaluation includes adequate rabbit models (segmental defect and PLF model), while advanced testing uses canine, sheep, and non-human primates as a final step before clinical trials.

#### *3.1. Initial Evaluation in Rodents*

#### 3.1.1. Ectopic Models

Rodent ectopic models have been extensively used for the initial evaluation of novel osteoinductive therapies. They might be also used for investigating the biology of ectopic bone induction and the formation of a bone organ or ossicle, including bone and bone marrow [31,32,39,48–57,71,73,76,86,87,93–105]. Rodent ectopic models (Tables 1 and 2) are further subdivided according to the species (mouse, rat) and the implantation site (subcutaneous or intramuscular). Implantation under the skin (Figure 2A–D) or into the muscle does not affect the bone formation outcome, and the bone formation occurs in the first two weeks following implantation of an osteoinductive device [76,86,87]. The later time points are needed for the evaluation of the bone longevity and maintenance of the ectopic bone structure. Molecular and cellular events during the cascade of bone formation can be evaluated using microCT/nanoCT and histological analyses. Immunohistochemistry,

flow cytometry, gene profile microarrays, and single-cell RNA sequencing are among other analytical techniques used for unraveling the mechanism of ectopic bone formation.

**Figure 1.** Classification of animal models from mouse to NHP based on the stage of preclinical development, namely as initial, intermediate, and advanced evaluation of osteoinductive devices. E—ectopic, SD—segmental defect, PLF posterolateral spinal fusion models, and L—level of lumbar transverse processes.

#### **Table 1.** Mouse ectopic and bone defect models.

#### **Table 2.** Rat ectopic and bone defect models.

**Figure 2.** Histological sections and gross anatomy of newly formed bone induced by rhBMP6 in a rat subcutaneous assay (**A**–**D**); rabbit (**E**,**F**) and sheep (**G**,**H**) posterolateral spinal fusion model. (**A**,**B**) On day 7 following implantation, endochondral bone formation occurs at the peripheral site of ABGS, while 14 days (**C**,**D**) after implantation, newly formed bone is present throughout the implant containing rhBMP6 lyophilized on allograft mixed with ABC. Gross anatomy of newly formed bone between transverse processes in rabbit (**E**) and sheep (**G**) PLF model. Histological sections through fusion mass in rabbit (**F**) and sheep (**H**) PLF model. Histological sections were processed undecalcified and stained by Von Kossa (**A**,**B**), Goldner (**C**,**D**,**F**), and hematoxylin and eosin (**H**). (Modified from [74,88], respectively.)

Murine models are initially used to unravel the potential mechanism of action of various signaling pathways and genes or proteins in bone induction or enhancement. However, due to limited translation of mouse to human bone regeneration and disease outcome, the rat is therefore a more suitable model when testing functional outcomes [106–108].

#### 3.1.2. Bone Defect Models

Mouse or rat bone defects are the initial orthotopic models to evaluate the osteoinductive properties of novel therapies and the osseointegration of newly formed bone with adjacent native bone. There are two main bone defect models in rodents: a calvarial critical-

size defect and segmental femoral defect. In the calvarial critical-size defect, circular bone defects are created in the mouse (3–5 mm) [109–118] and rat (4–8 mm) [29,55,58–60,66,77] calvaria (Figure 3A; Tables 1 and 2), while segmental defects of the long bone are typically created in the femur, both in mice (2–3 mm) [119–121] and rats (6–10 mm) [5,33–35,67,78,79] (Figure 3B; Tables 1 and 2) and filled with tested osteoinductive material. The development of a reproducible non-union model in the mouse is demanding, and, in contrast to rat non-union models, mouse non-union models are sparse [122]. The main shortcoming of this model is a relatively small defect size compared to clinically relevant proportions. Moreover, it is difficult to obtain a full stabilization of the fracture, therefore resulting in increased callus formation. Methods of evaluation include analyses of radiological images (CT/microCT), histological and histomorphometric analyses, and biomechanical testing, which might be conducted at the end of the observation period (Figure 4) [5,29,33–35,55,58–60,66,67,77–79,109–121,123].

**Figure 3.** First row: segmental defect (SD) models performed on various bones depending on the species; (**A**) calvarial, (**B**) femoral, (**C**) ulnar (or radial), and (**D**) tibial defect. Second row: posterolateral lumbar fusion (PLF) is conducted between adjacent transverse processes; the figure shows differences in the anatomy of rabbit (**E**), sheep (**F**), and human (**G**) lumbar spine.

**Figure 4.** BMP doses (µg), observation period (weeks), and sample size (*n*) used in the initial phase —rodent models. Data is presented through box and whisker plots in which dots represent all individual values from studies listed in Tables 1 and 2. E—ectopic model, C—calvarial, and F—femoral defect.

#### *3.2. Intermediate Evaluation in Rabbits*

#### 3.2.1. Segmental Defect Model

Potential therapeutical solutions for segmental bone defect restoration have been extensively evaluated in rabbits [36,41,42,61,68–70,80–84,87,124–126] (Table 3), and a defect can be created in the femur, radius, or ulna (Figure 3B,C). Regardless of the chosen anatomical site, the typical defect size is 15–20 mm. In published work, the defect was bridged with a broad range of delivery systems containing up to 150 µg of BMPs. The observation

period was typically 6–12 weeks. Few studies evaluated bone formation at earlier time points (2 and 4 weeks) or for a prolonged period (24 weeks) (Figure 5, 1st column).

