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Biomaterials in Orthopedics

  • Ismet KoksalEmail author
Living reference work entry
  • 227 Downloads

Abstract

Biomaterial technology offers several alternatives according to their biological, chemical, and physical properties. Within this respect, tremendous advances in the biomaterials field have been carried out in the last 50 years. In this chapter, biomaterials, biocompatibility and material deterioration, and the criteria for implant material selection in practice are discussed.

Keywords

Polylactic Acid Bone Cement Orthopedic Implant Orthopedic Application Calcium Phosphate Ceramic 
These keywords were added by machine and not by the authors. This process is experimental and the keywords may be updated as the learning algorithm improves.

Introduction

Over the past few decades, the ambition for longer and qualified life has led to an increase in the average age of the world population. While it was 29 years in 1950s, it reached to 37.3 years in 2000 and it is expected to be 45.5 years by 2050. Therefore, the number of middle-aged population that undergoes orthopedic surgery is subsequently increasing. This fact presents the need for advanced biomaterials and implant technologies (Nag and Banerjee 2012). It is essential to understand the mechanical, chemical, and biological properties of materials, both for the design of an orthopedic implant and for the proper selection of an implant on the side of surgeon.

Classification of Biomaterials

The material intended to perform the function of an organ or a tissue partly or completely in living organisms is called biomaterial, and biocompatibility , nontoxicity, inertness, appropriate mechanical properties, weight, and density are the desired properties for these materials. Furthermore, it is important that biomaterials are easily reproduced, fabricated, and sterilized and that they maintain their properties throughout their shelf lives (Patel and Gohil 2012; Parida et al. 2012). They can be used for replacement of diseased and damaged part (artificial hip joint), assisting in healing (sutures, bone plates), improving function (pacemaker), and correcting functional abnormalities, diagnosis (probes and catheters), and treatment (catheters and drains). They can be generally classified according to material types (metals, ceramics, polymers, composites, natural materials), duration of use (temporary, short term, long term), or usage areas (in vivo, in vitro).

Metals

Metallic biomaterials are important for the reconstruction of hard tissue due to their conductivity, biological compatibility, mechanical properties (high tensile and compression modulus, appropriate elasticity, and plasticity deformations), resistance to environmental factors such as sterilization and implantation, and reasonable cost (Patel and Gohil 2012; Niinomi et al. 2012). However, they have some disadvantages like low corrosion resistance and high density (Hermawan 2012). Metallic biomaterials can be grouped as (1) stainless steel (ss) alloys, (2) cobalt (Co)-chrome (Cr) alloys, (3) titanium (Ti) and its alloys, (4) costly alloys, and (5) others (Hermawan 2012). Comparing with the other metals, SS alloys are preferable since they are cheaper; they have higher strength, elasticity, and hardness; and they are easily processed. However, they have some limitations such as corrosion and low bone adhesion (Niinomi et al. 2012). Co–Cr alloys have been in use in dentistry and artificial joints for many decades due to their high corrosion and wear resistance (Hermawan 2012; Niinomi et al. 2012). Ti has good tensile strength and corrosion resistance. Therefore, it is preferred for implant applications (Hermawan 2012; Niinomi et al. 2012). Costly metals such as gold (Au), silver (Ag), and platinum (Pt) have good castability, ductility, and resistance to corrosion (Hermawan 2012). Aforementioned “others” group includes tantalum (Ta), amorphous alloys, and biodegradable metals that are used as implants (Hermawan 2012). Porous Ta is used for bone ingrowth in orthopedic applications such as hip and knee arthroplasty, spinal surgery, and bone graft substitutes since it has good chemical stability and biocompatibility (Niinomi et al. 2012). Amorphous alloys display better corrosion and wear resistance, tensile strength, and fatigue strength than its crystalline counterparts (Hermawan 2012). For the purpose of improvement of MR susceptibility, zirconium-based alloys such as Zr–Nb (zirconium–niobium) and Zr–Mo (zirconium–molybdenum) have been developed (Niinomi et al. 2012).

