Properties and characterization of bone repair materials

Properties and characterization of bone repair materials

4

Kendell M. Pawelec1, Ashley A. White2, Serena M. Best3 1University of Michigan, Ann Arbor, MI, United States; 2Lawrence Berkeley National Laboratory, Berkeley, CA, United States; 3University of Cambridge, Cambridge Centre for Medical Materials, Cambridge, United Kingdom 

4.1   Introduction To perform successfully in vivo, an implant material must possess particular and well-controlled properties suited to each individual application. Biomaterials for bone applications are used in a variety of settings, from orthopedic to dental, but in all cases biomaterials characterization is key to ensuring that implants can perform their intended function. For example, in major load-bearing situations, such as joint replacement, mechanical strength may be most highly emphasized, while in small bone defects the chemistry and structure of the material may be more important for triggering the formation of new bone tissue. While emphasis may be placed on one area, say strength, the situation is not so simple as to be able choose the strongest available material and implant it. We must first understand which microstructural properties influence that strength, and how to control them, and we must consider how susceptible that material is to alterations in the body’s environment. What may seem like the ideal material could fail catastrophically when placed in the body. The better choice may be a weaker material that is inherently less susceptible to degradation in vivo, but whose microstructure can be altered to make it stronger, and thus the more successful implant material for that application. The important things to understand are (1) how to manipulate a material’s characteristics through changes in chemistry, microstructure and processing; (2) the relevant techniques to characterize the material’s properties thoroughly; and (3) the interrelationship between different levels and types of material properties. Fig. 4.1 illustrates this concept, showing how mechanical and architectural and physiological effects interact with one another and come together, to determine the in vivo success of the material. Indeed, all of these properties are interrelated, and it is necessary to consider these associations in order to control the performance of an implant material. This chapter is intended as an introduction to material properties and characterization for non-materials specialists, who seek to form a view of the most important material properties for an implant. Most of the chapter focuses on the characteristics needed to describe materials for bone repair: mechanical properties, architectural and microstructural properties, and physiological effects, summarized in Table 4.1. Mechanical properties cover a broad spectrum of tests which can be useful for different materials classes, from tensile tests to viscoelastic characterization. Architectural and microstructural Bone Repair Biomaterials. https://doi.org/10.1016/B978-0-08-102451-5.00004-4 Copyright © 2019 Elsevier Ltd. All rights reserved.

66

Bone Repair Biomaterials Architectural and microstructural properties

Tension

Compression

Stress, σ (Mpa)

Mechanical properties

Biological Chemistry coatings

Charge

Topography

In vivo performance

Strain, ε (m/m)

Bulk properties

Physiological effects

Figure 4.1  The in vivo performance of biomaterials depends on interrelated properties.

properties include the composition and grain structures of metals and ceramics, concepts such as porosity and permeability, and a description of polymer characteristics. Physiological effects are those which a material is expected to be exposed to after implantation in the body, which carry important implications for the longevity and biological response of the material. Finally, in the last section, the different classes of materials are compared, highlighting their relative strengths and weaknesses.

4.2  Mechanical properties When selecting a material for use as an implant, one of the major considerations is its required mechanical performance, in the particular skeletal application. Bone is particularly sensitive to its mechanical environment, and thus, mechanical properties are important on many length scales. Macroscopically, implants must withstand loading and movement, in the range of mega-pascals. However, mechanics at the microscopic level heavily influence cellular responses, such as differentiation toward bone cells (osteogenesis) and mineralization. It has been found that mechanical stiffness between 11 and 25 kPa can lead to the up-regulation of osteoblastic markers and, at times, spontaneous osteoblastic differentiation [1–3]. This highlights the importance of considering multiple length scales when characterizing mechanical behavior. This section will define the most important mechanical criteria and how they are evaluated. Table 4.2 offers a comparison of mechanical data for several types of replacement materials, along with those of bone.

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Table 4.1  Key material properties and relevant characterization techniques Property

Characterization techniques

References

Tensile test Compression test 3- and 4-point bend test Calculation from stress-strain curve Calculation from strength test Single-edge notched beam Indentation Cyclic stress test Constant stress test Rheology

[4] [4] [5] – – [6] [4,7] [8] [9] [10]

Mechanical Tensile strength Compressive strength Flexural strength Elastic (Young’s) modulus Ductility Toughness Hardness Fatigue Creep Viscoelasticity

Architectural and microstructural Composition

Crystal structure Grain size Surface area Porosity Permeability Polymer molecular weight Thermal transition temperatures

Rietveld analysis IR spectroscopy X-ray fluorescence spectroscopy X-ray diffraction Scanning electron microscopy Mercury porosimetry Nitrogen adsorption - BET Nitrogen adsorption - BJH Microcomputed tomography (μCT) Fluid flow Computation from percolation theory Gel permeation chromatography Differential scanning calorimetry Thermomechanical analysis

[11] [12] [13] [14] [15] [16] [17,18] [19] [20] [21] [22] [23,24] [25] [25]

Hip joint wear simulator Multi-directional motion wear test DC/AC/impedance electrochemistry Absorption spectroscopy Emission spectroscopy Auger electron spectroscopy X-ray photoelectron spectroscopy Fourier transform IR spectroscopy Raman spectroscopy Contact angle measurement Zeta potential Atomic force microscopy Scanning tunneling microscopy Interferometry

[26] [27] [28] [29] [29] [30] [30] [31] [31] [32] [31] [32]

Physiological environment Wear Corrosion potential Dissolution/degradation rate Surface chemistry

Surface charge Surface roughness

[33]

68

Table 4.2  Mechanical properties for common bone repair materials Material

Tensile/bending strength (MPa)

Yield strength (MPa)

Compressive strength (MPa)

Elastic modulus (GPa)

Toughness (MPa-m1/2)

Elongation to fracture (%)

Bone (cortical) Bone (cancellous) Stainless steel Ti alloys CoCr alloys Hydroxyapatite Glass-ceramic AW Alumina

50–150 10–20 480–1000 900–1200 400–1900 115–120 220 280–600

30–70 – 200–800 830 450–1600 – – –

160–250 23 – 450–1850 480–600 350–400 1080 4500

4–30 0.2–0.50 190–210 110–120 210–250 80–110 118 350

2–12 – 20–95 44–66 120–160 1.0 2.0 3–6

0–8 – 20–55 18 10 – – 0

Zirconia

800–1500

–

1990

210–220

6.4–10.9

–

Polyethylene (incl. UHMWPE) PMMA collagen (film/ scaffold)

23–48

21–28

20

0.6–2.2

–

350–525

35–80 –

54–73 –

80 10–100 × 10−6

2.2–4.8 2–700 × 10−6

0.7–1.6 –

0.5–6 –

Bone Repair Biomaterials

PMMA, poly(methyl methacrylate); UHMWPE, ultra-high molecular weight polyethylene. These are general values, meant to represent a range of processing techniques. Note that data on collagen structures includes reports in both the dry and hydrated states for films and scaffolds. Data compiled from various tables in Refs. [32,34–37].

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4.2.1  Tension and compression: the stress–strain curve Of the many mechanical tests available, the most common are uniaxial tension and compression tests; both yield information on a variety of properties in the form of a stress-strain curve [4]. To perform these tests, the specimen is fitted between two plates and a load is applied either to squeeze the specimen together (compression) or to pull it apart (tension). Fig. 4.2(a) shows a schematic of this procedure. As load is applied at a specified rate, the deformation of the specimen is measured, or vice versa. To normalize the data for specimen geometry, load and deformation are converted to stress (σ), calculated as the load divided by the cross-sectional area of the specimen, and strain (ε), the change in length of the specimen divided by its original length. By plotting stress against strain, we can get a good picture of the material’s behavior and work out some of its mechanical properties. Fig. 4.3 shows a sample plot with examples of the shapes of curves for several different types of material behavior. In the first region of a curve, the shaded box, the material’s behavior is elastic. Stress and strain are directly proportional to one another and deformation is recoverable. In this regime atoms in the material’s structure are displaced only slightly by stretching of atomic bonds. The second part of the curve (if present) is the plastic region. It is defined as the region were stress and strain are no longer proportional, and permanent deformation occurs as whole arrays of atoms move to a new location in the crystal structure by breaking and reforming atomic bonds. From the curves, we can obtain the following pieces of data: • Elastic (Young’s) modulus, E: Often referred to as the stiffness, it is the slope of the linear, elastic region of the stress-strain curve. The steeper the slope, the higher the Young’s modulus and the higher the stiffness of a material. Some materials, many polymers for example, do not have a linear elastic portion (Fig. 4.3(d)). In this case either a tangent modulus is used, by drawing a line tangent to the curve at a specified stress and using its slope, or a secant modulus is taken from the slope of a line drawn from the origin to a specified point. • Yield Strength, σys: This is the stress at the point when elastic deformation ends and plastic deformation begins. Often the yield strength is cited as the “tensile strength” of a material,

(b)

F

Tension

Compression

(a)

(c)

F

F b d

L

R

Figure 4.2  Schematics of mechanical strength tests. (a) Tensile and compression tests with a traditional dog-bone sample. (b) Flexural strength test. (c) Diametral compression test.

