architecture of the scaffolds, significantly affects nutrient diffusion, cell adhesion, and matrix deposition of the regenerated tissues. Scaffolds have to be carefully designed to match specific mechanical, mass transport, and biological requirements; however, customizing the architecture to better suit these requirements remains a challenging issue.

Computer-aided designs (CAD) and computational simulations using finite element analysis (FEA) have played a major role in the reduction of in vitro and in vivo experimental efforts and costs. Furthermore, the accuracy of FEA results is acceptable to predict a design optimization before the manufacturing. An overview of this design paradigm is reported in **Figure 3**.

The images in STL file are directly processed with computer-aided manufacturing (CAM) in order to print the specimens in layer. Selective laser melting (SLM) techniques have shown capability of 3D cellular structure production through printing of layered structures based on CAD (**Figure 4**) [19–21]. During SLM, metallic powder is melted by a scanning laser beam [21], featured by excellent reproducibility and efficiency [22]. The limitations of SLM are the quality of the print which depends on the process setting parameters. Poor setting parameters such as layer thickness, laser-beam spot, laser power, and hatching space contribute to incomplete melting of the metal powders. Furthermore, the minimum scale that can process is limit the lattice thickness and pores in hundreds micrometer. Although, the scale limitation of the print is larger than nano-scale of the real bone, these SLM-processed structures are adequately to allow bone ingrowth (>300 μm).

The methodological approaches for the design of scaffold architectures are classified into the architecture formed by the repetition of unit cells (cellular structures) or that consisting of lattices (lattice domain). This collection of unit cells is also known as computer-aided system for tissue scaffolds (CASTS). Chua et al. [23, 24] and Cheah et al. [25] have developed a parametric library of scaffold structures and an algorithm to automate the entire process of matching desired anatomical shapes in order to reduce the time-consuming processes. **Figure 5** shows examples of the unit cells used in CASTS.

Limmahakhun et al. [26] emphasized that the internal architectures of scaffolds play a crucial role in determining the overall mechanical and biological performances. It is interesting to note that the mechanical response from a cellular structure depends on the loading modes. The cubic structure [0° ± 90°] has much higher compression stiffness but lower shear and torsion stiffness than that in octahedron-type structures [±45°]. The pillar octahedron combined with the strut characteristics featured along 0° and ±45° demonstrates high compression, shear, and torsional stiffness (26.94, 18.93, and 9.52 MPa, respectively). To this end, for axially loaded applications, such as bone implants to fix long-bone defects, the cubic structure

**Figure 3.** Flowchart steps in the design of tissue-engineering scaffolds.

architecture of the scaffolds, significantly affects nutrient diffusion, cell adhesion, and matrix deposition of the regenerated tissues. Scaffolds have to be carefully designed to match specific mechanical, mass transport, and biological requirements; however, customizing the architec-

**Figure 2.** Diagram of material-porogen mixtures of different porosities which are filled into a mold layer by layer to

Computer-aided designs (CAD) and computational simulations using finite element analysis (FEA) have played a major role in the reduction of in vitro and in vivo experimental efforts and costs. Furthermore, the accuracy of FEA results is acceptable to predict a design optimization before the manufacturing. An overview of this design paradigm is reported in

The images in STL file are directly processed with computer-aided manufacturing (CAM) in order to print the specimens in layer. Selective laser melting (SLM) techniques have shown capability of 3D cellular structure production through printing of layered structures based on CAD (**Figure 4**) [19–21]. During SLM, metallic powder is melted by a scanning laser beam [21], featured by excellent reproducibility and efficiency [22]. The limitations of SLM are the quality of the print which depends on the process setting parameters. Poor setting parameters such as layer thickness, laser-beam spot, laser power, and hatching space contribute to

ture to better suit these requirements remains a challenging issue.

78 Scaffolds in Tissue Engineering - Materials, Technologies and Clinical Applications

**Figure 3**.

form laminated multiporous layers.

**Figure 4.** SLM processing techniques. Reprinted from Ref. [31] with permission from Elsevier.

**Figure 5.** Unit-cell replications of bone tissue engineering scaffolds. Reprinted from Ref. [26] with permission from Elsevier.

is the best choice because of its high stiffness under compressive force. High fatigue life under compression loading was observed in cubic unit cells [19]. On the other hand, when subjected to a combined loading (compression, shear, and torsion), such as femoral hip implants, the pillar octahedron is an appropriate architecture as it provides the greatest shear and torsional stiffness and high compressive stiffness. Note that the pillar octahedral structure also has the greatest accumulated stress during the stress relaxation test [26].

Unit geometry directly controls the tortuosity, a measure of flow distance that travels in the porous materials [27]. The strut intersection in the middle of octahedral types diverts the fluid from the straight direction and slows down the flow rate (**Figure 6b**). The low flow rate when pipetting a cell suspension onto the pillar octahedral shapes would favor cell adherence (**Figure 6d–f**), while the high flow rates of cubic and truncated octahedral shape contribute to cell deposition at the bottom of the well. The greater cell proliferation of the pillar octahedron could be attributed to the greater surface areas of these polyhedral structures that are printed with 3D printing techniques, compared with the cubic and octahedron (**Figure 6a**). Together with its better cell proliferation rate, pillar octahedral structure has demonstrated balanced mechanical and biological properties.

CAD-based methods [28, 29] and implicit surface modeling (ISM) methods [18, 30] are common to create the cellular structures with porosity-graded structures, since the internal architectures are fully controlled with no hanging edge [18, 31]. For the CAD-based methods, graded structures of the pillar octahedral scaffolds could be attained by varying the architectural parameters, such as lattice diameter (t) and unit size (L) (**Figure 7**).

The graded cellular bone scaffolds with desired porosity and pore size on each location could be done using manual or automatic algorithm processes. The primitive unit cell that is recommended by the author's works is the pillar octahedral shape [26]. Our findings proved that this unit has balanced mechanical and biological performances for bone tissue regenerations. **Figure 8** shows an example of pillar octahedral unit cells with different porosities and graded patterns; axially and radially graded patterns, fabricated with cobalt chromium (CoCr) alloys

**Figure 6.** SEM images showing surface roughness of (a) cubic and (b) octahedral unit cells, (c) smaller pores (40 μm) on the surface after printing on the overhanging features, (d) cell adhesion and (e) and (f) cell coverage after 4 and 7 days. Reprinted from Ref. [26] with permission from Elsevier.

using SLM techniques. Manually mating the unit cells inside the graded cellular bone scaffolds could be fully controlled with no error, however it is time-consuming.

On the other hand, ISM allows scaffold architectures to be easily described using a single mathematical equation, with freedom to introduce different pore shapes and architectural

**Figure 7.** Assembling of different unit porosities.

is the best choice because of its high stiffness under compressive force. High fatigue life under compression loading was observed in cubic unit cells [19]. On the other hand, when subjected to a combined loading (compression, shear, and torsion), such as femoral hip implants, the pillar octahedron is an appropriate architecture as it provides the greatest shear and torsional stiffness and high compressive stiffness. Note that the pillar octahedral structure also has the

**Figure 5.** Unit-cell replications of bone tissue engineering scaffolds. Reprinted from Ref. [26] with permission from

Unit geometry directly controls the tortuosity, a measure of flow distance that travels in the porous materials [27]. The strut intersection in the middle of octahedral types diverts the fluid from the straight direction and slows down the flow rate (**Figure 6b**). The low flow rate when pipetting a cell suspension onto the pillar octahedral shapes would favor cell adherence (**Figure 6d–f**), while the high flow rates of cubic and truncated octahedral shape contribute to cell deposition at the bottom of the well. The greater cell proliferation of the pillar octahedron could be attributed to the greater surface areas of these polyhedral structures that are printed with 3D printing techniques, compared with the cubic and octahedron (**Figure 6a**). Together with its better cell proliferation rate, pillar octahedral structure has demonstrated balanced

CAD-based methods [28, 29] and implicit surface modeling (ISM) methods [18, 30] are common to create the cellular structures with porosity-graded structures, since the internal architectures are fully controlled with no hanging edge [18, 31]. For the CAD-based methods, graded structures of the pillar octahedral scaffolds could be attained by varying the architec-

The graded cellular bone scaffolds with desired porosity and pore size on each location could be done using manual or automatic algorithm processes. The primitive unit cell that is recommended by the author's works is the pillar octahedral shape [26]. Our findings proved that this unit has balanced mechanical and biological performances for bone tissue regenerations. **Figure 8** shows an example of pillar octahedral unit cells with different porosities and graded patterns; axially and radially graded patterns, fabricated with cobalt chromium (CoCr) alloys

tural parameters, such as lattice diameter (t) and unit size (L) (**Figure 7**).

greatest accumulated stress during the stress relaxation test [26].

80 Scaffolds in Tissue Engineering - Materials, Technologies and Clinical Applications

mechanical and biological properties.

Elsevier.

**Figure 8.** Cellular structures and graded cellular structures (FGMs) built by pillar octahedral unit (L and t are unit and strut size, respectively). Reprinted from [31] with permission from Elsevier.

features, including pore-size gradients. Although, automatically processing time based on the algorithm is more convenient compared to the CAD-based method, the algorithm that could perfectly serve the required graded internal architectures of the scaffolds with no internal defect is more complexity. While many of the ISM are available, both of the Schwarz's Diamond and Schoen's Gyroid shapes are preferable in promoting cell migration and tissue ingrowth [18]. Furthermore, these techniques are suitable to create the graded cellular structures.

### **3. The effects of graded cellular bone scaffolds**

Designing an architecture of the scaffolds that mimics the complex structure of bone tissue is the new generation of bone tissue engineering [32]. Scaffolds with graded cellular structures behave similar to the bone that is replaced. Furthermore, porosity gradation affects the types of cell regeneration and promotes the environment to be more suitable for cell functions. While the biological cell interactions are directly influenced by the porosity gradation, the mechanical performance which is affected by porosity and cellular structures could be modified to match the desired functionalities. The improved tissue regeneration rate of graded cellular structures led the research community to find the possible models of bone tissue engineering. However, such model has not been developed yet.

### **3.1. Biological cell interactions**

features, including pore-size gradients. Although, automatically processing time based on the algorithm is more convenient compared to the CAD-based method, the algorithm that could perfectly serve the required graded internal architectures of the scaffolds with no internal defect is more complexity. While many of the ISM are available, both of the Schwarz's Diamond and Schoen's Gyroid shapes are preferable in promoting cell migration and tissue ingrowth [18]. Furthermore, these techniques are suitable to create the graded

**Figure 8.** Cellular structures and graded cellular structures (FGMs) built by pillar octahedral unit (L and t are unit and

Designing an architecture of the scaffolds that mimics the complex structure of bone tissue is the new generation of bone tissue engineering [32]. Scaffolds with graded cellular structures behave similar to the bone that is replaced. Furthermore, porosity gradation affects the types of cell regeneration and promotes the environment to be more suitable for cell functions. While the biological cell interactions are directly influenced by the porosity gradation, the mechanical performance which is affected by porosity and cellular structures could be modified to match the desired functionalities. The improved tissue regeneration rate of graded cellular structures led the research community to find the possible models of bone tissue engineering. However, such model has not been

cellular structures.

developed yet.

**3. The effects of graded cellular bone scaffolds**

strut size, respectively). Reprinted from [31] with permission from Elsevier.

82 Scaffolds in Tissue Engineering - Materials, Technologies and Clinical Applications

Different layers of the tissue perform different roles in maintaining the organ functions. Each cell type has individual specific functions and is important for tissue formations. Bone, cartilage, and ligament are formed simultaneously as a bone and joint structure, as a result of tissue formation by osteoblast, chondrocyte, and fibroblast cell-types, respectively. The graded cellular bone scaffolds are critical for multitissue regenerations such as bone and cartilage tissues [33] and ligament-bone interface [34]. These cell types have obviously different environments, thus scaffolds should be tailored with different pore sizes and porosities. For example, fibroblasts (cell size 20–50 μm) can span void spaces up to 200 μm [35]. On the other hand, the preferred pore diameter for osteoblast (cell size 20–30 μm [36]) is 100–350 μm [37].

Beside the pore size, porosity and pore interconnection also facilitate bone ingrowth. The scaffold with porosity >90% and pore interconnection promotes more cell infiltration, proliferation and extracellular matrix deposition, since it has a better flow mechanic for nutrient and waste transportation [38] and allows cellular signals between interconnecting networks [39]. Furthermore, the graded cellular structures also increase the fluid permeability and flow, which enhances the cell diffusion throughout the whole scaffolds [40, 41].

