Exploring the future of metallic implants: a review of biodegradable and non-biodegradable solutions
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Yahya Ahmed
and
Wasif Farooq
Abstract
As advancements in medical technology continue to evolve, the demand for innovative implant materials has become increasingly vital for enhancing overall experience of patients. Traditional non-biodegradable implants, while effective, often necessitate removal through invasive and costly surgical interventions, leading to significant clinical challenges. To address these issues, the development of biodegradable materials has gained prominence due to their ability to gradually degrade and be absorbed by the body, presenting a compelling alternative to permanent implants. This review examines both biodegradable and non-biodegradable metallic implants, focusing on key aspects such as biocompatibility, mechanical properties, and degradation kinetics. Furthermore, it explores the applications of these materials across various medical fields, emphasizing their potential to improve patient care. This review aims to bridge the gap between laboratory innovations, clinical practices, and industrial applications by summarizing current research. It offers valuable insights for researchers, clinicians, and industry professionals, contributing to the ongoing dialogue regarding the future of implant technology and advancing the understanding of material selection for diverse medical applications.
1 Introduction
Implants have revolutionized healthcare by providing tailored solutions for various medical conditions (Nouri and Wen 2015). In orthopedics, implants such as joint replacements and bone fixation devices restore mobility and stability in patients with fractures or musculoskeletal disorders. Figure 1 illustrates the various parts of the body that commonly utilize biomaterials. Cardiovascular implants like stents, pacemakers, and artificial heart valves are critical in managing heart diseases, improving blood flow, and regulating electrical activity. Dental implants offer long-term tooth replacement solutions, enhancing function and aesthetics. Beyond structural support, implants are used in drug delivery systems, enabling localized, controlled medication release, improving therapeutic outcomes while minimizing side effects. Additionally, implants play a vital role in reconstructive surgeries, helping repair tissues or replace organs, thus enhancing survival and quality of life (Nouri and Wen 2015).
Utilization of biomaterials in various anatomical regions.
The evolution of biomaterials has further expanded the application of implants by enhancing biocompatibility, reducing rejection risks, and improving its durability (Awasthi et al. 2021; Guglielmotti et al. 2019; Insuasti-Cruz et al. 2022). In regenerative medicine, implants stimulate tissue growth and regeneration. Biodegradable implants are advantageous for temporary applications like bone fixation and tissue scaffolds, as they degrade naturally, eliminating the need for removal surgery. Non-biodegradable implants, on the other hand, offer long-term solutions for permanent conditions, providing durability and mechanical strength but may require revision surgeries due to risks like infection or wear.
Implants are primarily fabricated from metals, polymers, ceramics, and natural materials such as silk, collagen, and fibrin, with composite biomaterials combining these elements for enhanced performance. While gold was once extensively used, researchers now focus on developing cost-effective materials that deliver superior performance without adverse effects. The effectiveness of implants, particularly in aging populations, has driven significant demand for biomedical implant materials. However, implant failure remains a concern due to corrosion, wear, and the mismatch between the implant’s properties and the human bone’s Young’s modulus. Thus, biomaterials must meet stringent criteria, including biocompatibility, osseointegration, corrosion and wear resistance, and mechanical reliability (Hussein et al. 2015; Pandey et al. 2020).
Specific types of implants serve distinct functions: dental implants, commonly made of titanium, provide a permanent solution for tooth replacement; cardiovascular implants, such as pacemakers and stents, manage heart conditions using titanium and biocompatible polymers; and orthopedic implants like hip and knee replacements, made from metals, ceramics, or polymers, restore joint function. Prosthetic limbs, constructed from lightweight materials like alloys and carbon fiber, offer personalized solutions for individuals with limb loss. The exploration of both biodegradable and non-biodegradable materials, as well as hybrid approaches, holds promise for the development of multifunctional, patient-specific implants that improve medical outcomes.
The primary objective of this review is to provide a comparison between biodegradable and non-biodegradable implants, focusing on their material properties, clinical applications, and inherent limitations. It aims to analyze recent advancements in both categories, with particular attention to their performance in key medical areas such as orthopedics, cardiovascular, and dental applications. Furthermore, the review seeks to explore emerging trends, challenges, and potential future directions for the development of next-generation implants that promise enhanced patient outcomes, biocompatibility, and functionality.
The scope of the paper includes a detailed examination of the fundamental principles and design considerations for both biodegradable and non-biodegradable implants. This will encompass an exploration of the materials used, degradation mechanisms, and mechanical performance. The review will provide an in-depth analysis of clinical applications, supported by case studies, comparing the effectiveness of these implant types in various medical contexts. Additionally, it will address current challenges, such as regulatory hurdles, biocompatibility concerns, and long-term performance, while highlighting innovative solutions like smart implants and hybrid designs.
1.1 History of biomaterials
Biomaterials, as we know them today, did not exist a few decades ago, and terms like “biomaterial” were not widely used. Medical devices were not specifically manufactured, apart from external prostheses such as spectacles, fracture fixation devices, limbs, and orthodontic devices. There were no organized official approval processes or feasible knowledge of biocompatibility, and undoubtedly no courses related to biomaterials were available. However, basic biomaterials have previously been utilized, yielding unsatisfactory and ambiguous outcomes (Ratner and Zhang 2020). The history of dental implants is rich and varied, with different civilizations and eras contributing to their development. Gold was among the earliest engineering materials utilized, due to its excellent mechanical properties, making it a preferred material for early implants (Saini 2015). Ancient civilizations like the Phoenicians and Mayan civilizations used gold wire and ivory teeth for tooth correction. The 18th century saw widespread dental work with pioneers like Pierre Fauchard and John Hunter (Awasthi et al. 2021). The Industrial Revolution introduced stainless steel for teeth deformities, and chromium and cobalt were utilized in the fabrication of subperiosteal implants. Various bioimplants produced in the 1950s had an extremely low success rate, with limited comprehension of sterilization and biocompatibility. Primarily, the initial investigations into metals concentrated on chemical theories to illustrate bioreaction. In 1829, a study was conducted on various metallic materials, including platinum, gold, lead, and silver, in dogs. Platinum emerged as the most successful material. In 1886, nickel-plated sheet steel, along with nickel-plated screws, was thoroughly studied to produce bone fixation plates (Ratner and Zhang 2020). Figure 2a and b shows the timeline leading to the discovery of implants.
Flowchart showing the growth of implants over the years. (a) An illustration of a timeline depicting the discovery of implants from 2500 BC to 600 AD. (b) A timeline depicting the discovery of implants between the 18th century and 20th century (reproduced with permission from Taylor & Francis).
The clinical trial process for nondegradable and degradable implants is a multi-phased endeavor that typically spans several years. The timeline for clinical trials of such implants typically follows a structured progression, beginning with a preclinical phase lasting 1–3 years, during which laboratory and animal testing are conducted to assess biocompatibility, mechanical strength, and overall safety (www.Accessdata.Fda.Gov 2024). This is followed by Phase I clinical trials, lasting 6 months to 2 years, involving a small group of healthy volunteers or patients to evaluate initial safety and biological responses (ClinicalTrials.Gov 2024). Next, Phase II clinical trials, spanning 1–3 years, focus on a larger patient cohort to assess efficacy, optimize dosage, and identify potential side effects. Phase III trials, which may take 2–5 years, involve large-scale studies aimed at confirming efficacy, monitoring adverse effects, and comparing the implant’s performance with existing treatments. Upon completion of these trials, the regulatory review and approval process, lasting 1–2 years, begins. This involves submitting clinical trial data to regulatory bodies such as the FDA or EMA for thorough assessment (Clinical Trials Guidance Documents | FDA 2024). If approved, the implant can be marketed and used in clinical practice. However, post-marketing surveillance is ongoing to monitor the implant’s long-term safety and effectiveness (Medical Devices | FDA 2024).