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Bone Repair Biomaterials

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Figure 4.3  Representative stress-strain graphs comparing tensile behavior for different types of materials: (a) brittle materials, such as ceramics (b) moderate modulus and low yielding typical of metals (c) a distinct yield point, visible in many polymers and (d) nonlinear elastic behavior, often observed for elastic polymers. The elastic region is shaded, and the plastic region (if applicable) is in white. The elastic modulus (E) is shown schematically in (c). Methods for calculating the yield point are also illustrated: intercept from an offset line (b), secant modulus (d) and tangent modulus (d). since it represents the highest practical strength of a material for many applications, in which plastic deformation would render the material useless. When the yield point cannot be precisely identified, as is the case with many metals which undergo a gradual transition from elastic to plastic, a line parallel to the elastic slope can be constructed at a specified strain (usually 0.002). The yield strength is then taken as the stress where the parallel line meets the curve. Fig. 4.3(b). • Ultimate tensile strength,σuts: This is the highest stress a material experiences during a tensile test. It may or may not be the stress at which the material fails (compare curves in Fig. 4.3). • Failure Strength, σf: This is the stress at catastrophic failure. It may or may not correspond to the ultimate tensile strength. • Work of fracture (toughness): Taken as the area under the entire stress-strain curve, from origin to failure, this is a measure of the ability of a material to absorb energy up to fracture. Note that this is a slightly different concept from values most often cited as “toughness,” which are actually the fracture toughness and a measure of a material’s ability to resist crack propagation (see Section 1.3). • Ductility, percentage elongation (%EL), or percentage reduction in area (%RA): Ductility is a measure of the amount of plastic deformation at fracture. Very generally, this can be seen by the shape of the stress-strain curve. The curve in Fig. 4.3(a) undergoes extremely little

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plastic deformation, whereas the rest of the curves display considerably more. Values of ductility can be given as %EL or %RA according to the following formulae:

% EL = % RA =

lf − l 0 l0

× 100 %

A0 − Af A0

× 100 %

(4.1) (4.2)

  In these equations, lf and l0 are the final and initial length, respectively, and Af and A0 are the final and initial cross-sectional areas, respectively. A material with a low value of ductility (0%–5%) is said to be brittle, whereas a material with a high value is said to be ductile.

The four different curves displayed in Fig. 4.3 are typical of different types of materials. Fig. 4.3(a) displays the high modulus and brittle behavior common to ceramics and bone cement. No plastic deformation occurs in this case. Fig. 4.3(b) shows a material with a moderate modulus, a fair degree of yielding, and the blurred elastic plastic transition common to metals and alloys. The more distinct yield point and yielding behavior of Fig. 4.3(c) is common to polymers. Fig. 4.3(d) displays the typical behavior of elastic polymers, such as rubber. These are very general curve shapes and classifications, at best, but should give some idea of how varied material behavior can be. Indeed, the shape of a curve may differ considerably for the same material, depending on the processing conditions or testing environment. It should also be noted that traditional tensile tests report the macroscopic properties of the sample. To examine mechanical properties on the microscale, especially suitable for thin films and coatings, techniques such as nanoindentation should be applied [7]. While tensile tests work well for most metals and polymers, ceramics and hydrogels pose more of a problem. It is difficult to prepare specimens with the right geometry and nearly impossible to grip them properly. In some cases, a compression test can provide the information needed and can more closely approximate the mechanical forces acting on the material in situ. In general, the compressive modulus (analogous to the Young’s modulus) is reported.

4.2.2  Other strength tests While tension and compression tests are most often performed, other mechanical tests can provide useful data for a desired application. In particular, for materials which cannot undergo a uniaxial tensile test, it is necessary to find other ways to test the tensile properties of materials.

4.2.2.1   Flexural strength The flexural, or bending strength, is often tested with a three-point or four-point bend test, Fig. 4.2. When the load is applied to rod-shaped specimens, with either rectangular or circular cross-sections, the ends of the specimen bend upwards, forcing the midpoint lower. This places the top of the specimen in compression and the bottom in tension. Since the tensile strength of most brittle ceramics is approximately onetenth that of the compressive strength, it can be assumed that the specimen fails due

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Bone Repair Biomaterials

to tension and thus represents a good value for tensile strength. Care must be taken to ensure the specimen has broken as close to the center as possible, and always between the contact points rather than toward the outside of the contact points. The formulae for calculating bending strength for a three-point bend are shown below, both for rectangular (Eq. 4.3) and circular cross-sections (Eq. 4.4):





σfs = σfs =

3FL 2bd2

(4.3)

FL πR3

(4.4)

Here, F is the load, L is the distance between contact points, b and d are the width and depth of a rectangular cross-section specimen, respectively, and R is the radius of a circular cross-section specimen. The four-point bend test may be preferred over the three-point bend test for brittle ceramics. One cause for concern with the three-point test is that failure might be initiated prematurely on the top face by a flaw generated by the contact point at the center of the specimen. As the center is also where the peak load is located, this is expected to be the failure location under normal conditions. In the case of the four-point bend test, the bending moment is constant between the top two contact points. In this case, however, premature failure caused by a test-induced flaw should be easily distinguishable from a correct result where the specimen breaks between the two contact points. It should be noted, as well, that four-point bend tests tend to give lower strength values for ceramics. Since the load is spread over a wide area, there is more of a probability of encountering failure-inducing flaws, such as micropores. A more detailed description of the procedure and calculations can be found in ASTM Standard C1161 and Ref. [5].

4.2.2.2  Diametral compression Another option for testing tensile strength is the diametral compression test, also known as the Brazilian disc test [38]. This test is done on a round tablet or short, circular cross-section rod and can be particularly practical for brittle ceramics. In this test, a specimen is placed on its side so that its curved surface is touching the plates of a tensile tester (Fig. 4.2(c)). A compressive load is then applied, causing the specimen to crack vertically through its center. Though the name and technique are misleading, this is a measure of tensile strength. A crack is initiated in the center of the sample by a tensile load, which acts perpendicular to the applied load. The tensile strength is then calculated as shown:



σ=

2F πDt

(4.5)

Where F is the applied load, D is the specimen diameter, and t is its thickness. One of the advantages to this test is that the results are independent of the quality of the specimen surface.

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With both the bend test and the diametral compression test, it must be noted that the results are size dependent. The larger the specimen used, the more volume is placed under stress, which means the probability of small cracks or flaws leading to catastrophic failure increases. Since brittle ceramics are particularly prone to problems with microcracks and pores (discussed more in relation to toughness), the scatter in data is much higher than for other classes of material. Thus, statistical methods are important to make sensible use of data. Weibull statistics are the most useful for ceramic materials and are described in further detail by Wachtman et al. [39].

4.2.3  Other mechanical properties 4.2.3.1   Hardness Hardness tests are a measure of resistance to indentation and are notable for being fast, easy, and nondestructive. A force is applied to an indenter, such as a steel ball or diamond pyramid, and the resulting size or depth of the indentation in the surface of the material is measured using a microscope. There are several different types of hardness tests, varying in the shape and material of the indenter and the scale of the load applied. Some of the more common tests are Brinell and Rockwell to measure macrohardness, Knoop to measure microhardness and Vickers to measure both macro- and microhardness. Nanoindentation can also be performed with a Berkovich test. Specific hardness formulae match each type of hardness test [7]. Lower numbers indicate the material is easily scratched or dented, whereas higher numbers indicate more resistance to indentation. Hardness is often related to other properties, such as tensile strength. For steels, cast irons and brass, hardness and tensile strength have a linear relationship [34].

4.2.3.2  Shear and torsion Sometimes a material may be subjected to transverse loads (shear) or twisting motions (torsion). For shear, imagine a cube, fixed at the bottom. A force is then applied across the top, parallel to the top face, distorting the cube into a parallelepiped. Shear stress may be calculated in a similar way to longitudinal stress, but with a modification for the angle. The shear stress, τ, is equal to the shear force divided by the initial area across which the shear force is applied. The shear strain, γ, is equal to the tangent of the angle through which the deformation occurs. Shear modulus, G, is then τ divided by γ. Like tensile deformation, shear deformation can also be elastic, in small quantities, or plastic in larger quantities. Torsion involves a twisting motion about the long axis of a specimen. In this case, imagine a rod, fixed at the bottom, subjected to a torque around its top face. The top will be deformed by an angle Φ relative to the bottom. Shear stress is then a function of the applied torque and shear modulus is a function of the angle of twist. Poisson’s ratio is a way of relating the deformation of a material along different axes. For example, a tensile specimen experiences an elongating strain along the z axis, reducing the cross-sectional area of the specimen proportionally and reducing

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Bone Repair Biomaterials

the strain in the x and y directions. If the material is isotropic along the x-y face then εx = εy. Poisson’s ratio, υ, can then be represented as: ν=−



εx εz

=−

εy εz

(4.6)

The ratio can be useful for relating shear and elastic moduli by the equation:

E = 2G (1 + ν)

(4.7)

4.2.3.3  Fracture toughness As mentioned previously, the general toughness of a material, in terms of its ability to absorb energy as it fractures, can be calculated by the area under a stress-strain curve. However, fracture toughness is a slightly different concept. Materials can have many microscopic defects, such as microporosity, and macroscopic defects, such as cracks. Applied stress is concentrated at the edges of these defects or crack tips, lowering the stress needed to break the material. If a material is tough, then it can deform plastically as the crack propagates, slowing or stopping the crack’s motion, or it may have a second phase hindering crack growth. However, if a material is brittle, then the concentrated stress causes the crack to propagate quickly and catastrophically. There are several ways to measure fracture toughness, one of the most common being the single-edge notched beam [6]. This consists of a setup like the three-point bend test, wherein a notch as thin as possible is placed on the bottom face of the rod, where the specimen is in tension. If done properly, this notch will serve as a crack initiator and its propagation will cause the failure of the material, Fracture toughness, KIc, is calculated by:

√ KIc = Yσ πa

(4.8)

Y is a geometrical factor, σ is the stress at which catastrophic failure occurs and a is the crack length. (In the case of an internal crack, a represents the radius of the crack or one-half the crack length.) Toughness can also be measured indirectly, by indenting the sample with a load that is high enough that cracks emanate from the corners of the indent. Measuring the length of the cracks enables an estimate of the toughness to be calculated.

