Topical Review

Recent advances in bio-medical implants; mechanical properties, surface modifications and applications

Published 12 September 2022 © 2022 IOP Publishing Ltd
, , Citation Mohammed Zwawi 2022 Eng. Res. Express 4 032003DOI 10.1088/2631-8695/ac8ae2

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2631-8695/4/3/032003

Abstract

The demand for bio-medical implants has significantly increased to treat different medical conditions and complications. The latest research in medical and material science is paving the path for the new generation of biomedical implants that mimic the natural bone and tissues for enhanced biocompatibility. A bio-medical implant must be bio-compatible, non-toxic and bioactive. The main reasons for implantation are ageing, overweight, accidents and genetic diseases such as arthritis or joint pain. Diseases such as osteoporosis and osteoarthritis can severely damage the mechanical properties of bones over time. Different materials including polymers, ceramics and metals are used for biomedical implants. Metallic implants have high strength and high resistance to corrosion and wear. Biocompatible metallic materials include Ti, Ta, Zr, Mo, Nb, W and Au while materials such as Ni, V, Al and Cr are considered toxic and hazardous to the body. Bioresorbable and degradable materials dissolve in the body after the healing process. Mg-based metallic alloys are highly degradable in the biological environment. Similarly, different polymers such as Poly-lactic acid (PLA) are used as bio-degradable implants and in tissue engineering. Biodegradable stents are used for the slow release of drugs to avoid blood clotting and other complications. Shape memory alloys are employed for bio-implants due to their unique set of properties. Different surface physical and chemical modification methods are used to improve the interfacial properties and interaction of implant materials with the biological environment. This review explains the properties, materials, modifications and shortcomings of bio-implants.

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1. Introduction

The importance of biomedical implants has increased significantly to treat bone fractures, heart conditions and other medical complications. The need for implants is increasing to cater ever increasing medical complexities. Cutting-edge research in medical science is leading us to the latest generation of biomedical implants that mimic the natural bone and tissues for enhanced biocompatibility. Biomedical implants integrate various mechanical, physical, biological and chemical properties. Biomedical implants must be bio-compatible, non-reactive and must have the functionality of natural tissues. The increase in demand for biomedical implants requires the development of highly advanced and reliable implants. Ageing, overweight, accidents and genetics are the main reasons behind diseases such as arthritis or joint pain that require implants to replace the diseased bone and hard tissues. The weakening of bones and knee joints leads to severe pain, inflammation, and loss of functionality. Moreover, diseases such as osteoporosis and osteoarthritis severely decrease the mechanical properties of bones over time. Osteoarthritis is caused as a result of a decrease in bone density, imbalance in hormones, trauma, accident and change in morphology of bones [1]. Biomaterials and implants are employed for vascular stents, dental implants, knee replacement, bone implants and tissue engineering. Various materials including polymers, ceramics and metals are used for biomedical implants. First-generation metallic implants such as Co–Cr, Ti and stainless-steel-based biomaterials are widely used due to good compatibility, higher corrosion resistance and better mechanical properties. The use of stainless steel dates the back to early 1900s. Steel implant is alloyed with Cr to overcome the problem of rusting and corrosion. The stress cracking of stainless steel implants is triggered by corrosion due to the body fluid and tensile stress. This cracking leads to the leaching of Co, Cr and Ni into the body causing different allergies and infections. Cr can cause issues in the kidney, blood and liver. The hazardous effects of stainless steel implants make them unfit for medical implants in long term. One of the most important parameters in designing and selecting an implant is the reactions of the body to a foreign object. Several reactions take place with the implant and surrounding tissues interface and these reactions determine the success and biocompatibility of the implant. Biocompatible materials include Ti, Ta, Zr, Mo, Nb, W and Au while materials such as Ni, V, Al and Cr are considered hazardous to the body [2]. While designing the implants for load-bearing applications properties such as fatigue strength, elongation, elastic modulus and tensile strength are kept in mind. The higher strength of the bone can lead to the shielding effect and prevent the transfer of stress to the adjacent bone leading to bone resorption and damaging the bone tissues. The implant should have properties close to the natural tissues to avoid any complications regarding load and stress management. Latest biomaterials are also known as bioactive materials because of a favourable response of the host tissues towards the implant. Bioactive materials undergo osseointegration with the perforation of the host tissue into the implant. Over time, the host tissues completely replace the implant material with the release of soluble by-products. The bioactive response of a material is controlled by the pore size and roughness of the surface, growth of host tissues and supply of nutrients. Various ceramic materials are used as bioactive materials such as porous sol-gel foam and poly lactic acid-filled bioglass. A similar bioactive response can also be achieved through the addition or coating of bioactive proteins such as bone morphogenic protein (BMP) to the surface of the porous ceramic implant [3].

The review paper explains various mechanical properties required for biomaterials by comparing elastic modulus, biocompatibility, and corrosion and wear resistance for different biomaterials. The elastic modulus of biomaterial must have to be near the modulus of natural bone to avoid serious complications and stress shielding effect. Different metallic, ceramic and polymeric non-degradable and degradable implant materials are discussed. Moreover, materials that are used in stents and for drug delivery along with design parameters are discussed. Similarly, a detailed analysis is presented for materials used in tissue engineering and scaffolds. The use of surface modification techniques, surface functionalization and nanotechnology to enhance the properties and overcome the shortcomings of biomaterials are analyzed.

2. Properties of implant materials

2.1. Modulus of elasticity

Natural bone has a modulus of elasticity between 10 to 40 GPa. Various techniques are employed to determine the modulus of elasticity of various materials. The modulus of elasticity for different materials is listed in table 1. The modulus of elasticity for the implant materials should be close to natural bone. A high modulus of elasticity leads to the shielding effect. The shielding effect occurs due to insufficient load transfer through the bone and the natural bone fails to gain sufficient strength to bear the stress and load [4]. Low stress and load can result in bone thickness less than required and loss of bone mass. Moreover, the shielding effect can adversely affect the implants through loosening and other complications. That is why, implants will low modulus of elasticity are preferred to uniformly distribute the stress and load across bone and implant, prolonging the life cycle for the implant as well. Beta-phase titanium alloys are considered to have a low modulus of elasticity [5].