2 Types of biomaterials
Biodegradable and non-biodegradable biomaterials are used in synthesizing biomedical implants such as orthopedic implants, dental implants, cardiovascular implants and wound closure tools. Types of materials differ according to the application, the main considerations in implant/biomaterial selection revolve around factors such as mechanical strength, to withstand physical loads; biocompatibility, ensuring that the implant does not trigger adverse immune responses; durability, for long-term functionality in permanent implants; and degradability, especially in temporary implants meant to break down after healing. Other important factors include corrosion resistance, toxicity, and the implant’s ability to integrate with surrounding tissues, which significantly influence both patient outcomes and implant durability.
Medical materials science plays a pivotal role in advancing biomedical devices and implants, with material selection being crucial for optimizing clinical outcomes. The hierarchical structure depicted in Figure 3 categorizes the primary materials used in medical applications into four groups: metals, ceramics, polymers, and composites. Each of these material types exhibits unique properties that make them suitable for specific medical uses. Understanding their advantages, limitations, and applications is essential for developing safe and effective medical devices.
Flowchart enumerating the applications, benefits, and drawbacks of various materials used to manufacture bioimplants.
2.1 Metals
Metals play a critical role in biomedical applications, especially in devices such as joint prostheses, bone plates, dental implants, wires, and staples. Their high strength, durability, and wear resistance make them indispensable for load-bearing applications where mechanical stability is paramount. Metals such as titanium, stainless steel, and cobalt-chromium alloys are particularly favored due to their ability to withstand high mechanical stresses and their ductility, which allows for the manufacturing of complex and precise designs. These characteristics make them essential in orthopedic and dental procedures where the materials must endure prolonged use without significant wear or deformation.
However, metals have inherent limitations, particularly in terms of corrosion resistance and biocompatibility. Over time, exposure to the physiological environment can cause metal implants to corrode, releasing metal ions that may lead to inflammatory responses, allergic reactions, or even toxicity. This poses significant challenges for long-term implantation, as the buildup of corrosion products can compromise both the functionality of the implant and the health of surrounding tissues.
Metals and alloys such as titanium, cobalt-chromium, magnesium, and stainless steel are widely used in the medical field due to their superior mechanical properties, including fatigue resistance, biocompatibility, and high wear resistance. These materials are used in manufacturing medical devices like hip and knee replacements, dental implants, bone plates, and screws, highlighting their versatility in a range of medical applications. According to reports, substantial investments, amounting to nearly £2.5 million, have been made in medical apparatus composed of metals and alloys (Zhang et al. 2021).
For biomedical implants, metals must meet stringent criteria to ensure they provide reliable, long-term performance without compromising patient safety. Specific design and material requirements must be followed, especially concerning biocompatibility, mechanical strength, and corrosion resistance. Among the metallic biomaterials, titanium, cobalt-chromium, and stainless-steel alloys remain the most employed due to their proven reliability in medical use (Rohman 2014). These metals are extensively used in applications ranging from fracture fixation devices like bone plates and screws to dental wires, braces, and joint replacements, where high strength and wear resistance are critical for success.
2.2 Ceramics
Ceramics are primarily utilized in medical coatings, tools, and equipment due to their inert nature and excellent biocompatibility. In applications such as bone grafts, dental crowns, and hip replacements, ceramics like alumina, zirconia, and bioactive glass offer superior corrosion resistance, making them suitable for long-term implantation in contact with biological tissues. The latest developments in bio-ceramics have led to the emergence of a translucent monolithic zirconia variety that simplifies tooth preparation while offering commendable mechanical properties (Aisyah et al. 2020; Ghodsi and Jafarian 2018). However, ceramics are brittle and have low impact strength, which limits their use in high-stress applications where mechanical strength and toughness are critical. Additionally, the manufacturing of ceramic implants is expensive and complex, requiring specialized techniques that further limit their widespread application.
2.3 Polymers
Polymers are commonly used for soft tissue applications such as artificial veins, tendons, and arteries. Their versatility, ease of manufacture, and low density make them ideal for applications where flexibility is required. Polymers such as polyethylene with an extremely high molecular weight, polypropylene, poly (glycolic acid), polyurethane (PU), polyvinyl acetate, polyethylene, and polytetrafluoroethylene (PTFE) can be engineered to mimic the properties of natural tissues, making them suitable for reconstructive surgeries and tissue scaffolds. Among them, poly (lactic acid) is the most widely used synthetic polymer (Wei et al. 2018). However, polymers tend to degrade more quickly than other materials, limiting their use in long-term applications. This limitation can be advantageous in biodegradable implants, but it poses challenges when longevity and durability are needed.
2.4 Composites
Composites combine materials from different categories to harness the advantages of multiple properties in one implant, offering solutions for applications such as artificial joints, heart valves, and dental implants. Combining these materials aims to enhance mechanical strength, hardness, and fatigue. In polymer matrix implants, adding ceramic reinforcements like hydroxyapatite (HA), calcium phosphates, or bio-glass improves biocompatibility and the modulus of elasticity of the base material. This enhances the mechanical properties of the implant, bringing it closer to those of natural human bone, which may diminish stress shielding (Festas et al. 2020). Carbon-fiber-reinforced polymers and metal-ceramic composites provide enhanced mechanical strength, corrosion resistance, and biocompatibility. Despite their promise, composites are challenging to manufacture consistently due to the variability in the mechanical properties of the constituent materials. This inconsistency can lead to variations in implant performance and presents significant challenges in maintaining quality control during production.
2.5 Key considerations in implant selection
The selection of materials for medical applications depends on balancing their mechanical, biological, and chemical properties. Metals, while excellent in mechanical performance, may be unsuitable for applications where biocompatibility is a priority. Ceramics, with their exceptional biocompatibility and corrosion resistance, are ideal for long-term implants but are unsuitable for high-impact conditions due to their brittleness. Polymers are preferred for applications requiring flexibility and biodegradability, although their mechanical strength and longevity are limited. Composites offer a tailored approach by combining the advantages of different materials, though they introduce challenges in consistency and reliability. The key to optimizing medical implants lies in selecting the right material for each specific application, ensuring that both clinical and functional requirements are met.
Metal implants have long been utilized in medical applications for repairing hard tissue due to their exceptional mechanical properties. Titanium alloys, for instance, are extensively employed for facial tissue repair. Biomechanical design is a key consideration for metallic implants, necessitating careful attention to their mechanical characteristics, including stiffness, strength, resistance to wear and fatigue, and corrosion resistance. Figure 4 shows the schematic illustration of the essential criteria for bioimplants. Due to their critical importance, Significant research is underway to enhance biocompatibility and refine mechanical properties. Microbial corrosion should not be overlooked, as it can greatly impact the success of implants in load-bearing locations. This issue is of utmost importance due to its influence on the longevity and functionality of implants. It is imperative to thoroughly consider the effects of microbial corrosion in the development and maintenance of dental and orthopedic implants (Yang et al. 2017). Figure 5 presents some commonly encountered corrosion issues in bio-implants. Uniform corrosion usually occurs on the surface of dental implants & orthopedic implants. Pitting corrosion occurs in implants where body fluids are trapped. Galvanic corrosion occurs when two implants of different compositions are in contact with each other, for example, hip or knee joint implants. Stress corrosion cracking is generally found in implants that have a complex design, like, for example, fixation screws etc.
A schematic diagram showing the different criteria of bioimplants.
Schematic diagrams showing some corrosion mechanism types in bioimplants.
3 Non-biodegradable implants
Non-biodegradable implants are designed to provide long-term or permanent solutions for a wide range of medical applications, offering sustained mechanical support and stability. Unlike biodegradable implants, they remain in the body indefinitely unless surgically removed. Below is a comprehensive discussion of the common materials, applications, and challenges associated with non-biodegradable implants.