Since different pore sizes affect the type of cell formations, graded cellular scaffolds potentially produce the multi-tissue grafts on the bone-ligament interface by following the phenotypic gradients that exist at the natural ligament-bone interface. In addition to the graded features in architectures, graded features in terms of different material compositions also promote biologic functions like bone materials. Two materials with spatially graded fraction of polymer and hydroxyapatite (HA), fabricated with co-electrospinning techniques, are found to be metabolically active from the study of rat bone marrow stromal cell cultures [34]. Gene expression of bone morphogenic protein-2 and osteopontin was elevated on mineral-containing regions as compared to regions without mineral, which confirmed osteoblastic phenotypic maturation of this polymeric-HA graded scaffolds by day 28.

Computational study has been utilized to optimize the best porosity distribution in functionally graded scaffolds for bone tissue engineering [42]. Based on the porosity distribution law, the graded scaffolds with tri-linear law promote larger amounts of bone formation compared to the models of bi-linear, linear, and constant laws. Alternatively, the more the complexity of porosity distribution laws (i.e., with increasing number of coefficients, Ai), the better the scaffold geometry can be tailored. A larger number of design variables increases the probability that optimizes a geometry to match the specific boundary and loading conditions.

The effect of the loading conditions appears more critical. Boccaccio et al. [42] showed that the loading conditions are essential in determining optimal porosity distribution. For a pure compression loading, it was predicted that the changes of the pore dimension are marginal, and using a graded cellular bone scaffold allows the formation of amounts of bone slightly larger than those obtainable with a homogeneous porosity scaffold. For a pure shear loading, instead, bone formations of graded cellular bone scaffolds are significant compared to a homogeneous porosity scaffold. While increasing the pore diameters leads to an increased value of the scaffold Young's modulus, increasing a porosity distribution law makes a scaffold generate larger amounts of bone formations. Graded porosity characteristics contribute to optimized loading distributions, which enhances sensing signal to maximize osteoblastic cellular activities, as termed as mechanobiologic signal.

#### **3.2. Mechanical performances**

A scaffold should have the mechanical properties sufficient to maintain integrity until the new tissue regeneration. Bioceramics (hydroxyapatite, bioglass, etc.) and metals (titanium, tantalum, cobalt chrome, etc.) have been commonly used as a biomaterial for bone tissue regeneration. The limitation of bioceramics is their brittleness and contributes to easily break after replacements. While greater stiffness, endurance, and strength of metals are preferred, the higher stiffness of metals shields the stress distributed to the adjacent bone tissue and leads to bone resorption called "stress shielding." Therefore, the mechanical properties of the scaffolds should match that of the native tissue to both prevent stress shielding and give proper mechanical performances.

Despite the types of materials used for fabrications, the relative modulus of cellular materials (*E*/*Es* ) has a power law relation to the relative density (*ρo* /*ρs* ), based on Gibson and Ashby model [43]:

$$\frac{E}{Es} = \text{ } \text{ } \text{ } \text{ } \text{ } \text{ } \frac{\rho o}{\rho s} \text{ }\text{ }\tag{1}$$

where *E* and *Es* are the moduli of base and cellular materials and *ρo* and *ρs* are the density of base and cellular materials. This relationship depends on the unknown coefficients (ϕ and *n*) which are a factor of each cellular unit.

The cellular structures with various relative densities can be created by varying the diameter of the beam thickness (**Figure 7**). The moduli and strength of those cellular scaffolds can be predicted upon the relative density, as shown in an example of cobalt chrome (CoCr) cellular structures (**Figure 9**). It is clear that the elastic modulus (stiffness), yield stress, and ultimate compressive strength of cellular structures increase with decrease in the porosity. The CoCr cellular structures of pillar octahedron have stiffness and compressive strengths between 2.33–3.14 GPa and 113–523 MPa, respectively, which are comparable to those of cortical bone tissues (2.73–17 GPa [44] and 100–150 MPa [45]). In addition, these CoCr cellular structures also demonstrated a greater energy absorption (24.6–116.86 MJ/m3 ) than bone tissues.

Unfortunately, the quality of the SLM-built is a major concern that affects the scaffold reproductions. Although FEA could predict the modulus of cellular scaffolds, the accuracy of the prediction is low [46]. Poor correlation between FEA and physical testing results from an incomplete melting of the metal powders. Internal voids of the printed scaffolds increase the stress concentration, and crack may occur, weakening the scaffold constructs (**Figure 10**). Completely selective-melting process in each layer of the manufacture plays a main role to minimize such error of the prediction.

**Figure 9.** Normalized elastic modulus (a), normalized yield strength (b), normalized ultimate compressive strength (c) of CoCr cellular structures vs. relative density. Reprinted from [31] with permission from Elsevier.

value of the scaffold Young's modulus, increasing a porosity distribution law makes a scaffold generate larger amounts of bone formations. Graded porosity characteristics contribute to optimized loading distributions, which enhances sensing signal to maximize osteoblastic

A scaffold should have the mechanical properties sufficient to maintain integrity until the new tissue regeneration. Bioceramics (hydroxyapatite, bioglass, etc.) and metals (titanium, tantalum, cobalt chrome, etc.) have been commonly used as a biomaterial for bone tissue regeneration. The limitation of bioceramics is their brittleness and contributes to easily break after replacements. While greater stiffness, endurance, and strength of metals are preferred, the higher stiffness of metals shields the stress distributed to the adjacent bone tissue and leads to bone resorption called "stress shielding." Therefore, the mechanical properties of the scaffolds should match that of the native tissue to both prevent stress shielding and give

Despite the types of materials used for fabrications, the relative modulus of cellular materi-

*Es* = *ϕ* (

base and cellular materials. This relationship depends on the unknown coefficients (ϕ and *n*)

The cellular structures with various relative densities can be created by varying the diameter of the beam thickness (**Figure 7**). The moduli and strength of those cellular scaffolds can be predicted upon the relative density, as shown in an example of cobalt chrome (CoCr) cellular structures (**Figure 9**). It is clear that the elastic modulus (stiffness), yield stress, and ultimate compressive strength of cellular structures increase with decrease in the porosity. The CoCr cellular structures of pillar octahedron have stiffness and compressive strengths between 2.33–3.14 GPa and 113–523 MPa, respectively, which are comparable to those of cortical bone tissues (2.73–17 GPa [44] and 100–150 MPa [45]). In addition, these CoCr cellular structures

Unfortunately, the quality of the SLM-built is a major concern that affects the scaffold reproductions. Although FEA could predict the modulus of cellular scaffolds, the accuracy of the prediction is low [46]. Poor correlation between FEA and physical testing results from an incomplete melting of the metal powders. Internal voids of the printed scaffolds increase the stress concentration, and crack may occur, weakening the scaffold constructs (**Figure 10**). Completely selective-melting process in each layer of the manufacture plays a main role to

are the moduli of base and cellular materials and *ρo*

*ρo* \_\_\_ *ρs*) *n* /*ρs*

), based on Gibson and Ashby

) than bone tissues.

and *ρs*

(1)

are the density of

) has a power law relation to the relative density (*ρo*

also demonstrated a greater energy absorption (24.6–116.86 MJ/m3

cellular activities, as termed as mechanobiologic signal.

84 Scaffolds in Tissue Engineering - Materials, Technologies and Clinical Applications

**3.2. Mechanical performances**

proper mechanical performances.

\_\_*E*

which are a factor of each cellular unit.

minimize such error of the prediction.

als (*E*/*Es*

model [43]:

where *E* and *Es*

The mechanical properties of titanium alloys fabricated with SLM have been extensively studied under static and fatigue loadings [19, 21, 47–51]. Although titanium cellular structures have similar stiffness as bone tissues, fatigue is still a technical issue, which mainly resulted from poor design of the internal architectures [19, 52]. Naturally, the implant is expected to have similar properties as bone. Structural gradient has been observed in bone tissues, depending on required functions, such as load-bearing capacity and biological properties [5]. However, under axial compression, there is no benefit of graded cellular bone scaffolds over a uniformed bone scaffold with the same relative density [31]. **Figure 11** shows that CoCr graded cellular structures have similar elastic modulus, yield strength, and compressive strength as the uniformed cellular structures. CoCr cellular scaffolds and graded cellular scaffolds with the porosity ranging from 40 to 70% exhibit elastic modulus around 2.7–3.1 GPa, which is in the same level of bone stiffness [44, 53]. Therefore, mechanical properties of the cellular bone scaffolds are mainly related to the relative density rather than the porosity-graded characteristics.

The opposite findings were noticed with the bioceramic graded cellular scaffolds. Wang et al. [17] studied a novel calcium polyphosphate bioceramic scaffold with a graded pore structure similar to the bimodal structure of cortical and cancellous bones (**Figure 12**). The compressive strength of porosity-graded calcium polyphosphate (PG-CPP) scaffolds was better than that of homogeneous calcium polyphosphate (H-CPP) scaffolds, which was significant (p < 0.05) at each time point. This fact is also noted that, after 28 days of degradation, the compressive strength of PG-CPP scaffolds was even greater than that of primary H-CPP.

**Figure 11.** Relationships of compressive mechanical comparisons between graded cellular strcutrues (FGMs) and cellular structures. The lines represent the upper and lower predictive values of the cellular structures. Reprinted from [31] with permission from Elsevier.

#### Graded Cellular Bone Scaffolds http://dx.doi.org/10.5772/intechopen.69911 87

The mechanical properties of titanium alloys fabricated with SLM have been extensively studied under static and fatigue loadings [19, 21, 47–51]. Although titanium cellular structures have similar stiffness as bone tissues, fatigue is still a technical issue, which mainly resulted from poor design of the internal architectures [19, 52]. Naturally, the implant is expected to have similar properties as bone. Structural gradient has been observed in bone tissues, depending on required functions, such as load-bearing capacity and biological properties [5]. However, under axial compression, there is no benefit of graded cellular bone scaffolds over a uniformed bone scaffold with the same relative density [31]. **Figure 11** shows that CoCr graded cellular structures have similar elastic modulus, yield strength, and compressive strength as the uniformed cellular structures. CoCr cellular scaffolds and graded cellular scaffolds with the porosity ranging from 40 to 70% exhibit elastic modulus around 2.7–3.1 GPa, which is in the same level of bone stiffness [44, 53]. Therefore, mechanical properties of the cellular bone scaffolds are mainly related to the relative density rather than the

86 Scaffolds in Tissue Engineering - Materials, Technologies and Clinical Applications

The opposite findings were noticed with the bioceramic graded cellular scaffolds. Wang et al. [17] studied a novel calcium polyphosphate bioceramic scaffold with a graded pore structure similar to the bimodal structure of cortical and cancellous bones (**Figure 12**). The compressive strength of porosity-graded calcium polyphosphate (PG-CPP) scaffolds was better than that of homogeneous calcium polyphosphate (H-CPP) scaffolds, which was significant (p < 0.05) at each time point. This fact is also noted that, after 28 days of degradation, the compressive

**Figure 11.** Relationships of compressive mechanical comparisons between graded cellular strcutrues (FGMs) and cellular structures. The lines represent the upper and lower predictive values of the cellular structures. Reprinted from

strength of PG-CPP scaffolds was even greater than that of primary H-CPP.

porosity-graded characteristics.

[31] with permission from Elsevier.

**Figure 12.** Bioceramic graded cellular scaffolds with different graded porosities: (a) porous scaffold with a hollow center, (b) two graded porous layers, and (c) three graded porous layers [17]. Reprinted with the permission of a creative commons attribution license.

It is still controversial whether the graded cellular scaffolds can improve the mechanical properties of the constructs compared to homogeneous scaffolds. The bioceramic scaffolds with a porosity-graded structure in this study have a much better mechanical property. PG-CPP gives a prolonged deflection besides the elastic deformation which indicates that the porositygraded scaffolds exhibit a different fracture behavior from that of the homogeneous scaffolds. Furthermore, PG-CPP exhibits nonbrittle fractures whereas the H-CPP fractures catastrophically. Hence, it may be inferred that the mechanical properties of the porosity-graded scaffold can be substantially improved by a graded porous structure [17]. To end this, types of materials used for fabrications, internal structures of the scaffolds, and the design of graded scaffolds play an important role to control the mechanical properties of the graded cellular scaffolds.