Table 1. Mechanical properties of different metallic biomaterials.

MaterialsManufacturing MethodMechanical properties
  E (GPa) σUTS (MPa) σ0.2 (MPa)H (HV)epsilonmax (%)Reference
Ti-6Al-4VCasting Forming1109768473465.1[11]
 Laser Melting109126711104097.28[12]
CP-Ti10555048015[12]
Ti-24Nb-4Zr-8Sn4683070015[2]
316 L SS190480–1290480–69012–40[13]

2.2. Bio compatibility

Materials that are used for the implants are required to be biocompatible, nontoxic and should not trigger any allergic reaction and other medical complications. The success and life period of any implant are dependent on the interaction of the body environment with the implant material. Other factors for biocompatibility of an implant are responses of the immune system and degradation properties of the implant. Implants materials are divided into bioresorbable, bioactive and bio tolerant materials [6]. Bioactive materials are preferred due to their strong affinity and interaction with the body environment. The reactions and responses of the body at the implant/tissue interface dictate the integration of the implant. Successful implants are compatible and bio-active. Some of the issues related to the responses of the body are thrombosis which is the coagulation, clotting or adhesion of blood cells with the implant material and the formation of fatty layer or tissues around the implant material [7].

2.3. Fatigue strength

Fatigue strength and resistance are one of the most crucial parameters to determine the life of the implant. Bone implants undergo cyclic loading due to body motion that can result in plastic deformation. The stress concentrates on manufacturing defects and inhomogeneities leading to the failure of the implant. Fatigue strength and resistance are determined through bending, torsion, tension, and compression. Ti-based alloys have better fatigue resistance and strength [8].

2.4. Corrosion and wear resistance

Implant materials are required to have high resistance to corrosion and wear. Wear and corrosion debris and ions are responsible for some very severe complications such as inflammation, toxic and allergic reactions. Attack by the immune system and release of wear debris are shown in figure 1 [9]. The lifetime of any implant is determined by the wear, corrosion and abrasion resistance. Moreover, low wear resistance can cause the loosening and failure of the implant [10]. Different mechanical properties of metallic biomaterials are discussed in table 1.

Figure 1. Refer to the following caption and surrounding text.

Figure 1. Wear debris for implant material. Failure of implant is mainly due to this wear debris that are released to surrounding tissues and result in bone resorption. Bone resorption results in implant loosening and ultimate failure [9].

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3. Implant materials

3.1. Metallic implants

Metallic implants are preferred due to their excellent mechanical properties, easy availability and low manufacturing costs. In early age, gold and silver were used in dental and bone implants. The implant material must have high corrosion resistance, high compatibility with tissues, and non-toxic and comparable mechanical properties with hard tissues. The metallic implants are graded based on their compatibility and toxicity levels in the body. A range of elements can be found in the human body for various functions as listed in table 2 [11]. The excess of these elements can cause serious complications.

Table 2. The role of various elements in body and effects of high concentration.

ElementsFunction in bodyEffects of high concentration
CuVarious enzymes contain Cu to regulate the transportation of iron in body.Excessive concentration of Cu can cause severe damage to liver. A genetic disease, Wilson disease, causes the accumulation of excess Cu in body.
FeUsed for the transportation of blood in the body.High concentration of Fe in the blood causes liver damage and are responsible for cardiac issue and diabetes.
MnMn helps in the formation of various connective tissues and bone development. Moreover, Mn is important of metabolism of fats and carbohydrates. Mn regulates the sugar level in blood.Deficiency of Mn leads to underdevelopment and deformity in bones while high amount of Mn leads to the poor absorption of Fe.
MoBody uses Mo to process genetic materials and proteins. Moreover, it is used in break down of various drugs and toxic substance inside the body.Excess of Mn causes under development, increase in uric acid and diarrhea issues. Increased uric acid causes muscle and joint pain.
CoCobalt is one of the main continent elements of vitamin B12. It is used in the production of red blood cells and antibacterial compounds to prevent infection.High concentration of Co in body causes damage to heart muscles leading to congestive heart failure.
CrCr helps to regulate the blood sugar level and to control the diabetes type II.Excess of Cr in body can lead to issues with digestion and low sugar in blood. Moreover, excessive Cr can damage liver, kidney and various nerve tissues.

3.1.1. Stainless steel implants

Most used stainless-steel implant is the surgical grade 316LSS. 316LSS offers a variety of mechanical, physical, and chemical properties. The materials for this implant grade are easily available with lower costs. Stainless steel contains various alloying elements such as Mo, Cr, Mn, Cu, Ni, Si and C. Stainless steel is highly corrosion resistant due to formation of chromium oxide layer preventing oxidation. Stainless steel has ability leach metallic ions such as Cr, Ni and Fe that are harmful for body and can cause serious health issues [12]. Different surface modification techniques are utilized to overcome the negative effects associated with stainless steel. Coatings of ceramic, polymer and composites are applied to metallic implants to prevent metallic leaching. FDA-approved polymer PCL due to biocompatibility and high mechanical strength. As PCL is hydrophobic, a combination of the hydrophilic polymer is used for the coating to enhance the prefoliation and cell adhesion [13]. In 316LSS, Ni can be added to further enhance the corrosion resistance. The presence of carbon in stainless steel causes Cr to oxidize at grain boundaries leading to localized pitting or corrosion. Stainless steel is extensively used in bio-medical implants and trauma procedures. 316LSS is used in knee, hip and ankle replacement implants. Stainless implants have better mechanical and tribological properties and are bio-compatible with biological environment. Due to the metallic leaching issues for stainless steel, the implants are mostly used short term purposes [14]. Loading conditions for implants can be dynamic and static. Dynamic loading conditions involve fatigue strength of the implant and most implants fail because of fatigue failure. The biological environment in the body leads to pitting and corrosion that severely affects the fatigue strength of an implant. Corrosion is undesirable as it reduces the fatigue strength of the implant [15].