**Figure 5.** BMP doses (µg), observation period (weeks), and sample size (*n*) used in the intermediate and advanced phases, which include rabbit, dog, sheep, and non-human primate models. Data are presented through box and whisker plots in which dots represent all individual values from studies listed in Tables 3–6. SEG—segmental defect models; PLF—posterolateral spinal fusion model.

#### **Table 3.** Rabbit segmental defect model.

#### **Table 4.** Rabbit PLF model.

#### **Table 4.** *Cont*.

**Table 5.** Dog and sheep segmental defect model.

**Table 6.** Sheep and non-humane primate PLF models.

**Table 6.** *Cont*.

#### 3.2.2. Posterolateral Spinal Fusion (PLF) Model

Rabbit is the most commonly used species for the evaluation of the efficacy and safety of promising therapeutical solutions for achieving PLF [5,6,37,38,40,62–64,74,85,86,127–129]. The transverse processes of lumbar vertebrae are exposed, and an osteoinductive device is implanted bilaterally between adjacent transverse processes (L4-L5 or L5-L6) [127]. Transverse processes should be decorticated before the implantation to promote osseointegration of newly formed bone with native bone [86]. In the majority of previous studies, the BMP dose was up to 1000 µg and was delivered on various carriers (Table 4). The spinal fusion outcome was evaluated 6 weeks following surgery, and the majority of rabbit PLF studies had an observation period of fewer than 10 weeks. Few studies had a prolonged observation period (>10 weeks), but later time points might be important to determine the survival and long-term maintenance of newly induced bone [6,86], which is clinically relevant in patients undergoing PLF surgery (Figure 5, 1st column).

Methods of evaluation in rabbit segmental defect and PLF models are similar and include segmental mobility testing, radiological methods (x-ray and CT/microCT), histological (Figure 2F), histomorphometric analyses, and biomechanical testing [5,6,37,38,40,62– 64,74,85,86,127–129].

#### *3.3. Advanced Evaluation of Bone Regeneration Therapies*

#### 3.3.1. Dog and Sheep Segmental Defect Model

Dog and sheep segmental defect models are used for advanced evaluation of bone regeneration therapies. In dogs, the defect (20–25 mm) is created in the radius or ulna (Figure 3C) [130–135]. Applied doses of BMPs were in the range between 100 and 650 µg, which is higher compared to the rabbit model. Moreover, the typical observation period (12–24 weeks) was also prolonged compared to the rabbit model (Figure 5, 2nd and 3rd columns).

Tibial segmental bone defects in sheep (Figure 3D) were recently developed to evaluate novel bone regeneration therapies in conditions mimicking the size and biology of segmental bone defects in the clinics [2,136–140]. Moreover, there are two subtypes of this model: a fresh defect (FD) and biologically exhausted defect (BED), the latter mimicking a patient with a non-union [2]. Following the creation of a large defect (30 or 45 mm) in the sheep tibia in the FD model, a polymethyl-methacrylate spacer is inserted to induce the formation of the Masquelet membrane. Six weeks following the creation of the defect, an osteoinductive device was inserted after the removal of the spacer (FD model). In the BED model, the defect is in the first instance left untreated leading to a non-union. Subsequently, debridement of the non-union or fibrotic tissue ingrowth (BED model) is performed, followed by implantation of a spacer for 6 weeks, and then, finally, after removal of the spacer, an implant is inserted. BMP doses applied in this model ranged from 344 to 3800 µg, while the typical observation period was up to 16 weeks [2]. Although the osteoinductive device containing BMP6 on a carrier achieved bridging in FD (30 mm), it was found that larger, biologically exhausted defects appear to require a cell-based implant together with BMP to achieve proper clinically relevant bridging (Table 5) [2]. Importantly, defects were mechanically well stabilized with a circular external fixator according to the Ilizarov technique.

#### 3.3.2. Sheep PLF Model

The sheep PLF model is highly translatable to clinics because the size of the lumbar vertebrae of the sheep is comparable to humans. However, only a few preclinical studies have been conducted on this model [3,65,88] (Table 6). Sheep PLF may be conducted at a single level or as a multisegmental procedure. Moreover, it may be performed with or without instrumentation [88]. The observation period and applied BMP doses in this model were typically significantly longer/higher than in studies on small animals: the follow-up period was up to 6 months with a BMP amount up to 10 mg (Figure 5, 3rd column). Methods of evaluation included X-ray monitoring, microCT evaluation, histological analyses (Figure 2H), and biomechanical testing [3,65,88].

#### 3.3.3. Non-human Primate (NHP) PLF Model

Non-human primates are the most similar animal species to humans, both anatomically and genetically. However, only a few studies were conducted using NHP PLF [14,63,141] (Table 6), primarily due to ethical and economic reasons. In these studies, the goal was to achieve a single-level fusion between adjacent lumbar transverse processes, which are anatomically similar to humans. The applied BMP2 doses (3–12 mg), as well as observation period (24 weeks), were comparable to the sheep PLF model (Figure 5, 4th column).