Ceramics

Ceramics including silicates, metallic oxides, carbides, and various refractory hydrides, sulfides, and selenides are refractory, polycrystalline compounds, and they are generally inorganic. Ceramics can be used as a part of the musculoskeletal system, dental, and orthopedic implants (i.e., femoral head and acetabular cup of hip implants) (Best et al. 2008; Guven 2010; Patel and Gohil 2012). Bioceramics can be classified into three basic types: bioinert ceramics, bioactive ceramics that chemically bind to biologic tissue, and bioresorbable ceramics that degrades by the time in the body. Alumina (Al2O3), zirconia (ZrO2), and carbon are examples of bioinert ceramics. Synthetic hydroxyapatite (HA) (sintered at high temperature) and glass ceramics are in bioactive class. Calcium phosphate ceramics are in bioresorbables. Alumina was the first ceramic widely used in clinic. Owing to its low coefficient of friction, minimal wear rate, good abrasion resistance, strength, and inert behavior, it has been used in orthopedics (Best et al. 2008; Patel and Gohil 2012). Zirconia is an important biomaterial due to its mechanical properties such as good strength and fracture toughness (high crack and twist resistance), and therefore, it can be used in total hip replacement (THR) as femoral heads and in stem applications (Guven 2010; Patel and Gohil 2012). Carbon bioceramics have good compatibility and fatigue resistance, and because of the similarity between mechanical properties of carbon and bone, it can be used in orthopedic implants, but their brittleness and low tensile strength inhibits their load-bearing usage in this field. HA {Ca10[PO4]6[OH]2} is the most widely used synthetic calcium phosphate ceramic which has the ability to bind to the biological tissue, thus providing bone ingrowth (Best et al. 2008). Such reasons make it useful for coating on materials such as Ti since it does not have appropriate strength and toughness for load-bearing applications. The bioactivity property makes them preferable for rapid bone repair after surgery (Best et al. 2008). The biodegradable ceramics are used in orthopedic applications and the speed of degradation is inversely proportional to the decreasing Ca/P ratio in some calcium phosphate ceramics that makes them useful for implantation (Ratner et al. 2012). Ca3(PO4)2has the porosity structure, which is similar to spongy bone, and the total porosity determines the mechanical behaviors such as tensile and compressive strength and fatigue resistance (Yoruc and Sener 2012).

Polymers

Polymers are huge molecules, formed from joining of small molecules known as monomers (Sarsılmaz and Sarsılmaz 2003; Nag and Banerjee 2012; Patel and Gohil 2012; Niinomi et al. 2012; Parida et al. 2012). Polymers’ properties, including strength, softness, hydrophilic, hydrophobic, flexible, porosity, and non-porosity, offer compliance with the various organs (Sarsılmaz and Sarsılmaz 2003). Polymers can be categorized as natural and synthetic polymers. Synthetic polymeric biomaterials range from hydrophobic, non-water-absorbing materials (polyethylene (PE), polypropylene (PP), polytetrafluoroethylene (PTFE), and poly(methyl methacrylate) (PMMA)) to somewhat more polar materials (poly(vinyl chloride) (PVC), and nylons), to water swelling materials (poly(hydroxyethyl methacrylate) (PHEMA)) and, beyond, to water-soluble materials (poly(ethylene glycol) (PEG or PEO)) (Ratner et al. 2012). PMMA, which is a good core material, has similar elastic modulus to that of bone and fills the area between bone and implant, enabling force transfer from the surface of the prostheses to the bone. PMMA has good resistance to compression; however, it has weak shear strength. In addition, it has also many disadvantages, such as causing cell destruction at high temperatures (90 °C’) in the bone-cement surface and accelerating the blood flow that leads to itching. UHMWPE (ultra-high-molecular-weight polyethylene) has been preferred due to its suitable physical, chemical (inert), and mechanical properties in orthopedics. Inertness causes slow degradation, which prevents the emission and excretion of particles. PP resembles to PE, but it is stiffer. It has also high chemical resistance and tensile strength. PTFE (Teflon) is a hydrophobic material, which has an excellent lubrication (Sarsılmaz and Sarsılmaz 2003).