3.1.2. Mg based alloy implants

Mg is one of the essential minerals for the human body. Conventional implants require a secondary procedure for the removal that is quite painful and can cause post-procedure complications. To solve these issues, bio-degradable Mg-based alloy implants are used. Mg-based implants can degrade over time without any harmful by-products. Moreover, Mg-based alloys have properties such as higher strength-to-weight ratio, good machineability and castability. Mg-based implants have higher yield strength with a low Young's modulus that is close to the natural bone [16]. The difference in Young's modulus between implant material and natural bone makes the implant material carry more load leading to stress shielding. These stress shielding can lead to further complications such as inflammation, implant loosening and skeletal thickening. The mechanical properties of Mg-based implants prevent them from shielding stress in bone implants. Mg aids in different metabolic and biological reactions in the body. Mg maintains the structural integrity for proper bone healing. Iron in pure form shows biodegradability behaviour but corrosion issues make it inefficient and ineffective. Even with the extensive use of Mg-based implants, some fundamental issues are still unaddressed. These issues include a fast degradation rate and loss of mechanical properties at certain pH levels [17].

The performance, corrosion behaviour, and properties of implants are defendant on the microstructure of implants. The microstructure is controlled through processing methods, alloying elements, design selection, impurities and heat treatments. To overcome these complications and shortcomings, Mg is used with different alloys such as Zn, Sr, Ca, and Sn. Similarly, Mg-based glasses and hybrid implants with ceramic and polymer coatings are used for tissue engineering and drug delivery applications [18]. Shortcomings associated with Mg-based implants are mostly related to wear properties. Moreover, the degradation of Mg is required to be optimized to enhance the functionality of Mg-based bone implants. The degradation of Mg can lead to the formation of cavities causing inflammatory response and infection. The high degradation rate of Mg releases more gas and degradation debris. Improved resistance to degradation can overcome cavity formation and release of gases. The addition of organic coating to Mg-based implant can further optimize the rate of degradation [19].

3.1.3. Titanium alloy implants

Titanium-based alloys are selected for implants due to low density, high strength, inertness, high corrosion resistance, enhanced biocompatibility, and low modulus. Ti-based alloys were developed for the aviation industry due to their excellent corrosion resistance. High biocompatibility with the biological environment made it suitable for implants. The modulus of Ti is much lower than that of steel and Co-based implants, making it a more suitable candidate for implants. Apart from implants, Ti is also used in other medical applications like wheelchairs, artificial legs and limbs. The strength of Ti-based alloys is near to that of stainless-steel alloys, but the density is significantly less. Ti-based alloys are used in dental implants, knee replacement, hip replacement, bone screws and rods, heart valves, pacemakers, and surgical instruments. Ti-based alloys have shortcomings such as lower shear strength making them less desirable for certain medical applications. Moreover, the high coefficient of friction for Ti-based alloys can release the wear debris and particles leading inflammatory response from the body causing pain and discomfort [9, 20].

3.1.4. Co alloy implants

Co-based alloys were employed in the manufacturing of various aeroplane components for the first time. Mo, C, Ni, W and Cr are used as alloying elements with Co base material. Co-based alloys have exceptional mechanical properties, higher wear, and corrosion resistance at high temperatures. In the early 1940s, Co-based alloys were used for the first time for dental implants. Now Co-based alloys have been used in several different medical applications such as knee and orthopaedic implants. Co-based alloys have higher values for modulus than natural bone which can lead to stress shielding causing bone resorption. The function of different alloying elements is shown in table 3 [7] and toxicity profile of materials are shown in table 4.

Table 3. Functions of various alloying elements.

Alloying elementFunction
MoIncrease corrosion resistance and strength
CIncrease casting ability and wear resistance
NiIncrease strength and corrosion resistance
WIncrease the overall strength but decrease corrosion resistance
CrIncrease corrosion and wear resistance

Table 4. Toxicity profile for metal-based implants.

Implant metalIndicators for toxicityToxic concentrationImplant areaReferences
CoChange in thyroid functions, allergic dermatitis, cardiomyopathy, polycythemia>5.0 μg l−1 Dental framework, shoulder, stents, knee and hip implants[20]
CrLiver and kidney complications, respiratory issues, ulceration, dermatitis>15 μg l−1 Dental implants, knee, shoulder, hip and spine implants[21, 22]
MoChange in testosterone level, acute psychosis, neuro complications, seizures, increase in uric acid>3 μg l−1 Dental implants, knee, shoulder, hip and spine implants[22, 23]
NiHypersensitivity, vertigo, headache, changes in vision, acute sensation, vomiting>10 μg l−1 Plates, rods, pins, or screws for bone fixation, and dental implants[22, 24]
TiYellowish nail syndrome>10 μg l−1 Dental implants, knee, shoulder, hip and spine implants[24, 25]
FeCardiomyopathy, gastrointestinal symptoms, arthritis, cirrhosis,>5 mg L−1 Stents and Ortho fixation implants[22, 24, 26]
MnPsychosis, Parkinson symptoms, headache>5 μg L−1 Ortho fixation implants[24, 27, 28]
AlTremors, Alzheimer, hepatic complications, anemia, dyspraxia, osteomalacia, dysarthria>30 μg L−1 Ortho and dental fixation implants[22, 24, 29]
CuHemolysis, Renal and hepatic complications, cardiac issues, gastrointestinal symptoms>0.14 mg dL−1 Bone fixation healing by protection against bacteria[22, 24]