4.2.3.4  Fatigue When a material is placed under continuous cyclic stresses, it may fail by a mechanism called fatigue. Repetitive loading and unloading, like that applied to an orthopedic hip implant during walking, can create microscopic cracks that propagate little by little with each subsequent cycle. Because cracks and scratches concentrate stresses locally, it will take a much lower stress than the normal failure stress to propagate the crack and eventually cause failure. Thus, the fatigue strength may be as little as one-quarter to one-third the material’s tensile strength.

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Fatigue is tested by a cyclic tension or bending test at different maximum stresses [8]. The endurance limit is determined by the stress that causes little to no failure after 106 to 108 cycles. The fracture surface is normally perpendicular to the direction of the applied tensile stress. The fatigue behavior of a material is sensitive to environment, which is especially important when considering implants which will be operating at physiological conditions (37°C). Corrosion or other degradation of the material, and the cycle rate, will affect the endurance limit. In a phenomenon called corrosion fatigue, corrosion forms pits, which act as stress concentrators, causing cracks to initiate and propagate with each cycle. In some cases, processing conditions may also contribute to fatigue. Small scratches or grooves caused by machining can lower the endurance limit of a material, thus it is always best to have a smooth polished specimen.

4.2.3.5   Viscoelastic properties Many polymers, particularly amorphous polymers like plastics and rubber are susceptible to a phenomenon called viscoelastic creep. (While metals and ceramics may also experience this phenomenon, they only do so at temperatures far exceeding body temperature.) “Viscoelastic” refers to a type of behavior in between elastic and viscous behavior. With elastic behavior, the material stretches immediately with applied load and immediately recovers upon removal of the load, like a rubber band. A viscoelastic material, on the other hand, has a delayed deformation upon application of stress. After removal of the stress, the material slowly recovers the strain, but never completely. The behavior is also subject to the rate of strain. If the strain is fast, the material behaves more elastically; if slow, the material behaves more viscously. Viscoelastic creep, then, occurs when a constant load slowly deforms a material. This property can be significant for some polymers even at temperatures as low as body temperature and at modest stresses, far below the yield stress. It has been suggested that cells respond to the viscoelastic properties of materials, using it to regulate processes such as stem cell differentiation [40]. To change the creep behavior in polymers, the microstructure can be altered. In general, the higher the degree of crystallinity a polymer has (Section 3.5), the less susceptible it is to creep. Creep is evaluated by applying an instantaneous load and measuring strain as a function of time [9]. An extension of creep is the study of viscoelasticity, utilizing rheology, which is most often applied to liquids, gels, and colloids. The rheological properties of a material can be extremely important for materials which evolve over time, such as hydrogels, natural polymer networks, and calcium phosphate setting bone cements [41]. In addition, rheological techniques can be used to evaluate the chemistry and structure of polymers, which effect biological response [42]. The most common rheological measurement is viscosity (η), which can be described in terms of shear stress (σ) and shear rate ( γ˙ ):

η=

σ γ˙

(4.9)

During testing, a force or velocity is applied to a sample, and the shear stress and shear rate are measured, based on the geometry of the testing apparatus. Common examples

76

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Figure 4.4  Viscosity and modulus are used to characterize viscoelastic materials and phase changes (a) Viscosity, as a function of strain rate, for a 1 wt% collagen slurry. Additives can shift the viscosity up or down [42]. (b) Schematic of the storage (G′) and loss (G″) modulus during gelation. When G′ dominates, the material has solid-like properties.

of testers which allow constant shear rate through the entire sample are: narrow gap concentric cylinders and small-angle, cone-and-plate geometry. One thing to note is that the experimental setup may affect the findings [43]. For a more in-depth treatment of apparatus and experimental design, readers are referred to [10]. Viscosity can be affected by many things, such as temperature and composition of the sample. In the case of ideal liquids, which exhibit Newtonian behavior, the viscosity is constant with respect to shear rate. Under everyday conditions, many materials fit into this category; however, all materials become non-Newtonian at high enough shear rates. In general, most become shear thinning, or the viscosity becomes lower at higher shear rate, seen on a plot of viscosity versus shear rate, Fig. 4.4(a). This can impact injectable materials, like calcium phosphate cements [41]. Importantly, the viscoelastic nature of liquids leads to a time delay in responding to the applied force. To characterize this, an oscillatory test is often used, where the stress or strain is swept over a range of frequencies. At high frequencies, solid-like, elastic behavior dominates (in phase) and at low frequencies, liquid-like behavior (out of phase) is observed, Fig. 4.4(b). Viscoelastic materials are therefore defined by two moduli: a storage-modulus (G′, solid-like) and a loss-modulus (G″, liquid-like). When G″ is larger than G′, liquid-like behavior dominates, and vice versa. This is important for applications which rely on phase transitions, such as gelation or cement setting. For example, in practical applications, gelling of a polymer network is measured as the time it takes for G′ to become the dominate moduli [44].

4.3  Architectural and microstructural properties The architecture and microstructure of a material contributes a great deal to its overall properties, influencing its mechanics and the way in which the body interacts with the material. Pore size, grain characteristics (in the case of crystalline materials) and phase distribution also affect material properties. In the case of amorphous or

Properties and characterization of bone repair materials

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semicrystalline polymers, small changes in attributes like molecular weight, thermal transition temperatures and cross-linking can have a profound effect on the behavior of a material. In many materials systems, it can be extremely difficult to decouple the mechanics, chemistry and architecture, as a change in one is often accompanied by a change in the other [40,45].

4.3.1   Phase composition The chemical and crystallographic identities of the phases present in a material, along with their relative amounts, distribution and orientation, determine the material’s intrinsic properties. Therefore, it is very important to be able to identify these phases precisely. Phases may be distinguished in a number of different ways. If crystalline, the structure will have unique lattice parameters and an arrangement of atoms identifiable by X-ray diffraction (XRD). In the case of polycrystalline materials, Rietveld analysis applied to XRD data can determine the relative amounts of phases present in a specimen [11]. Infrared (IR) spectroscopy can be used to identify the chemical bonds present, their relative quantity, and any changes in the character or quantity of the bonds over time [12]. This technique is particularly useful for polymers, which are not crystalline. The locations of different phases in a material can sometimes be determined by microscopy. Grains in metals can usually be seen in light microscopy, as different phases, in an alloy, will reflect light differently. Scanning electron microscopy (SEM), with an attached X-ray beam, can be used for energy dispersive spectroscopy (EDS) to pinpoint an elemental analysis of a specific location in the specimen. This method is not as sensitive for determining elemental ratios as some other methods. When knowing the exact composition of a material is very important, such as in the case of calcium phosphates which have different dissolution behavior depending on their Ca/P ratio, more sensitive methods like X-ray fluorescence (XRF) spectroscopy are a better choice [13].

4.3.2   Grain structure The term “grain” refers to a crystal in a polycrystalline material. Normally these grains are randomly oriented, coming together at areas of atomic mismatch, known as grain boundaries. Boundaries are areas of high energy, making them more chemically reactive than their surroundings. Heat treatments cause grains to grow due to a reduction in boundary energy. Since grain boundaries inhibit dislocation motion, as grains grow, mechanical properties are decreased. Additionally, fine-grained ceramics tend to have lower porosity since pores are removed by the easy vacancy transport that can occur along grain boundaries. A smaller grain size also improves properties such as surface wear [46], as well as strength and fracture toughness in ceramics and metals [47,48]. At the same time, however, increased reactivity with a smaller grain size, and increased grain boundary area, has been found to increase corrosion of metal implant materials and dissolution of bioceramics in vivo [49,50].