Composites

Composites are composed of a matrix and natural fibers (Guven 2010; Patel and Gohil 2012). Material properties like stiffness, strength, and weight can be controlled in composite structures by changing its constituent materials, distribution, content, and interaction among them (Hermawan 2012). An example of composite biomaterial is the human bone that is composed of the low elastic modulus organic matrix and the high elastic modulus mineral fibers (Hermawan 2012). Other examples of biocomposites are orthopedic implants with porous structures, dental filler, and bone cement composed of PMMA and UHMWPE (Hermawan 2012; Patel and Gohil 2012; Parida et al. 2012). The use of biocomposites in orthopedics depends on their high strength and low elastic modulus (Guven 2010; Patel and Gohil 2012). Subsequent to the failure of a biocomposite material implant, fibers or particles can be exposed to the surrounding tissue. Since biocomposites are composed of two or more materials, they have a high probability of causing adverse tissue reactions (Ratner et al. 2012).

Natural Materials

Natural materials would not cause adverse effects such as toxicity, infection, etc., because of their similarity with the macromolecules in biologic tissues. They have some drawbacks like degrading at high temperatures, being immunologic, and having high costs due to manufacturing difficulties. The wide variety of natural polymers relevant to the field of biomaterials includes plant materials such as cellulose; animal materials such as tissue-based heart valves and sutures, collagen, heparin, and hyaluronic acid; and other natural materials such as deoxyribonucleic acid (DNA) (Ratner and Bryant 2004).

New Uses of Biomaterials in Orthopedics

Biodegradable Materials

The need for re-surgery, which increases the cost and patient morbidity, can be overcome by the use of biodegradable materials (Brar et al. 2009). Biodegradation process is the breaking down of biomaterials naturally either without enzymes and/or with enzymes. The rate of degradation should allow for tissue healing process, and the conformance between these two rates can be controlled by the hydrolytic bond, block copolymers, hydrophilic and hydrophobic groups, and crystallinity. The residuals of degradation should be nontoxic (Piskin 1995).

Biodegradable Polymers

Biodegradable (bioresorbable, bioabsorbable) polymers like polylactic acid (PLA) and polyglycolic acid (PGA) can be used in orthopedic applications such as non-load-bearing small bones, joint fracture fixation, arthrodesis, tendon, and ligament repairs (Eglin and Alini 2008). Furthermore, they can be used as scaffold materials for cell growth (i.e., chondrocyte). Biodegradable polymeric materials with their complete absorption time and mechanical properties loss time are shown in Table 1 (Kontakis et al. 2007).
Table 1

Biodegradable polymeric materials’ absorption and mechanical properties loss time

Material

Complete absorption time

Mechanical properties loss time

PGA

4–7 weeks

36 weeks

SR-PGA

3 months

1 month

6–12 months

PLLA

>5 years

 

SR-PLLA

5–6 years

Reduction to cortical bone levels in 36 weeks

>5 years

P(D/L)LA 70/30

2–3 years

18–36 weeks

PLA/PGA (PLGA) 80/20

1–2 years

6–8 weeks

1–1.5 years

P(D/L)LA 96/4

2 years

 

PDS

2 months

 

Biodegradable Metals

Biodegradable metals can be identified as magnesium (Mg) alloys, pure iron (Fe) and Fe alloys, and metallic glasses developed from both Mg- and Fe-based alloys (Hermawan 2012). The fracture toughness of Mg is higher than ceramic biomaterials, while the elastic modulus and compressive yield strength of Mg are closer to those of human bone than other metals (Brar et al. 2009; Hermawan 2012). In addition, Mg stimulates bone growth and healing processes (Brar et al. 2009). However, the corrosion resistance of Mg is far inferior to those of Ti and Co–Cr alloys. They were firstly used in an orthopedic application as a pure Mg plate along with gold-plated steel nails, but the corrosion of that implant was so quick that it degraded in only 8 days. The corrosion resistance of Mg alloys for biomedical applications has to be improved since the time for degradation is still short (Hermawan 2012). Purification of Mg decreases the corrosion rate as it also decreases the yield strength. Therefore, its usage in orthopedics and other load-bearing applications is limited. This problem can be overcome by adding some other alloying elements essential to human body, microstructure design, processing, and/or an additional coating (Brar et al. 2009; Hermawan 2012). For temporary implant applications like stents, pure Fe and Fe alloys such as Fe–Mn and Fe–Mn–Pd have been developed as biodegradable metals. In biodegradable metals, transportation and elimination of degradation products should be considered carefully in the design procedure (Hermawan 2012; Niinomi et al. 2012).