3.1 Common non-biodegradable materials
3.1.1 Titanium alloys
Titanium alloys suitable for orthopedics, orthodontics, cardiovascular, and reconstructive purposes must possess excellent resistance to corrosion, high biocompatibility, and exceptional mechanical characteristics, such as a low density and Young’s modulus. Given their superb biocompatibility and commendable resistance to corrosion, high specific weight, and lack of allergic reactions, titanium alloys are increasingly popular in biomedical applications (Simona Baltatu et al. 2019). Some critical factors to consider when developing biocompatible titanium alloys are low elastic modulus and toxicity levels (Verma 2020). Titanium is the only biomaterial suitable for bone integration among various metallic biomaterials due to its bioactive behavior. The leading cause of this behavior can be attributed to hydrated titanium oxide on the surface, which enables the incorporation of phosphorus and calcium (Simona Baltatu et al. 2019). Various alloying elements in titanium alloys contribute to diverse mechanical and microstructural properties. The properties of the titanium alloy can be tailored to meet specific requirements by incorporating several alloying elements. The alloying elements incorporated in titanium can be classified into three groups:
Alpha (α) stabilizers: Al, O2, N2, C
Beta (β) stabilizers: Si, Mo, Nb, Ta, Mn, Fe, Cr, W, Co, Ni, Cu, V
Neutral elements: Th, Sn, Zr, Hf, Ge.
Some examples of titanium alloys with different alloying elements are given in Figure 6.
Flow chart illustrating some of the critical alloying elements in Ti alloys for bioimplants.
3.1.1.1 Corrosion behavior of Ti alloys
The corrosion behavior of each titanium alloy is distinct and determined by the weight percentage of titanium and the alloy elements. As a material frequently employed in biomedical applications, Ti6Al4V alloy is regarded as the benchmark against which all others are measured. In the subsequent section, the corrosion characteristics of specific titanium alloys are elaborated. The main distinction between Ti6Al4V and Ti6Al4V ELI (extra-low interstitial) is that the ELI exhibits an enhanced fracture toughness due to its minimal oxygen content of 0.13 percent. An investigation was conducted by Tamilselvi et al. (2006) to analyze the corrosion properties of Ti6Al4V (ELI) and Ti6Al7Nb in SBF (Simulated bodily fluid) over multiple time intervals. As the time interval grew from 0 to 360 h, the Ecorr of Ti6Al7Nb alloys improved while it decreased. As the immersion duration increased, a stable surface film appeared on the alloy. While the behavior of TiAl4V was comparable, the current density of Ti6Al7Nb was significantly lower. The researchers concluded that the diminished current density values in Ti6Al7Nb, in contrast to Ti6Al4V, can be attributed to the passive film’s niobium content (Tamilselvi et al. 2006). Stango et al. (2018a) investigated the electrochemical characteristics of Ti6Al4V by utilizing the EIS technique and immersion in SBF solution. The obtained values are presented in Table 1. The data provided in Table 1 shows important insights into the corrosion behavior of untreated and hydroxyapatite (HAp)-coated Ti6Al4V alloys. In terms of open circuit potential (Ecorr), the untreated Ti6Al4V sample exhibits a more negative value compared to the HAp-coated. This suggests that the untreated alloy has a higher tendency to corrode, while the HAp-coated alloy, with its less negative Ecorr, demonstrates improved resistance to corrosion initiation. Although both samples share similar Tafel slope values for the anodic and cathodic reactions, indicating similar electrochemical reaction mechanisms, there is a significant difference in the corrosion current density (Icorr). The untreated sample has a significantly higher Icorr compared to the HAp-coated sample. This reduction of approximately 69 % in Icorr for the coated sample indicates the effectiveness of the HAp coating in slowing down the corrosion process. The results indicate that the HAp coating provides a physical barrier, reducing the interaction between the alloy and the corrosive environment. This is further reflected in the lower Icorr value and the shift toward a nobler Ecorr, suggesting that the coating enhances the corrosion resistance of the alloy. Consequently, these findings highlight the benefits of using HAp coatings, especially in applications where the corrosion resistance of the material, such as in biomedical implants, is a critical factor.
EIS parameters obtained for Ti6Al4V in SBF (Stango et al. 2018a).
| Sample | Ecorr (mV vs SCE) | βa (mV) | βc (mV) | Icorr (µA/cm2) |
|---|---|---|---|---|
| Ti6Al4V (untreated) | −214 | 179 | 146 | 1.575 |
| Ti6Al4V (HAp coated) | −162 | 179 | 146 | 0.492 |
In addition to testing uncoated (bare) Ti6Al4V alloy, they also examined it with a hydroxyapatite coating. It was determined that the untreated Ti6Al4V sample had a charge transfer resistance value (Rct) of 29,012 Ω cm−1. In contrast, the HAP-coated Ti6Al4V alloy displayed a significantly higher Rct of 85,279 Ω cm−1 Gouda et al. (2022) also investigated the corrosion behavior of the Ti–14Mn biomedical alloy in Ringer solution, mimicking body fluid composition at 37 °C. The potentiodynamic polarization (PDP) curves shown in Figure 7 revealed consistent linear behavior indicative of passive film formation, primarily TiO2.
The potentiodynamic polarization curves of (14Mn) alloy in ST, 10 % CR, 30 % CR, and 90 % CR conditions (Gouda et al. 2022).
The authors concluded their study by implying that HAP coating enhanced the resistance to corrosion of the Ti6Al4V alloy and made it more desirable than the uncoated Ti6Al4V for biomedical purposes. Sasikumar and Rajendran (2017) researched the electrochemical characteristics of Ti6Al7Nb and Ti5Al2Nb1Ta alloys. In SBF, both untreated and heat-treated alloy types were evaluated. After carefully examining their potentiodynamic graphs, it becomes evident that the Ti6Al7Nb alloy undergoes a significant transition towards a nobler energy level. Furthermore, a substantial decrease in corrosion current density is seen in comparison to the Ti5Al2Nb1Ta alloy that underwent heat treatment.
Additionally, subsequent immersion testing conducted seven days later demonstrated that the current density value of the heat-treated alloys was significantly reduced when compared to that of the untreated alloys (Sasikumar and Rajendran 2017). The corrosion characteristics of Ti–13Mo–7Zr–3Fe (TMZF) alloy, which comprises metastable β and both α + β phases, Ti–35Nb–7Zr–5Ta (TiOsteum), and Ti6Al4V ELI were assessed individually in a 5 M HCl solution and a 0.9 percent NaCl solution, respectively (Atapour et al. 2011; Stango et al. 2018b). Table 2 presents Ecorr and Icorr values for different titanium-based alloys and their phases, including TMZF in both α + β and metastable β phases, TiOsteum, and Ti–6Al–4V ELI. These materials are frequently evaluated for their corrosion resistance in biomedical applications due to their use in implants. The Ecorr values provide insight into the thermodynamic tendency of each material to corrode, with more negative values indicating a higher likelihood of corrosion initiation. Among the samples, TMZF in the α + β phase exhibits the most negative Ecorr value of −421 mV, suggesting the greatest susceptibility to corrosion. Conversely, TiOsteum shows the least negative Ecorr at −292 mV, indicating a higher resistance to corrosion initiation. TMZF in the metastable β phase and Ti–6Al–4V ELI show intermediate values of −343 mV and −380 mV, respectively. This trend highlights that TiOsteum possesses a more noble electrochemical behavior compared to the other alloys, potentially making it a superior candidate for environments requiring enhanced corrosion resistance.