Although both homogeneous and porosity-graded metallic bone scaffolds show a reasonable mechanical strength, the benefits of cellular graded structures optimized the functions required for such scaffold. Metallic cellular structures have lighter weight than solid metals, since an introduction of pores inside the materials. The stiffness of cellular structures can be tailored to match that of bone, according to the design and density of the scaffold's architectures. The stiffness of the cellular bone scaffolds is related to the porosity, following a nonlinear relationship as reported by Gibson and Ashby [43]. Therefore, the cellular bone scaffolds share more of the stress to an adjacent bone tissue than the solid counterpart due to the lower stiffness of the scaffolds.

Basically, the cellular structural implants with low stiffness properties reduce the peri-prosthetic bone-stress shielding, yet increase the bone-implant interface failure. **Figure 13** shows that more peri-implant bone-stress shielding occurs with the high-stiffed implant, and finally increases higher risks of the implant loosening. The initial stability of a femoral stem is necessary for biological bony ingrowth, which can be secured by minimizing relative micromotion

**Figure 13.** Influences of graded cellular structures on bone resorption. Reprinted from Ref. [31] with permission from Elsevier.

at the bone-implant interfaces. Excessive micromotions (>150 μm) allow fibrous connective tissue to grow, which prevents bone ingrowth between the contact surfaces and leads to aseptic loosening and failure of the implant [54–56]. Although homogeneous cellular implants increase stress sharing to peri-prosthetic bone due to lower construct stiffness, the greater interface failure due to excessive micromotions (>150 μm) adversely causes an initial implant instability and inhibits bone osseointegration.

According to computational study [57], the porous femoral stem with uniform relative density of 50% is approximately three times more flexible than the titanium stem. This implant can qualitatively simulate the behavior of an implant made out of tantalum foam using the welldefined cellular structures. The amount of bone resorption and the interface failure index of this stem are about 34% and 2.87, respectively, and the interface failure is maximum (0.71) at the edge of the proximal region. Compared to the solid titanium implant, the amount of bone resorption decreases by 50%, whereas the maximum interface failure increases about 40%. This shows that a decrease in the implant stiffness with uniform porosity distribution aiming at reducing bone resorption has the undesirable effect of increasing the risk of interface failure at the proximal region.

Kuiper and Huiskes [58] showed that the use of a graded material in an orthopedic stem can lead to a reduction of both stress shielding and bone-implant interface failure. The bone resorption and interface failure of graded cellular stems are 16% and 1.15, respectively [57]. The peak value of the local interface failure is 0.25. Compared to the titanium stem, both the amount of bone resorption and the peak of interface failure were decreased by 76 and 50%, respectively. With respect to the uniformly distributed cellular implant, the decrease in bone resorption and interface failure peak is of 53 and 65%, respectively. A graded cellular implant with optimized relative density distribution is thus capable of reducing concurrently both the conflicting objective functions. In particular, bone resorption reduces as a result of the cellular material which makes the implant more compliant; the interface stress and micromotions, on the other hand, are minimized by the optimized gradients of cellular material.

**Figure 14.** Strain (a) and stress (b) distributions along the proximal, middle, and distal parts of the tube surfaces under compression loads. Reprinted from Ref. [31] with permission from Elsevier.

Furthermore, for designing the scaffolds with functionally graded structures to mimic the graded structures of the host bone, the stress sharing to the adjacent bone is increased around 50% compared to the uniformed cellular bone scaffolds [31]. The degree of proximal bonestress sharing depends on a porosity-graded orientation. In case of intramedullary implants, such as femoral stems, the scaffolds with porosity gradient along a longitudinal plane present the maximum stress distribution to the proximal bone (**Figure 14**).

Graded cellular implants are functionally designed to match the mechanical and morphological properties of bones. While the graded metallic cellular implants exhibit mechanical properties similar to the uniformed cellular implants, the graded one is more capable of preventing bone-stress shielding and promoting larger amounts of bone ingrowth to the implants. The optimized designs of biomimetic graded cellular implants are complex depending on the functional objectives of the constructs.

### **4. Conclusion**

at the bone-implant interfaces. Excessive micromotions (>150 μm) allow fibrous connective tissue to grow, which prevents bone ingrowth between the contact surfaces and leads to aseptic loosening and failure of the implant [54–56]. Although homogeneous cellular implants increase stress sharing to peri-prosthetic bone due to lower construct stiffness, the greater interface failure due to excessive micromotions (>150 μm) adversely causes an initial implant

**Figure 13.** Influences of graded cellular structures on bone resorption. Reprinted from Ref. [31] with permission from

According to computational study [57], the porous femoral stem with uniform relative density of 50% is approximately three times more flexible than the titanium stem. This implant can qualitatively simulate the behavior of an implant made out of tantalum foam using the welldefined cellular structures. The amount of bone resorption and the interface failure index of this stem are about 34% and 2.87, respectively, and the interface failure is maximum (0.71) at the edge of the proximal region. Compared to the solid titanium implant, the amount of bone resorption decreases by 50%, whereas the maximum interface failure increases about 40%. This shows that a decrease in the implant stiffness with uniform porosity distribution aiming at reducing bone resorption has the undesirable effect of increasing the risk of interface failure

Kuiper and Huiskes [58] showed that the use of a graded material in an orthopedic stem can lead to a reduction of both stress shielding and bone-implant interface failure. The bone resorption and interface failure of graded cellular stems are 16% and 1.15, respectively [57]. The peak value of the local interface failure is 0.25. Compared to the titanium stem, both the amount of bone resorption and the peak of interface failure were decreased by 76 and 50%, respectively. With respect to the uniformly distributed cellular implant, the decrease in bone resorption and interface failure peak is of 53 and 65%, respectively. A graded cellular implant with optimized relative density distribution is thus capable of reducing concurrently both the conflicting objective functions. In particular, bone resorption reduces as a result of the cellular material which makes the implant more compliant; the interface stress and micromotions, on the other hand, are minimized by the optimized gradients of cellular

instability and inhibits bone osseointegration.

88 Scaffolds in Tissue Engineering - Materials, Technologies and Clinical Applications

at the proximal region.

Elsevier.

material.

The graded cellular bone scaffolds show logical concepts for bone tissue engineering. Cellular structures with graded pore sizes and porosities could mimic the graded structure in bones [5]. The advantage of graded cellular structures over the uniformed cellular structures is that the former provide more realistic environment for biological and mechanical functions. For stress-sharing orthopedic applications, the axially graded cellular structure demonstrated balanced mechanical performances and maximized proximal stress transfers around a peri-implant material [31]. Instead of using the uniformed cellular structures, we believe incorporating graded cellular structures in a structure like femoral prosthesis that will improve the load distribution in adjacent bones, greater bone osteointegration, and optimum nutrient permeability of the components [17].

### **Author details**

Sakkadech Limmahakhun1 and Cheng Yan2 \*

\*Address all correspondence to<EMAIL_ADDRESS>
1 Department of Orthopaedic Surgery, Faculty of Medicine, Chiang Mai University, Thailand

2 School of Chemistry, Physics and Mechanical Engineering, Queensland University of Technology, Australia

### **References**

provide more realistic environment for biological and mechanical functions. For stress-sharing orthopedic applications, the axially graded cellular structure demonstrated balanced mechanical performances and maximized proximal stress transfers around a peri-implant material [31]. Instead of using the uniformed cellular structures, we believe incorporating graded cellular structures in a structure like femoral prosthesis that will improve the load distribution in adjacent bones, greater bone osteointegration, and optimum nutrient permeability of the components [17].

and Cheng Yan2

\*Address all correspondence to<EMAIL_ADDRESS>
90 Scaffolds in Tissue Engineering - Materials, Technologies and Clinical Applications

Medicine. 2002;**13**(12):1245-1249

\*

1 Department of Orthopaedic Surgery, Faculty of Medicine, Chiang Mai University,

2 School of Chemistry, Physics and Mechanical Engineering, Queensland University of

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**Author details**

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### **Clinical Application of Macroporous Ceramic to Promote Bone Healing in Veterinary Clinical Cases** Provisional chapter Clinical Application of Macroporous Ceramic to Promote

DOI: 10.5772/intechopen.70341

Pedro Olivério Pinho, José Miguel Campos, Carla Mendonça, Ana Rita Caseiro, José Domingos Santos, Ana Colette Maurício and Luís Miguel Atayde Pedro Olivério Pinho, José Miguel Campos, Carla Mendonça, Ana Rita Caseiro, José Domingos Santos, Ana Colette Maurício

Bone Healing in Veterinary Clinical Cases

Additional information is available at the end of the chapter and Luís Miguel Atayde

http://dx.doi.org/10.5772/intechopen.70341 Additional information is available at the end of the chapter

#### Abstract

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Autogenous cancellous bone is the most effective material in promoting rapid healing and still considered the "gold standard" for evaluation of bone graft substitutes. The harvesting process to collect autologous bone is associated with complications and its availability is limited. Allogenic bone is another alternative with osteoconductive properties, and it act as a structural graft when applied in defects of long bones, but some disadvantages are also associated. The development of the bone grafts substitutes has gained tremendous popularity over the last two decades. Osteoconductive materials act as scaffolds were cells from the surrounding tissues with osteogenic capacities can lay new bone, and may be produced using different types of agents, such as bone products, ceramics, bioactive glasses, collagen, polymers, and composites. Bonelike® is produced by the incorporation of P2O5–CaO glass-based system within a hydroxyapatite matrix. Bonelike® Poro consists of polygonal granules with 2000–2800 μm and 4000–5600 μm of diameter with pore sizes range from 100 to 400 μm. This chapter will focus on the different techniques were this ceramic synthetic bone substitute was used to promote bone regeneration with special attention in both experimental and clinical cases of veterinary orthopaedics in dogs and cats, horses and ruminants, including results obtained with Bonelike®.

Keywords: bone regeneration, bone graft, bone substitutes, synthetic bone substitutes, orthopaedics, veterinary, clinical cases

© 2017 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

© The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and eproduction in any medium, provided the original work is properly cited.

### 1. Introduction

The bone healing process shares many similarities with soft tissues' healing, but, in contrast, the bone is the only tissue that has the capacity of healing without scar formation. Nevertheless, incomplete healing may occur, and the tissue other than the bone may be found at the healing site. Bone healing depends on adequate vascular supply and stability of the bone fragments. Proper bone healing can only occur after restoration of mechanical stability to achieve an ideal biomechanical environment. Many clinical situations may require additional osteosynthesis surgical procedures to acquire the biomechanical stability and immobilization necessary for correct bone regeneration and future functional recovery. Intrinsic mechanisms present unique histological characteristics that appear isolated or in association, depending on the bone fragment mobility [1–4]. The amount of bone callus produced depends on the stability of the fracture sides and usually increases with fracture instability. Spontaneous healing of complete fractures often occurs with highly unstable fragment ends and high interfragmentary strain (deformation occurring at the fracture site relative to the size of the gap), to a limit of 2% strain [1, 4]. Fracture healing under restricted motion induces an initial reduced amount of bone callus formation. This type of healing relies on fracture configuration and the implant's rigidity and may be achieved by external coaptation of the fracture or after gliding implant fixation with intramedullary pins and nails. With these surgical reconstructive methods, the amount of the callus produced is highly variable and dependent on the fracture configuration and the rigidity of the frame used [1, 3]. When the fixation is performed with a bone plate, the amount of callus formation will be different if the plate is not applied on the tension side of the bone, if the reduction is not perfect or when the plate lacks rigidity. In stable fractures using a rigid plate for osteosynthesis, there is significantly decreased callus formation between the bone fragments [1, 2]. This phenomenon is called 'primary healing', referring to the direct filling of the fracture site with the bone, without formation of significant callus (periosteal or endosteal). Healing under these conditions occurs by direct osteonal proliferation with interdigitation of bone fragments providing a very stable union. In opposition when the bone ends are separated by a gap inferior to 0.01 mm and the interfragmentary gap is less than 2%, the primary osteonal reconstruction results in direct formation of lamellar bone, oriented in the normal direction [1, 3]. This is termed contact healing and is initiated by osteoclasts from osteons near the fracture line. Another context for direct bone healing is observed in gaps from 800 μm to 1 mm and interfragmentary strain less than 2%, and gap healing is designated. Here, bone union and Haversian remodellation are separated by sequential process steps. Fracture site is filled directly by intermembranous bone formation, but the newly formed lamellar bone is oriented perpendicular to its long axis, and, later on, it undergoes secondary osteonal reconstruction (remodellation). In this way, the fracture repair depends on two important events: the rate of new bone growth (that replaces the stiffness and strength of the bone) and the strength and duration of the implant (maintenance of its function until it collapses and fails) [1–4]. When new bone formation is compromised and/or a risk of implant failure is expected, promotion of fracture healing is indicated, in order to enhance new bone formation, allowing for corrected bone structure formation and organization [1]. Fractures with impaired bone repair mechanisms or fixation failure will result in a non-union condition. Adequate bone repair depends on four crucial elements: an osteoconductive matrix, an osteoinductive signal, an osteogenic cell acting in response to that signal and an adequate blood supply [1, 3, 4]. There are several clinical situations where it is important to promote and enhance bone healing process in order to restore the original bone structure and function, both in human and veterinary medicine. These include traumatic injuries or tumour resections with substantial and irregular bone loss, gap filling following corrective osteotomy, arthrodesis or arthroplasty, spinal fusion, non-union or delayed bone union, metabolic diseases and local or systemic disease in aged patients. In these situations, bone regeneration is compromised, and the bone defect exceeds the intrinsic biological restoration mechanisms. These clinical cases occasionally result in unsatisfactory outcomes and are a challenging scenario to orthopaedic surgeons. When facing such clinical problems, different treatment strategies can be used to improve new bone formation, avoiding the formation of bone with inferior quality to the original [1, 3, 5–7].