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Bone Repair Biomaterials

Typically, grain size is measured from micrographs showing the grain structure. In some cases, a sample may have to be polished and etched with acid, which preferentially attacks grain boundaries, to reveal the grain structure. By drawing a line across a micrograph and counting the number of times it intersects a grain boundary, the grain size can then be calculated by comparing this result with the scale bar.

4.3.3  Porosity The porosity of a material contributes greatly to both its mechanical properties and the way in which it behaves in vivo, and it can span many length scales, Fig. 4.5 [53]. For monolithic devices, which must function in load-bearing situations, porosity is often viewed as a detriment, since pores act as stress concentrators, decreasing mechanical properties. Even small pores (<10 μm), arising from incomplete densification during sintering of a ceramic powder, or defects from casting a metal, are detrimental to mechanical properties. On the other hand, small pores, less than 10 μm, have been linked to bone formation in vivo [54]. Also, osteoblasts preferentially attach to areas with high surface roughness and microporosity, on calcium phosphate scaffolds [55]. However, in varying the microporosity of ceramics, other factors are also affected, such as increased solubility of ions from the less crystalline surface and differences in pore characteristics. (b)

(a)

10 µm

100 nm (d)

(c)

50 µm

15 µm

Figure 4.5  Porosity can span several length scales within the same material, as seen for (a and b) poly(styrene)-block-poly(acrylic acid) (PS–PAA) scaffolds and (c and d) hydroxyapatite scaffolds. Both structures contain (a and c) large scale pores, on the order of several hundred microns, and (b and d) microsized pores within the walls. Differences in architecture are also illustrated: (a and b) isotropic and (c and d) anisotropic. Scale bar: (a) 10 μm, (b) 100 nm, (c) 50 μm and (d) 15 μm. (a and b) Adapted from George PA, Quinn K, Cooper-White JJ. Hierarchical scaffolds via combined macro- and micro-phase separation. Biomaterials 2010;31(4):641–47. https://doi. org/10.1016/j.biomaterials.2009.09.094, with permission from Elsevier. (c and d) Adapted from Deville S, Saiz E, Tomsia AP. Freeze casting of hydroxyapatite scaffolds for bone tissue engineering. Biomaterials 2006;27(32):5480–89, with permission from Elsevier.

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Conversely, for applications where biological interaction is desired, such as scaffolds for tissue engineering, porosity is a necessity, with the increased surface area providing greater environmental interaction. To promote cellular infiltration and bone ingrowth, pores must be interconnected and between 100 and 400 μm in diameter [56]. Interconnectivity of scaffold structures is related to how easily cells or particles can circulate through the structure. The minimum size of the pore interconnects needed for bone ingrowth is controversial. Within titanium and ceramic implants, pore interconnects of greater than 50 μm diameter encourage more bone ingrowth than pores with smaller pathways, with a sharp increase in gone ingrowth when the interconnect size increases to 100 μm [57–59]. While it promotes cellular infiltration, an increased surface area, due to pores, may also affect the degradation time of the material, and is something to bear in mind when designing scaffolds. Like porosity, scaffold architecture can also influence the biological activity of the scaffold. Architecture falls into two categories: isotropic (having no innate directionality) or anisotropic, Fig. 4.5. Most native tissues are anisotropic, and bone is no exception. Bone tissue responds to scaffold architecture, with the features of the scaffold dictating where bone growth occurs [60,61]. While scaffold architecture does not affect the final mineral content, the mechanical strength of the newly formed bone increases significantly in anisotropic scaffolds compared to isotropic [62]. Together, porosity and architecture control where and when bone formation in scaffolds occurs. It is important to note that both aspects of scaffold structure are tied together, along with the mechanical and chemical properties. Changes to the scaffold architecture will affect interconnectivity and permeability, as well as the mechanical strength, and induce a measure of anisotropy to all properties. The size and distribution of pores can be measured by a variety of techniques. Mercury porosimetry characterizes porosity by forcing mercury into the pores of a material. The pressure required to intrude mercury into the pores is inversely proportional to the size of the pores. Using this technique can give a complete set of information about pore size, distribution, and surface area and bulk and skeletal density [16]. However, mercury porosimetry can only be used with stronger samples, which are able to withstand the pressures involved in mercury intrusion. It is also an end point test. Microcomputed tomography (μCT) is a nondestructive method for imaging a 3D structure. Based on the attenuation of X-rays in the materials, 3D maps of the structure are constructed, which can be further analyzed for characteristics such as: porosity, pore size, distribution of phases, and anisotropy. For materials with low attenuation, such as polymers, it can be difficult to separate the material from the background air, and scan resolution is limited by the size of the sample. While nanoscopic features are out of reach for most machines, μCT can accurately reach resolutions which are submicron. Another option, for more delicate materials with fine porosity, is nitrogen adsorption, in which gas molecules are physically adsorbed onto a solid surface. Using BET theory, one can use the data from nitrogen adsorption to calculate the specific surface [18], measured in units of m2/g. A green (unsintered) ceramic powder may have a surface area of hundreds of square meters per gram, while a sintered powder may be in the low tens of square meters per gram. Pore size distributions can also be calculated from nitrogen adsorption data with BJH theory [19]. Adsorption mechanisms are described in more detail in Lowell et al. [17].

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4.3.4  Permeability In porous materials, describing the pore size or porosity is not sufficient for describing the flow of nutrients and cells into the scaffold interior. Nutrient perfusion has been shown, in many cases, to be the key factor limiting scaffold effectiveness, and leading to necrotic areas within the scaffold center [63]. In bone tissue, low permeability has been directly correlated to a greater probability of nonunion between grafts and native tissue, and a poor healing response [64]. Perfusion is related to features such as the interconnection size between the pores and the tortuosity of the porous structure [65]. Permeability, or the ability of fluids to flow through an object, is a parameter which captures and reflects many features of porosity. It can either be measured directly, or computed, for example, through percolation theory, Fig. 4.6. Measuring permeability directly generally involves flowing a liquid with known density and viscosity through the structure and measuring one of three things: (1) the time required for a known volume to pass through the scaffold, Fig. 4.6(a), (2) the volume of liquid passing through for a given time, or (3) the pressure drop across the scaffold while maintaining a constant rate of fluid flow [21]. There is no standardized protocol for measuring permeability, making results difficult to compare between studies.

Figure 4.6  Permeability measurements describe the flow of nutrients and cells through a porous structure (a and b) Permeability measured via flow in anisotropic scaffolds: (a) schematic of the test setup, based on the time for a known volume to flow through the scaffold, and (b) permeability as a function of pore size. (c) Using percolation theory, the diameter of the particle which could traverse an infinitely large porous structure can be determined (dc). (a and b) Adapted from Pawelec KM, van Boxtel HA, Kluijtmans SGJM. Ice-templating of anisotropic structures with high permeability. Mater Sci Eng C 2017;76:628–36. https://doi. org/10.1016/j.msec.2017.03.142, with permission from Elsevier. (c) Adapted from Ashworth JC, Mehr M, Buxton PG, Best SM, Cameron RE. Cell invasion in collagen scaffold architectures characterized by percolation theory. Adv Healthc Mater 2015;4:1317–21. https://doi. org/10.1002/adhm.201500197.

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In anisotropic scaffolds, with linear architecture, permeability scales with the pore size [66]. In addition, it has been found that, in cases where the isotropic pore size does not vary significantly with percent porosity, intrinsic permeability (k) was related to porosity via a logarithmic relationship, where Φ was percent porosity, and a and b constants [67]:

(4.10)

k = aebϕ

One way to overcome the variability in analytic techniques, for measuring permeability, is to turn toward computer modeling to establish new metrics. One such metric is the percolation diameter, derived from percolation theory, which is defined as the diameter of the smallest object which is able to pass through an infinitely large sample of a given porous structure, Fig. 4.6(c) [68]. Determining the percolation diameter (dc) relies on a relationship between virtual object diameter (d) and maximum accessible distance from a chosen surface (L):

L ∝ (d − dc )

− 0.88



(4.11)

Unlike many metrics to determine interconnectivity, the percolation diameter is not dependent on the sample size tested. However, like all computational methods, calculations are limited by the computing power available.

4.3.5  Polymer structure and properties Starting from the smallest structural level, characteristics like molecular weight, tacticity, chain configuration, degree of polymerization, and cross-linking can affect the behavior of polymers. Properties such as crystallinity, melting temperature and glass-transition temperature are determined by these molecular-level characteristics. These, in turn, affect mechanical properties and in vivo response.

4.3.5.1  Molecular-level structural characteristics During the polymerization process of synthetic polymers, single, repeating units (mers) combine to form long chains of varying lengths, Fig. 4.7(a). The molecular weight then represents the average total molar mass for one chain in a polymer. There are several ways in which to calculate molecular weight, including a number-average molecular weight and a weight-average molecular weight [23]. The degree of polymerization is a related concept and represents the number of mers in a polymer chain. During synthesis, the degree of polymerization can be followed over time, using IR spectroscopy, to measure the chemical bonds present and their relative quantity over time [12]. Molecular weight can be measured experimentally by osmotic pressure, gel permeation chromatography and light scattering, among other techniques [23,24]. The higher the molecular weight, the more rigid is the polymer. Soft waxes or resins have molecular weights on the order of 1000 g/mol, while solid, hard polymers have molecular weights on the order of 1–100 × 104 g/mol. One of the most widely used polymers in orthopedic applications, ultra high molecular weight polyethylene (UHMWPE), has a molecular weight of 4 × 106 g/mol.