Corrosion potential and corrosion current densities of TMZF (α + β phase and metastable β), TiOsteum & Ti-6Al-4V (Atapour et al. 2011).
| Material used | Ecorr (mVSCE) | icorr (nA/cm2) |
|---|---|---|
| TMZF (α + β phase) | −421 (±12) | 29 (±15) |
| TMZF (metastable β) | −343 (±83) | 20 (±10) |
| TiOsteum | −292 (±6) | 12 (±5) |
| Ti–6Al–4V ELI | −380 (±65) | 31 (±13) |
Corrosion current density Icorr, which directly correlates with the corrosion rate, further emphasizes the superior performance of TiOsteum. With an Icorr of 12 nA/cm2, it exhibits the lowest corrosion rate among the materials tested. In contrast, TMZF in the α + β phase and Ti–6Al–4V ELI display higher Icorr values of 29 nA/cm2 and 31 nA/cm2, respectively, indicating a faster corrosion rate. TMZF in the metastable β phase also demonstrates a lower Icorr, which suggests a moderately improved corrosion performance relative to its α + β phase counterpart, but not as significant as TiOsteum. In conclusion, these results indicate that TiOsteum, with the least negative Ecorr and lowest Icorr, provides superior corrosion resistance when compared to the other tested materials. TMZF in the metastable β phase also shows improved corrosion behavior over the α + β phase, although it does not perform as well as TiOsteum. Ti–6Al–4V ELI, a commonly used alloy in biomedical implants, demonstrates corrosion behavior comparable to the TMZF α + β phase, but both exhibit lower resistance compared to TiOsteum. Furthermore, the active-to-passive transition was not observed in any of the four alloys, according to the report (Atapour et al. 2011). TiO had the most significant activation time, according to the findings of OCP tests, but Ti6Al4V had the shortest activation time in 5 M HCl testing. Tests of anodic polarization on Ti6Al4V ELI and TMZF (α + β phase) alloys unveiled active-to-passive transitions. On the contrary, metastable β was identified in TiO steum and TMZF as exhibiting spontaneous passive behavior. Hydrogen evolution rates on TiO steum and Ti6Al4V were discovered to be roughly double those of both types of TMZF alloys during the cathodic polarization test. The cathodic process has, nevertheless, little effect, as evidenced by the dissimilar corrosion rates and passivation characteristics of these two alloys. The weight loss test results indicated that TiO steum experienced the least weight loss due to corrosion, while Ti6Al4V underwent the significant weight loss (Atapour et al. 2011; Stango et al. 2018b).
3.1.2 Stainless steel
Stainless steel is a corrosion-resistant alloy primarily composed of iron, chromium, nickel, and other elements. The amount of chromium must be at least 10.5 % by weight for it to be classified as stainless steel. Additional alloying elements in these alloys include nickel, manganese, molybdenum, and niobium. Stainless steels are utilized in numerous industrial sectors, including the biomedical industry, due to their exceptional combo of mechanical characteristics and corrosion resistance. However, because of their interaction with the environment, every material has its limitations; stainless steels are no exception; for example, austenitic stainless steel contains nickel, and when an implant made of austenitic stainless steel is used, nickel can be released over a while into the body which can lead to adverse reactions for patients that are allergic to nickel (Hedberg and Odnevall Wallinder 2016). A typical potentiodynamic polarization curve of stainless steel, as depicted in Figure 8, illustrates the presence of pitting and general corrosion. In contrast, the microstructure of different grades of stainless steel is illustrated in Figure 9. Figure 10 presents the flow chart illustrating the effect of alloy elements on the properties of 304 stainless steel.
Corrosion graph showing uniform corrosion and pitting.
An illustration showing the stainless-steel grades and their microstructures.
Schematic indicating the different effects of various alloying elements on SS-304.
3.1.2.1 Corrosion behavior of steels
The study conducted by Mutlu and Oktay (2015) focused on the localized corrosion characteristics of AISI 316 austenitic stainless steel and a biomedical-grade TiNbCu alloy, with particular attention to their behavior in simulated body fluid (SBF), specifically Hank’s solution. The research explored the effect of varying pH levels on the corrosion properties of these materials, as pH is a critical factor in simulating physiological conditions. In the human body, pH can fluctuate due to localized factors such as inflammation, infection, or tissue degradation, making it important to understand how biomaterials behave under such conditions. The authors’ investigation of AISI 316 stainless steel, widely used in biomedical applications due to its corrosion resistance and biocompatibility, revealed significant changes in electrochemical behavior when exposed to different pH levels of Hank’s SBF solution. As the pH of the solution decreased, representing more acidic conditions, the corrosion current density Icorr increased significantly. This indicates an acceleration in the corrosion process, as Icorr is directly related to the corrosion rate; a higher Icorr reflects a faster degradation of the material. Simultaneously, the corrosion potential Ecorr dropped to more negative values, signifying an increased thermodynamic tendency for the material to corrode. This shift suggests that the material becomes more susceptible to corrosion in more acidic environments, which is critical in scenarios such as chronic inflammation, where the pH of the surrounding tissue may become more acidic.
These findings, presented in Figure 11, provide clear evidence of the strong relationship between pH and the corrosion behavior of AISI 316 stainless steel. The observed increase in Icorr and drop in Ecorr with decreasing pH values highlights the material’s susceptibility to corrosion in acidic conditions, which could lead to localized corrosion in biomedical implants. Localized corrosion, such as pitting or crevice corrosion, is particularly concerning in stainless steels because it can lead to rapid material degradation, compromising the integrity of implants and causing potential failure. Moreover, the comparison of AISI 316 stainless steel with the TiNbCu alloy offers important insights into material performance in biomedical applications. Titanium-based alloys, particularly those like TiNbCu, are known for their superior corrosion resistance and biocompatibility. In contrast to AISI 316, TiNbCu alloys are often more resistant to localized corrosion due to the formation of a stable and protective oxide layer on their surface. The study by Mutlu and Oktay (2015) underscores the limitations of AISI 316 in corrosive environments like those simulated by Hank’s SBF solution with decreasing pH, suggesting that alternative materials like TiNbCu may provide better performance in situations where localized corrosion is a concern. The observed increase in Icorr and drop in Ecorr with decreasing pH values emphasize the need for careful consideration of material selection for biomedical implants, especially in scenarios where the pH may become more acidic. While AISI 316 remains a widely used material, its susceptibility to localized corrosion under acidic conditions highlights the potential advantages of alternative materials like TiNbCu, which may offer enhanced corrosion resistance and durability in biomedical applications. These findings have important implications for the design and lifespan of implants, especially in complex physiological environments.
Tafel curves of AISI 316 stainless steel in Hank’s simulated body fluid at varying pH levels (7.4, 5.0, 2.5) (Mutlu and Oktay 2015) (reproduced with permission from Taylor & Francis).
The study conducted by Köse (Köse 2018) sought to evaluate the corrosion characteristics of AISI 420 stainless steel in simulated body fluid environments over varying periods. AISI 420 stainless steel is commonly used in medical devices and implants due to its mechanical properties and corrosion resistance, but welding processes can significantly affect its performance in corrosive environments. In this investigation, three samples were tested: one untreated base sample and two laser-welded samples, denoted as B1 and B2. The corrosion performance of each sample was assessed using the weight loss method, which is a widely used technique to quantify material degradation in corrosive environments by measuring the mass loss over time. The base sample served as a control, while the two laser-welded samples underwent different welding conditions. B1 was welded using a 3500-W laser, while B2 was welded at a higher power of 4,000 W. The primary objective of the study was to determine how these different welding processes affected the corrosion resistance of AISI 420 stainless steel in a simulated body fluid environment, which mimics the conditions that materials experience when implanted in the human body. Laser welding, although beneficial for its precision and control, can alter the microstructure and surface properties of the material, potentially influencing its susceptibility to corrosion.
