Feature Article

Athletes, especially those competing at the Olympic level, are keenly aware of the damage that high-intensity training can do to the body, especially to bones. Novel biomaterials technology may help fractures to heal better and faster, thereby helping athletes to renew their training regimen and sports activity more quickly.

When it comes to sports-related bone injuries as opposed to diseased tissues, new biomaterials should be formulated to serve as a template, stimulating and facilitating new tissue growth quickly during the healing and recovery process. Before looking at new developments in biomaterials, we will discuss basic fractures and fracture mechanics and the fundamentals of bone deformation at the nano and micro level.

Basic fractures and fracture mechanics

Figure 1: In the bone fracture shown here, compressing stress (left) combines with tensile stress on the opposite side.

Figure 1 shows a bone fracture. A force applied to one side of the bone leads to bone fracturing. At least two types of stress have been applied: compressing stress at and towards the damage point on the left, where the force is applied, and tensile stress on the opposite side, which forces the opening of the fractured bone. Shearing stresses also contribute to the fracture, as indicated by the irregular fracture surfaces.

Fracture mechanics recognises three basic modes of fracture (Figure 2). Mode I is the opening fracture, mode II is known as in-plane shearing and mode III is out-plane shearing. Of the three fracture modes, the most serious is mode I. Normally, bone-related damage and/or fractures often involve a combination of two or three modes, as indicated in Figure 2. In addition, there are different types of stresses, including tensile, compressing and shearing forces.

What determines if a bone is going to fracture or not? The concept of fracture toughness KIC, in fracture mechanics, is defined by the following equation:

KIC = σc (πa)^1/2    eq (1)

σc is critical failure stress for a given defect size a. This means that fracture stress and defect size determine the fracture toughness of a material. Failure stress in human bone varies approximately from 1.5 MPa.m^1/2 for a weak bone up to 3.5 MPa.m^1/2 or slightly higher for a healthy strong bone at a normal loading rate ranging from 0.55 to 2.75 MPa.m^1/2s^-1 (reference 1). The higher the KIC, the tougher the bone. Older people tend to have a lower KIC than younger people.

Figure 2: There are three basic modes of fracture; bone-related damage often involves a combination of two or three of them.

Defects always exist. However, it is the largest defect that is a decisive factor for a given stress. Defect a in equation (1) also represents a crack developed from a defect. Figure 3 plots the maximum defect (or crack) size as a function of critical stress sc applied for a given fracture toughness KIC. For example, at applied stress of say 100 MPa, a defect greater than 72 µm will lead to failure if a bone has a fracture toughness KIC of 1.5 MPa.m^1/2 (the blue line in Figure 3); the other two bones have a fracture toughness KIC of 2.5 and 3.5 MPa.m^1/2 and are safe (the red and green lines in Figure 3). They will fracture at a much higher stress—166 and 233 MPa, respectively, at the same 72-µm defect size.

Deformation speed effect on bone fracture

Figure 3: The maximum defect (or crack) size is plotted as a function of critical stress sc applied for a given fracture toughness KIC.

Bone deformation (Figure 4) is represented by a typical stress s and strain ε curve. Mechanical properties change with deformation speed, shown here in fracture stress; Young’s modulus, which is defined by the initial linear slope; and total energy up to fracture Ec, which is the energy of integration under the stress and strain curves up to fracture:

Ec = ∫σdε    eq (2)

Reported failure stresses vary with fracture speed. At a lower speed, bone will not fracture until stresses above 100 MPa are reached; at high speeds, fracture could occur with stresses as low as 50 MPa.², ³ In addition, less total energy is required to fracture bone at a high speed. Fracture toughness will have the same effect, i.e., a higher speed means lower fracture toughness. In sports, accidents often occur unexpectedly at high speed. Therefore, any measure that can lead to a reduction in the impact that speed causes will greatly and effectively minimise the probability of bone fracture.

In the following two sections, we will discuss the fundamentals of time-dependent behaviours of bone fracture that are relevant to biomaterials.