#### *3.4. Anatomical Characteristics of the Species*

#### 3.4.1. Segmental Bone Defect

The general anatomy of long bones (Figure 3B–D) of species discussed in this review is similar, and obviously, the greatest difference among them is size. Differences in the bone size reflect the segmental defect created in each species. The length of the long bone segmental defect in mice (3 mm) or rats (5–8 mm) is small and created in the femur, the largest bone in rodents [5,33–35,67,78,79,119–121,123,142]. Long bones are significantly larger in rabbits/dogs, and segmental defects (15–20 mm in rabbits and 20–25 mm in dogs) are created in the femur, radius, or ulna [36,41,42,61,68–70,80–84,87,124–126,130–135]. Sheep segmental bone defects are usually created in the tibia and, due to their size (30–45 mm), are considered clinically relevant because the defect size compares to those in patients [2].

#### 3.4.2. PLF

Posterolateral spinal fusion (PLF) in preclinical studies is conducted in the lumbar portion of the spine. Although the basic anatomical features of lumbar vertebrae are similar among species discussed in this review, they differ in size and proportions of the different parts of the vertebrae. Rabbits (Figure 3E) and sheep (Figure 3F) have long transverse processes compared to the size of the vertebral body, while humans (Figure 3G), as an adaptation to erect posture and bipedal locomotion, have large bodies and short transverse processes. Importantly, transverse processes in rabbits are slanted and oriented anteriorly (Figure 2E or Figure 3E). On the other hand, the transverse processes in sheep (Figure 2G or Figure 3F) and humans (Figure 3G) are horizontal. The distances between the transverse processes are relatively short in rabbits (20–30 mm), while they are comparable in sheep and humans (40–50 mm).

#### **4. Appropriate Bone Morphogenetic Proteins Dosing**

The selection of the appropriate dose for each indication is one of the most challenging steps in the design of experiments. In this review, we compared doses used in various models in mice, rats, rabbits, sheep, and NHP. In the mouse and rat ectopic models, BMP doses were typically up to 25 µg [31,32,39,48–57,71,73,76,86,87] (Figure 4, 1st column, 1st row) per implant, while they were increased in the rat bone defects (up to 50 µg) [5,29,33–35,55,58–60,66,67,77–79,123] (Figure 4, 2nd column, 1st row). BMP doses used in rabbits were up to 150 µg in segmental bone defects [36,41,42,61,68–70,80– 84,87,124–126,143] (Figure 5, 1st column, 1st row) and up to 300 µg in rabbit PLF procedures [5,6,37,38,40,62–64,74,85,86,127–129] (Figure 5, 1st column, 1st row). Tested BMP

doses in dogs were higher and were typically in the range between 100 and 650 µg [130–135] (Figure 5, 2nd column, 1st row). Moreover, in the sheep and NHP models, BMP doses were significantly higher: in the sheep, the doses were between 500 µg and up to 4 mg [3,65,88] (Figure 5, 3rd column, 1st row), while in the NHP, the doses were up to 12 mg [14,63,141] (Figure 5, 4th column, 1st row).

#### **5. Duration of the Observation Period**

The bone induction process is significantly faster in small animals compared to higher species animals and humans. Therefore, observation periods were significantly shorter in rodent ectopic and bone defect models than in studies on sheep and non-human primates (Figures 4 and 5, 2nd row). The average observation period in mouse and rat ectopic models was 3–4 weeks [31,32,39,48–57,71,73,76,86,87] (Figure 4, 1st and 2nd columns, 2nd row). However, depending on the purpose of the study, observation periods in these studies might vary from a few days (studies on the mechanism of bone induction) to several months (bone longevity). Typical observation periods in the rat calvarial defect and femoral segmental defect models are slightly prolonged and last 5 and 10 weeks, respectively [5,29,33–35,55,58–60,66,77–79,123] (Figure 4, 2nd column, 2nd row). The observation period in segmental defect studies was up to 12 weeks in rabbits. [36,41,42,61,68–70,80–84,87,124–126,130–135,143] (Figure 5, 1st column. 2nd row). On the other hand, spinal fusion outcome in the rabbit PLF model was typically evaluated after 6 weeks [5,6,37,38,40,62–64,74,85,86,127–129] (Figure 5, 1st column, 2nd row). However, to study longevity or resorbability of compression-resistant matrices, the follow-up period might be prolonged. As expected, a longer observation period in dogs and sheep was up to 12 months [3,65,88,130–135] (Figure 5, 2nd and 3rd columns, 2nd row), while in NHP studies, it was 6 months [14,63,141] (Figure 5, 4th column, 2nd row).

#### **6. Sample Size**

Defining an appropriate sample size is a prerequisite for obtaining valid conclusions from each study. Moreover, the appropriate size of the sample is affected by several parameters, including experimental design and purpose of the study as well as expected differences among experimental groups. The sample size in the majority of reviewed studies here was 5–10 per group regardless of the animal species or model (Figures 4 and 5, 3rd row). Moreover, there is a consensus in published work that the minimal number of animals per experimental group is four. However, a few animals might die during surgery or follow-up periods due to reasons non-related to the tested osteoinductive therapy; therefore, at least five animals per experimental group should be included.