The findings, as depicted in Figures 12 and 13, clearly show that the base sample exhibited superior corrosion resistance compared to both laser-welded samples. This enhanced corrosion resistance of the base metal can be attributed to its unaltered microstructure, which retains the inherent passivity and protective oxide layer that stainless steel forms to resist corrosion. In contrast, the laser-welded samples, particularly B1 and B2, showed significantly higher weight loss, indicating accelerated corrosion rates. The increased susceptibility to corrosion in the welded samples is likely due to the thermal effects of the welding process, which can induce changes in the material’s grain structure, create heat-affected zones (HAZ), and potentially lead to the formation of micro-cracks or defects that compromise the material’s passive oxide layer. Moreover, the study highlights the effect of laser power on corrosion performance. B1, which was welded at a lower power of 3,500 W, displayed less corrosion compared to B2, which was welded at 4,000 W. This suggests that higher laser power exacerbates the thermal impact on the material, resulting in more severe microstructural changes and greater susceptibility to corrosion. The higher energy input in the B2 sample likely caused more extensive melting and re-solidification of the material, increasing the likelihood of defects such as porosity or grain boundary sensitization, which weaken the material’s corrosion resistance.
Graph showing the weight loss of the base metal and heat-treated B1 & B2 samples in SBF for 28 days (Köse 2018).
The corrosion rate (mm/y) of heat-treated B1 and B2 samples immersed in SBF and composed of base metal is depicted in the graph (Köse 2018).
These results emphasize the importance of carefully controlling welding parameters when fabricating components from AISI 420 stainless steel, especially for biomedical applications where corrosion resistance is a critical factor in ensuring the longevity and safety of implants. While laser welding offers precision and minimal heat input compared to traditional welding methods, this study indicates that even small variations in welding parameters can significantly impact the corrosion performance of stainless steel. The differences in corrosion resistance between the base metal and welded samples also underscore the need for post-weld treatments or protective coatings to restore or enhance the corrosion resistance of welded components, particularly in environments that mimic the human body.
The investigation by Köse (2018) highlights the critical relationship between welding processes and corrosion performance in AISI 420 stainless steel. The base sample, with its intact microstructure, demonstrated superior corrosion resistance compared to the laser-welded samples, which were more vulnerable to corrosion due to the thermal effects of the welding process. Furthermore, the study shows that higher laser power results in more pronounced corrosion susceptibility, emphasizing the importance of optimizing welding parameters to balance mechanical integrity and corrosion resistance. These findings have significant implications for the fabrication of medical devices and implants, where both durability and biocompatibility are essential for long-term performance in the human body.
Claudio Zanca and colleagues conducted a study on the corrosion characteristics of uncoated and coated SS 304 with three distinct biomedical coatings [hydroxyapatite (HAp), collagen (CL), and chitosan (CS)] in an SBF environment. Over 21 days at 37 °C, SS304 samples’ resistance to corrosion was assessed in SBF, both with and without the application of a coating. The samples were coated with galvanic deposition. The values of Ecorr and icorr, reported in Table 3, were ascertained using Tafel’s curves, as depicted in Figure 14. The data presented in Table 3 compares the corrosion behavior of SS 304 stainless steel in three different conditions: uncoated, coated with a hybrid system (HS) and chitosan (CS), and coated with an additional layer of curcumin-loaded liposomes (CL). The uncoated SS 304 sample shows a more negative Ecorr, indicating a greater susceptibility to corrosion initiation. Additionally, its corrosion Icorr is the highest among the samples, reflecting a relatively fast corrosion rate. These results highlight the vulnerability of uncoated SS 304 in corrosive environments where it lacks protection. When SS 304 is coated with HS and CS, a significant improvement in corrosion resistance is observed. The Ecorr becomes less negative, suggesting a reduced tendency for corrosion initiation. Furthermore, the corrosion current density is substantially lower, indicating a slower corrosion rate. This enhancement is likely due to the protective properties of the HS and CS coating, which forms a barrier that shields the SS 304 surface from the corrosive medium. Chitosan, in particular, is known for creating a protective, adhesive layer, which contributes to the material’s improved corrosion resistance. Adding curcumin-loaded liposomes (CL) to the HS and CS coating further enhances the material’s corrosion protection. The corrosion potential shifts to a more positive value, indicating that the coated sample is even less prone to corrosion. Additionally, the corrosion current density decreases slightly compared to the sample coated only with HS and CS, reflecting a further reduction in the corrosion rate. The presence of curcumin-loaded liposomes likely contributes to this improvement due to their potential antioxidant properties, which help to neutralize reactive species in the corrosive environment, offering additional protection to the SS 304 substrate.
Ecorr and Icorr values of all three SS 304 specimens (Zanca et al. 2021).
| Sample | Ecorr (V versus Ag/AgCl) | Icorr (A/cm2) |
|---|---|---|
| SS 304 (uncoated) | −0.161 | 3.92E-06 |
| SS 304 (coated with HS and CS) | −0.017 | 8.51E-07 |
| SS 304 (coated with HS, CS and CL) | 0.071 | 8.22E-07 |
Tafel curves of SS304, SS304 coated with HA, CS, and SS 304 coated with HA, CS, and CL immersed in SBF (Zanca et al. 2021).
The results demonstrate a progression in corrosion resistance as the coating system becomes more complex. The uncoated sample exhibits the poorest performance, while the sample coated with HS, CS, and CL provides the highest level of corrosion protection. This study highlights the effectiveness of multifunctional coating systems in significantly improving the electrochemical stability of SS 304, especially in environments where corrosion resistance is essential for the long-term durability and safety of the material. Such advanced coatings offer significant potential in extending the lifespan of stainless steel in applications where high-performance corrosion resistance is critical. The researchers reported that improvements in resistance to corrosion and biocompatibility of SS 304 resulted from the application of the coatings. Furthermore, they advised the implementation of galvanic deposition as a cost-effective approach to attaining these advantages (Zanca et al. 2021). Yeganeh et al. (2021) investigated the corrosion behavior of 17-4 PH SS that was both wrought and selectively laser melted (SLM). For one hundred hours, each sample was submerged in Ringer’s solution. The corrosion results derived from the PDP test, depicted in Figure 15, indicate that the corrosion current density of the SLM sample is around one-tenth lower than that of the wrought sample. These values are detailed in Table 4. Table 4 examines the corrosion performance of two types of stainless steel samples: one produced through traditional wrought processing and the other via selective laser melting, a modern additive manufacturing technique. The electrochemical data highlight differences in corrosion behavior between these fabrication methods, focusing on corrosion potential and corrosion current density. The wrought sample exhibits a more negative Ecorr indicating a higher susceptibility to corrosion initiation. This suggests that the wrought material is thermodynamically more prone to corrosion in aggressive environments. In contrast, the SLM sample demonstrates a less negative Ecorr, implying improved corrosion resistance and a lower likelihood of corrosion initiation. This shift in corrosion potential can be attributed to the refined microstructure and homogeneity typically associated with SLM processing, which enhances the material’s protective oxide layer.
PDP graph of wrought and SLM 17-4 PH SS in Ringer’s solution (Yeganeh et al. 2021) (reproduced with permission from AIP Publishing).
Corrosion values of wrought and SLM 17-4 SS samples acquired from the PDP test (Yeganeh et al. 2021).
| Sample type | Ecorr (mV) | Icorr (μA cm−2) |
|---|---|---|
| Wrought | −216 | 2.6 |
| SLM | −176 | 0.25 |
Additionally, the corrosion current density Icorr which directly correlates with the rate of corrosion, is significantly lower for the SLM sample compared to the wrought material. The lower Icorr for the SLM sample indicates a much slower corrosion rate, further underscoring its superior corrosion resistance. This improved performance can be linked to the unique microstructural characteristics imparted by the SLM process, such as finer grain size and a more uniform distribution of alloying elements, which contribute to enhanced passivation and reduced corrosion activity.
The findings demonstrate that SLM-processed stainless steel exhibits better electrochemical stability and corrosion resistance compared to its wrought counterpart. These results are significant in the context of material selection for applications in harsh environments, particularly where long-term durability and corrosion resistance are critical. The superior performance of SLM samples suggests that additive manufacturing techniques like SLM may offer significant advantages over traditional processing methods in producing stainless steel components for industries such as aerospace, medical implants, and chemical processing. Furthermore, after 100 h of immersion, the improved corrosion resistance of SLM steel might have resulted from the diminished formation of local galvanic cells.