Fundamentals of nano and micro levels of bone deformation and fracture

Figure 4: Bone deformation is represented by a typical stress s and strain ε curve.

The structure of bone has been well documented. It mainly consists of two basic components: collagen and hydroxyapatite (HA) nano crystals. Collagen is made up of three polypeptide strands, approximately 300 nm long and 1.5 nm in diameter (about 1/1000000 the thickness of a human hair), and forms aggregates such as fibrils. HA is a nano crystal platelet approximately 1 to 2 nm thick, 10 to 50 nm wide and 30 nm long.⁴ These two elements are the basic building blocks of human bone from the nano and micro to the macro scale. The volume fraction of collagen, assigned as fc, is given by eq 3:

fc = (ρ- ρh)/(ρc-ρh)    eq (3)

where bone, collagen and HA densities are ρ, ρc and ρh, respectively.

Taking collagen density ρc = 1.19 g/cm³ (reference 5) and HA crystal ρh = 3.155 g/cm³, ⁶ the volume fraction of collagen varies as a function of bone density (Figure 5). In a typical range of compact bone density between 1.9 and 2.0 g/cm³, collagen is a predominantly continuous phase acting as a matrix (> 50% by volume). It is easy to imagine how collagens, which have a fibre-like structure, and platelet-shaped HA are organised. HA crystals are densely packed and, in very large numbers, occupy the spaces between collagen fibres.

Figure 5: The contribution of collagen by volume varies as a function of bone density.

Collagen is an organic polymer (polypeptide) with visco-elastic properties, a unique characteristic of polymeric materials. This structure determines the time-dependent fracture behaviour—the higher the speed of an accident, the more brittle the bone fracture. This is critical to determine whether or not a bone is tough and strong. A higher proportion of collagen in a bone means greater toughness. On the other hand, HA is a rigid filler, providing bone with the required stiffness and strength. In 1 mm³ of bone, there can be 200 million to 1 billion nano HA platelets, which are individually distributed within the collagen matrix. It is clear that all mechanical property variations are rooted in deformation and fracture of collagen fibres and HA at the nano scale first and then at the micro and millimetre scales. Collagen plays a dominant role with the capacity to deform at much larger scales than HA. Importantly, collagen fibres sharing deformation, as most polymers do, determine the fracture toughness of the bone. In sports injuries, the damage and fracture of the bone’s nano and micro structures are predominantly relevant to collagen, and HA is of secondary importance. This is different from other diseased bones. A fundamental understanding of this will guide future development of biomaterials for sports injuries.

Development of new biomaterials
Having discussed the basic nano and micro structure of collagen and HA and mechanical and fracture mechanical properties, the question is what kind of biomaterials are best suited to attend to the biological needs of sports injuries?

Biomaterials for sports injuries should be formulated differently than materials for diseased bones. Ideally, the biomaterials can act as templates and stimulate and facilitate regeneration of new bone. A formulation should have properties and a composition similar to the bone’s constituent materials. Additionally, the new biomaterials should be bioresorbable, acting as a template that gradually dissolves in the body. (Permanent replacements were discussed in a previous publication.)⁷

Collagen fibres are the key materials that are predominantly damaged in sports injuries. Therefore, the new biomaterial system should contain at least a biopolymeric material that will serve as a template for the required mechanical properties and be able to induce and facilitate collagen regeneration. In addition, the system should be able to promote new HA formation and new collagen regeneration. For example, a combination of polyester, such as poly(α-hydroxy acids), and calcium phosphate can do the job (this does not preclude the use of other suitable materials). Polyesters are commonly used biodegradable materials that act as a template for collagen; appropriate calcium phosphates are also highly bioactive, which facilitates new HA formation.

One other important materials factor to consider is control of the speed of bioresorbable processes, bearing in mind the regeneration speed of biological collagen and HA. This requirement is based on the fact that different people (bone density and age variations) have different injuries and, thus, need different healing mechanisms. Controlling the speed at which the new biomaterials work, with the requisite bioactivity, is an important materials factor to consider when developing biomaterials for sports injuries and other orthopaedic applications.