#### **7. Study Outcomes**

In Tables 1–6, it was not possible to describe the study outcomes due to non-comparable scoring grades for healing or spinal fusion experiments. The prerequisite in reporting the outcome of bone defect and spinal fusion studies is a clearly described success rate as the percentage of successfully rebridged defects or fused spine segments, respectively. Moreover, the method (radiological images, mobility testing) used to determine rebridgment/fusion should be clearly described. Surprisingly, in a large number of published studies, the success rate was not explicitly described. Several authors used their own scoring grades instead of standardized binary outcomes (successful or unsuccessful rebridgment/fusion). However, even when the binary outcome was used, the determination of successful rebridgment/fusion differed among authors. For example, a few authors determined success rate only on X-ray images without microCT, histology, and biomechanical testing. We suggest that successful rebridgment/fusion should also be determined with microCT, histological sectioning, and biomechanical testing.

The experimental outcome of osteoinductive therapies using rodent ectopic models should be determined by microCT and histology. MicroCT analyses provide information on newly formed bone volume expressed as bone volume (BV) or bone volume/tissue volume ratio (BV/TV). Additionally, if the tested osteoinductive device contains ceramics, microCT analyses might be used to determine the amount of residual ceramic matrix. Moreover, microCT analyses allow the determination of structural properties of newly formed bone by calculating trabecular parameters (trabecular number, trabecular thickness, trabecular separation). The structural properties of newly induced bone should also be analysed by histology and histomorphometry to determine the volume of the bone and remaining carrier/matrix.

#### **8. Conclusions**

Due to the large socioeconomic burden of degenerative diseases of the spine and segmental defects of long bones, there is an imminent need for the development of novel osteoinductive therapeutic solutions [1,22]. However, until now, none of the osteoinductive devices have been approved for use in PLF and large segmental defects in patients. A broad range of bone regeneration strategies have been proposed and tested in different animal models. A vast majority of these studies have been conducted in rats and rabbits, leading only to the initial and intermediate steps of preclinical testing, and despite claiming positive results, only a few have been further tested in sheep and NHP models. Infuse™, a BMP2-containing osteoinductive device, has been approved for use in ALIF and acute tibial fractures but has also been used in various off-label indications. However, numerous side effects related to high BMP dose and a large release from the bovine collagen as a carrier have been reported. Therefore, there is a need for an osteoinductive device that would be efficacious at lower doses of BMP delivered on a carrier with a prolonged BMP release. There is some hope that novel engineered BMPs or innovative delivery systems for BMPs may reduce the required therapeutic doses. A novel ABGS containing rhBMP6 within autologous blood coagulum was evaluated in preclinical studies, and in exploratory clinical trials (high tibial osteotomy, distal radial fracture, and posterolateral interbody fusion), it was proven safe and efficacious at relatively low BMP6 doses [73,74,76,86–88,91,92].

**Author Contributions:** Conceptualization, N.S. and S.V.; Writing—Original Draft Preparation, N.S. and S.V.; Writing—Review and Editing, N.S., N.I., D.M., F.P.L., and S.V.; Visualization, N.I. and N.S.; Supervision, S.V., F.P.L., and D.M.; Funding Acquisition, S.V., F.P.L., and D.M. All authors have read and agreed to the published version of the manuscript.

**Funding:** This program was funded by the FP7 Health Program (FP7/2007-2013) under grant agreement HEALTH-F4-2011-279239 (Osteogrow), H2020 Health GA 779340 (OSTEOproSPINE), and European Regional Development Fund—Scientific Center of Excellence for Reproductive and Regenerative Medicine (project "Reproductive and regenerative medicine—exploration of new platforms and potentials," GA KK.<IP_ADDRESS>.0008 funded by the EU through the ERDF).

**Institutional Review Board Statement:** Not applicable.

**Informed Consent Statement:** Not applicable.

**Acknowledgments:** Special thanks to Jack Ratliff (Ratliff Histology Consultants LLC) for the excellent preparation of the undecalcified histology sections presented in this paper.

**Conflicts of Interest:** S.V. is a cofounder of perForm Biologics and a coordinator of the EU HORI-ZON 2020 grant OSTEOproSPINE funding clinical studies of the new drug for bone repair (patent WO2019076484A1).

#### **References**

#### *Case Report*

### **Design Techniques to Optimize the Sca**ff**old Performance: Freeze-dried Bone Custom-Made Allografts for Maxillary Alveolar Horizontal Ridge Augmentation**

#### **Felice Roberto Grassi 1 , Roberta Grassi 2 , Leonardo Vivarelli 3 , Dante Dallari 3 , Marco Govoni 3 , Gianna Maria Nardi 4 , Zamira Kalemaj <sup>5</sup> and Andrea Ballini 1,6,7, \***

Received: 29 January 2020; Accepted: 16 March 2020; Published: 19 March 2020

**Abstract:** The purpose of the current investigation was to evaluate the clinical success of horizontal ridge augmentation in severely atrophic maxilla (Cawood and Howell class IV) using freeze-dried custom made bone harvested from the tibial hemiplateau of cadaver donors, and to analyze the marginal bone level gain prior to dental implant placement at nine months subsequent to bone grafting and before prosthetic rehabilitation. A 52-year-old woman received custom made bone grafts. The patient underwent CT scans two weeks prior and nine months after surgery for graft volume and density analysis. The clinical and radiographic bone observations showed a very low rate of resorption after bone graft and implant placement. The custom-made allograft material was a highly effective modality for restoring the alveolar horizontal ridge, resulting in a reduction of the need to obtain autogenous bone from a secondary site with predictable procedure. Further studies are needed to investigate its behavior at longer time periods.