3.1.5. Biodegradable alloys

Ti, Co, and stainless steel-based metallic implants are long-lasting and inert. The healing process can take up to 6 months and these implants must be removed after healing to avoid complications and various side effects. Long-term complications of metallic implants include inflammation, infection, metallic ions leaching and stress shielding. Bioresorbable and degradable materials dissolve in the body after the healing process avoiding the complication of metallic implants. Moreover, metallic implant can cause stress shielding effect that makes bone weaker than intended. Mg-based alloys are highly degradable in the biological environment. These degradable alloys can form air pockets in tissues, rapid pitting, and premature degradation leading to the failure of the implant. Addition of alloying elements such as calcium, aluminium, lithium and different rare earth materials aid in controlling the degradation rate of Mg and improving the mechanical properties. In addition, coating, grain refinement and fast cooling techniques are applied to Mg-based alloys to improve mechanical strength and degradation rates [21]. Mg reacts with water in the body which leads to corrosion and pitting causing premature failure of implants. Mg2+, H2 and OH are produced because of corrosion. OH is soluble and increases the pH level of the body. Mg-based implants are highly biocompatible with the biological environment. Mg ions can balance the increase in pH level caused by OH ions. Both these ions combine to form a highly biocompatible implant and tissue interface. H2 is a highly diffusible gas but excessive production can lead to certain complications such as pressure build-up and can affect bone growth. The degradation mechanism works usually through a corrosion path. Over time, Mg-based implant corrodes leaving the vacant space for the growth of bone and soft tissues. Apart from bone implants, Mg-based alloys are also used in degradable stents that have a life of up to 1 year. These stents are helpful to avoid long-term side effects [22]. The basic properties and mechanical characteristics of degradable implants are explained in table 5.

Table 5. Degradation behavior, mechanical properties, and biological response for different materials.

MaterialDegradation characteristicsMechanical properties and shortcomingsBiological responseReferences
IronInert/very slowIrregular corrosionInflammatory response and release of wear debris[24, 25]
MgHigh degradation rate that can fail implant before completion of healing process.Sufficient strength of implant, Irregular implant pitting, rapid corrosion.Release of gas can exert pressure on surrounding tissues.[26, 27]
TiInertCan bear high loads and have longer life cycle.Ti alloys are bioactive.[21, 28]
Zinc based AlloysSlow degradation that can exceed the time for healing.Optimal strengthNon inflammatory response[29]
PolymersControllable degradationCannot be used in load bearing application, high flexibility, Implants can swell due to hydrophilicity.Inflammation and production of hydrolysis by products[30]

3.1.6. Polymer based implants

The need for bio medical implants and devices has been increasing day by day. To meet the ever-increasing demands, various materials including polymers are being studied for conclusive outcomes. Metallic implants have several disadvantages and require several revision surgeries to avoid the adverse effects of the implants. Moreover, metallic implants can cause allergic reactions, stress shielding, implant loosening, release of toxic ions, reduced strength and corrosion [23]. To overcome these issues, different natural and synthetic polymer materials are used in implants and drug delivery systems. Natural polymer-based implants have additional benefits such as degradable properties. Polymers are used for bone and dental implants, rods, pacemaker, tissue engineering, stents, and plates. Polymeric materials exhibit properties such as high strength, inertness, good compatibility, high chemical resistance, low density, good elasticity, high thermal stability, high durability, and toughness. A combination of two or more polymers can greatly enhance the properties and overcome the shortcomings of polymer implants [24]. Different polymers are used in the coating of metallic implants to improve the interface properties, and compatibility with the biological environment and to overcome the shortcomings of implant materials. Modern fabrication techniques and 3d printing have made polymeric implants cost effective and more biocompatible. These techniques can be used to produce versatile implants for different applications ranging from drug delivery to tissue repairs. New approaches and innovative polymeric implants are being developed for bone repairs and tissue engineering [25]. Polymer materials that are used in bio-medical applications are cellulose [26], collagen [27], deoxyribonucleic acid [28], polyvinyl chloride (PVC) [24], polypropylene (PP) [29] and polymethyl methacrylate (PMMA) [30]. The properties of polymers such as polylactic acid (PLA) can be tailored for biodegradation with controlled degradation [31].

3.1.7. Ceramic-based implants

Ceramic-based implants are used in a variety of applications in orthopaedic and dental implants. The ceramic implants interact with the biological environment in different ways such as physical attachment with tissues, implant fixation and growth of bone tissue into the porous ceramic implant. Calcium phosphate or hydroxyapatite (HA) is one of the most used bioactive ceramic materials for implants. Calcium-based ceramic materials are further classified into different categories based on the ratio between calcium and phosphate. The ratio is an important factor for hydroxyapatite to make it bioactive for implants. Hydroxyapatite can be used in different non-load bearing implants and are degradable properties. When degradable ceramic implants like calcium-phosphate are exposed to water and a certain temperature, it leads to the formation of HA phase [32]. These degradable implants are temperature sensitive causing issues in the fabrication of such implants. Furthermore, HA is also coated on various metallic implants through plasma spraying and laser coating techniques. These coatings make the underlying material more biocompatible, improve interfacial properties, and prevent degradation. HA is employed as bone cement for artificial joints. Similarly different other ceramic materials such as SiO2, P2O5, CaO and Na2O are used for similar applications [33].

Alumina (Al2O3) is commonly used for load-bearing applications such as orthopaedic and dental implants. Al2O3 is bioinert and biocompatible with high wear and corrosion resistance. Due to excellent biocompatibility, Al2O3-based implants do not require cementing materials for fixation. The fine grain size of Al2O3 further improves fatigue resistance and strength for long-term use [34]. Even with all these properties, Al2O3 implants have a high elastic modulus typically > 300 GPa contributing to the shielding effect and making the implant ineffective for long-term healing. With a low range of elastic modulus, zirconia addresses the shortcomings of alumina. Ceramic materials are extensively employed in femoral heads in hip implants [35]. Alumina and zirconia high great wear resistance but low fracture toughness in comparison with metallic implants leads to fractures of the femoral head and other medical complications. The use of ceramic femoral head and metallic stem create interfacial interaction that adds to more complications. These complications can be minimized with proper interfacial and design optimization [36].