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100 80 60 40 20 0

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Figure 4.7  Synthesis of synthetic and natural polymers occurs via different mechanisms (a) Synthetic polymers are built of individual mers which react to form polymer chains. (b) Natural polymers, made from amino acids, can assemble in stages, as exemplified by collagen, which undergoes fibrillogenesis to form fibers that can then be stabilized via cross-linking. (c) Importantly, for natural polymers, greater cross-linking can reduce the ligands available for cell adhesion and spreading, as seen with HT1080 cells on collagen films (100% cross-linking refers to 1.15 g EDC and 0.276 g NHS per gram collagen). (c) Adapted from Malcor J-D, Bax DV, Hamaia SW, Davidenko N, Best SM, Cameron RE, Farndale RW, Bihan D. The synthesis and coupling of photoreactive collagen-based ppeptide to restore integrin reactivity to an inert substrate, chemically-crosslinked collagen. Biomaterials 2016;85:65–77. https://doi.org/10.1016/j.biomaterials.2016.01.044, with permission from Elsevier.

Synthetic polymer chains can be linear, branched, or combinations of the two. Linear chains are simply long chains of monomers joined end-to-end. Branched chains, on the other hand, occur when shorter chains grow off the longer chain. In an extreme example, a network polymer consists of many chains and many branches, bonded together at various points by cross-links. Tacticity refers to the arrangement of side groups around a chain. An isotactic arrangement has all side groups on the same side, a syndiotactic alternates sides, whereas an atactic arrangement is completely random. When more than one type of monomer is present in the polymer structure, it is called a copolymer, for example poly (lactic-co-glycolic acid) (PLGA). The way in which the two types of monomers are mixed in a copolymer can also be described in three ways: alternating (the monomers alternate back and forth), block (several monomers of each type join together and then alternate with a group of another monomer type), and random (completely random arrangement of monomer types). The various components, of the polymer’s chemical structure, can be determined experimentally by a wide array of characterization techniques. IR spectroscopy can identify the types of bonds present, while nuclear magnetic resonance can determine chemical groups. Wide-angle X-ray scattering determines the local structure of semicrystalline polymers, while small-angle X-ray scattering can identify if a polymer is multi-phase, a copolymer, or an ionomer. Natural polymers differ from synthetic polymers, in that the building blocks are invariably either amino acids or sugars for proteins and polysaccharides, respectively. Synthesis is regulated by cells, and the specific composition and structure is determined by the species of origin, tissue type, and age of the individual [69,70]. In vertebrates, collagen is, by far, the most abundant structural protein, and it forms most of

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the protein component in native bone tissue [71]. The assembly of collagen fibers, a process known as fibrillogenesis, is tightly controlled in vivo, and highlights some of the ways natural protein polymers differ from synthetic, Fig. 4.7(b). Individual peptide chains (procollagen) are formed within the endoplasmic reticulum of the cell and subsequently modified [72]. Three chains come together to form a triple helix, called a fibril. Individual fibrils are tightly wound together to form fibers, whose size is further regulated by their biological environment [69]. An added feature of in vivo fibrillogenesis, is that the process allows cells to orient collagen fibers according to loading in the tissue, making the precise structure difficult to recreate in vitro.

4.3.5.2  Degree of crystallinity Arrangements of polymer chains which allow for high packing, for example a linear chain with small side groups in an isotactic configuration, can create ordered polymer structures. Secondary bonding can then occur between adjacent chain segments and regions of crystallinity can form. The complexity of the chain affects this, including the size of side groups, branching, and the tacticity. While both isotactic and syndiotactic arrangements allow for crystallization, atactic structures cannot crystallize. Owing to the size and complexity of all polymer chains, it is virtually impossible for them to be completely crystalline. Thus, polymers are either amorphous or semicrystalline. The degree of crystallinity can be determined by the density of the material. The more crystalline, the more compact the structure is and thus, the denser the polymer is. In general, ductile semicrystalline polymers have a crystallinity of about 50%, whereas very brittle ones are 90%–95% crystalline. The degree of crystallinity can also be influenced by processing conditions. A slower cooling rate can allow for more crystallization, as the chains have more time to arrange themselves in an orderly fashion. Along with the more ordered, dense structure comes an improvement in both tensile modulus and tensile strength.

4.3.5.3   Thermal transitions Many polymers undergo important thermal transitions. A melting temperature (Tm) is only present in semicrystalline polymers. It occurs when the solid material with ordered, aligned chains turns into a randomly oriented viscous liquid. Due to the range of molecular weights in the structure, polymer melting actually takes place over a range of a few degrees rather than at one, clearly defined temperature. Like the degree of crystallinity, chain properties, like the size of side groups and ease of rotation, affect the melting temperature. The greater the packing, the higher the Tm. Branching, on the other hand, decreases Tm, by creating defects in crystallinity. The rate at which the polymer is heated also affects the Tm; a higher rate results in a higher melting temperature. Natural polymers have a denaturation temperature, rather than a melting temperature. At this temperature, the bonds which stabilize the proteins into a distinct 3D shape are disrupted. As protein shape is often key to performing its biological function, denaturation effectively changes the biological properties. An example of this is the transition from collagen to gelatin upon heating. Both have the identical peptide sequence, but due

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to the triple helix structure of collagen, lost in gelatin, different cell adhesive sequences are available to cells [73,74]. The denaturation temperature depends on the specific peptide sequence, and the amount of stabilization present, often in the form of cross-linking. The glass-transition temperature (Tg) marks the point at which a polymer goes from rubbery behavior to a rigid solid as it cools. As temperature decreases, the motion of large segments of polymer chains is reduced. All types of polymers experience a glass-transition. The same molecular characteristics that affect Tm affect Tg in a similar way. The two temperatures are typically linked, as well; a change in one affects the other by the relationship Tg = 0.5–0.8Tm (K). Several techniques can measure thermal transition temperatures in polymers. Differential scanning calorimetry (DSC) is the most widely used. DSC monitors heat flow to determine transition temperatures (if present). Other techniques include thermomechanical analysis and dynamic mechanical analysis and are described in more detail elsewhere [25].

4.3.5.4  Strengthening polymers While polymers do not have a grain structure in the same way that metals and ceramics do, they can be strengthened in several different ways. As previously mentioned, increasing the crystallinity and molecular weight of synthetic polymers will strengthen them. In the same way that grain boundaries impede dislocation motion, anything in a polymer’s structure that will impede the slippage of segments of molecular chains will strengthen it. Drawing, extending a polymer in one direction, strengthens semicrystalline polymers, in the drawing direction, by aligning the molecular chains in the structure through uniaxial force. Cross-linking involves forming strong bonds between molecules previously only linked by weak Van der Waals forces. These strong bonds then stop chains from moving. Cross-linking is particularly important for natural polymers, which tend to have lower mechanical properties than their synthetic counterparts. While many methods for cross-linking natural polymers exist, they can have a deleterious effect on the biological activity of the polymer, by altering cell adhesion sites or removing them entirely, Fig. 4.7(c) [71,75]. This is especially true of collagen and gelatin, which are often used in bone repair scaffolds. Various strategies have been advanced for dealing with this problem, including: cross-linking at sites not involved in adhesion, reintroducing attachment sites, and utilizing custom designed peptide linkers [76–78].

4.4  Physiological effects When designing and testing materials for bone repair, it is important to bear in mind that implants will be operating within a physiological environment: a corrosive saline solution at 37°C. In addition, on top of normal interactions with cells and tissues, implants must also survive immune responses, including large changes in pH and tissue inflammation, which occur immediately following implantation or are triggered by degradation of the implant material. This section will discuss the properties that

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influence in vivo reactions such as corrosion, dissolution, degradation, cell interaction with surfaces, and biomineralization. Materials are characterized into three basic categories, depending on their interaction with living tissue. Bioinert materials do not interact with tissues and are instead encapsulated by a fibrous capsule. Bioactive materials, on the other hand, do interact and bond directly to cells and are incorporated into native tissue over time. In many cases, bioactive materials are also bioresorbable, completely dissolving over time.