**Keywords:** geometry optimization of scaffolds; allograft; block bone grafts; custom made bone; design techniques for scaffold; precision and translational medicine

#### **1. Introduction**

Implant-supported rehabilitation of the edentulous ridge requires adequate volume and integrity of the alveolar bone [1].

Bone resorption in the maxillary ridge, due to trauma, pathology or tooth loss, frequently results in a knife-edged deformity, which complicates implant placement and stabilization, particularly in the posterior jaw [2–4]. Grafting with allograft bone has been documented to be a useful tool in reconstructing jaw anatomy, [5] restoring esthetics [6], and providing biomechanical support for the placement of dental implants [7].

Clinically, the most suitable banked bone allografts are fresh-frozen (FFBAs), freeze-dried bone allografts (FDBAs), and demineralized freeze-dried (DFDBAs) [8], although in oro-maxillofacial surgical interventions FDBAs and DFDBAs are the most used.

Frozen bone is accessible for human receivers after at least 6 months of quarantine at −80 ◦C [9] and no additional preparation is required. Moreover, the osteoinductive proteins are preserved [10].

Strict guidelines for tissue harvesting and storing at −80 ◦C make the risk of primary infections and antigenicity reasonably low [11]. In addition, the reduction of water content in frozen bone by the freeze-drying process further decreases potential microbial contaminations [12].

Frozen bone and relative freeze-dried/demineralized products are accessible as cancellous granules/blocks, corticocancellous granules/blocks, and cortical granules or chips. Once defrosted, frozen bone has handling qualities similar to fresh bone [8] while freeze-dried allografts have the additional advantage of storage at room temperature.

Bone density can be measured with high reproducibility by means of Cone-Beam-Computed-Tomography (CBCT) scans [13]. Further techniques, such as orthopantomography, do not assure fitting precision in density determination [13].

The aim of the present translational study was to evaluate the clinical success of horizontal ridge augmentation in severely atrophic maxilla (Cawood and Howell class IV) using custom made bone harvested from the tibial hemiplateau of cadaver donors.

#### **2. Materials and Methods**

A 52-year-old woman presented compromised anterior maxillary ridges, who presented for the placement of dental implants, was included in this pilot study. Written and verbal information was given to the patient before enrollment, and written informed consent was obtained. The study was conducted in full accordance with the World Medical Association Declaration of Helsinki on experimentation involving human subjects, as revised in 2008.

Inclusion criteria was a horizontal severely atrophic maxilla (Cawood and Howell class IV), needing a bone grafting procedure prior to implant placement. Exclusion criteria were established according to Venet et al. (2017) [12]. Plaque index score was maintained ≤25% throughout the study [14].

#### *Graft Sample Blocks*

The processing was performed on corticocancellous bone blocks obtained from a proximal tibial epiphysis, in the anatomical region between the articular surface of the tibial plateau and tuberosity.

Human allogeneic bone blocks were collected from cadaveric donors, stored at −80 ◦C and processed in the accredited public non-profit Musculoskeletal Tissue Bank of IRCCS Istituto Ortopedico Rizzoli (Bologna, Italy), authorized by the Italian National Transplant Center for the collection, processing, and distribution of human musculoskeletal tissue [15], and registered in the European Tissue Establishment list (code IT000096). The choice of the block to be machined involved considering dimensions slightly greater than the machining area defined with the graft design.

After thawing, each block was fixed with clamps on a special stainless-steel table in a GMP-Class A (Good Manufacturing Practice) Clean Room environment. Then, the table was fixed inside the CNC (Computer Numerical Control) milling machine (model Bright, Delta Macchine, Rieti, Italy). After tool fixing, a 3 mm diameter ball mill, the execution of the machining trajectory was started and accurately monitored by the operator.

At the end of the processing, the block was removed and finished, the supports were broken by hand and sharp edges, and rough corners were hand-refined using sterile rasps. Then, grafts were cleaned with organic solvents, washed with sterile water, and freeze-dried (VirTis Genesis 25, SP Scientific, Warminster, PA, USA). Finally, grafts were individually wrapped in triple pack. Microbiological

sampling was performed during the processing of grafts and after the lyophilization protocol in order to exclude microbial contamination and declare tissues suitable for implantation.

According to the Guidelines of the Italian National Transplant Center, lyophilized bone grafts are preserved at room temperature for a maximum of 5 years.

*Materials* **2020**, *13*, x FOR PEER REVIEW 3 of 11

#### *Graft Design*

[16–18].