3.1.3 Cobalt-chromium alloys
Cobalt-chromium alloys have exhibited exceptional wear resistance compared to alternative metallic biomaterials; thus, they are the material of choice for implants subjected to substantial friction and wear. Molybdenum is the prototypical element alloyed with cobalt and chromium; nickel, tungsten, and iron are further alloying elements. A range of procedures can be employed to produce cobalt-chromium alloys, as demonstrated in Table 5. These techniques include casting, ball milling, pressure bed fusion (PBF), and direct energy deposition (DED) (Acharya et al. 2021; Balamurugan et al. 2008). As seen in Figure 16, implant applications are currently limited to F562 and F75 (Figure 17).
Different techniques are available for the manufacturing of cobalt-chromium alloys.
| Manufacturing technique | Working process | Advantages | Disadvantages |
|---|---|---|---|
| Ball milling (Jamkhande et al. 2019) | Particle reduction of bulk powder takes place in the ball mill container, where heavy balls exert high mechanical energy on the powder. | It is very helpful for producing high-quality nanoparticles from the bulk powder. | The milling process takes a lot of time, which leads to high energy consumption. The powder can get contaminated because of the steel balls. |
| Powder bed fusion (Singh et al. 2019) | PBF is an additive manufacturing method in which a laser beam scans each layer of the powder bed at a precise speed to melt and fuse it to the solid material under the powder bed. Different PBF techniques are selective laser sintering, selective laser melting, direct laser metal sintering, and the electron beam method. | PBF can be used for a wide variety of materials, such as titanium, stainless steel, etc., and it is the best technique when it comes to the production of intricate products such as bioimplants because of its excellent precision. | The process is slow and takes a long time to complete. The power consumption of PBF is also quite high. |
| Direct energy deposition (Kannan et al. 2021) | DED is another additive manufacturing technique in which metallic powder supplied to the focus of a laser beam gets melted and deposited onto the substrate. | Grain structure can be controlled in DED; therefore, this technique can also be used in complex bioimplants. | Post-processing is required before the finished product can be used. |
| Casting (Dhanani et al. 2017) | Molten metal is poured using a sprue into a hollow mold made of heat-resistant material. The poured molten metal solidifies inside the mold cavity, designed according to the required shape. | Cost-effective because the processing cycle duration is short, and assembly is relatively inexpensive compared to other manufacturing techniques. | Problems with casting include inaccurate dimensions due to contraction, errors that can occur in material composition, etc. |
ASTM recommended Co-Cr alloys for implant applications.
OCP values of the CoCrMo alloys in saline solution, Hank’s solution, and E-MEM + FBS (Hiromoto et al. 2005) (reproduced with permission from Elsevier).
3.1.3.1 Corrosion behavior of Co-Cr alloys
In their research, Hiromoto et al. (2005) explored the influence of varying molybdenum content on the corrosion resistance of Co-Cr-Mo alloys, with specific weight percentages of 6 wt% and 8 wt% molybdenum. Two distinct weight ratios, denoted as 6Mo-1 (50 %) and 6Mo-2 (88 % forging), were employed to further assess how mechanical processing impacts the alloy’s corrosion resistance. Additionally, an ASTM F75-92 reference alloy (Co29Cr6Mo1Ni) was used to provide a comparative baseline. The researchers then conducted potentiodynamic polarization and EIS tests and obtained the polarization curves presented in Figure 18 and the EIS parameters shown in Table 6. The study aimed to investigate the electrochemical properties of these alloys in biologically relevant environments, including a cell culture medium (E-MEM + FBS), a 0.9 % NaCl saline solution, and Hank’s solution, which mimic physiological conditions. The results demonstrated that in saline solution, the open circuit potential Icorr values for the 6Mo-1, 6Mo-2, and 8Mo alloys were slightly lower than that of the ASTM F75-92 reference alloy. This suggests that the molybdenum-enhanced alloys, while generally corrosion resistant, may exhibit a marginally higher tendency toward corrosion initiation in saline environments compared to the reference alloy. However, in Hank’s solution, the Ecorr values across all alloys were similar, indicating comparable corrosion resistance in this medium. When exposed to the E-MEM + FBS medium, the Ecorr values for all alloys increased slightly, implying improved electrochemical stability and a reduced likelihood of corrosion initiation in the more complex cell culture environment, which simulates in vivo conditions more closely. Furthermore, the study revealed significant differences in transpassive behavior, an important aspect of alloy performance in corrosive environments. The transpassive potential in saline solution was calculated to be around 0.5 V, indicating the onset of a high-current density corrosion phase. This value was lower in Hank’s solution, around 0.3 V, suggesting that the corrosion protection mechanisms, such as passivation, may be less effective in Hank’s solution compared to saline. The current density in the transpassive region for Hank’s solution ranged from 0.3 to 0.8 V, which was higher than that observed in saline, indicating that more aggressive corrosion occurred in Hank’s solution during transpassivation. In E-MEM + FBS, the transpassive potential was even lower at 0.25 V, but the current density in the transpassive region between 0.25 and 0.53 V was higher than that in both Hank’s and saline solutions. This suggests that while the potential for transpassive corrosion initiation is lower in E-MEM + FBS, once initiated, the rate of corrosion is higher, likely due to the presence of organic components and proteins in the medium that can interact with the alloy surface. These findings highlight the complex interaction between alloy composition, forging process, and the corrosion environment. Molybdenum, as an alloying element, is known for its ability to enhance corrosion resistance, particularly by contributing to the formation of stable passive oxide layers on metal surfaces. However, the subtle differences in Ecorr values and transpassive behavior between the alloys with varying molybdenum content suggest that the protective benefits of molybdenum are nuanced and can be influenced by the specific environmental conditions. Moreover, the higher current densities observed in the transpassive regions for Hank’s solution and E-MEM + FBS underline the importance of considering the medium’s composition when evaluating corrosion performance. Hank’s solution, which contains a balance of ions and salts, and E-MEM + FBS, which mimics physiological fluids with added proteins and nutrients, can provoke different corrosion mechanisms compared to saline solution. Similarly, Mohammad Moradi et al. (2022) expanded upon this knowledge by investigating a novel Co-Cr-MoNbCu alloy with added copper (4.5 wt%) and niobium (0.2 wt%) for enhanced corrosion resistance in simulated body fluid (SBF). Their potentiodynamic polarization and electrochemical impedance spectroscopy (EIS) tests revealed that the inclusion of copper and niobium significantly improved the alloy’s resistance to pitting and crevice corrosion, which are common failure modes in biomedical implants. The polarization curves showed lower corrosion current densities for the optimized Co-Cr-MoNbCu alloy, confirming that the alloying elements contributed to the formation of a more stable and protective passive layer. The combination of potentiodynamic polarization and EIS data offers a comprehensive understanding of the corrosion mechanisms at play in these alloys. EIS parameters, including charge transfer resistance and solution resistance, further corroborated the superior performance of the Co-Cr-MoNbCu alloy. These findings suggest that the strategic incorporation of alloying elements like copper and niobium, alongside molybdenum, can yield materials with significantly enhanced corrosion resistance, suitable for long-term use in biomedical applications. These studies collectively emphasize the importance of alloy composition and processing techniques in determining the corrosion performance of Co-Cr-Mo alloys in biological environments. The influence of molybdenum content, forging ratios, and additional alloying elements like copper and niobium is demonstrated, providing critical insights for the design of next-generation biomedical materials. The variations in electrochemical behavior across different testing media also underscore the necessity of evaluating alloy performance under conditions that closely mimic the intended application environment, particularly for materials used in biomedical implants where long-term corrosion resistance is paramount.