One option is to employ water soluble bioglass (also called bioresorbable glass). The first generation of bioglass developed in the late 1960s by Larry Hench proved to be both biocompatible and bioactive. Since then, many types of bioactive glasses have been developed; indeed, this is still an important research area even today. Here the main concerns are the control of bioresorbable speed at use. New inorganic controlled release technology (CRT) has recently been developed at Ceram, which is based on water soluble glass technology. Figure 6 is an example of this technology that demonstrates the water soluble bioglass degradation (and, at the same time, release of the active ingredient) that can be controlled precisely through formulation design. In reality, water soluble bioglass of this kind can be made with a half-shelf life of minutes to months to suit different needs in medical applications. Further development of the new inorganic CRT will make it possible to develop a series of biomaterial systems that can be tailored to individual needs for sports injuries, taking into consideration that sports injuries need different treatment and healing processes.

Figure 6: Water-soluble bioglass degradation and release of active ingredients can be precisely controlled by means of formulation design.

Since 1969, HA has been one of the most important biomaterials for medical applications. Increasing the bioactive properties of HA can speed up the healing process for sports injuries. A new technology has been investigated at Ceram to substitute different elements (in pairs or greater than two) into the HA crystal structure (mx-HA). This new technology makes it possible, at least, to change HA’s surface charge potential and surface characteristics. A series of new HA-based biomaterials with substituted elements in its crystals, with surface properties and, therefore, biological properties that are different from pure HA, have been synthesised. In one example, the mx-HA has demonstrated greater bioactivity than pure HA.⁸ Taking advantage of the more bioactive mx-HA, it is now possible to formulate biomaterials that increase the speed of HA regeneration as well as collagen and optimise outcomes for sports injuries and other medical needs.

In summary, the goal is to develop new biomaterials for sports injuries and similar applications in order to help the body regenerate new healthy tissue. A fundamental understanding of mechanical properties and fracture mechanics and the nano and microstructure of collagen and HA-based materials is key to new biomaterials development. In addition, a combination of new bioresorbable inorganic CRT, such as water-soluble bioglass, and biopolymers to stimulate collagen growth, will be beneficial in helping new HA crystal grow into the collagen fibre matrix and, thus, accelerate the healing of bone-related sports injuries.

A lot of work remains to be done in this area, but it is hoped that bone-related fractures, in the future, will heal more quickly and more effectively, and help athletes return to training in record, if not Olympic, time.

References:
1.    ASTM E399 Standard Test Method for Plane-Strain Fracture Toughness of Metallic Materials.
2.    R. Simpson, J.D. Currey, D. Hynd, “The Effect of Strain Rate on the Mechanical Properties of Human Cortical Bone,” J. Biomech. Eng. 130, 1 (2008).
3.    A. Ural, et al., “The Effect of Strain Rate on the Mechanical Properties of Human Cortical Bone,” J. Mechanical Behaviour of Biomedical Materials, 7, 1021–1032 (2011).
4.    C. Burger, et al., “Lateral Packing of Mineral Crystals in Bone Collagen Fibrils,” Biophysical Journal, 95, 4, 1985–1992 (2008).
5.    I. Jager, P. Fratzl, “Mineralized Collagen Fibrils: A Mechanical Model with a Staggered Arrangement of Mineral Particles,” Biophysical Journal, 79, 4, 1737–1746, (2000).
6.    Ceram materials characterisation results.
7.    Xiang Zhang, “Advances in Man-Made Materials for Orthopaedics,” OrthoTec, 2, 3 (2011).
8.    P. Jackson, B. McCarthy, “Advances in Hydroxyapatite,” Med-tech Innovation, 1, 6, 26-28.
 

Xiang Zhang, PhD, is
Principal Consultant,
Division of Medical Materials and Devices, Ceram,
Queens Road, Penkhull, Stoke-on-Trent, Staffordshire ST4 7LQ, UK  |  Tel. +44 1782 764 428
enquiries@ceram.com
www.ceram.com/healthcare