Digital Imaging and Communications in Medicine (DICOM) data of the maxilla (Figure 1) were acquired by a Cone Beam Computerized Tomography (CBCT) scanner (Orthophos XG 3D Ceph, Dentsply Sirona Italia Srl, Roma, Italy) and imported into the 3DSlicer software (www.slicer.org) [16–18]. *Graft Design*  Digital Imaging and Communications in Medicine (DICOM) data of the maxilla (Figure 1) were acquired by a Cone Beam Computerized Tomography (CBCT) scanner (Orthophos XG 3D Ceph, Dentsply Sirona Italia Srl, Roma, Italy) and imported into the 3DSlicer software (www.slicer.org)

**Figure 1.** Cone-Beam-Computed-Tomography (CBCT) axial (a) and cross-sectional (b1,b2) images before surgical procedure. **Figure 1.** Cone-Beam-Computed-Tomography (CBCT) axial (**a**) and cross-sectional (**b1,b2**) images before surgical procedure.

After setting a threshold for the automatic selection of the areas delimiting the cortical bone, a

manual analysis for each slice was performed to correct potential errors. Starting from the selected areas, an automatic procedure generated the 3D model as a stereolithography (STL) file. The graft design was performed by Rhinoceros ver. 4 (www.rhino3d.com) following these steps: (a) placement of a parallelepiped in the position where the volume increase was required; (b) Boolean subtraction between the parallelepiped and the STL model of the patient's anatomy; (c) After setting a threshold for the automatic selection of the areas delimiting the cortical bone, a manual analysis for each slice was performed to correct potential errors. Starting from the selected areas, an automatic procedure generated the 3D model as a stereolithography (STL) file.

revision of the model for the manufacture by a 3-axis milling machine to obtain a L-shape section. This procedure was performed for each graft. The design of the grafts has been repeatedly validated and subsequently corrected following the surgeon's indications, up to the final model (Figure 2). The graft design was performed by Rhinoceros ver. 4 (www.rhino3d.com) following these steps: (a) placement of a parallelepiped in the position where the volume increase was required; (b) Boolean subtraction between the parallelepiped and the STL model of the patient's anatomy; (c) revision of the model for the manufacture by a 3-axis milling machine to obtain a L-shape section. This procedure was performed for each graft.

The design of the grafts has been repeatedly validated and subsequently corrected following the surgeon's indications, up to the final model (Figure 2).

#### *Trajectory Planning and Graft Manufacturing*

The planning of machining trajectories was performed using Rhinoceros by the Rhino-CAM 2 plug-in (MecSoft Corporation, Irvine, CA, USA). Each graft was positioned in the center of the machining CAM area with the larger flat area facing down to keep the cortical portion of the tissue intact, in order to enhance the graft resistance during processing and avoid breaking during implantation. Bone bridges have been added to the CAD design to maintain fixation during processing and removed by hand at the end of the graft manufacturing procedure. The working area was defined according to the design's dimensions and the diameter of the milling tool.

The tool-path trajectories have been programmed, subsequently simulated and modified to obtain the best result (Figure 3). Then, the definitive trajectories have been exported as G-Code coding specific for the milling machine used.

*Materials* **2020**, *13*, x FOR PEER REVIEW 4 of 11

**Figure 2.** (a) Render of 3D bone donor site harvested from a cadaver tibial hemiplateau. (b1–b3) Render of 3D reconstruction of patient's anatomy (yellow) and designed bone grafts (white). (c) Final block for clinical use. **Figure 2.** (**a**) Render of 3D bone donor site harvested from a cadaver tibial hemiplateau. (**b1**–**b3**) Render of 3D reconstruction of patient's anatomy (yellow) and designed bone grafts (white). (**c**) Final block for clinical use.

*Materials* **2020**, *13*, x FOR PEER REVIEW 5 of 11

*Trajectory Planning and Graft Manufacturing*

custom allograft was freeze-dried.

*Surgical Procedure*

**Figure 3.** Tool-path trajectory planning: (a) added bone bridges; (b) work area; (c) graft design; (d) tool-path. The machining operation was composed by two different phases: horizontal roughing and **Figure 3.** Tool-path trajectory planning: (**a**) added bone bridges; (**b**) work area; (**c**) graft design; (**d**) tool-path.