4.4.1   Metallic corrosion Within the physiological environment, which contains ions, organic substances and dissolved oxygen, metallic implants have the tendency to corrode. Corrosion leads to reduced structural integrity and can contribute to by-products that adversely affect biological functions [79]. Metals vary in their tendency to corrode, depending on their electrode potential. This potential can be measured by AC, DC, or impedance electrochemistry [28]. Some metals, such as Au and Pt, are inert, whereas metals like Cr, Co, Al, Zn, and Ti are quite reactive and easily corrode. While aqueous corrosion in the body can be a hindrance, a similar electrochemical reaction with the oxygen in air, called dry corrosion, can be an advantage. Upon exposure to air in an ambient environment, some metals immediately form an adherent oxide layer on their surface, just a few nanometers thick. Since corrosion is a surface phenomenon, this oxide layer acts as a passivating layer, preventing the transport of metallic ions and electrons between the metal implant and body fluids. This serves to prevent both aqueous corrosion and leaching of any potentially irritating or toxic ions, such as Ni2+, Cr3+, or Co2+. Certain metals, Al, Cr, and Ti, are so reactive in air that they are almost always included as components in alloys to ensure an oxide layer will form. A nitric acid treatment on stainless steel is another common way to create this layer. Once in the body, the passivation oxide layer does not always completely protect the metallic implant from corrosion [79]. First, mechanical stresses may cause the oxide film to wear off. While it can reform, biological factors may affect its ability to do so. Biological macromolecules, such as proteins, cells, and bacteria can upset the equilibria of corrosion reactions, by altering the local pH, affecting electrode potential and affecting the amount of oxygen available to reform and maintain the layer. Weak points in the oxide layer can also cause what is known as pitting corrosion. This forms holes in the oxide layer that cannot be reformed. Fretting corrosion is a cyclic process in which the passive oxide layer is continuously removed and reformed, gradually wearing away at the surface of the implant. Finally, stress corrosion can occur in the presence of an applied load. It reacts by attacking the oxide layer at its weakest and forming small cracks that grow in length and depth over time, eventually leading to device failure.

4.4.2   Ceramic dissolution Ceramics are not subject to corrosion as metals are. The strongly directional interatomic bonds in the structures mean that large amounts of energy are required for their disruption. Hence, the reason passivating oxides on metals, such as Al2O3, ZrO2, TiO2,

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protect metallic implants. Rather than degrade, ceramics may tend to dissolve in aqueous environments, depending on their chemical composition and microstructure. If the dissolution rate can be controlled, implants can be designed to dissolve at the same rate as native tissue integration. Bioceramics (discussed in more detail in a later chapter) include Al2O3, ZrO2, calcium phosphates, calcium carbonates and bioglasses, ranging on the scale of in vivo behavior. Factors such as chemical composition (i.e., Ca/P ratio for calcium phosphates), porosity and even mechanical stress play a role in the rate and extent of degradation. The more chemically similar a material is to bone, the more likely the body is to recognize it and bond to it or interact with it in some way. Indeed, complete degradation of a ceramic is often planned and serves as a reservoir of useful ions (Ca2+, CO3 2 −, Na+, PO3 2 −) as the body rebuilds. Degradation is related to the surface area, thus porous scaffolds will dissolve more quickly, and lose mechanically stability earlier, than an equivalent dense ceramic. The dissolution rate of bioceramics can be predicted by immersing the material in an aqueous solution, like simulated body fluid, and monitoring its weight loss over time. More specific information about the dissolution products can be learned by performing spectroscopic analysis, such as inductively coupled plasma mass spectrometry (ICP-MS), on the liquid as the material degrades [29]. Care should be taken in selecting the specific technique to be used, however, as different techniques have different levels of sensitivities to specific elements. While this method can give a general idea of how the ceramic might perform in the body, a host of other factors, such as the presence of cells and proteins, local pH changes caused by inflammation and mechanical stresses, will play a role in the degree of dissolution, which cannot be simulated.

4.4.3  Polymer degradation Polymers degrade in the body through physiochemical processes [24]. A common form of degradation occurs via hydrolysis, where polymer chains spread apart and swell as biomolecules and water move into the structure. This process can cause the scission of any susceptible molecular functional groups, leading to the breakdown of mechanical properties. Material properties that combat this process are a high molecular weight and a high degree of cross-linking and crystallinity. Not all synthetic polymers are susceptible to hydrolysis, though: polyethylene (PE), used in joint replacement prostheses, and polymethyl methacrylate (PMMA), used as bone cement, are two examples. Where degradation may occur, it must be incorporated into the implant design. In some cases, a small amount of swelling, caused by hydrolysis, may be beneficial in locking a polymer into position. The process of degradation leads to a change in implant geometry and mechanical properties over time. In many cases, polymer degradation lowers the Tg, which can be problematic if the polymer Tg is close to that of body temperature already. Lowering the Tg to less than 37°C, may cause a strong polymer to become rubbery and weak. It is important to match degradation rate of the polymer to the regeneration of healthy tissue. As with ceramics, the degradation rate can be evaluated by immersing the material in phosphate buffered saline or simulated

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body fluid, and analyzing the content of the liquid over time by spectroscopy, and the mass loss of the polymer [29]. These tests do not always take into consideration other factors that can degrade polymer properties in vivo: adsorption of biomolecules, such as proteins, as well as calcification and other forms of mineralization. Natural polymers, designed to be degraded and recycled in vivo, are often sensitive to enzymes, such as proteases, which can be elevated in the presence of cells, and thus degrade faster than in vitro tests would predict. Polymer degradation also has secondary effects, which should also be considered. Degradation products may cause a local alteration of pH, affecting cellular response or further altering the rate of implant degradation [24]. Low-molecular weight compounds can also leech from the polymer, such as plasticizer, which can weaken or embrittle the polymer and lead to faster dissolution. Unlike other classes of materials, polymer degradation is very sensitive to processing conditions and sterilization procedures. Polymers particularly prone to hydrolysis, for instance, should be kept moisture-free during processing and storage to prevent any premature degradation. Special care must be taken with sterilization techniques, each of which has advantages and disadvantages. When using heat sterilization, the temperature must be kept below Tg, at all times, to prevent changes in morphology. During chemical sterilization procedures, on the other hand, a polymer may absorb the chemicals and later release them into the body. Additionally, irradiation can cause bond ruptures, which can lower the molecular weight, or can create new cross-links, both of which will affect the mechanical properties [80]. It can also generate free radicals that may interact with oxygen in the body to create unwanted by-products.

4.4.4   Wear Materials which are continuously rubbed against one another, a scenario most often observed in joint replacements, will experience a gradual loss of matter at the interface. The matter lost, in the form of particulate debris, contributes a great deal to orthopedic device failure and is the main factor limiting the long-term performance of joint replacement prosthesis [79]. This process, known as wear, may occur by abrasion, adhesion, or fatigue. In the case of abrasion, a rougher or harder surface gradually wears away a softer surface. With adhesion, a softer material may wear off and transfer onto the harder material, creating a film. Surface fatigue, on the other hand, gradually creates debris particles as changing cyclic stresses are repeatedly applied to the contact area. Subsequently, wear particles may remain between the surfaces of the two interacting materials, and may be transferred from one surface to another, or they may be lost from the system. Generally, it is the softer material which wears more quickly: a polymer in the case of a metal-polymer interface or the metal in the case of a metal-ceramic interface. Wear resistance can be improved by altering the microstructure and chemical structure of the material. For example, a higher molecular weight, degree of branching, and cross-linking all improve the wear of polymers, while a fine, narrowly distributed grain size gives better wear properties in ceramics. The key problem with wear particles is related to the biological effects of the particle. Wear particles, from both metals [79] and polymers [26], may be irritating to

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Chemistry

Biological coatings

Charge

Topography

Bulk properties Figure 4.8  Material surfaces can be modified in several ways: chemistry, biological coatings, charge, and topography.

the body’s tissues, causing an inflammatory reaction. In the case of hip prostheses, these wear particles can induce an inflammatory response which results in local bone loss around the implant [81]. Metallic particulates are especially undesirable, causing toxicity to the surrounding tissue. Wear can be tested by several methods, including the hip joint wear simulator [26], the multidirectional motion wear test [27] and the flat-on-ring wear test [82].

4.4.5  Surface properties Many biological reactions to implants are driven by the material’s surface, as this is the first part of an implant the body’s environment encounters. This is especially important where nanomaterials are concerned, due to their large surface area to volume ratio. Manipulation of surface properties to optimize biological behavior is an active area of research [56,83]. While not entirely elucidated, some general trends in the types of properties that affect cell behavior have emerged related to: surface chemistry, biological coatings, surface charge, and topography, Fig. 4.8. Special methods must be employed to characterize surface properties, as the surface has a different reactivity compared with the bulk material. It also involves such a small amount of material that particularly sensitive equipment is needed and, additionally, the surface is easily contaminated by air or other substances.

4.4.5.1  Surface chemistry The effects of chemistry, or composition, at the material’s surface is one of the most well studied. The composition and chemical structure affect whether cells attach to the surface, resorb the implant material or form a capsule around it. All three types of behavior could be useful, depending on the situation. For applications such as hip and knee replacements, robust cell attachment is important for stability; bioceramics, particularly calcium phosphates, are noted for their chemical similarity to the inorganic constituent of bone. Thus, coatings of hydroxyapatite (HA) and tricalcium phosphate (TCP) have been shown to enhance in vivo response to metallic implant materials [84,85]. In addition, to coatings, implanting ions favorable to the body, within the surface of a material, can change the binding behavior of cells. For example, silicate or zinc substituted hydroxyapatite delay osteoclast formation, while strontium enhances the maturation of osteoblast cells, both of which can lead to greater bone formation [86,87].