1. Introduction

The bone healing process shares many similarities with soft tissues' healing, but, in contrast, the bone is the only tissue that has the capacity of healing without scar formation. Nevertheless, incomplete healing may occur, and the tissue other than the bone may be found at the healing site. Bone healing depends on adequate vascular supply and stability of the bone fragments. Proper bone healing can only occur after restoration of mechanical stability to achieve an ideal biomechanical environment. Many clinical situations may require additional osteosynthesis surgical procedures to acquire the biomechanical stability and immobilization necessary for correct bone regeneration and future functional recovery. Intrinsic mechanisms present unique histological characteristics that appear isolated or in association, depending on the bone fragment mobility [1–4]. The amount of bone callus produced depends on the stability of the fracture sides and usually increases with fracture instability. Spontaneous healing of complete fractures often occurs with highly unstable fragment ends and high interfragmentary strain (deformation occurring at the fracture site relative to the size of the gap), to a limit of 2% strain [1, 4]. Fracture healing under restricted motion induces an initial reduced amount of bone callus formation. This type of healing relies on fracture configuration and the implant's rigidity and may be achieved by external coaptation of the fracture or after gliding implant fixation with intramedullary pins and nails. With these surgical reconstructive methods, the amount of the callus produced is highly variable and dependent on the fracture configuration and the rigidity of the frame used [1, 3]. When the fixation is performed with a bone plate, the amount of callus formation will be different if the plate is not applied on the tension side of the bone, if the reduction is not perfect or when the plate lacks rigidity. In stable fractures using a rigid plate for osteosynthesis, there is significantly decreased callus formation between the bone fragments [1, 2]. This phenomenon is called 'primary healing', referring to the direct filling of the fracture site with the bone, without formation of significant callus (periosteal or endosteal). Healing under these conditions occurs by direct osteonal proliferation with interdigitation of bone fragments providing a very stable union. In opposition when the bone ends are separated by a gap inferior to 0.01 mm and the interfragmentary gap is less than 2%, the primary osteonal reconstruction results in direct formation of lamellar bone, oriented in the normal direction [1, 3]. This is termed contact healing and is initiated by osteoclasts from osteons near the fracture line. Another context for direct bone healing is observed in gaps from 800 μm to 1 mm and interfragmentary strain less than 2%, and gap healing is designated. Here, bone union and Haversian remodellation are separated by sequential process steps. Fracture site is filled directly by intermembranous bone formation, but the newly formed lamellar bone is oriented perpendicular to its long axis, and, later on, it undergoes secondary osteonal reconstruction (remodellation). In this way, the fracture repair depends on two important events: the rate of new bone growth (that replaces the stiffness and strength of the bone) and the strength and duration of the implant (maintenance of its function until it collapses and fails) [1–4]. When new bone formation is compromised and/or a risk of implant failure is expected, promotion of fracture healing is indicated, in order to enhance new bone formation, allowing for corrected bone structure formation and organization [1]. Fractures with impaired bone repair mechanisms or fixation failure will result in a non-union condition. Adequate bone repair depends on four crucial elements: an osteoconductive matrix, an

96 Scaffolds in Tissue Engineering - Materials, Technologies and Clinical Applications

Autogenous cancellous bone grafting (ACBG) is used in veterinary orthopaedic surgery as a bone void filler to improve bone healing in the treatment of bone defects in low-grade fractures in both mechanical and biological assessment score, arthrodesis, delayed unions or non-unions, enhancement of fracture healing, periprosthetic coating, spinal fusion procedures and void filler of bone gaps resulting from fractures, osteotomy, ostectomy, arthrodesis or tumour resection [1, 7]. The 'gold standard' when evaluating bone graft substitutes is still considered to be ACBG. Cancellous bone graft provides osteoconductive properties, acts as a scaffold for osteoprogenitor cells and delivers viable cells without the risk of immune reactions or infectious disease transmission. However, its use is associated with some limitations including the need of an additional surgery for harvesting cancellous bone, donor-site morbidity and limited amount of bone graft. The last point can result in an insufficient amount to completely fill the defect which may implicate the need for harvesting from more than one donor site. In humans, the second surgery for autologous bone graft collections is associated with 25% of morbidity, with major complications occurring in 3–4% of the patients. Complications include pain, sepsis, stress fractures, intraoperative haemorrhage, increased anaesthetic and surgical times and limited supply [8, 9]. Limited supply is more critical in small dogs and toy breeds, in cats or in animals that have been previously submitted to bone graft harvest. In those cases, the harvesting from the humerus may be the best option, because higher amounts of bone can be collected when comparing to the tibia while also presenting accelerated healing and complete restoration cancellous bone. Another drawback is its lack of strength of ACBG, hampering its use as a structural graft [1, 10]. Cortical allograft could be an alternative, which provides structural strength along with osteoinductive and osteoconductive properties. Nevertheless, the transmission of infectious diseases and adverse immune reactions are important risks to consider [11].

In the case of human patients, demineralized bone matrix (DBM) is the most common source of partially purified bone-inducting factors used. During the demineralization process, allogenic bone is chemically sterilized to preserve its osteoinductive properties from the original bone collagen network and molecular signalling. Commercial forms of DBM are also available for canine patients, where immune reactions are reduced by the removal of the periosteum, cartilage and bone marrow and by freezing process [6, 7, 12–14].

In the last decades, the well-known disadvantages of autografts and allografts have encouraged the development of bone synthetic substitutes to be employed as bone grafts, reflecting in the increase of the clinical application of these types of biomaterials. An ideal bone substitute should be biocompatible/safe; should be resorbable, with a similar mechanical resistance as the cortical bone; should have osteoconductive, osteoinductive and osteogenic properties; and should be easily handled and sterilized. The bone substitute should not cause any adverse systemic or local reaction; should provide a favourable environment to be colonized by blood vessels, cells and growth factors; and should be obtained at a low cost [7, 9].

These biomaterials are available for use in veterinary orthopaedics, alone or in combination with other strategies, such as osteoinductive bone morphogenetic proteins (BMPs) or cellbased treatments [6, 7, 11, 14]. They are composed of different materials with osteoconductive properties, such as bone products, ceramics, polymers and composites. The clinical application of each material is defined by its properties, which depend on composition and physical characteristics (e.g. granulometry, shape, pore size and interconnective porosity). The latter is determined by the manufacturing technique used [7, 9]. The synthetic bone substitutes offer an ideal substrate for bone cell colonization and consequent new bone formation [7].

Osteoconductive materials are grouped in two main categories. The first category includes ceramic-based bone substitutes, and the second category includes polymer-based bone graft substitutes (less commonly used). The most popular ceramic-based bone substitutes are calcium sulphate, bioactive glass and calcium phosphate. Calcium sulphate was the first material used as a bone graft substitute in clinical field, where it showed to be user-friendly, inexpensive, readily available and stable for filling-in bone defects, without a negative effect in bone healing. However, it presents fast absorption rates, leading to the loss of mechanical properties before equivalent new bone formation, constituting its main disadvantage and limiting its use in relevant clinical cases. Bioactive glass was designed as a bone graft for dental applications. Its use in orthopaedic surgery seems to be limited by its brittleness, radiopacity (which compromises radiographic evaluation of bone healing) and prolonged resorption times. Compared to bioactive glass, calcium phosphate is less radiopaque with faster reabsorption, but its osteointegration rate is highly variable, depending on its crystal size and stoichiometry. Tricalcium phosphate compounds are available in different presentations, including tricalcium phosphate, hydroxyapatite and a combination of the two.

Regardless of the mineral detail of its composition, the final formulations/presentation forms are of extreme relevance, concerning both its external shape and internal architecture. External shape and granule size will determine its suitability for particular lesion applications, considering effective size and ease of access for implantation. Its architecture is determinant for the resolution of the bone defect since bone ingrowth is dependent on the pore size. According to Ragetly et al., a minimum pore size of 100 μm is required for bone ingrowth and for optimal promotion of ingrowth pores should have a granulometry between 300 and 500 μm [7, 9, 15].

### 2. Preclinical trials

In the last decades, the well-known disadvantages of autografts and allografts have encouraged the development of bone synthetic substitutes to be employed as bone grafts, reflecting in the increase of the clinical application of these types of biomaterials. An ideal bone substitute should be biocompatible/safe; should be resorbable, with a similar mechanical resistance as the cortical bone; should have osteoconductive, osteoinductive and osteogenic properties; and should be easily handled and sterilized. The bone substitute should not cause any adverse systemic or local reaction; should provide a favourable environment to be colonized by blood vessels, cells and growth factors; and should be obtained at a low

98 Scaffolds in Tissue Engineering - Materials, Technologies and Clinical Applications

These biomaterials are available for use in veterinary orthopaedics, alone or in combination with other strategies, such as osteoinductive bone morphogenetic proteins (BMPs) or cellbased treatments [6, 7, 11, 14]. They are composed of different materials with osteoconductive properties, such as bone products, ceramics, polymers and composites. The clinical application of each material is defined by its properties, which depend on composition and physical characteristics (e.g. granulometry, shape, pore size and interconnective porosity). The latter is determined by the manufacturing technique used [7, 9]. The synthetic bone substitutes offer an ideal substrate for bone cell colonization and consequent new bone forma-

Osteoconductive materials are grouped in two main categories. The first category includes ceramic-based bone substitutes, and the second category includes polymer-based bone graft substitutes (less commonly used). The most popular ceramic-based bone substitutes are calcium sulphate, bioactive glass and calcium phosphate. Calcium sulphate was the first material used as a bone graft substitute in clinical field, where it showed to be user-friendly, inexpensive, readily available and stable for filling-in bone defects, without a negative effect in bone healing. However, it presents fast absorption rates, leading to the loss of mechanical properties before equivalent new bone formation, constituting its main disadvantage and limiting its use in relevant clinical cases. Bioactive glass was designed as a bone graft for dental applications. Its use in orthopaedic surgery seems to be limited by its brittleness, radiopacity (which compromises radiographic evaluation of bone healing) and prolonged resorption times. Compared to bioactive glass, calcium phosphate is less radiopaque with faster reabsorption, but its osteointegration rate is highly variable, depending on its crystal size and stoichiometry. Tricalcium phosphate compounds are available in different presentations, including tricalcium phosphate, hydroxyapatite and a combination

Regardless of the mineral detail of its composition, the final formulations/presentation forms are of extreme relevance, concerning both its external shape and internal architecture. External shape and granule size will determine its suitability for particular lesion applications, considering effective size and ease of access for implantation. Its architecture is determinant for the resolution of the bone defect since bone ingrowth is dependent on the pore size. According to Ragetly et al., a minimum pore size of 100 μm is required for bone ingrowth and for optimal promotion of ingrowth pores should have a granulometry between 300 and

cost [7, 9].

tion [7].

of the two.

500 μm [7, 9, 15].

A series of regulatory steps are essential in the development of biomaterials, requiring its physicalchemical characterization; in vitro validation, and at an intermediate stage, in vivo suitability; safety; and performance assessment: biocompatibility, osteoconduction, osteointegration and osteoinduction [16]. In sight of the ethical issues associated to the use of animals for experimental purposes, extensive efforts aim at refining in vitro methods as valid alternatives. Unfortunately, current in vitro models still underachieve in replicating the tissue response of a live animal to a bone substitute [16–18].