3.1.8. Shape memory alloys

Ti-based shape memory alloys such as TiNi, TiZr, TiNiAg and TiNbSn are used in medical implants due to some unique mechanical properties. Shape memory alloys are super elastic with shape memory effects along with high recovery strain, high strength, and low modulus of elasticity [2]. The properties of shape memory alloys are based on reverse martensitic transformation [37]. Due to these outstanding properties, shape memory alloys are employed in multifunctional applications. TiNi alloys are highly bio-compatible and show much better corrosion resistance than Co-based alloys. Thierry et al [38] compared the functionality of TiNi and stainless-steel stents. TiNi stent had few thrombi in comparison with stainless steel stent. TiNi exhibits biocompatible behaviour in ex vitro and in vivo studies. Low stiffness and high strength of shape memory alloys make them suitable for bone implants, joint replacement, stents and tissue engineering. The shape memory effect enables the growth of the tissue along or into the porous implant. TiNi based shape memory alloys are better than other similar alloys due to better low stiffness and low strains. Super elastic and shape memory properties are dependent on the chemical composition of TiNi. Even a slight change in composition can lead to the formation of different phases with different final properties and transformation hysteresis behaviour. Fu et al [39] analyzed the preparation of TiNi shape memory alloy films through co-sputtering with composition Ti50Ni50. At this composition, the film had very low residual stresses at room temperature. By decreasing or increasing at% of Ti to 47% or 53% respectively, the residual stresses increased significantly due to an increase in intrinsic stresses. Different research studies employed various alloying elements such Nb, Ag, Fe, Cu and Zr to control transformation hysteresis properties and behavior of shape memory alloys. Addition of Nb to TiNi alloys widens the transformation hysteresis for enhanced shape memory effects and improve elastic properties. The addition of Ag leads to antibacterial properties. Ni in TiNi alloy is responsible for the release of ions leading to corrosion of the implant. Moreover, these ions can trigger an allergic response and lead to severe medical complications. Surface treatments and coatings can cause a significant decrease in the formation of these ions [40].

4. FDA approval process for implants

FDA approval of any medical device or implant takes around 12 years till final approval. The process involves designing the implant from concept design to final approval. The approval time has significantly increased from 8 years in the 1960s to 12 years. The approval process is categorized into different stages as under.

  • –  
    Development and research phase
  • –  
    Clinical research
  • –  
    Applications
  • –  
    FDA post-marketing review

Development and research phases start from concept design to final design that can take around 1 to 3 years. Clinical research includes phases I, II and III. Clinical research can take up to 5 to 10 years. During clinical research, medical devices and implants are tested phase-wise by increasing risk factors for the patient. Phase I clinical trials include healthy subjects to determine the safety profile and any possible side effects. Once, phase I trials are completed successfully, the trail is moved to phase II. Phase II clinical trials are based on the effectiveness and efficacy of implant or medical device. Phase II study aims to collect data on the effectiveness of the implant and medical device. The safety of implants and short-term side effects are studied. After completion of phase II trials, the new medical device or implant is tested against a similar implant available on the market. After passing all three phases, the implant is deemed fit for use. Long terms effects and post-marketing effects of the implant or medical device are studied in the phase IV phase. Medical implants and stents are classified into classes II and III of the FDA list [41].

5. Cardiovascular stents

5.1. Drug eluting stents

Drug-eluting stents are revolutionary development to overcome issues related to heart diseases. Heart diseases involve blockage or loss of functioning of valves and main arteries. Coronary artery disease is a result of these complications. The build-up of plaque or cholesterol or fats causes complete or partial artery blockage. Stents are primarily used to widen and open these narrow arteries [42]. Stents are implanted to strengthen the weak arteries to rule out the collapse of artery walls. Early stents were metallic and were made up of stainless steel, Ni-Ti and CrCo. But over time these stents have the probability to fail due to blockage and thrombosis. It is estimated that around 15% of stents lose considerable functionality due to in-stent restenosis. To avoid these issues, drug-eluted stents are used. These stents slowly release the drug to overcome the issues of inflammation, thrombosis and immune system response [43]. Moreover, these drugs induce a bio-active response for cell prefoliation and biocompatibility. Early drug-eluting stents are made up of three parts; a metallic stent, a polymeric coating for drug loading, and active drug agents to induce a favourable response. These early drug eluting stents have short comings related to the response of body with additional polymeric coating, late-stage thrombosis, stent displacement, issues with constant and required drug release. Further advancements have led to the development drug-eluting stents with controlled release of anti-proliferative active drug agents. These drug agents are slowly released from the polymeric coating. The polymeric coating consists of three parts. The first layer is parylene layer that is used for joining other subsequent layers with the metallic stent. The second layer contains the drug agents for the controlled release [43]. The last layer controls the release mechanism of drug agents. The most important factor to consider is the release of a timely and enough amount of drug. Due to trigger response and inflammation with the use of polymeric layers, polymer-free drug-carrying stents are made. Drug carriers and molecules are attached to stent material with a mechanism to control the release rate of the drug. Techniques and methods used in the manufacturing of polymer-free stents are shown in figure 2 [44].

Figure 2. Refer to the following caption and surrounding text.

Figure 2. Different manufacturing techniques that are being used in the manufacturing of polymer free stents.

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5.2. Biodegradable polymeric stents

Biodegradable polymeric stents have gained much attention due to their degradation behaviour and enhanced biocompatibility. Polylactic acid is one of the most used polymers for stents. Polymers have lower mechanical properties than metallic stents which is why polymer stents have the greater volume to compensate for these shortcomings. Apart from polymers, Mg based metallic stents are biodegradable too. Drug eluted biodegradable polymeric stents have several advantages over metallic stents. Much of studies on polymeric stents focus the response of body and probable complications. Non-degradable materials in stents can cause severe complication in later stages. These complications such as blood-clotting and inflammation are main causes for stent failure. The long-term use of biomaterials requires biocompatibility as well as the detailed understanding of the host response to foreign objects [45]. The use of biodegradable coatings on metallic are proven to be fruitful in controlled release of drug agents and avoiding serious complications. Some of the issues that these polymer coated metallic stents are the unusual high stresses on polymeric layer due to expansion of metallic frame. Biodegradation is a surface and bulk erosion process. The examples of biodegradable polymers are poly carboxyphenoxy hexane-sebacic acid [46], poly fumaric acid-sebacic acid [47], poly imide-sebacic acid [48], poly ortho ester [49], Poly lactic acid [50], PGA [51], PLGA [52], poly oxaamides [48], poly(p-dioxanone) [53], and poly trimethylene carbonate [54].