parallel finishing. The tool used for each phase was the same, a fluted 3 mm ball mill with two teeth, hard metal k10 material (tungsten carbide). In each phase rotational velocity of the tool was 5300 RPM and feed rate was 200 mm/min, that corresponds to a feed per tooth of 19 μm and a tangential velocity of 50 m/min (0.833 m/s). Horizontal roughing was performed as a full immersion vertical milling, with a machining depth of 3.5 mm, leaving a stock margin of 0.3 mm. Parallel finishing was a vertical milling, with a machining depth depending on stock margins and parallel steps of 0.7 mm (Figure 3d) leaving a variable stock margin between 0 mm and 0.15 mm (a sufficient dimensional accuracy for oral implant surgery). This machining strategy, performed by few passages, reduced the machining time and the bone heating during the process. Although temperature was not monitored during the fabrication of these two custom grafts, our machining parameters, especially the small feed per tooth, allow a local temperature between 40 °C and 60 °C for cortical bone and lower values for cancellous bone. According to Krause [19], these parameters do not cause the denaturation of proteins and enzymes that occurs for temperature above 70 °C. As reported by Van Isacker et al. [20], frozen bone can be handled and reshaped like normal bone and is fully workable, but freeze-dried bone, unless rehydrated in saline, is brittle like ceramic, and is not fully workable. The machining operation was composed by two different phases: horizontal roughing and parallel finishing. The tool used for each phase was the same, a fluted 3 mm ball mill with two teeth, hard metal k10 material (tungsten carbide). In each phase rotational velocity of the tool was 5300 RPM and feed rate was 200 mm/min, that corresponds to a feed per tooth of 19 µm and a tangential velocity of 50 m/min (0.833 m/s). Horizontal roughing was performed as a full immersion vertical milling, with a machining depth of 3.5 mm, leaving a stock margin of 0.3 mm. Parallel finishing was a vertical milling, with a machining depth depending on stock margins and parallel steps of 0.7 mm (Figure 3d) leaving a variable stock margin between 0 mm and 0.15 mm (a sufficient dimensional accuracy for oral implant surgery). This machining strategy, performed by few passages, reduced the machining time and the bone heating during the process. Although temperature was not monitored during the fabrication of these two custom grafts, our machining parameters, especially the small feed per

full-thickness crestal incision with two vertical releasing incisions (Figure 4a).

Local anesthesia was obtained by infiltrating articaine (4% containing 1:100,000 adrenaline). The exposure of the three-dimensional aspect of the bone defects was achieved through a

The clinically sized, anatomically shaped, custom-made bone block was placed in position strictly overlapping the underlying alveolar crest and fitted securely to the residual bone (Figure

Thus, to execute a better processing procedure and to obtain a further reduction in temperature, machining was performed on a frozen bone block. At the end of the machining operation, the tooth, allow a local temperature between 40 ◦C and 60 ◦C for cortical bone and lower values for cancellous bone. According to Krause [19], these parameters do not cause the denaturation of proteins and enzymes that occurs for temperature above 70 ◦C.

As reported by Van Isacker et al. [20], frozen bone can be handled and reshaped like normal bone and is fully workable, but freeze-dried bone, unless rehydrated in saline, is brittle like ceramic, and is not fully workable.

Thus, to execute a better processing procedure and to obtain a further reduction in temperature, machining was performed on a frozen bone block. At the end of the machining operation, the custom allograft was freeze-dried. *Materials* **2020**, *13*, x FOR PEER REVIEW 6 of 11

#### *Surgical Procedure* 4b1). The recipient site was weakened with multiple micro-holes to enhance bleeding from the trabecular bone [12,21,22]. Rigid fixation of the scaffold to the residual crest was obtained by means

Local anesthesia was obtained by infiltrating articaine (4% containing 1:100,000 adrenaline). The exposure of the three-dimensional aspect of the bone defects was achieved through a full-thickness crestal incision with two vertical releasing incisions (Figure 4a). of a titanium mini-screw (1.5 mm width, 8 mm length) (Tekka by Global-D, Lyon, France) (Figure 4b2) [12]. The grafted area was closed with a pulley suture for proper flap adaptation and to avoid any tissue strangulation by an absorbable 4.0/5.0 suture material. Sutures were removed 14 days postoperatively (Figure 4c) and no prosthetic device was admitted in the following 90 days, with

particular care in domiciliary oral hygiene procedures.

**Figure 4.** The custom-made freeze-dried bone allograft (FDBA) was positioned in the deficit area and stabilized with a screw. (a) Anatomical site prior surgical procedure and clinical view of the defect. (b1,b2) Screw positioning on a custom-made FDBA. (c) Sutures. Preceding to the second phase, supplementary CBCT scans were performed to evaluate grafts **Figure 4.** The custom-made freeze-dried bone allograft (FDBA) was positioned in the deficit area and stabilized with a screw. (**a**) Anatomical site prior surgical procedure and clinical view of the defect. (**b1,b2**) Screw positioning on a custom-made FDBA. (**c**) Sutures.

The clinically sized, anatomically shaped, custom-made bone block was placed in position strictly overlapping the underlying alveolar crest and fitted securely to the residual bone (Figure 4b1). The recipient site was weakened with multiple micro-holes to enhance bleeding from the trabecular bone [12,21,22]. Rigid fixation of the scaffold to the residual crest was obtained by means of a titanium mini-screw (1.5 mm width, 8 mm length) (Tekka by Global-D, Lyon, France) (Figure 4b2) [12].

The grafted area was closed with a pulley suture for proper flap adaptation and to avoid any tissue strangulation by an absorbable 4.0/5.0 suture material. Sutures were removed 14 days postoperatively (Figure 4c) and no prosthetic device was admitted in the following 90 days, with particular care in domiciliary oral hygiene procedures.

Preceding to the second phase, supplementary CBCT scans were performed to evaluate grafts gain (Figure 5).

**Figure 5.** CBCT scans were done after 9 months to define implant placement width and length, on (a) axial and (b1,b2) cross-sectional images. After a 9 month healing period, micro-screws were removed (Figure 6a,b), and 3.7 mm in diameter for a 10 mm in length dental implants (iRES SAGL, Mendrisio, Switzerland) were placed, and sutures performed (Figure 6c). No supplementary grafting procedure was required in the current investigation [22].