Potentiodynamic polarization curves of Co-Cr-Mo obtained in SBF (Mohammad Moradi et al. 2022) (reproduced with permission from Elsevier).
EIS test results of Co-Cr-Mo and optimal alloy in SBF (Mohammad Moradi et al. 2022).
| Material | Ecorr (V versus Ag/AgCl) | Icorr (A/cm2) | ipassive (A/cm2) | Rs (Ω cm2) | Rct (kΩ cm2) |
|---|---|---|---|---|---|
| Co-Cr-Mo | 0.22 | 0.001 | 0.008 | 10 | 3.4 |
| Co-Cr-Mo-4.5Cu-0.2Nb | 0.28 | 0.001 | 0.008 | 10 | 4.3 |
The electrochemical performance of Co-Cr-Mo and its variant alloy, Co-Cr-Mo modified with Cu and Nb, was systematically analyzed to assess their corrosion resistance as shown in Table 6. The more positive Ecorr value exhibited by the Cu and Nb-modified alloy suggests that this material is less prone to spontaneous corrosion in the environment tested, a desirable trait for materials used in applications where prolonged exposure to corrosive agents is expected, such as in medical implants or oil and gas industry components exposed to aggressive fluids. Both alloys demonstrate similar corrosion current densities. This suggests that while the baseline and modified alloys have comparable rates of anodic dissolution, the overall corrosion process is not necessarily faster or slower for either alloy under the same conditions. However, corrosion resistance is also heavily influenced by the alloy’s ability to form and maintain a passive layer on its surface, which serves as a protective barrier against further degradation. This is where the addition of Cu and Nb plays a significant role. The charge transfer resistance (Rct) is a parameter closely associated with the formation and stability of the passive layer. A higher Rct value, as observed in the Cu and Nb-containing alloy, indicates that the passive film formed on the surface of this alloy is more resistant to breakdown, thereby offering enhanced protection against localized corrosion processes such as pitting and crevice corrosion. This is particularly relevant in applications that demand long-term material stability in chloride-rich environments, which are notorious for initiating localized corrosion. Despite the improvement in Rct, the passive current density (ipassive) remains consistent between the two alloys. This suggests that while the passive film on both alloys offers similar resistance during steady-state conditions, the modified alloy’s passive film may be more resilient in resisting breakdown or repassivating quickly after any localized attack. The enhancement in the passive behavior could be attributed to the combined effects of Cu and Nb. Copper is known for its ability to inhibit corrosion, possibly through the formation of protective copper oxides, while Nb contributes to refining the grain structure of the alloy, enhancing the stability and uniformity of the passive layer. The solution resistance (Rs) was comparable between the two alloys, indicating that external factors such as the conductivity of the electrolyte had a negligible impact on the electrochemical performance of the materials. This reinforces the interpretation that the observed improvements in corrosion resistance are predominantly due to the intrinsic properties of the alloy’s composition, rather than external experimental conditions. The inclusion of Cu and Nb in the Co-Cr-Mo alloy enhances its overall corrosion resistance, as evidenced by a more positive Ecorr and higher Rct. These findings suggest that the modified alloy has a superior ability to form and maintain a protective passive layer, which is critical for materials exposed to corrosive environments. The implications of these results are significant, as the modified alloy could find applications in industries where corrosion resistance is paramount, such as in medical devices, marine engineering, and the oil and gas sector, where materials must withstand both uniform and localized forms of corrosion. The enhanced performance of the Co-Cr-Mo-Cu-Nb alloy opens new avenues for the development of high-performance materials with tailored corrosion resistance for use in harsh environments. The optimal alloy exhibits a higher resistance to corrosion than the Co-Cr-Mo alloy. This is evident from the higher Ecorr and Rct values observed in the optimal alloy compared to the Co-Cr-Mo alloy. This increase in corrosion potential and charge transfer values indicates an enhanced corrosion resistance in the optimal alloy (Mohammad Moradi et al. 2022).
3.2 Issues and challenges of non-biodegradable bioimplants
Design errors, manufacturing defects, or metallurgical issues with the material itself may all contribute to the failure of an implant. Existing manufacturing and metallurgical difficulties might, nevertheless, merely contribute marginally to the occurrence of implant failures. Metals are exposed to the saltwater environment once they have entered the body, a condition that can result in corrosion. An example of an area prone to fretting corrosion is the region surrounding the screw head in contact with the apertures in the plate, where oxygen levels are reduced. Material degradation caused by corrosion may result in an untimely failure of the device. Upon removal of the plate and screws after the healing of the fracture, discernible corrosion appears on the metal surface (Nunamaker 2019). Figure 19 illustrates a schematic of failure due to fretting.
Schematic showing the occurrence of fretting corrosion.
Recent research has frequently associated the short-term durability of artificial joints with stress shielding, a property of bioimplant materials whose Young’s modulus is a determining factor. In the case of titanium alloys, their wear resistance is relatively low, which limits their use in bearing surfaces. On the other hand, stainless steel also has shortcomings, such as low impact resistance, i.e., several small dents and scratches on the bearing surfaces due to small and hard particles hitting it. This results in an escalation of surface roughness, which amplifies the wear rate and generates substantial quantities of wear debris.
Recently, many industrial manufacturers and researchers have been concerned about the harmful complications associated with the implantation of metal-on-metal artificial joints and hip prostheses because they produce large amounts of metal ions when they start to wear, and these ions are likely to cause cancer or allergic reactions in patients after a few years after implantation (Shen et al. 2018). In dental applications, the incompatibility of the modulus of elasticity of the human jaw (about 10–20 GPa) and titanium implants (about 110 GPa) due to a significant difference of about 100 GPa is a problem that is always present in almost all forms of implants. This large difference in Young’s modulus causes the stress shielding effect that, over time, loosens the implant and causes it to fail. The shortcomings of metallic materials cause complications in the design of implants, which can lead to irreparable implants. For example, a dental drill cannot be used for restoration to drill a rigid titanium implant body. To facilitate reimplantation, it is critical to eliminate the adjacent tissues in conjunction with the implant. Prolonged contact with the intricate external oral environment can potentially induce bacterial infection in the tissues encompassing dental implants. Plaque accumulation on the dental implant neck has been demonstrated in several clinical cases; this, in conjunction with other toxins, destroys the osseointegration interface established by the surrounding tissues. Moreover, this will induce inflammation in the adjacent tissues, ultimately leading to the termination of the osseointegration interface; this constitutes a significant factor in the implant’s failure. (Jiang et al. 2020). It has been reported that, unlike biodegradable scaffolds, using inert metallic scaffolds for tissue regeneration is somewhat challenging, and post-processing is required to modify and control the surface chemistry (Xiao et al. 2018).
4 Biodegradable implants
Biodegradable implants represent a significant advancement in medical technology by offering temporary support, reducing the need for follow-up surgeries, and minimizing long-term complications associated with permanent devices. These implants are designed to break down naturally in the body to allow the surrounding tissue to heal without the risk of long-term foreign body reactions. In the last few years, there has been a lot of research done about biodegradable metals that are used for medical implants on account of their remarkable biocompatibility and capability to degrade over time, which is especially useful for many biomedical applications in which the bio-implant is needed temporarily. Biodegradable implants, like magnesium alloy implants, degrade after a specific time and are expelled from the human body via urination. This eliminates the necessity for an additional surgical procedure to remove the implant once it has fulfilled its intended purpose, thus making biodegradable implants not only patient-convenient but also cost-effective. This is the fundamental reason there is a paramount focus on biodegradable implants.
Non-biodegradable implants have been utilized in various orthopedic and dental applications for over a century. However, a new era of bioimplants has been ushered in with the introduction of biodegradable metals like Zirconium, magnesium, and calcium. When implementing biodegradable implants, the rate of corrosion or degradation is a critical consideration. In some cases, the implant must be removed from the body after a brief duration, and a rapid degradation rate is required. On the contrary, an implant characterized by a reduced degradation rate would be more appropriate if an extended presence within the body is needed. This is why numerous studies have examined the degradation rates of different biodegradable implants. Below is a detailed discussion on the common biodegradable materials, their degradation rates and applications observed in various studies.