Biodegradable Ceramics

In 1920, tricalcium phosphate (TCP) was suggested as a bioabsorbable material to fill bone defects. Subsequently, the first strong and tough alumina (Al2O3) has been designed for hip joints and patented. Synthetic calcium phosphate (CaP) ceramics and zirconia (ZrO2) was pointed out as other options to TCP and alumina (Eliaz 2012). The main characteristic of biodegradable ceramics is that they are designed to help in the self-repair processes in the living organism for a given period and then resorbed. Thus, problems associated with performance of a synthetic material in the living body will be avoided. The real challenge in the design of biodegradable ceramics is to adapt their degradation kinetics to living tissue formation, which is normally slower (Vallet-Regi et al. 2012). Additionally, as resorption takes place, there is a decrease in the mechanical properties of the ceramics, which could be problematic to the role assigned to the biodegradable ceramic. After roughly 100 years of controlled clinic use, the in vivo and in vitro degradation mechanisms of ceramic materials are still largely unknown. Crack propagation mechanisms (low temperature degradation, shocks, wear, and dissolution–reprecipitation) and the in vivo environment remain to be firmly established (Eliaz 2012). There are also other biodegradable ceramics that are discussed in further sections.

PEEK

PEEK is a colorless, linear aromatic organic polymer that belongs to the family of polyaryletherketone (PAEK). Its properties, including toughness, superior strength, biocompatibility, ideal imaging properties, optimal modulus, excellent chemical resistance, and the ability to repeatedly be sterilized without the degradation of its mechanical properties, make it one of the most adaptable biomaterials for long-term implants. While metals/ceramics are preferred for hard tissue applications, polymers are suitable for soft tissue applications (Patel and Gohil 2012). The modulus of implantable-grade PEEK can be adapted in contrast to metals and ceramics and therefore prevent the stress shielding effect (Kurtz and Devine 2007). With the fiber reinforced PEEK, it is possible to enhance its mechanical strength and physical performance. It was already used in the finger joints, hip and femoral bone replacements, bone screws, and pins.

Synthetic Ligaments

Synthetic ligaments have been used for several decades in the reconstruction of various ligaments. Eight hundred fifty-five prosthetic ligaments, which have been tracked for 15 years, were failed in a range of 40–78 % due to the tissue reactions, wear debris, and mechanical limitations (Vunjak-Novakovic et al. 2004). Permanent prostheses were designed to have high initial and constant strength characteristics and an increased fatigue resistance to overcome continuous cyclic joint loading. Scaffolds for tissue-engineered ligaments must be designed not only to have high initial strength but also to have the ability to transfer the load to the maturing tissue in a gradual manner. Gradual load-bearing transfer provides required mechanical properties and also allows for autogenous cell and tissue ingrowth. Synthetic ligaments have several advantages such as not requiring tissue harvesting; reduced surgery time (30 min approximately); minimum damage to the skin, capsula, tendons, and bone; immediate recovery and excellent functionality; possibility of day surgery; short rehabilitation sessions (4–6 weeks); easier revisions with narrower tunnels; and also they do not have the disadvantages of auto- or allografts (Senni 2013).