#### *Statistical Evaluation*

gain (Figure 5).

The outcome values were analyzed using the *t-test* for paired samples for pre–post differences with time as the factor and IBM Statistical Package for the Social Sciences (SPSS Inc. Version 21.0, 4b2) [12].

gain (Figure 5).

particular care in domiciliary oral hygiene procedures.

Chicago, IL, USA) software to detect significant differences between pre-test and post-test scores obtained from bone volume analysis and measurements of variation in ridge thickness from CBCT images at 9 months after allograft insertion. **Figure 4.** The custom-made freeze-dried bone allograft (FDBA) was positioned in the deficit area and stabilized with a screw. (a) Anatomical site prior surgical procedure and clinical view of the defect. (b1,b2) Screw positioning on a custom-made FDBA. (c) Sutures. Preceding to the second phase, supplementary CBCT scans were performed to evaluate grafts

*Materials* **2020**, *13*, x FOR PEER REVIEW 6 of 11 4b1). The recipient site was weakened with multiple micro-holes to enhance bleeding from the trabecular bone [12,21,22]. Rigid fixation of the scaffold to the residual crest was obtained by means of a titanium mini-screw (1.5 mm width, 8 mm length) (Tekka by Global-D, Lyon, France) (Figure

The grafted area was closed with a pulley suture for proper flap adaptation and to avoid any tissue strangulation by an absorbable 4.0/5.0 suture material. Sutures were removed 14 days postoperatively (Figure 4c) and no prosthetic device was admitted in the following 90 days, with

**Figure 5.** CBCT scans were done after 9 months to define implant placement width and length, on (a) axial and (b1,b2) cross-sectional images. **Figure 5.** CBCT scans were done after 9 months to define implant placement width and length, on (**a**) axial and (**b1,b2**) cross-sectional images. *Materials* **2020**, *13*, x FOR PEER REVIEW 7 of 11 After a 9 month healing period, micro-screws were removed (Figure 6a,b), and 3.7 mm in diameter for a 10 mm in length dental implants (iRES SAGL, Mendrisio, Switzerland) were placed, and sutures performed (Figure 6c). No supplementary grafting procedure was required in the

**Figure 6.** Nine months after the surgical procedure, the implant was positioned. (a) Soft tissues before surgical procedure. (b) Surgical flap is elevated. (c1) Implant was positioned and (c2) sutures performed. (d) Frontal view of the final rehabilitation. *Statistical Evaluation*  **Figure 6.** Nine months after the surgical procedure, the implant was positioned. (**a**) Soft tissues before surgical procedure. (**b**) Surgical flap is elevated. (**c1**) Implant was positioned and (**c2**) sutures performed. (**d**) Frontal view of the final rehabilitation.

#### **3. Results**

with time as the factor and IBM Statistical Package for the Social Sciences (SPSS Inc. Version 21.0, Chicago, IL, USA) software to detect significant differences between pre-test and post-test scores obtained from bone volume analysis and measurements of variation in ridge thickness from CBCT images at 9 months after allograft insertion. A patient presenting atrophic anterior maxillae was selected to participate in the study. The patient received 2 custom-made allografts. The healing period was uneventful.

The outcome values were analyzed using the *t-test* for paired samples for pre–post differences

**3. Results**  A patient presenting atrophic anterior maxillae was selected to participate in the study. The patient received 2 custom-made allografts. The healing period was uneventful. At the crestal level, the patient met the inclusion criteria of ridge width around 2 mm [22]. The titanium fixation mini-screws detached and bleeding from the bone graft were detected, demonstrating revascularization of the site.

> At the crestal level, the patient met the inclusion criteria of ridge width around 2 mm [22]. The titanium fixation mini-screws detached and bleeding from the bone graft were detected,

demonstrating revascularization of the site.

Moreover, the regenerative surgical procedure went well. In fact, the custom-made allograft scaffolds perfectly fitted in the bone anatomy and were therefore easily adapted to the bone defects during surgery, secured by titanium mini-screws (Figures 4–6). This excellent matching of the size/shape helped the surgeon to reduce the operation time [23]. Moreover, all implants were inserted with fitting primary stability (Figure 6d).

Paired comparisons showed a significant mean increase in ridge thickness of 5.0 ± 0.55 mm from the preoperative measurement to the postoperative measurement (p < 0.01).

The bone resorption that occurred during the incorporation period corresponds to 7.8% (95% confidence interval (CI): (6.4%, 9.2%)) of the measured postoperative ridge thickness.

#### **4. Discussion**

Maxillary ridge augmentation relates to procedures designed to correct a thin alveolar ridge. For dental implant placement, adequate bone volume is necessary, unfortunately, bone volume is not always adequate, particularly in elderly patients.

Through the current improvements in planning and manufacturing software and hardware, graft customization can be combined with the implant dentistry digital workflow with the potential patients to become part of daily clinical practice [24].

The reduction of risk factors such as infections caused by contamination of surgical site, or contamination of the graft during preoperative preparation, can be obtained with the manufacturing of a custom sterile graft [23]. The correct design of a custom graft, with also a reduction of surgical time, can be obtained with a proper planning of surgical procedure.