5.3. Design parameters for stents

The factors that are needed to be considered while designing the stents are; the dimensions of a stent, expansion mechanism, strength, condition of blockage or arty walls, compression and expansion tolerance and flexibility of a stent [55]. The stent material must be highly compatible with the biological environment, bioactive, corrosion resistant, and of higher strength. Ti-based stents are more commonly used. Modified surfaces of Ti stents with nanostructures help in biocompatibility to reduce the blood clotting around the stent. For drug-eluting stents, drug agents and polymeric or metallic coating should be compatible with metallic frames [56]. In the case of the biodegradable stent, the studies should be carried out to know the potential side effects and to assess the safety. Maintaining structural integrity and strength is an important factor to consider in the case of polymeric stents. A comprehensive detailed analysis should be carried out for every component of a stent with host tissues to determine the efficiency and performance of any stent [57]. For degradable stents, controlling the degradation rate for the required period is a crucial factor. Extensive clinical trials and in vitro studies must be performed to check the efficacy and safety of stents. Moreover, after the implantation follow ups are required for patients to ensure the functioning of stent [58].

6. Tissue engineering

Tissue engineering has progressed significantly with the development of new materials and techniques. Tissue engineering can maintain, restore and improve various damaged tissues and organs. Substitute tissues and reconstruction techniques are applied for such tissues. Tissue engineering techniques involve the implantation of tissues into the body, tissue growth with growth substances and the use of an external scaffold for tissue growth [59]. Scaffolds represent a major portion of tissue engineering and recent research studies have shown increased interest in designing and fabricating scaffolds. A Scaffold is a three-dimensional porous material used to enhance cell interaction and adhesion, differentiation and prefoliation through sufficient transport of nutrients for cell growth and controlled biodegradation for cell growth without or with minimum side effects. Important properties of polymeric scaffolds are microstructure, size and shapes of pores, degradation rate and strength [60]. Biological and natural scaffold materials are derived from animals and humans while synthetic scaffolds are of polymeric materials. Scaffold material can be non-degradable or degradable. Materials such as bioactive ceramic materials, polymers, glasses and silicates are used as scaffolds in tissue engineering. But ceramics, silicates and glass/ceramic materials have insufficient biodegradability and biocompatibility that limit the use of these materials in tissue engineering. The stages of biodegradability in polymeric scaffolds are explained in table 6 [61].

Table 6. Stages for biocompatibility in polymeric scaffolds.

StagesRemarks
HydrolysisJust after implantation, biodegradable material starts to absorb water from the surrounding tissues.
DepolymerizationHydrolysis results in loss of molecular mass.
Loss of massPolymer chains start to break. Molecular mass and strength are lost.
Monomer dissolutionSmall particles assimilate and soluble monomer ions are produced.
ResorptionBy products produced during degradation are expelled from body through lungs and kidneys.

6.1. Natural and synthetic scaffold polymers

The issues of biocompatibility and biodegradability in various materials can be overcome by synthetic and natural polymeric materials for scaffolds. The properties of polymeric material are dependent on the arrangement of macromolecules, structure and composition. Polymeric scaffolds are divided into natural, synthetic and non-degradable synthetic polymers. Natural polymers are fully biodegradable and are bioactive with enhanced interaction with the biological environment [62]. Examples of natural polymeric scaffolds are gelatin [63], collagen [64], silk [65], keratin [66], myosin [67], actin [68], polysaccharides [69], and fibrinogen [70]. The properties of synthetic polymeric scaffolds are dependent on mechanical characteristics, porosity, and degradation time. These polymers are easy to produce with less cost than natural polymers. Synthetic polymers have tunable and controllable mechanical and physical properties. Most used synthetic materials are PLA, PLGA and PGA [71].

6.2. Types of scaffolds

6.2.1. Porous scaffold

Porous scaffolds contain homogenous-sized pores that are interconnected to form a network. This network of pores is very useful in tissue engineering and is employed in various applications such as the growth of tissues and bone. Porous geometry allows the cells to interact with the outer biological environment. Sponges and foam structured scaffolds are considered more stable than meshed structures [72]. But still the applications for sponges and foam structures are somehow limited due to the presence of open spaces. Some of the advantages of foam-based scaffolds are the effective proliferation of cells for tissue growth, avoiding a large cluster of cells due to pore size, providing extra surface area for extracellular matrix and transport network for nutrient and other important minerals to growth sites. The synthesis method for foam dictates the orientation of pores [73]. Interconnected pores are required for the development of blood vessels and nerve growth. The latest methods involve the three-dimensional fabrication of scaffolds by joining membranes to form the desired shapes. Pore size is more dependent on the size of cells and tissues. Properties such as pore size, surface area, surface/volume ratio, porosity and crystallinity are controllable through fabrication techniques. These porous scaffolds are very effective in the controlled release of drugs. Materials used for porous scaffold materials are PLA [74], PLLA [75], PDLLA [76], PCL [74], PLGA [77], PGA [78] and PBT [79]. Effective fabrication methods to control the pore size and porosity are particulate leaching method, solvent casting, and electrospinning [60].

6.2.2. Hydrogel-based scaffold

Hydrogel-based scaffolds are developed for enhanced growth of new cells and tissues. Latest applications and development in designing for degradable scaffolds have increased the importance of hydrogel-based scaffolds in bio and tissue engineering. Hydrogels are made from natural macromolecules that are suitable for drug delivery and tissue engineering. Natural macromolecules in hydrogels contribute to enhanced biocompatibility, controlled degradation, and cellular interactions. Hydrogels have limitations in mechanical properties in comparison with synthetic scaffolds but the structure of hydrogels is very similar to the components in body [60]. Gels are generally formed with the covalent interactions in the structure. Hydrogels made from synthetic materials have better control over pore size, structure, functions, and other mechanical properties. Hydrogels meet different criteria for the tissue growth that includes degradation rate, mechanics, biocompatibility, and cell adhesion [80]. The degradation rates must be in line with different cellular and biological processes of the body. The optimization of degradation rates enhances the cell growth. Degradation rates for these hydrogels are required to be tunable to meet different requirements. The degradation rate is controlled through the chemistry and structures of hydrogels. Hydrogels are employed for drug delivery, wound dressing and in bone regeneration. Hydrogels promote the diffusion of nutrients, cell migration, and angiogenesis. Natural hydrogel-based polymers are collagen, fibrin, HA, chitosan, gelatin and alginate while synthetic polymers used for hydrogels are PLA, derivates of PEG and PVA [81].