Host Response, Biocompatibility, and Material Deterioration

Human body is a corrosive environment for foreign materials that are not compatible with the biological tissues, and therefore, they create a host response such as inflammation, fibrous tissue development, and infection (Ratner and Bryant 2004; Nag and Banerjee 2012). The biggest threat is deep late infection, which occurs some months after the surgery. Tissue injury due to implantation can affect the morphology, function, and phenol type of the cells. The change in biological environment due to the implants can alter the functionality of surrounding tissues, e.g., bone loss due to the stress-shielding effect (Nag and Banerjee 2012). Biocompatibility is the most important factor which Williams defines as: “the ability of a biomaterial to perform its desired function with respect to a medical therapy, without eliciting any undesirable local or systemic effects in the recipient or beneficiary of that therapy, but generating the most appropriate beneficial cellular or tissue response in that specific situation, and optimizing the clinically relevant performance of that therapy” (Williams 1999). Biomaterials can be classified into five groups according to the biocompatibilities such as bioactive (e.g., HA), biotolerant (e.g., SS(stainless steel)), bioinert (e.g., Ti, Co–Cr), bioresorbable (e.g., PLA), and toxic. Biomaterials have to be strictly tested and proven to be biocompatible before the clinical use with general property assessments (tensile, compressive, shear strength, elasticity modulus, thermal properties, chemical properties, etc.), in vitro assessments (in lab conditions), and in vivo assessments (real tissue reaction). Aforementioned assessments can be performed via physiochemical, mechanical, and biological tests such as cell toxicity, systemic toxicity, thrombus, osmotic brittleness, lethality, intracutaneous toxicity, skin irritation, and hemolysis (FDA CDRH 2013). Besides the host response, there is an implant response that causes delay in the wound healing process. The cells around the implant can completely reject it, leading to chronic inflammation and then removal of the biomaterial. This situation may cause structural failure or migration of the implant (wear, fracture, stress shielding). The implant breaks down in two ways: mechanically (crack, wear, stress cracking, fracture) and biologically (corrosion, calcification, oxidation, dissolution) (Peksen and Dogan 2011; Ratner et al. 2012). Biological breakdown of biomaterials can be seen as polymer degradation, ceramic dissolution, and metal corrosion. These effects can lead to mechanical breakdown (Peksen and Dogan 2011). Wearing or fracture of implants can cause debris formation. Dynamic loading implants (hip, knee, etc.) cause a constant friction between two articulating parts, which can lead to the release of small particles. Stress shielding occurs where unequal distribution of the load between the implant and the bone exists. To prevent such failures, several precautions can be taken such as evaluation of biomaterial, appropriate design, fabrication, and sterilization (Nag and Banerjee 2012).

Selection Criteria for Orthopedic Implants

It is complicated and needs experience to select the most appropriate implant for a specific patient, since it requires knowing both the patient and material/implant properties. The survival of the implant depends on patient properties (such as chronic, systemic, metabolic diseases, cigarette or alcohol usage, drug treatment, host factors like difficult anatomical settling, infection irritation, lack of vascularization), surgery, and implant factors (such as inadequate implant properties and surgery techniques) (Peksen and Dogan 2011; Patel and Gohil 2012). Incorrect selection of an implant can lead to an early failure and subsequently a revision surgery. Orthopedic implant deterioration depends on the implant type and the mobility of surgery region (Peksen and Dogan 2011). Therefore, it is important to take into account the region properties such as mobility and loading while selecting the appropriate implant. Biocompatibility refers to a set of properties such as noncarcinogenic, non-pyrogenic, nontoxic, nonallergenic, blood compatible, and noninflammatory (Nag and Banerjee 2012; Patel and Gohil 2012). Biofunctionality is important in physical and mechanical perspectives that each implant needs to fulfill different functions of biologic tissues. For instance, in bone replacements, load transmissions and stress distribution are important, while in artificial knee joints, articulation comes into prominence (Patel and Gohil 2012). Selecting the appropriate implant for a specific orthopedic application can be achieved by satisfying the mechanical affinity between material and natural tissue. Some of the most important mechanical properties of biomaterials are tensile strength, yield strength, elastic modulus, corrosion and fatigue resistance, surface finish, creep, and hardness (Nag and Banerjee 2012; Patel and Gohil 2012). Yield strength parameter gives the information about the load-bearing capability of the implant. For instance, in joint replacement surgeries, a high load-bearing capability is needed, and therefore, an appropriate alloy with high yield strength has to be used. The elastic modulus (young’s modulus) describes the stiffness of the material. Implant modulus and natural biologic tissue modulus should be close to each other to prevent the adverse effects like bone mass loss and increased osteoporosis (Abbaszadeh et al. 2009). This is called stress-shielding effect, which consequently causes loosening and failure of implant (Patel and Gohil 2012). Corrosion is a degradation mechanism that implants release undesirable nonbiocompatible metal ions by chemical reactions and cause implant weakening and loosening. High wear resistance is crucial especially for orthopedic implants, because the low wear resistance or high coefficient of friction results in implant loosening since wear debris are biologically active and cause an inflammatory response that lead to the destruction of the healthy bone which supports the implant. Furthermore, friction corrosion results in noncompatible metallic ions’ release. Long fatigue life (fatigue strength) is related to the response of the material to the repeated cyclic loads and depends on the microstructure of biomaterials, which can be altered by the processing and heat treatment techniques. Fatigue fracture can cause implant loosening, stress shielding, and ultimate implant failure which is reported for hip prostheses (Peksen and Dogan 2011; Patel and Gohil 2012). When the bone implant interface starts to fail, a fibrous capsule develops and makes more relative motion between the implant and the bone under loading. As a result, patients feel pain and then undergo a revision procedure.