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Fully characterizing the surface composition of a material may be tricky, as traditional methods for determining composition, such as XRD, may penetrate too far into the material. Techniques which have been used include: auger electron spectroscopy (AES) and X-ray photoelectron spectroscopy (XPS) [30]. In addition, vibrational spectroscopy is a powerful tool, such as Fourier transform infrared spectroscopy (FTIR) and Raman spectroscopy, which can distinguish between a wide variety of materials, such as cellular membranes based on lipid composition [31,88].

4.4.5.2   Biological coatings A particular class of surface chemistry modifiers are cell adhesion ligands, antibiotics and growth factors, in the form of peptides or proteins. The majority of work, thus far, has centered on adhesion ligands, as these have profound effects on biological responses, independent of substrate stiffness, altering cell spreading, proliferation, migration, and stem cell differentiation [89–91]. The most well known class of adhesive ligands is the RGD (Arginylglycylaspartic acid) peptide sequence, found most often in fibronectin, fibrinogen, and gelatin. Several factors can affect the how cells respond to ligands, although specific interaction depends on the cell type. The global density of ligand influences the signaling effectiveness [89]. In addition, the spacing and clustering of integrins directs the formation of focal adhesions at the surface, changing the forces generated by the cell, which regulate processes such as differentiation [74]. Like any protein, the conformation of the peptide ligands also modifies cellular response. In the case of RGD ligands, cyclical peptides trigger different integrin expression than linear RGD [74]. As many natural polymers have some intrinsic adhesion ligand, they are most often used to modify synthetic polymers or ceramics, which are less bioactive. However, natural polymer structures can also benefit from an increasing concentration of ligand, especially after undergoing processes, such as cross-linking, which tend to alter sites where adhesive ligands sit [77]. Ligands can be absorbed or chemically cross-linked to the surface. Often their effect on cellular response is measured by changes in cellular adhesion or spreading.

4.4.5.3   Surface charge Changes in the surface charge affect cellular spreading and affinity for the surface of a material [92]. Charge makes a surface more conducive to tissue integration, with both a net positive or net negative charge shown to promote osteoblast adhesion [92–94]. Thus, bone formation is increased on wettable, hydrophilic surfaces [95]. The most common metric for quantifying surface charge of flat substrates is via wettability, or the affinity of a surface for liquids. Wettability is most often measured via contact angle [31,32]. Simply, a drop of water or other liquid is placed onto the surface of the material and the angle the edge of the droplet makes with the surface is measured. If the water droplet spreads out, the surface has good wettability and is said to be hydrophilic. If the water droplet beads up, it has poor wettability and is hydrophobic. If a surface has a net negative or positive charge it will be hydrophilic, while if it has a neutral charge it will be hydrophobic. For materials in suspension or emulsion, the zeta potential can be utilized to quantify the surface charge [31].

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4.4.5.4  Surface topography Topography, the sum of the three-dimensional features at the surface, is present at a number of length scales, from the nanometer range to millimeter scales. At every scale, evidence suggests that cells are sensitive to topographical features, but specific response is determined by the type of cell. Topography is an extremely powerful tool for controlling biological behavior, as it can influence cellular proliferation, metabolism, matrix synthesis, and differentiation [83,96]. By some estimates, surface topology may be more influential over the rate at which bone is formed at the material surface, than oxide thickness or microstructure [97]. Topography, especially grooved structures, can effectively align cells and increase migration in a particular direction. Surface roughness, while less ordered, can also affect cellular adhesion, proliferation and phenotype [98]. To encourage cellular attachment, porosity and roughness are more biologically favorable than smooth surfaces [99]. However, when additional roughness interferes with stability, or alters the density of adhesive ligands, it is undesirable [45]. The complex interaction between topography and cellular response has encouraged some groups to create “topographical libraries” to catalog topography and corresponding cellular behavior [100]. A good overview of the surface roughness of a material can be obtained by SEM. However, techniques such as atomic force microscopy (AFM) and scanning tunneling microscopy (STM) give a more sensitive, quantitative measurement on the nanoscale regime [32]. For large-scale topographical features, in the micro- to millimeter range, interferometry techniques can be very helpful, as they are noncontact and have a high degree of accuracy [33].

4.4.6  Biomineralization A large component of bone tissues is the mineral itself. In the body this is deposited as highly oriented nanoscale crystals along a collagen fiber matrix. Materials for bone repair, which are meant to interact with the native environment and lead to the production of healthy bone tissue, must encourage mineralization. These materials are classified into several types. Osteoinduction refers to the ability of a material to stimulate differentiation of a cell toward an osteoblast lineage which will lay down mineral and is often demonstrated by implantation in soft tissue environments, such as muscle. Osteoconduction is an ability to allow the deposition of mineralized tissue on the surface. The end goal of most implants is osseointegration: a direct bond between the implant surface and bone tissue, without a fibrous layer in between [101]. Osseointegration can alleviate problems with aseptic loosening of implants, which is an important cause of orthopedic implant failure. The ability to stimulate mineralization is linked to the availability of calcium and phosphate at the material interface. Ceramics, especially glass ceramics, are believed to stimulate biomineralization by rapidly dissolving at the surface and then reprecipitating through the action of bone forming cells. Composites mimicking the composition of bone, for example biomineralized collagen/gelatin, report increased mineralization compared to those without a mineral component [102]. Increasing mechanical stiffness has also been linked to increased mineralization, an effect observed even with hydrogels and natural polymer structures, which are order of magnitudes less stiff than

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ceramics or metals [2,45,103]. These properties are assessed from in vivo tests, and often quantified with histology, fluorochrome staining of bone deposition, and μCT, which can distinguish the mineral phase from surrounding soft tissue [85].

4.5  Comparing material classes The previous three sections have identified and discussed the mechanical, chemical, and architectural properties, along with the physiological effects that are important to the in vivo performance of a material. Several materials classes have been alluded to, which will now be compared, in terms of their overall performance and suitability, for different bone repair requirements. More in-depth discussions of these material classes will be covered in later chapters.

4.5.1   Metals Metals possess a great many attributes which make them highly suitable for bone repair applications which experience large loads, such as orthopedic implants. Compared to other classes of materials, summarized in Table 4.2, metals possess excellent tensile strength; they far surpass polymers and strength is markedly higher than any of the bioceramics, except for zirconia. While not as strong in compression as ceramics, metals are moderately ductile with much higher toughness, which is about 20 times greater than that of ceramics. Although not listed in the table, metals are fairly hard, have a reasonable fatigue life compared with other materials, and are easily machined. If selected properly, they also have good corrosion resistance and do not elicit adverse biological responses. Their ability to alloy also means their properties can be specifically tailored for a particular application. For these reasons, many hip and knee implants have metal components. However, metals possess several drawbacks. First, while those used in biological applications are not toxic in bulk, they also do not interact with body tissue in a synergistic way, as other materials do. While bulk metals are biocompatible, metallic wear particles are often toxic. This fact is driving a move away from metal-on-metal implants at articulating surfaces [79]. Also, while good mechanical properties can be beneficial, the strength and stiffness of metals exceed those of native bone. As a metabolically active tissue, bone tends to remodel itself according to the loads it experiences, a theory formalized as Wolff’s Law. In cases where an implant with superior mechanical properties is used in a major load-bearing orthopedic application, the metal bears greater stress than the bone, causing the bone to remodel itself accordingly, and eventually resorbing. Efforts to address these points have driven the use of porous metallic scaffolds in recent years [104].

4.5.2   Ceramics Ceramics, like metals, have many mechanical properties which make them attractive for implants: high wear resistance, low coefficient of friction, and high stiffness. Ceramics have very high compressive strengths, but comparatively low tensile strengths, Table 4.2. Where the tensile strength is low, such as hydroxyapatite (HA),

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their use is limited in load-bearing applications. In addition, a low toughness, 20 times less than metals, is also limiting for the use of ceramics. In particular the toughness of HA, the bioceramic most similar in composition to the inorganic component of bone, is far below that of cortical bone. Although bone contains a considerable amount of the inorganic “ceramic” phase, it is, in fact, a composite, with collagen fibrils which run longitudinally through it, thus greatly improving its toughness. These fibers prevent the type of catastrophic failure that plagues biological glasses and calcium phosphates. The high stiffness of ceramics might also pose a problem in a load-bearing situation owning to modulus mismatch with surrounding bone tissue. The biological response to ceramics is very dependent on the composition. Ceramics such as alumina and zirconia, are bioinert, having neither an adverse nor beneficial effect on the surrounding tissue. Others, such as calcium phosphates, can bond directly to tissues and demonstrate osseointegration, due to similarities with the inorganic portion of bone itself. There are many phases of calcium phosphates, varying in their mechanical properties and degradation rates. HA, for example, will not degrade in vivo, while TCP dissolves readily over time, releasing ions which can be incorporated into new bone growth. Bioactive glasses and glass ceramics, including Bioglass, developed by Hench et al. [105], and apatite-wollastonite (A-W) glass-­ ceramic, are a special class of resorbable bioceramics. They encompass a wide range of biological behavior, depending on the exact proportions of CaO, SiO2, Na2O and P2O5. Their behavior is driven by the dissolution of ions from the glass, which can then reprecipitate at the surface to form a strong bond with surrounding bone [106]. With excellent biocompatibility, the adoption of ceramics as bone repair biomaterials is limited by their poor mechanical properties. Ceramics with low toughness, like HA and bioactive glasses, are used only in non-load-bearing applications: as powders to fill in bony defects, middle ear implants, and alveolar ridge reconstruction, or as coatings on orthopedic, dental, and maxillofacial prosthetics. The significantly higher mechanical properties of certain glass ceramics enable their use in higher load-bearing situations in porous, granular, and even bulk forms, mainly in the spinal area. Alumina and zirconia, on the other hand, with their superior mechanical properties compared with other bioceramics, can be used in situations requiring both strength and wear resistance, such as femoral heads and acetabular cups.