The preparation and conduction of preclinical studies are bound to a number of sequential phases. First, the proposed materials ought to be tested in noncritical-sized defects, allowing for the preliminary assessment of multiple material samples' behaviour in defined in vivo conditions. This feature is very important on the initial screening of a biomaterial's in vivo behaviour, enabling the choice of the chemical composition and format of the biomaterial with the most potential for more challenges [19–21]. Furthermore, noncritical-sized defect ensures fast and reliable healing process, allowing for the observation of the various stages of bone healing and biomaterial degradation [21, 22]. Once the ideal composition is chosen and biocompatibility is analyzed, the material may be used directly in clinical trials, or if found necessary, in increasingly critical defects, to confirm the results obtained in the first approach with noncritical defects, and to determine its limits of efficiency and performance [21].

One of the fundamental aspects when choosing the most adequate animal models to test bone replacement materials is the size of the bone defect amenable to assess [23, 24]. A critical-sized bone defect (CSBD) is defined as the smallest bone defect that will not heal spontaneously during the lifetime in a particular bone and species of animal [23, 25]. In a more detailed description, a CSBD has been referred to as a defect that has less than 10% bony regeneration during the lifetime of the animal or duration of the experiment [26, 27]. Although the smallest size that creates a defect designated as 'critical' is not a well-established concept, it has been defined as a segmental bone deficiency of length exceeding 2–2.5 times the diameter of the affected bone [28]. Some animal studies suggest that CSBD in sheep could be approximately three times the diameter of the diaphysis. Therefore, a critical defect in long bone cannot simply be defined by its size, but may also be dependent on the species phylogenetic scale, the location in the skeleton, the surrounding soft tissue envelope and the load bearing on the affected limb.

Bone colonization of macroporous biphasic calcium phosphate (MBCP) ceramics implanted in different sites (femur, tibia and calvaria) on a critical-sized defect in two animal models (rats and rabbits) showed bone ingrowth in all MBCP-implanted sites but with distinct rates. Bone colonization appeared statistically higher in the femur of the rabbits (48.5%) compared to the tibia (12.6%) and calvaria (22.9%) sites. As such, the comparison of results between animals or different bone defect locations is subject to bias, so a well-conducted study and fully validated animal models are essential in the development of new synthetic bone substitutes [29]. Furthermore, the host's age, metabolic and systemic conditions and comorbidities also affect the defect's healing potential [30].

Attempts to repair a CSBD only lead to the formation of fibrous connective tissue rather than the bone [31]. For practical purposes, if there is no more than 30% of the mineralization area after 52 weeks, lifelong incomplete bone healing is to be assumed [32]. The incapacity of natural healing when left untreated represents the negative control, so that the osteogenic potential of the material being tested can be considered unequivocal. Furthermore, CSBD should heal with appropriate treatment, and the autologous cancellous bone grafts are still named as the gold standard or positive control. Any new treatment based on bone tissue engineering should be tested and compared with these two landmarks [25]. Given the clinical targets of such applications (mostly aged or health-impaired patients), animals should be skeletally mature, in order to avoid misleading results deriving from the superior potential of the young animals to regenerate bone defects [23].

Recently, the osteoinductive ability of porous calcium phosphate ceramics was studied in four animal species through the implantation of cylinders of hydroxyapatite/tricalcium phosphate (HA/TCP) (in the proportion of 60 and 40%, respectively) in dorsal muscles in dogs, rabbits, rats and mice. After 1 year, the implants were removed, and histopathology tests were conducted with haematoxylin/eosin staining and Masson's trichrome staining to observe the new bone tissue formation. The study concluded that the material is biocompatible and biologically safe (no tumour or any atypical cells were present) and would be considered as a potential for bone substitute. Apart from the rat's groups, there was new bone and bone marrow tissue development in large amounts. The osteogenic ability of the implant was superior in mice, followed by dog and rabbit [33].

Small rodents, rabbits, dogs and small ruminants are the most popular model species for bone regeneration studies, bracing several of the target species for veterinary clinical applications. Allografts and bone graft substitutes have not been fully evaluated in cats, and for that reason, Dorea et al. compared the efficacy and safety of a Bioglass® with that of autogenous and allogeneic cancellous bone graft in this species. Four defects in the lateral diaphyseal cortex of the femur with a diameter of 4.0 mm were created in each animal. One hole was filled with autogenous CBG, another with allogeneic CBG and a third one with Bioglass®. The fourth defect was left unfilled. The healing process was monitored every 2 weeks by X-ray. After 6 weeks, cats were euthanized, and the resolution of the defects was appreciated. The study indicated an acceptable bone regeneration in all defects. Although with a slower healing rate, Bioglass® showed to be an acceptable alternative to ACBG in cats [34].

### 3. Clinical cases

The use of synthetic bone grafts, in veterinary medicine, has been increasing in the last years, but even so its application is substantially lower than in human medicine. This is mainly due to the costs involved in the use of biomaterials, but nowadays with the development of society and consequent attention given to the animals, there was an increase in the availability of owners to invest in the treatment of their animals. Besides the sentimental value of certain animals, it is not to be despised the growth of the economic valuation of certain types of animals.

### 3.1. The use of synthetic bone grafts in small animal clinical cases

Attempts to repair a CSBD only lead to the formation of fibrous connective tissue rather than the bone [31]. For practical purposes, if there is no more than 30% of the mineralization area after 52 weeks, lifelong incomplete bone healing is to be assumed [32]. The incapacity of natural healing when left untreated represents the negative control, so that the osteogenic potential of the material being tested can be considered unequivocal. Furthermore, CSBD should heal with appropriate treatment, and the autologous cancellous bone grafts are still named as the gold standard or positive control. Any new treatment based on bone tissue engineering should be tested and compared with these two landmarks [25]. Given the clinical targets of such applications (mostly aged or health-impaired patients), animals should be skeletally mature, in order to avoid misleading results deriving from the superior potential of

Recently, the osteoinductive ability of porous calcium phosphate ceramics was studied in four animal species through the implantation of cylinders of hydroxyapatite/tricalcium phosphate (HA/TCP) (in the proportion of 60 and 40%, respectively) in dorsal muscles in dogs, rabbits, rats and mice. After 1 year, the implants were removed, and histopathology tests were conducted with haematoxylin/eosin staining and Masson's trichrome staining to observe the new bone tissue formation. The study concluded that the material is biocompatible and biologically safe (no tumour or any atypical cells were present) and would be considered as a potential for bone substitute. Apart from the rat's groups, there was new bone and bone marrow tissue development in large amounts. The osteogenic ability of the implant was

Small rodents, rabbits, dogs and small ruminants are the most popular model species for bone regeneration studies, bracing several of the target species for veterinary clinical applications. Allografts and bone graft substitutes have not been fully evaluated in cats, and for that reason, Dorea et al. compared the efficacy and safety of a Bioglass® with that of autogenous and allogeneic cancellous bone graft in this species. Four defects in the lateral diaphyseal cortex of the femur with a diameter of 4.0 mm were created in each animal. One hole was filled with autogenous CBG, another with allogeneic CBG and a third one with Bioglass®. The fourth defect was left unfilled. The healing process was monitored every 2 weeks by X-ray. After 6 weeks, cats were euthanized, and the resolution of the defects was appreciated. The study indicated an acceptable bone regeneration in all defects. Although with a slower healing rate,

The use of synthetic bone grafts, in veterinary medicine, has been increasing in the last years, but even so its application is substantially lower than in human medicine. This is mainly due to the costs involved in the use of biomaterials, but nowadays with the development of society and consequent attention given to the animals, there was an increase in the availability of owners to invest in the treatment of their animals. Besides the sentimental value of certain animals, it is not

the young animals to regenerate bone defects [23].

100 Scaffolds in Tissue Engineering - Materials, Technologies and Clinical Applications

superior in mice, followed by dog and rabbit [33].

3. Clinical cases

Bioglass® showed to be an acceptable alternative to ACBG in cats [34].

to be despised the growth of the economic valuation of certain types of animals.

Some isolated case reports in small animals using ceramic-based bone graft substitutes have been published in the last decades.

The application of β-tricalcium phosphate (TCP) has already stepped out of preliminary preclinical assays to veterinary patients' applications. Izumisawa et al. successfully solved pes varus in two miniature Dachshunds using a wedge of synthetic β-TCP to fill the gaps created by tibial corrective open osteotomies. According to the authors, 2 months after the surgery, the edge of osteotomies was integrated with the bone. The bone plates and screws were removed after 4 months, and, by then, the TCP wedges were completely resorbed and the osteotomy bone was remodelled. In both cases the use of synthetic bone graft TCP avoided the need of a second surgery to harvest autologous cancellous bone graft. In both cases, the postoperative angles were corrected and maintained during the follow-up period, and, morphologically, the body of the tibia in the affected hind limb nearly equalled that found in healthy limbs. Dogs were able to walk a few days postoperatively, and the authors concluded that the corrective transverse-opening osteotomy together with synthetic bone graft substitute β-TCP and veterinary T plate fixation is an effective method for the treatment of pes varus in small-breed dogs [35]. A single case of tarsal joint fusion with β-TCP and platelet-rich plasma (PRP) was reported by Hauschild et al., and ACBG, and hence the need of additional surgery to harvest the bone graft, was successfully avoided. The healing process was uneventful and was complete at 4 months after surgery, at which time the dog presented no signs of lameness [36].

Another report on the clinical use of β-TCP as synthetic cancellous bone graft in veterinary orthopaedics was presented by Franch et al. They retrospectively studied 13 clinical cases, where granules of β-TCP, with an irregular form and interconnected porous structure (without dead ends, mean porosity of 60% and mean pore size of 250 μm), were mixed with fresh blood and used as void filler in subcritical-sized bone defects in long bones. The β-TCP used in these clinical series was commercially available in sterile vials containing 2 g of granules with 99% of pure phase. The clinical cases are summarized in Table 1. All but one case achieved complete bone union, and radiographic bone ingrowth was at 100% in 10 cases, 90% in 1 case and 75% in another case. The publication reported excellent clinical results confirming the biocompatibility and usefulness of β-TCP as a synthetic bone graft for moderate to large subcritical bone defects with initially expected good biological conditions (blood supply, cellular activity, etc.), on which the main problem is to provide a structural scaffold to allow bone and capillary ingrowth and the healing of the defect [37]. The same author described the treatment of a distal radius atrophic non-union in a 1-year-old male Yorkshire terrier using a 3D-printed β-TCP scaffold with rhBMP-2 (TruScient®) to create a scaffold with the same shape as the defect. After the removal of the bone plate (10 months), load started to transmit along the bone axis, reducing the potential risk of stress protection. Eighteen months after surgery, the scaffold was no longer visible, and complete corticalization of the regenerated bone area was observed on computed tomography (CT) scan evaluation. Given the results, the author suggests that the 3D-printed β-TCP scaffold with rhBMP-2 is an excellent bone substitute, due to good osteoinductive properties given by rhBMP-2 complemented with good osteoconductive potential provided by the open-interconnected macroporosity from the β-TCP scaffold [38].