6.2.3. Fibrous scaffold

Scaffolds with nano-fibres can mimic the architecture and structure of biological cells and tissues at the nanoscale. Phase separation, electrospinning and self-assembly techniques are employed for the fabrication of fibrous scaffolds. Electrospinning is known to give the best results for fibrous scaffolds in tissue engineering. Only a handful of studies have explored the self-assembly and phase separation techniques for the fabrication of fibrous scaffolds. Fibrous scaffolds have a higher surface area, surface/volume ratio and interconnected micropores that favour cell adhesion, migration of nutrients and prefoliation. Fibrous scaffolds are used in several biomedical applications such as tissue engineering, bone regeneration, neural engineering, ligament growth and drug delivery. Natural polymer-based scaffold materials use collagen, chitosan and gelatin while other synthetic materials include PCL, PLA, PLGA and various other polymers. Polymer materials used for the fabrication are required to be modified through functionalization for some specific applications. Surface grafting, blending and coating are some of the commonly employed techniques for the functionalization of these polymers [82].

6.2.4. Polymeric-ceramic based scaffold

The inclusion of composite materials has made it possible to fine-tune the mechanical, physiological, and chemical properties of scaffolds through the control of morphology, the orientation of the reinforcing phase, and volume fractions. Ceramic materials used in the fabrication of polymer-ceramic composites are bio-active, biocompatible, inert, and degradable. Examples of inert ceramic materials include zirconia, alumina, and silicone nitrides. Ceramic materials have high resistance to corrosion and compression. Degradable ceramics are coralline, aluminium-calcium-phosphate and plaster of Paris. Some of the drawbacks of ceramic materials are higher brittleness, hard to fabrication, lower fracture toughness and high density. Polymer materials have properties such as flexibility but lack some other important mechanical characteristics such as stiffness and strength. Similarly, ceramic materials have limitations related to brittleness and fracture strength. Fabrication of polymer-ceramic-based scaffold can overcome the individual limitations for each material. The combination of polymer and ceramic can significantly improve the degradation properties of scaffolds [83]. The complications that arise during the development of polymer-ceramic scaffolds are maintaining interfacial stability and strength upon controlled degradation along with matching the degradation rate with that of the development of cells and tissues. Porous polymeric-ceramic-based scaffolds are used as substrate materials for bone regeneration and osteoconductive behaviour. Ceramic materials such as TCP, HAP and CP are used with a combination of PLA, PLLA, gelatin and collagen polymers to form a polymer-ceramic scaffold [84].

7. Use of nanotechnology in implants and bioengineering

Surface properties and chemistry play an important role in determining the compatibility of any implant material. Nanotechnology provides the opportunity to increase the surface area of implant material and to fine-tune the surface properties. The desired response from the host tissue and immune system is achievable by incorporating bioactive nanomaterials with bio-specific molecules. The interactions between the implant and host tissues occur at the nanoscale. The application of nanoscience and materials is determined by surface roughness, topography, and surface thermodynamics. Nanoscience and technology are used in different biomedical applications such as drug delivery and surface modifications. Nanoparticles are attached with bio-specific molecules to induce the required response upon interaction with host tissues. The very high surface area of nanoparticles provides more binding sites for the host tissues to attach. Metallic oxides such as iron nanoparticles are employed for diagnostic applications. These particles are used in MRI, ultrasonic techniques, and magnetic imaging [85]. Similarly, TiO2 nanoparticles are used in bone regeneration. ZrO2 nanoparticles are useful in dental applications and gold nanoparticles are used to check blood flow. Silver nanoparticles have antibacterial properties and are successfully employed in bio-sensing applications. Carbon nanotubes (CNTs) are used in drug delivery and cancer treatment. CNTs filled with drug agents are used to target the specific area in the body. CNTs are biocompatible with a minimum level of toxicity. The toxicity of nanomaterials is dependent on composition, dose, crystal structure, size range and particle shape. Interaction of nanomaterials with the biological environment can tell a detailed analysis of the toxicity of nanomaterials. Even after all the considerations, it is a very difficult task to completely analyze the toxicity effects of nanomaterials [3].

8. Surface modifications

Surface modifications are done to implant materials to improve the shortcomings and enhance the mechanical properties and compatibility of the implant with the biological environment. Various long-term complications are solved through surface modification and coating the implant with bio-active materials. Ti-based implants are biocompatible but still, the composition of the Ti implants is much different from the natural bone tissues. This can lead to osteointegration that is without any chemical bonding between Ti and natural tissues. Moreover, sub-surface layers of implant materials are affected by the low adhesion issues with oxide layers, work hardening and by low resistance to plastic shear. The surfaces of implants undergo abrasion wear and adhesion issue to the relative movement of bone and implant. This leads to the loosening and failure of the implant. The passive film formed on the top of metallic implants increases the corrosion resistance. But the complex environment of the body can dissolve or delaminate this oxide layer. The dissolution of this layer can lead to the production of toxic byproducts adding to medical complexities. To overcome these complications and shortcomings, various modification methods are used [86].