4.5.3  Synthetic polymers Synthetic polymers cover the entire spectrum of mechanical properties, depending on their degree of crystallinity, cross-linking, and molecular weight. This almost infinite variability is a key benefit when tailoring them for specific applications. However, it becomes vital to study polymer properties carefully. In most aspects, the mechanical properties of polymers, Table 4.2, do not compare well to those of bone. On one end of the spectrum, amorphous, rubbery polymers are soft and ductile; their low modulus allows them to extend hundreds of times their original length. Semicrystalline polymers have much higher moduli and lower extensibilities. A common trait in many synthetic polymers is their high toughness. The toughness is the result of yielding behavior, allowing the material to plastically deform and hinder the propagation

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of microcracks. This is not true of all bone repair polymers. PMMA, for example, exhibits brittle behavior. Given the large range of properties for polymers, they must be carefully selected so that the glass-transition temperature is above 37°C, to ensure they resist creep in physiological environments. Additionally, when used at articulating surfaces which experience friction, they should have sufficient cross-linking, molecular weight, and other chemical properties to resist wear. UHMWPE, used commercially for over 30 years, has excellent properties for the load-bearing surfaces in total knee and hip replacements, such as acetabular cups in hip prostheses. New trends aim to mitigate the negative immune responses to wear particles by incorporating therapeutics, such as vitamin E into the polymer, to be released as the material wears [107]. Another benefit of polymers is that they can be manipulated with relative ease into a variety of shapes: sheets, films, scaffolds, and fibers. Like other properties, polymers also exhibit a wide range of biological responses. Surface coatings, either with adhesion ligands or chemical groups, are often required to obtain surfaces which interact with living tissue. In many applications, polymers are designed to be biodegradable. The rate of degradation must be timed carefully to allow cells time to create their own matrix which can carry out native functionality, and further to make sure that the mechanical properties are not compromised before stabilization of the implant site has occurred. For bone implant materials, biodegradable polymers, owing to their weak mechanical properties, are often combined with ceramic particulate reinforcements to create a scaffold material with a controllable degradation rate.

4.5.4  Natural polymers and hydrogels Natural polymers and hydrogels, cross-linked networks which absorb large amounts of water without dissolving, are used as scaffolds or injectable materials for bone repair. While many natural polymers form hydrogels (collagen, gelatin, alginate), hydrogels can also be created from synthetic polymers, most notably, polyethelyne glycol (PEG). These materials are by far the weakest used for bone regeneration, Table 4.2. In native tissue, natural polymers, such as collagen fibers, support large mechanical properties. However, the process of harvesting natural proteins and creating a biomaterial implant removes the structural features which contribute to the mechanics. Thus, natural polymers are not meant for any application which sees high mechanical stresses or load-bearing. While macroscopically weak, on the microscale, these polymers can activate cellular machinery associated with mechanosensing, due to the arrangement of ligands on the surface [37,40]. One further drawback to polymers sourced from natural tissues is the concern over the immunogenicity and batch-tobatch variability. This has led to the rise of recombinant peptides, with a tailored and controllable amino acid sequence [45,103]. Despite lacking in mechanical properties, natural polymers and hydrogels remain attractive for their inherent biocompatibility, encouraging greater cellular attachment and matrix deposition compared to any other class of material. In addition, when the structure of hydrogels and scaffolds have high permeability, cells and vasculature can infiltrate through the scaffold without being blocked by adverse immune responses.

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By utilizing natural polymers which play a role in the body’s healing response, such as fibrin clots, proliferation can be increased, and the appropriate healing responses triggered [108,109]. As materials which are degraded naturally in physiological conditions, especially in the presence of cell enzymes, they make efficient vectors for the delivery of drugs and growth factors [110]. Several methods have been devised for overcoming the limitations imposed by mechanical properties. Utilizing methods such as cross-linking, networks can be strengthened, but care must be taken to avoid destroying the intrinsic structure and making the material unrecognizable to cells [77]. In many cases, natural polymers and hydrogels are used in combination with another material which can provide mechanical stability, while the hydrogel provides biocompatibility [108,111,112].

4.5.5  Composites As noted in the preceding sections, the material classes have some drawbacks, which limits their potential. Combining materials offers a way to better tailor the materials properties and biocompatibility of implants for specific applications. Simply stated, composite materials consist of two or more phases, distinguishable on the microstructural scale, whose constituents interact and result in a single material with unique properties. Natural design relies heavily on composites, such as bone itself, which consists of a matrix of a ceramic inorganic phase reinforced with fibrils of collagen. The overall properties of a composite are affected by several factors: the properties of the constituents, how well the phases are distributed, and in what manner, and the loading of the two phases with respect to one another. While the properties of the composite material might be a weighted average of the properties of the constituents, this is not always the case. Geometrical orientation of interfacial interactions may allow the phases to act synergistically, resulting in properties exceeding either of the constituent materials. A viable biocomposite must be: biocompatible, have good strength and fatigue properties, good toughness, and a modulus matching that of the material it is replacing. A final consideration in forming composites is, of course, how they will react in the body and how the body will interact with them. Often composites are constructed to have a reinforcing phase inside of a matrix, in the form of fibers (long or short) or particles. The aspect ratio distinguishes the two; if the ratio of the length to diameter of the phase is approximately the same, then it is a particle, whereas a large length to diameter ratio is representative of a fiber. Particle reinforcement provides isotropic properties, whereas aligned fibers result in anisotropic properties. To fully describe a composite, the type of reinforcement, its loading, how well it is distributed throughout the matrix, its orientation (aligned or random), and its size with respect to the matrix should be specified. Smaller reinforcements offer more surface area with which to interact with the matrix. If well dispersed, they also allow properties to be distributed throughout the material more homogeneously [113]. Despite advantages in mechanics, the biological response to all constituent materials must be examined. Determining which components to combine, into a composite material, requires examining materials properties to augment the inefficiencies of one with the

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strength of the other and vice versa. Nature’s example of natural bone has led to composites of ceramics and natural polymers. The ceramic strengthens the natural polymer matrix, which lends biocompatibility to the structure [45,114]. Another example is HAPEXTM [115], a composite composed of PE with particulate reinforcement of HA. The PE matrix has a low modulus, but a very high toughness. HA is a material with superb biological properties, but whose mechanical properties severely limit its in vivo use. Its modulus is higher than that of bone and its toughness is much lower. The composite of these two materials yields mechanical properties very similar to those of bone, with HA contributing bioactivity to the normally inert PE. Besides mechanical reinforcement, control over degradation may serve as the primary reason for creating a composite biomaterial. For example, combining a particulate phase, like HA or TCP, with a biodegradable polymer, such as poly(l-lactic acid) (PLLA), poly(lactic-co-glycolic acid) (PLGA) or poly(caprolactone) (PCL) allows the degradation rate of the polymer to be mediated. Degradation can be controlled by varying the loading of the reinforcing ceramic phase or the size of the reinforcing particles [113]. Fiber reinforcement is used to improve toughness and to give anisotropic properties to the composite. If fibers are long and aligned, mechanical properties should be improved along the longitudinal direction [112]. Unaligned long fibers or short fibers still offer a great deal of surface area with which to interact with the matrix while retaining isotropic properties. Fibers can improve toughness by bridging cracks or by fiber pull-out, wherein energy is absorbed by the frictional force required to pull the fiber out of place when fractured. In either case, the main goal of a fiber is to either stop crack propagation altogether or to elongate the path the crack must take to propagate across the material. Fibers can be constructed from many materials including synthetic polymers, natural polymers [112], and even carbon nanotubes [116].

4.6  Summary Materials properties play a large role in the design and fabrication of implants for bone repair. Characterizing a material requires understanding the mechanical properties, architectural and microstructural properties, and physiological effects. While these properties can be tailored to specific environments, it is important to note that they are all interrelated. Thus, altering one property will impact another and reflect on the in vivo behavior of the material and implant. It is therefore imperative to fully characterize any material used for bone repair, and to that end, this chapter has listed the common experimental set-ups and methods for testing each. By considering which properties are most important for each application, the type of material can be selected from several broad classes: metals, ceramics, synthetic polymers, natural polymers and hydrogels, and composites. As an area of active research, it is hoped that the available materials for bone repair continue to grow, allowing for better therapeutic outcomes for bone repair in the future.

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