Type of

Author

 Type of

No. of

Species Breed

 Sex-ageweight

Problem

Reduction device Bone defect grading

Results Consolidation

Percentage

Functional

new bone

recovery

formation

time (weeks)

Removal of the

Ref.

implant

macroporous

cases

ceramic

treatment

β-TCP

 Izumisawa

β-TCP wedge

1

 Canine Dachshund ♂-10m

 Pes varus and grade 3

Open-wedge

NCSD

 About 6

100

 Excellent

 4 months

[35]

osteotomy in the

distal tibia and

veterinary 1.5/2.0

T plate Open-wedge

NCSD

 About 6

100

 Excellent

 4 months

> osteotomy in the

distal tibia and veterinary 1.5/2.0

T plate Pancarpal

NCSD

 12

100

 Excellent

[37]

102 Scaffolds in Tissue Engineering - Materials, Technologies and Clinical Applications

arthrodesis with

2.7 DCP plate

lateral patellar luxation

shaped

2

Franch

Irregular

3

 Canine Shiba Inu ♀-7 y-12kg Traumatic carpal

hyperextension

interconnected

granules of β-

TCP (60%

porosity)

4

5

 Canine Crossbreed

♂-12y-10kg

 Fractures of the distal

radius (cranial bone

loss), ulna, III–IV–V

metacarpal bones and I

phalanx and

comminuted carpal

bone

6

7

 Canine German Shepherd

♂-7y-37kg

 Tibial fracture

crossbreed

8

9

 Canine English Bulldog

♂-10y-11kg

 Tarsometatarsal

Cross pinning

NCSD

 12

75

 Good

tarsometatarsal

arthrodesis

luxation

 Canine Belgium Shepherd

♀-8y-34kg

 Distal radius fracture

 3.5 DCP plate

 NCSD

 10

100

 Excellent

 Canine German Shepherd

♂-2y-38kg

 Comminuted mid-shaft

Interlocking nail

NCSD

 9

90

 Excellent

> with four cerclage

wires

3.5 DCP plate

 NCSD

 10

100

 Good

femoral fracture

fracture-luxation

 Canine Shiba Inu

♀-7y-12kg

 Traumatic carpal

Pancarpal

NCSD

 10

100

 Excellent

> arthrodesis with

2.7 DCP plate

Pancarpal

NCSD

 8 (radius)

100

 Excellent

> 12 (carpus)

arthrodesis, radial

and III metacarpal

fractures treated together with 2.7

DCP plate

hyperextension

et al. 2006

 Canine Dachshund ♂-9m

 Pes varus and grade 3

lateral patellar luxation

et al. 2005

Type of

Author

 Type of

No. of

Species Breed

 Sex-ageweight

Problem

Reduction device Bone defect grading

Results Consolidation

Percentage

Functional

new bone

recovery

formation

time (weeks)

Removal of the

Ref.

implant

macroporous

cases

ceramic

treatment

β-TCP

Segal and

β-TCP

 17 Canine Crossbreed

♂-7y-25kg

 Highly comminuted

IM pin through

2 CSBD 71mm

22–63

70% of

Successful

No due to the

[41]

femoral

long-term

associated costs

clinical

cage at 22

weeks and

outcome

100% at 63

weeks

segmental

femoral bone

defect and

27mm

segmental

radial defect

the cage and a

locking plate

fractures (femur, radius

and ulna) that fail to

previous fixation

Shani 2010

+ACBG (1:1)

and a titanium

mesh cage Contralateral

Petazzoni

Silicate

18 Feline Maine

bone

2016

calcium

Coon

♀-5y-5kg

 Non-union and bone

2.5 mm locking

CSBD 60mm

At 13 satisfactory

Excellent

6 weeks–18 weeks

[43]

104 Scaffolds in Tissue Engineering - Materials, Technologies and Clinical Applications

long-term

good; at 22 weeks

follow-up

implant fails and

revision surgery

was done; the plate

was removed 20

weeks later [46]

corresponding

healing. 22 weeks

non-union. 20

to 48% of mechanical

weeks post revision surgery

with ACBG healing of non-

union.

axis on the

sagittal plane

plate

loss

phosphate

widening

and transfer

and silicate

calcium

phosphate

CRM +

Massie

CRM

19 Canine Italian 20 Canine Rat Terrier 1y-7.1kg

21 Canine Golden 22 Canine English 23 Canine Schipperke 2y - 7.1kg

24 Canine Border 25 Canine Cocker 26 Canine Miniature

Poodle

2y - 3kg

 Dystrophic non-union

of R an L radius/ulna

Spaniel

10y - 12kg

Atrophic-oligotrophic

3.5 mm cortical

1 mm

 20

> screw and washer

+ 2.4 mm 4-hole

LCP

R 7-hole LC-DCP

2 mm

 8

> and L 8-hole LC-

DCP

non-union of L lateral

humeral condyle

Collie

9y - 16kg

Atrophic-oligotrophic

ESF hybrid

 10 mm

 10

> non-union of L tibia

Pointer

11y - 29.7kg

Atrophic-oligotrophic

5.5 mm cortical lag

1 mm

 16

> screw + 2.4 mm 8-

hole LCP

non-union of R lateral

humeral condyle Atrophic-oligotrophic

2.4 mm 10-hole

5 mm

 6

LCP

non-union of R radius/

ulna

Retriever

of the L tibia

12y - 24.3kg Dystrophic non-union

Greyhound

3Y-3kg

 Non-union of left and

2.0 mm 14-hole

R 65 mm L 2.5

R 9 and L 10

mm

LCP

right radius/ulna

non-union of R tibia

Atrophic-oligotrophic

2.0 mm 13-hole

32 mm

 10

LCP

3.5 mm 10-hole

16 mm

 10

LCP

rhBMP-2

et al. 2017 Table1.Descriptionofreportedclinicalcasesusingceramic-basedbonesubstitutesinsmallanimals,basedonavailableliterature.

NCSD, noncritical-sized

 bone defects; R, right.

#### Clinical Application of Macroporous Ceramic to Promote Bone Healing in Veterinary Clinical Cases http://dx.doi.org/10.5772/intechopen.70341 105

Treatment of long bones affected with large segmental defects is challenging in human and veterinary medicine [39]. Different surgical approaches have been developed, but most of these techniques present major complications. Cortical allografts are prone to infection and are rarely fully incorporated [11, 39]. Distraction osteogenesis should only be used in docile animals and cooperative owners because it needs a lot of adjustments and cleaning of the fixator besides having to be kept in place for prolonged periods that could lead to loss of stability and discomfort [40]. Osteoconductive materials and osteoinductive substances, such as canine-demineralized bone matrix, canine autogenous bone grafts and bone morphogenetic proteins, have been used for the treatment of bone defects and comminuted fractures, alone or in combinations with bioceramics [6, 12]. Another alternative treatment that has been used with success is the clinical application of titanium mesh cages [41], nevertheless with extreme variable outcomes [42].

One case report in a dog with large segmental femoral and radial bone defects that failed in the first attempt of surgical stabilization and finally treated with a combined approach with titanium mesh, β-TCP, ACBG into the defects and fractures was stabilized with locking plates. Titanium mesh allowed maintaining bone fragments and grafts within the defect being further, between 22 and 63 weeks, surrounded by a bony bridge. Clinical outcome was successful without visible lameness in pace, but it was visible, while running and reduction in range of motion with crepitation were noticed on the stifle joint. Those limitations could be associated with original trauma or complications of fixation methods. Long-term follow-up (greater than 1 year) shows a satisfactory active mobilization of the limb. In this case, the technique has been simple and should be an alternative for treatment of long segmental bone defects. However, further systematic clinical studies are needed in order to evaluate the efficacy, complications and spectrum of the clinical use of this method [41].

A case report of successful reconstruction of a long segmental tibial defect in a 5-year-old 5 kgspayed female Maine Coon cat were transverse distraction osteogenesis in the contralateral tibia was used to create free autograft for filling the defect. After fixation of the bone fragments with a locking plate, the autograft was transferred to the defect in the contralateral tibia, and the remnant space was filled with silicate calcium phosphate bone graft substitute. By 27 months, both tibias were healed; implants had been removed with an excellent functional outcome [43].

Combination of BMPs for augmentation of bone regeneration associated with bioceramics in the treatment of non-union fractures of long bones was mentioned in a review as a single case of non-union of the distal radius in a Pomeranian [44] and in a distal radius atrophic non-union in a 1-year old male Yorkshire terrier by Franch [45]. A recent prospective longitudinal cohort study of the use of compression-resistant matrix (CRM) immersed in rhBMP-2 in the treatment of 11 non-union long bone fractures in 9 dogs was evaluated, and treatment was successful with healing time from 7 to 20 weeks (median 10 weeks) and return to full or acceptable function in all dogs [46].

Non-weight-bearing bone lesions can also greatly benefit from the application of synthetic bone substitutes, such as those resulting from dental procedures. Boudrieau et al. published in 2004 one case report of treatment of severe mandibular malocclusion after partial mandibulectomy in a 14-month-old golden retriever using fenestrated, monocortical rib grafts with rhBMP-2 in a collagen-TCP sponge that showed new bone formation after 3 months with reconstruction of the defect. One year later, the bone was collected, and the new bone was robust with evidence of continued remodelling. No major complications were noticed, and bony remodelling was evident at 4-year follow-up [47].

Treatment of long bones affected with large segmental defects is challenging in human and veterinary medicine [39]. Different surgical approaches have been developed, but most of these techniques present major complications. Cortical allografts are prone to infection and are rarely fully incorporated [11, 39]. Distraction osteogenesis should only be used in docile animals and cooperative owners because it needs a lot of adjustments and cleaning of the fixator besides having to be kept in place for prolonged periods that could lead to loss of stability and discomfort [40]. Osteoconductive materials and osteoinductive substances, such as canine-demineralized bone matrix, canine autogenous bone grafts and bone morphogenetic proteins, have been used for the treatment of bone defects and comminuted fractures, alone or in combinations with bioceramics [6, 12]. Another alternative treatment that has been used with success is the clinical application of titanium mesh cages [41], nevertheless with extreme

106 Scaffolds in Tissue Engineering - Materials, Technologies and Clinical Applications

One case report in a dog with large segmental femoral and radial bone defects that failed in the first attempt of surgical stabilization and finally treated with a combined approach with titanium mesh, β-TCP, ACBG into the defects and fractures was stabilized with locking plates. Titanium mesh allowed maintaining bone fragments and grafts within the defect being further, between 22 and 63 weeks, surrounded by a bony bridge. Clinical outcome was successful without visible lameness in pace, but it was visible, while running and reduction in range of motion with crepitation were noticed on the stifle joint. Those limitations could be associated with original trauma or complications of fixation methods. Long-term follow-up (greater than 1 year) shows a satisfactory active mobilization of the limb. In this case, the technique has been simple and should be an alternative for treatment of long segmental bone defects. However, further systematic clinical studies are needed in order to evaluate the efficacy, complications

A case report of successful reconstruction of a long segmental tibial defect in a 5-year-old 5 kgspayed female Maine Coon cat were transverse distraction osteogenesis in the contralateral tibia was used to create free autograft for filling the defect. After fixation of the bone fragments with a locking plate, the autograft was transferred to the defect in the contralateral tibia, and the remnant space was filled with silicate calcium phosphate bone graft substitute. By 27 months, both tibias were healed; implants had been removed with an excellent functional

Combination of BMPs for augmentation of bone regeneration associated with bioceramics in the treatment of non-union fractures of long bones was mentioned in a review as a single case of non-union of the distal radius in a Pomeranian [44] and in a distal radius atrophic non-union in a 1-year old male Yorkshire terrier by Franch [45]. A recent prospective longitudinal cohort study of the use of compression-resistant matrix (CRM) immersed in rhBMP-2 in the treatment of 11 non-union long bone fractures in 9 dogs was evaluated, and treatment was successful with healing time from 7 to 20 weeks (median 10 weeks) and return to full or acceptable

Non-weight-bearing bone lesions can also greatly benefit from the application of synthetic bone substitutes, such as those resulting from dental procedures. Boudrieau et al. published in 2004

variable outcomes [42].

outcome [43].

function in all dogs [46].

and spectrum of the clinical use of this method [41].

According to the authors' knowledge, those are the first reports of the use of rhBMP-2 delivered via adsorption into a CRM for regenerating bone in chronic defect non-union fractures in dogs. Its radiographic appearance seems to have a normal bone density and was integrated to native bone. The combined surgical and regenerative strategies reported achieved predictable, timely reconstruction of defect non-union fractures in small-breed, older dogs. The use of rhBMP-2 should be done with caution due to its very potent effect that is very versatile and with a wide range of functions and dose dependent. Finally, the incorporation of a regenerative strategy into the surgical resolution of non-union fracture defects avoided the morbidity associated with autologous bone grafting and provides fast return to normal function [48].