8.1. Chemical modifications

Chemical modifications and reactions at the surface overcome the shortcoming of the implant materials and increase the adhesion of biological cells on the surface by providing more binding sites for easy prefoliation. Self-assembled molecules (SAMs) are commonly used to chemically modify the surface with the addition of thiol groups. SAMs form a bio-active and closely packed coating over a specific substrate surface. Silanes follow the same path as silicon-based substrates. Different functional groups can be incorporated just by changing the terminal groups in SAMs [87]. Different functional groups such as carboxyl, amino and hydroxyl are used to chemically modify the surface. All these functional groups are available from biological sources. In the latest studies, SAMs are extended by adding a second layer of molecules with different functions and properties. Implementation of more complex surface modification techniques is easing cell adhesion and prefoliation through different amino acids [88]. Hyaluronic acid (HA) is a linear polysaccharide. HA is employed in the lubrication of joints, tissue regeneration, and treatment of cancer and tumors. HA binds to the CD44 gene that acts as a docking or active site. HA binded with CD44 is used for imaging and drug delivery applications. HA is chemically modified by using the CD44 receptor. The binding of HA to CD44 can be controlled or altered for intended applications [89]. Another study modified HA using a multi-step approach for sulfation HA synthesis to change the binding specification of CD44 for HA. Modified HA targeted P-selectin along with CD44 to utilize these receptors for targeting cancer cells and tumors. Sulfated HA is used for duel targeting of p-selectin and CD44 with improved therapeutic effects in comparison with single targeting [90]. Another research study focused to find non-reabsorbable, biocompatible and low-cost dermal fillers. HA and polyacrylamide fillers are known to provide cost-effective solutions for cosmetic and plastic surgery. Even after eight months of using these fillers, the fillers did not undergo cell migration and cause any infection or necrosis at implant site. These filler materials are bio-active in nature to integrate with surrounding tissues. Polyacrylamide filler exhibited inflammatory reaction and biodegradation over longer period, but HA showed no such biodegradation and exhibited reduced inflammatory reaction. Further research is required to understand the efficacy and safety of these fillers [91].

8.2. Topographic modifications and surface coating

Topographic modifications and surface coating are meant to control the interaction of the biological environment and cell adhesion to the implant or any medical device. Topographic features play an important role in the determination of the biocompatibility of medical devices [86]. The surface of any implant may look flat with the naked eye but contains topographic features at a microscopic level. These microscopic features increase the roughness of the surface for better cell adhesion and improvement in various surface-based properties. The surfaces of implant materials can be altered through surface engineering to improve cell adhesion and prefoliation. Surfaces of implant materials can be modified according to the need. Some of the surface modification methods are the application of self-assembly coatings, functional groups and surface additives for binding sites [92]. The long-term performance of implants is dependent on implant/tissue interfacial properties. Similarly, mechanical shortcomings such as low corrosion and wear resistance affect the life of any implant. Issues of tribological properties can easily be addressed through the application of suitable coating. Surface properties are important for the successful integration of implant with natural tissues. Some of the surface modification techniques are sputtering, ion implantation, plasma spray coating, nitriding, boriding and carburizing [93]. These surface modifications improve the biocompatibility and shortcomings of implants. Nitrogen ions are implanted on metallic implants such as Ti through ion-implantation. Ions implantation improves the wear and corrosion resistance of Ti by many folds. Similarly, diamond-like carbon coatings is applied to attach biological molecules to the surface of the implant. Living cells attach and grow on these coating without any complications and immune response for biocompatibility [94].

8.3. Cell adhesion and functionalization

Cells are complex in functions. These cells interact with the surrounding environment through the number of receptors at the outer wall. The receptor response is created through the production of antigens This response triggers the chain of reactions in a cell. Specific receptors are developed for the biological environment. Integrins are a class of cell receptors that are found in adhesive proteins such as vitronectin, laminin and fibronectin. Integrins are formed due to the non-covalent bonding of alpha and beta subunits [95]. The radius of these receptors is between 4 to 6 nm. The attachment or detachment mechanism of these cells involves control by receptors and a feedback loop. At the end of adhesion protein, hair-like structures acts as a binding site for cell adhesion. Feedback allows more receptors to move towards the binding sites. Similarly, a strong propulsive force induces stress for the movement of the cells along the surface. Cells spread and perfoliate along the surface in this way. No adhesions of cells are not desirable and strong attachment of cells effectively renders the movement and prefoliation. The addition of a natural protein layer to the surface increases the surface reactivity and adhesion to the cells. The response of a body to any implant material is dependent on the protein interactions [96]. Cellular responses such as migration, proliferation and cell adhesion are affected by microstructural properties, interfaces, surface topology and surface chemistry of a biomaterial. Rough nano topology gives rise to osteogenic maturation. Similarly, an increased implant-to-bone ratio enhances the strength and efficiency of bonding energy and cell proliferation. Cell adhesion and proliferation are dependent on multiple processes and mechanisms. The interaction between tissue and implant, cell surface receptors, functional molecules deposited on the implants as well as interaction forces such as electrostatic and van der waal. Tissues around the implant react to different surface features. Metabolic and phenotypic activities influence the strength and healing process of the implant [97].

9. Conclusion and future directions

The field of medical science is progressing rapidly to incorporate ever-growing demands for biomedical implants to treat injuries and traumas with the development of new materials. With the introduction of three-dimensional fabrication techniques, various complex geometry implants can be fabricated. Various alloying elements are being studied for metallic implants for controlled degradation but toxicity analysis and biocompatibility studies for these alloying elements are necessary. The development of the latest surface modification techniques and coating improve the interfacial properties for efficient integration of implants in a biological environment. The chemistry of coating material, corrosion analysis, and fabrication techniques and parameters are needed to be well understood for efficient and accurate outcomes. The development of the latest composite materials will enhance the mechanical performance, and provide flexibility and new functionality to implant materials. The issue of non-degradable implants such as stress shielding, toxicity and harmful byproducts are being addressed through degradable implants. These degradable implants support the growth of tissues and self-sustainability. The focus for future biodegradable materials is to work on strategies for controlling impurity level, coatings and alloying elements, functionalization of implant materials, optimizing biodegradation rate at implant/tissue interface and developing new degradable materials. Designs for polymeric scaffolds are being optimized for the timely and desirable response in tissue engineering. Demand for replacements of tissues and organs will drive up the advance and efforts in tissue engineering. Polymeric scaffold materials can control various physical and chemical properties in tissue engineering. Future research studies and development are required to overcome the shortcoming of porosity, bioactivity, and mechanical properties. A wide range of materials and fabrication techniques are needed to be studied to counter these limitations.

Data availability statement

All data that support the findings of this study are included within the article (and any supplementary files).

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