When the implant is placed into the body, a host response occurred by the host organism (local and systemic) to the implanted material or device and a highly complicated period including events at cellular or tissue level between implant and biologic tissue begins. Furthermore, there can be metal allergy and carcinogenic effects in the patient’s body (Peksen and Dogan 2011; Patel and Gohil 2012). It was shown that bone cement has negative effects on complement system, phagocytosis of microorganisms that results in tissue death, and therefore fibrosis. This can result to implant failure. The survival of implant in vivo highly depends on the bond between the bone and the implant. Bone cement provides mechanically strong bond between bone and implant. Without cement use, this interaction can be provided by bone growth through the implant surface. The strength of this interaction depends on close contact and minimum motion at the interface. For instance, in the literature it was shown that more than 0.5 mm space and 150 μm motion between tissue and implant surface negatively affects the bone growth (Kuzyk and Schemitsch 2011). Besides these factors, the implant surface has a crucial role in implant survival. Implant materials cause a local effect in the host body that can trigger systemic effects. In orthopedic implantation, the use of bone cement alters the local tissue effect. After bone-cement application, a thin layer between bone and cement occurs which includes macrophages because of the mechanical trauma during surgery. The particles of materials can cause osteolysis by triggering the macrophages, which increase the synthesis of bone resorption factors, suppressing the functions of osteoblasts. As a result, implant loosens and other particles are released by rubbing. The primary response for these particles is the foreign material response. The further local effects of implants are the changes in bone geometry because of the changes in mechanical loading, which is known as stress shielding effect. This effect accelerates the particle release into the host body and therefore causes osteolysis. The particles released from metallic implants like Co, Cr, Ni, V, Ti, and Al can be toxic above normal limits and cause several diseases such as anemia, osteomalacia, cardiomyopathy, atopic dermatitis, or cancer. Metal allergy, which is rarely seen in orthopedic implant applications, can also be one of the reasons of implant loosening. The amount of particles released from materials depends on mechanical effects, wear properties of materials, and surgery techniques. This amount can be decreased by the appropriate material, implant design, and surgical technique. Long-term effects for metallic implants can be carcinogenic when an implant is applied without bone cement to a young patient after several years. Some in vivo studies show that Co, Cr, Ni, or SS can be carcinogenic (Peksen and Dogan 2011). Although there are some reported case studies about malignancies in patients with joint implants, still there are uncertainties about long-term effects (Cohen 2012).

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Copyright information

© Springer-Verlag Berlin Heidelberg 2014

Authors and Affiliations

  1. 1.Social Security InstitutionAnkaraTurkey

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