CRM is made of a collagen sponge with implanted granules of hydroxyapatite and tricalcium phosphate generating a semi-rigid framework that can resist to compressive forces in vivo, and it has been successfully used to fill bone defect in several studies [46, 48, 49]. In a prospective case series published by Arzi et al., CRM was used with rhBMP-2 in immediate reconstruction of segmental mandibulectomies (mandibular defects ≥5 cm) in four dogs for treatment of benign or malignant tumours. After tumour recession, the critical-sized bone defects were stabilized with titanium locking plates, and CRM, soaked with 0.5 mg/ml rhBMP-2 15 min before implantation at a volume equivalent to half of the calculated volume of CRM (with a half to three quarters of the mandibular height and a length 2 mm greater than the defect), was tightly implanted. Radiographic evaluations were made postoperatively at weeks 2, 4, 8 and 12 after surgery. In two clinical cases, a CT scan of the mandible was acquired 3 months after surgery. All dogs had proper occlusion after surgery and in the follow-up evaluation and returned to normal activity. After 2 weeks the entire defect site was covered with gingiva, and at 4 weeks, it was completely solid. There was no recurrence of the tumour or fractures during the controls at 2 and 3 months after surgery, and all owners reported an excellent quality of the life of their dogs [49]. The radiodensity of regenerated mandible increases throughout radiographic controls from postoperative radiographs to 4 weeks after surgery. The CRM scaffold had evidence of new bone formation connecting the adjacent mandible and smooth margins at 4 weeks post-surgery. At 8 weeks the scaffold continued to increase its radiodensity, and a mineralized union with the mandible was noticed. New bone formation and complete integration of CRM material with native mandible tissue were evident on CT image evaluation. The authors concluded that this surgical and regenerative approach achieved a rapid return to normal activity, with normal anatomy and occlusion, bone regeneration and re-established biomechanical function [49].

Verstraete published another case series in six dogs where CRM and rhBMP-2 were used as a regenerative approach to fill non-union mandibular fracture defects stabilized with locking titanium miniplate. All dogs were adopted from a shelter, and, apart from two dogs, it was not possible to know the duration of the non-union. In all cases the non-union defect was debrided and cleaned from fibrous and devitalized tissues and old implant materials. The volume of the defects was estimated with a CT with three-dimensional reconstruction, and an amount of CRM enough to fill a half to three quarters of the mandibular height with a length of 2 mm greater than the mandibular gap was prepared. The CRM was soaked with a volume of rhBMP-2 (0.5 mg/ml) corresponding to 50% of the CRM volume used to fill the defect. As an example, with a CRM piece of 2 cm length, 0.5 cm width and 1 cm height (2 0.5 1), the total CRM volume is 1 cm3 , corresponding to 0.5 ml of rhBMP-2 solution with 0.5 mg/ml concentration. An extra-oral approach was used to counter the plate in ventrodorsal position. The wedges of the non-union were debrided to remove sclerotic and devitalized bone and attached soft tissues, and the plate was then fixed to the bone with two to three locking titanium screws in each segment of the mandible and CRM was implanted. Radiograph follow-ups of the mandible were started immediately after surgery and 2, 4, 8 and 12 weeks after surgery. All dogs healed the soft tissues over the defects and with immediate to normal function and correct occlusion, and solid cortical bone formation was noticed within 3 months. There were no recurrence fractures on the follow-up period [48].

Synthetic bone graft (Bonelike®) is produced by the incorporation of P2O5–CaO glass-based system within a hydroxyapatite (HA) matrix. Bonelike® macroporous (BL® Poro) consists of polygonal granules with 2000–2800 and 4000–5600 μm of diameter with pore size range from 100 to 400 μm (Figure 1). Its osteoinductive and osteoconductive properties have been confirmed in experimental models of bone regeneration in sheep, have been used in clinical orthopaedic application and recently are being used in small animals [50–54].

We used Bonelike® in small granules and Bonelike® presenting macroporous structure (BL® Poro) in vivo in combination with ACBG (Figure 2(A)), PRP (Figure 2(B)) and rigid internal fixation with bone plates and screws to treat atrophic non-union of long bones in two adult dogs that were referred to the veterinary medicine teaching hospital, Vasco da Gama University, Coimbra, Portugal, for surgical correction of atrophic non-union that fails in at least one previous surgery. After clinical and orthopaedic examination, two orthogonal views of the affected and contralateral bone were obtained with a standard marker to calibrate magnification. Informed consent to use BL Poro® was obtained from all clients. Clinical cases were classified according to Weber-Cech terminology, as defect, dystrophic, necrotic or oligotrophic-atrophic non-union, and were labelled as infected if this was suspected due to either radiographic evaluation or appearance and tissue culture at surgery. For both clinical cases, autologous cancellous bone grafts were harvested from the uppermost proximal humerus as described by Innes [5].  The second dog had an oligotrophic non-union in the proximal radius and two non-unions in the ulnar diaphysis caused by a gunshot trauma. Revision surgery was performed to remove all the failed implants, debridement, ischemic bone fragments and other avascular soft tissues. Sclerotic and atrophic bone ends of the biologically inactive non-unions were osteotomized with a bone. Rongeur until bleeding is seen to expose medullary cavity and improve vascularity in a bone defect. Prior to copious flushing, samples of the removed tissues and deep wound swabs were collected

Clinical Application of Macroporous Ceramic to Promote Bone Healing in Veterinary Clinical Cases http://dx.doi.org/10.5772/intechopen.70341 109

possible to know the duration of the non-union. In all cases the non-union defect was debrided and cleaned from fibrous and devitalized tissues and old implant materials. The volume of the defects was estimated with a CT with three-dimensional reconstruction, and an amount of CRM enough to fill a half to three quarters of the mandibular height with a length of 2 mm greater than the mandibular gap was prepared. The CRM was soaked with a volume of rhBMP-2 (0.5 mg/ml) corresponding to 50% of the CRM volume used to fill the defect.  An extra-oral approach was used to counter the plate in ventrodorsal position. The wedges of the non-union were debrided to remove sclerotic and devitalized bone and attached soft tissues, and the plate was then fixed to the bone with two to three locking titanium screws in each segment of the mandible and CRM was implanted. Radiograph follow-ups of the mandible were started immediately after surgery and 2, 4, 8 and 12 weeks after surgery. All dogs healed the soft tissues over the defects and with immediate to normal function and correct occlusion, and solid cortical bone formation was noticed within 3 months. There were

Synthetic bone graft (Bonelike®) is produced by the incorporation of P2O5–CaO glass-based system within a hydroxyapatite (HA) matrix. Bonelike® macroporous (BL® Poro) consists of polygonal granules with 2000–2800 and 4000–5600 μm of diameter with pore size range from 100 to 400 μm (Figure 1). Its osteoinductive and osteoconductive properties have been confirmed in experimental models of bone regeneration in sheep, have been used in clinical

We used Bonelike® in small granules and Bonelike® presenting macroporous structure (BL® Poro) in vivo in combination with ACBG (Figure 2(A)), PRP (Figure 2(B)) and rigid internal fixation with bone plates and screws to treat atrophic non-union of long bones in two adult dogs that were referred to the veterinary medicine teaching hospital, Vasco da Gama University, Coimbra, Portugal, for surgical correction of atrophic non-union that fails in at least one previous surgery. After clinical and orthopaedic examination, two orthogonal views of the affected and contralateral bone were obtained with a standard marker to calibrate magnification. Informed consent to use BL Poro® was obtained from all clients. Clinical cases were classified according to Weber-Cech terminology, as defect, dystrophic, necrotic or oligotrophic-atrophic non-union, and were labelled as infected if this was suspected due to either radiographic evaluation or appearance and tissue culture at surgery. For both clinical cases, autologous cancellous bone grafts were harvested from the uppermost proximal humerus as described by Innes [5].  The second dog had an oligotrophic non-union in the proximal radius and two non-unions in the ulnar diaphysis caused by a gunshot trauma. Revision surgery was performed to remove all the failed implants, debridement, ischemic bone fragments and other avascular soft tissues. Sclerotic and atrophic bone ends of the biologically inactive non-unions were osteotomized with a bone. Rongeur until bleeding is seen to expose medullary cavity and improve vascularity in a bone defect. Prior to copious flushing, samples of the removed tissues and deep wound swabs were collected

orthopaedic application and recently are being used in small animals [50–54].

, corresponding to 0.5 ml of rhBMP-2 solution with 0.5 mg/ml concentra-

CRM volume is 1 cm3

no recurrence fractures on the follow-up period [48].

108 Scaffolds in Tissue Engineering - Materials, Technologies and Clinical Applications

Figure 1. Scanning electron microscope image of a Bonelike® Poro granule showing the interconnected pore architecture with macropores and micropores.

for bacterial culture and sensitivity antibiogram. Joints are aligned, bone segments are pulled apart trying to restore original length of the bone and the fracture was stabilized with a bone plate and screws maintaining a rigid fixation for a prolonged time with minimal discomfort to the animal during postoperative period. The bone defect created after cleaning and the fracture fixation was filled with a homogenous mixture of ACBG and BL® Poro in 1:1 proportion in a PRP gel to maintain the aggregation of the components. The mixture was applied gently avoiding rupture and consequent collapse of the macroporous architecture of BL® Poro (Figure 3). Soft tissues were closed routinely and contribute to maintain the bone graft in situ with created soft tissue envelope. All the animals returned to acceptable or good function and decrease lameness grades. All cases achieved bone union between 2 and 8 months without major complications [54].

Our research group also has been using Bonelike® and a BL® Poro mixed with fresh blood to fill voids in noncritical-sized bone defects in dogs and cats with good clinical and radiographical outcomes without any adverse local nor systemic reaction (Figures 4 and 5). A

Figure 2. BL® Poro granules (white) mixed with autologous cancellous bone graft (red) (A). Mixture of BL® Poro granules with autologous cancellous bone graft and platelet-rich plasma, ready for implantation (B).

Figure 3. Femoral bone defect after fracture bridging plate fixation (A). Careful filling of defect to avoid the collapse of BL® Poro macroporous architecture (B).

clinical case of an 8-year-old, 21 kg, female Collie with a mandibular fracture along the root of canine teeth. The loose teeth exodontia and improved dental occlusion with an external skeletal fixator where fixation pins were placed along the aboral bone surface avoiding tooth roots were performed. The pins were bended and fixed together with a mouldable-stage application of methyl methacrylate. After hardening of the acrylic, the bone gap was approached directly from the oral mucosa and filled with mixture of Bonelike® (spherical granules of 250–500 μm) and blood. The soft tissues were surgically routinely closed and healed uneventfully. On radiographic evaluation 5 weeks post-surgery, the osteointegration of the biomaterial was visible, and fracture line was no longer visible. The fixator was removed 4 months later, and there was no story of mandible retraction or other types of complications (Figure 6) (unpublished data).

Clinical Application of Macroporous Ceramic to Promote Bone Healing in Veterinary Clinical Cases http://dx.doi.org/10.5772/intechopen.70341 111

Figure 4. Immediate postoperative radiographs of the tibia of a 7-year-old 35 kg male german shepherd dog with a nonunion fracture stabilized with a 3.5 mm 12-hole broad DCP plate. Bone defect was filled with Bonelike® 250–500 μm, autologous cancellous bone graft and platelet-rich plasma: cranio-caudal view (A) and mediolateral view (B). Radiolucent lines were visible within the graft and between the graft and the edges of the bone defect.

clinical case of an 8-year-old, 21 kg, female Collie with a mandibular fracture along the root of canine teeth. The loose teeth exodontia and improved dental occlusion with an external skeletal fixator where fixation pins were placed along the aboral bone surface avoiding tooth roots were performed. The pins were bended and fixed together with a mouldable-stage application of methyl methacrylate. After hardening of the acrylic, the bone gap was approached directly from the oral mucosa and filled with mixture of Bonelike® (spherical granules of 250–500 μm) and blood. The soft tissues were surgically routinely closed and healed uneventfully. On radiographic evaluation 5 weeks post-surgery, the osteointegration of the biomaterial was visible, and fracture line was no longer visible. The fixator was removed 4 months later, and there was no story of mandible retraction or other types of complications

Figure 3. Femoral bone defect after fracture bridging plate fixation (A). Careful filling of defect to avoid the collapse of

Figure 2. BL® Poro granules (white) mixed with autologous cancellous bone graft (red) (A). Mixture of BL® Poro granules

with autologous cancellous bone graft and platelet-rich plasma, ready for implantation (B).

110 Scaffolds in Tissue Engineering - Materials, Technologies and Clinical Applications

(Figure 6) (unpublished data).

BL® Poro macroporous architecture (B).

Figure 5. Postoperative radiographic evaluation. Mediolateral view at 2 months of control showing complete bone union on the caudal edge of the defect and not on the cranio-distal edge; here, a radiolucent line was noticed but with a grey colour compatible with new bone formation (A). Cranio-caudal view at 2 months of control, complete bone bridge was observed in lateral and medial cortices with irregular pattern within the bone defect (B). Mediolateral view at 8 months of control (C) and cranio-caudal view at 8 months of control. In both radiographs completed, bone union and remodellation were visible without radiolucent lines in the bone defects (D).

#### 3.2. The use of synthetic bone grafts in equine clinical cases

Besides the majorly emotionally driven small animal clinics, other veterinary medicine fields are growing regarding the clinical application of the biomaterials, as economic implications are added to the scenario. For example, equines are the 'driving force' of an industry that moves enormous financial amounts, with competition prizes and price of some animals reaching astronomical values. With this set of constraints, the use of biomaterials in animals no longer restricts to preclinical trials carried out before their application in humans and is now part of the procedures available to veterinary clinicians. These techniques can be used in a number of clinical situations in veterinary medicine, where there is a bone defect, including fractures such as comminuted fractures, filling of bone cysts and arthrodesis [6, 55–59].

Equine patients are challenging in terms of orthopaedic pathologies and viable treatment options. While most of the small animal cases referred for orthopaedic treatment relate to fractures, and surgical reduction leads to expected full recovery, but even if this is not possible, they can survive with minimal complications with unsupported limb. Contrarily, in result of their weight, equine patients with unsupported limb have a high risk of developed severe secondary pathologies at contralateral limb, like laminitis. Other problems can surge in these equine patients; these animals cannot spend long periods lying down due to complications of myositis, nerve paralysis and decubital scars, so they weight injured limb for prolonged periods. For all those situations, fracture cases are seldom amenable of medical or surgical correction, often resulting in the decision of euthanasia. Other clinical situations are however indicated for grafting and/or biomaterial application, such as arthrodesis procedures.

Arthrodesis techniques were developed for treatment of debilitating osteoarthritis. These are also indicated for treatment of (stable) articular fractures, unstable joint injuries, septic arthritis and osteochondrosis. It is a surgical procedure used to promote the fusion between opposite bones in the joint, resulting in immobilization. Surgical arthrodesis involves the destruction of

Figure 6. Bone defect resulting from canine teeth exodontia and mandibular fracture along the root of canine stabilized with a methyl methacrylate external skeletal fixator (A); immediate postoperative image of the defect filled with Bonelike® 250–500 and blood. The biomaterial was noticeable due to its multiple radiopaque spherical structures inside the bone defect. Radiolucent lines were visible within the spheres and between the graft and the edges of the bone defect. Fracture line was visible near by the fixation pin (B); 5 weeks of radiographic control, loss of radiopaque appearance of the biomaterial and radiopaque fracture lines was no longer visible due to new bone formation (C).

the articular cartilage of the bones to promote a bone-to-bone contact and subsequent fusion. Then, the joint must be aligned and stabilized into a stable, weight-bearing position [60–63]. Synthetic bone grafts can be used in the arthrodesis technique as void filler, improving the bone contact, helping the joint stabilization and avoiding the autologous bone graft collection. Since synthetic bone grafts deliver important osteoconductive and osteoinductive proprieties, the healing process is improved and shortened [50, 52]. Surgical arthrodesis is used in clinical situations that it is not possible recover the joint. With this technique the joint is stabilized, relieving the pain and improving horse's quality of life. In this species, the proximal interphalangeal, the intertarsal and the tarsometatarsal joints are the joints where more arthrodeses are performed, due to these joints that only have little range of movements, and their permanent immobilization does not affect normal locomotion of the horse. In these cases, the goal of the surgical procedure is to return the horse to its activity. Although the permanent immobilization does not affect normal locomotion, this procedure can weaken the joint region, due to loss of shock absorption capacity of the joint, and fracture risk may be increased [64–66]. Arthrodesis technique is also used in joints that have important movements (like metacarpophalangeal joint), knowing that in these cases the horse will have a mechanical lameness. The use of this technique in these types of joints allows to decrease pain and to increase comfort and the use of the affected limb, thus preventing the appearance of lesions in the contralateral limb overhead [64–66].

3.2. The use of synthetic bone grafts in equine clinical cases

112 Scaffolds in Tissue Engineering - Materials, Technologies and Clinical Applications

as comminuted fractures, filling of bone cysts and arthrodesis [6, 55–59].

Besides the majorly emotionally driven small animal clinics, other veterinary medicine fields are growing regarding the clinical application of the biomaterials, as economic implications are added to the scenario. For example, equines are the 'driving force' of an industry that moves enormous financial amounts, with competition prizes and price of some animals reaching astronomical values. With this set of constraints, the use of biomaterials in animals no longer restricts to preclinical trials carried out before their application in humans and is now part of the procedures available to veterinary clinicians. These techniques can be used in a number of clinical situations in veterinary medicine, where there is a bone defect, including fractures such

Equine patients are challenging in terms of orthopaedic pathologies and viable treatment options. While most of the small animal cases referred for orthopaedic treatment relate to fractures, and surgical reduction leads to expected full recovery, but even if this is not possible, they can survive with minimal complications with unsupported limb. Contrarily, in result of their weight, equine patients with unsupported limb have a high risk of developed severe secondary pathologies at contralateral limb, like laminitis. Other problems can surge in these equine patients; these animals cannot spend long periods lying down due to complications of myositis, nerve paralysis and decubital scars, so they weight injured limb for prolonged periods. For all those situations, fracture cases are seldom amenable of medical or surgical correction, often resulting in the decision of euthanasia. Other clinical situations are however

indicated for grafting and/or biomaterial application, such as arthrodesis procedures.

Arthrodesis techniques were developed for treatment of debilitating osteoarthritis. These are also indicated for treatment of (stable) articular fractures, unstable joint injuries, septic arthritis and osteochondrosis. It is a surgical procedure used to promote the fusion between opposite bones in the joint, resulting in immobilization. Surgical arthrodesis involves the destruction of

Figure 6. Bone defect resulting from canine teeth exodontia and mandibular fracture along the root of canine stabilized with a methyl methacrylate external skeletal fixator (A); immediate postoperative image of the defect filled with Bonelike® 250–500 and blood. The biomaterial was noticeable due to its multiple radiopaque spherical structures inside the bone defect. Radiolucent lines were visible within the spheres and between the graft and the edges of the bone defect. Fracture line was visible near by the fixation pin (B); 5 weeks of radiographic control, loss of radiopaque appearance of the

biomaterial and radiopaque fracture lines was no longer visible due to new bone formation (C).

As examples of the use of synthetic bone grafts, it will be described in three clinical cases, where Bonelike® was applied in the arthrodesis of metacarpophalangeal joint, in the proximal interphalangeal joint and in the distal intertarsal joint together with tarsometatarsal joint.

The first clinical case is a newborn donkey, born with a severe metacarpophalangeal joint flexural deformity. The flexural deformity is so severe that it is impossible to manipulate and move the metacarpophalangeal joint and the dorsal surface of the fetlock supported the animal's weight. Due to this situation, the dorsal aspect of the fetlock presented an ulceration that communicated with the joint (Figure 7). It was decided, as the first approach, the surgical treatment, to quickly as possible resolve the situation and prevent the worsening of the complications [67, 68]. In this type of surgery, to relieve flexor forces, the tightened tendons can be cut [67, 69, 70]. In this clinical case, both the superficial flexor tendon and the deep flexor tendon were involved in the contracture of the joint, as well as the extensor digital tendon, which was displaced caudally, working as a flexor, further forcing the joint to adopt the flexor position (Figure 8(A)). Tenotomy of the three involved tendons was undertaken, in order to reposition the joint into a physiological alignment (Figure 8(B)). With this aggressive approach, the metacarpophalangeal joint stayed without support; so, a surgical arthrodesis of metacarpophalangeal joint was also performed to stabilize the joint in a physiological position. To do the arthrodesis, first the joint cartilage was removed with a curette to promote a bone-tobone contact between the metacarpal bone and the first phalange; after that the joint space was filled with Bonelike® spherical granules mixed with autologous blood. A plate with cortical screws was used to stabilize the joint in a physiological position (Figure 8(C and D)) [68]. Although arthrodesis is not commonly used on flexural limb deformities, it was decided to apply this technique in the described clinical case, considering the lack of structural support of the joint after the tenotomy of the three tightened tendons. This approach was previously described by [71] that performed arthrodesis in a llama, in a miniature horse and in a

Figure 7. Newborn donkey with a severe metacarpophalangeal joint flexion deformity. The donkey supported the right hind limb with the dorsal aspect of fetlock (A). Large cutaneous ulcer on dorsal aspect of fetlock (B).

miniature donkey, with severe bilateral congenital flexural deformities of the metacarpophalangeal joint. The arthrodesis of the fetlock joint was not aimed at full recovery for future athletic activity but rather having in mind the animal's life quality [68].

The second clinical case was an adult horse that suffered an accident that resulted in the laceration of the medial aspect of the pastern. With this laceration, a medial ligament that causes instability to the pastern joint and consequently lameness was damaged. In clinical examination, a biomechanical instability of this joint was observed, and in X-ray examination, radiologic signs of subluxation and an increase in the medial joint space can be seen. Towards that it was decided to do the arthrodesis of proximal interphalangeal joint [68]. To approach the proximal interphalangeal joint, an I-shaped incision was performed in the skin over the dorsal aspect of the joint, and then a Z-incision was performed over the digital extensor tendon to expose the proximal interphalangeal joint. During the procedure, it was confirmed that this joint was not stable and there was an increased articular gap between the first and the second phalange on the medial aspect of the joint (Figure 9(B)). The cartilage from proximal interphalangeal was removed using a curette, and the joint space (Figure 9(A)) was filled with Bonelike® (Figure 9(C)). Three lag screws 3.2 mm were placed from the first to the second

Clinical Application of Macroporous Ceramic to Promote Bone Healing in Veterinary Clinical Cases http://dx.doi.org/10.5772/intechopen.70341 115

Figure 8. Metacarpophalangeal joint arthrodesis. Extensor digital tendon (yellow arrow) displaced caudally (A). Tenotomy of extensor digital tendon (yellow arrows) and flexor digital tendon (red arrows) (B). Removal of joint cartilage with a curette (C). A plate and screws to stabilize the joint and fill joint space, using a syringe, with Bonelike® spherical granules (D).

phalange crossing the joint space. The head screws stayed in a shelf made previously with an osteotome (Figure 9(D)). On the follow-up X-rays performed after the surgery, an important bone proliferation, with evidence of bone fusion and gradual reduction of the joint space, was noticed (Figure 10) [68].

miniature donkey, with severe bilateral congenital flexural deformities of the metacarpophalangeal joint. The arthrodesis of the fetlock joint was not aimed at full recovery for future

Figure 7. Newborn donkey with a severe metacarpophalangeal joint flexion deformity. The donkey supported the right

The second clinical case was an adult horse that suffered an accident that resulted in the laceration of the medial aspect of the pastern. With this laceration, a medial ligament that causes instability to the pastern joint and consequently lameness was damaged. In clinical examination, a biomechanical instability of this joint was observed, and in X-ray examination, radiologic signs of subluxation and an increase in the medial joint space can be seen. Towards that it was decided to do the arthrodesis of proximal interphalangeal joint [68]. To approach the proximal interphalangeal joint, an I-shaped incision was performed in the skin over the dorsal aspect of the joint, and then a Z-incision was performed over the digital extensor tendon to expose the proximal interphalangeal joint. During the procedure, it was confirmed that this joint was not stable and there was an increased articular gap between the first and the second phalange on the medial aspect of the joint (Figure 9(B)). The cartilage from proximal interphalangeal was removed using a curette, and the joint space (Figure 9(A)) was filled with Bonelike® (Figure 9(C)). Three lag screws 3.2 mm were placed from the first to the second

athletic activity but rather having in mind the animal's life quality [68].

114 Scaffolds in Tissue Engineering - Materials, Technologies and Clinical Applications

hind limb with the dorsal aspect of fetlock (A). Large cutaneous ulcer on dorsal aspect of fetlock (B).

The third clinical case was a report of a horse with chronic and intermittent lameness of the right hind limb, with a slight biomechanical instability of the tarsal region during walking, posing clinical and radiological signs of bone spavin. After the confirmation of joint pain by clinical examination with the positive nerve blocks, it was decided to do the distal intertarsal joint together with tarsometatarsal joint arthrodesis. The surgical approach to those joints was made by an incision over the medial aspect of the tarsus perpendicular to the cunean tendon. The cunean tendon was transected, and joints were identified with two needles (Figure 11(A) and (B)). A 3.2 drill bit was inserted into the joint space of each joint and forced in three different directions, using a single entry point (Figure 11(C) and (D)) [68]. Drilling paths were filled with Bonelike® (Figure 11(E) and (F)). The objective of drilling the joint space was to destroy the cartilage and to promote the contact between the bones and subsequent bone fusion. Due to the osteoconductive proprieties of Bonelike®, the application of this bone graft