4.1 Common biodegradable materials
4.1.1 Magnesium and its alloys
Biodegradable implants represent a significant advancement in medical technology by offering temporary support, reducing the need for follow-up surgeries, and minimizing long-term complications associated with permanent devices. These implants are designed to break down naturally in the body, allowing the surrounding tissue to heal without the risk of long-term foreign body reactions. Increased magnesium ion levels in the human body have underscored the potential of magnesium alloys as a highly attractive alternative for implant materials, due to their favorable biocompatibility, biodegradability, and mechanical properties that closely mimic natural bone tissue. Magnesium releases corrosion products through urine that are both soluble and non-toxic within the body. This remarkable attribute of magnesium renders it the most optimal option for biotoxicity. In addition, magnesium alloys have the potential to improve many mechanical properties, including resistance to corrosion and wear (Ramalingam et al. 2020; Tsakiris et al. 2021). Mg-Al alloys are easily castable and possess superior mechanical characteristics and corrosion resistance. On the other hand, investigations indicate that neurodegenerative issues may be linked to magnesium-aluminum alloy degradation within the human body. This discovery is consistent with other research that has established aluminum in potable water as a neurodegenerative agent in human subjects. Mg-Zn alloys are well-suited for biomedical applications as zinc is a vital component for many physiological processes within the human body, making their degradation negligible. The Mg-Ca alloy exhibits considerable potential as a biomedical material on account of Ca’s pivotal role as a metallic component in bone regeneration. However, an excessive amount of Ca may reduce the alloy’s resistance to corrosion. Studies indicate that the concentration of Ca should remain below 1 % for optimal corrosion resistance. Table 7 illustrates various Mg alloys, their corresponding applications, and the effect that various alloying elements have on the mechanical properties (Chen et al. 2018). Presented in Table 8 are the commonly used Mg alloys and their mechanical properties, while Figure 20 presents the flow chart illustrating the effect of various alloying elements on Mg alloys.
Biocompatibility and practical uses of prevalent biodegradable magnesium alloys (Chen et al. 2018).
| Alloys | Biocompatibility | Applications |
|---|---|---|
| Mg-Ca | Healing of the bones | Orthopedic applications |
| Mg-Zn | Zn is an essential element that plays a critical role in bodily functions such as the senses of taste and smell | Repairing the intestinal tract with suture materials and bile |
| Mg-Sr | Promotes osteoblast maturation | Skeletal applications |
| Mg-Si | Formation of connective tissue and bone | Promotes bone repair |
| Mg-Al | Degradation of Al causes neurological disorders | Orthopedic applications |
| Mg-Y | Excellent biocompatibility | Stent and screw |
| Mg-Nd | Excellent biocompatibility | Stent and screw |
Mechanical properties of Mg alloys.
| Alloy | Elastic modulus (GPa) | Yield strength (MPa) | Tensile strength (MPa) |
|---|---|---|---|
| AZ91 | 55.3 | 215.6 | 343.3 |
| AM50 | 45.87 | 122.8 | 236.9 |
| Mg-Zn-Y-Nd | 229 | 159 | |
| AM60 | 45 | 130 | 240 |
| AZ31 | 45.30 | 203 | 294 |
| AZ61 | 43 | 105 | 130 |
| ZK60 | 44.8 | 237 | 312 |
| AZ80 | 39 | 275 | 329 |
Schematic indicating the different effects of various alloying elements on magnesium.
Recent studies have also focused on the development of hybrid alloys beyond the binary alloy/composite compositions. A study by Kumar et al. (Kumar and Pandey 2021) on the effect of ultrasonic assisted sintering on the mechanical and degradation properties of a hybrid composite of Mg-Nb-Zn-Ca alloy have shown that the chemical composition of the corrosion products from the EIS test of the alloy was similar to that of human bones, which indicates a favorable environment for the recovery of bone tissues.
4.1.2 Zinc and its alloys
Zinc, similar to magnesium, is regarded as one of the body’s most vital nutrients and elements. The daily recommended zinc intake for adults is around 10–15 mg. Zinc degrades slower than magnesium, so its dissolution is not considered a health risk. Consequently, Zn and its alloys are being developed as a new category of biodegradable materials for use in the manufacturing of bioimplants. Furthermore, zinc is showing promise as an alternative and competitor to magnesium and its alloys in the biomedical sector, particularly in the domains of orthopedic regeneration and cardiovascular therapy. Zinc in its pure form is, nevertheless, unsuitable for a wide range of applications due to its substandard mechanical properties, including but not limited to tensile strength and hardness. Alloying this substance with elements such as Ca, Sr, and even Mg is crucial for improving its mechanical properties and adjusting the degradation rate (Kabir et al. 2021; Kubásek and Vojtěch 2012). It has been reported that zinc alloyed with magnesium leads to enhanced strength, and zinc alloyed with Cu and Mn results in improved flexibility (Gui et al. 2019). Another study suggests that the nutrient alloying components, including Mg, Cu, Mn, Sr, and Ca, hold significant importance for human health and must be considered the primary and best alloying elements for biomedical purposes when zinc is being alloyed (Aghion et al. 2012).
Guo et al. (2020) investigated the corrosion performance of the Zn-Zr alloy was conducted in the corrected SBF. Three samples labeled as R1, R2, and R3 were utilized in the study using PDP, as shown in Figure 21, with varying zirconium weight percentages (0.5 %, 0.8 %, and 1.1 %, respectively). Based on the data and Figure 22, it is evident that the R2 alloy exhibits superior corrosion resistance in SBF. This is due to its high polarization resistance value and a corrosion potential higher than R1 and similar to R3, as shown in Table 9. The authors note that increasing the amount of Zr results in higher corrosion potential for the Zn-Zr alloy. They also concluded that after 90 days of implantation, the R2 alloy remained mechanically intact and exhibited outstanding biocompatibility. Mechanical characteristics of the commonly used zinc alloys are illustrated in Table 10. Figure 23 illustrates a schematic diagram indicating the different effects of various alloying elements on zinc.
Potentiodynamic graph of R1, R2, and R3 in corrected SBF (Guo et al. 2020) (reproduced with permission from Wiley).
Bar graph showing the corrosion rates of R1, R2, and R3 immersed in corrected SBF for different numbers of days (Guo et al. 2020) (reproduced with permission from Wiley).
PDP test results of R1, R2 & R3 (Guo et al. 2020).
| Material | Ecorr (V versus SCE) | Rp (kΩ cm2) |
|---|---|---|
| R1 | −1.222 | 1.833 |
| R2 | −1.210 | 2.201 |
| R3 | −1.208 | 1.970 |
Mechanical properties of Zn alloys.
| Alloy | Modulus (GPa) | Yield strength (MPa) | Tensile strength (MPa) |
|---|---|---|---|
| Zn-8Al | 85.5 | 206 | 255 |
| Zn-4Mn | 290 | ||
| Zn-1Mg-0.1Mn | 195 | 300 | |
| Zn-Cu-Ti | 159 | 241 | |
| Zn-Cu | 241 | 331 |
Schematic indicating the different effects of various alloying elements on zinc.
Dayanidhi et al. (Pathak and Pandey 2021) evaluated the in vitro corrosion behavior of zinc-hydroxyapatite and zinc-hydroxyapatite-iron composites to highlight their potential as biodegradable materials in biomedical applications. They employed various analytical methods, such as energy dispersive X-ray spectroscopy (EDX) and X-ray diffraction (XRD), to examine the corrosion products and surface morphology of the samples after immersion in simulated body fluid (SBF) for periods of 3, 15, and 31 days. The Results indicated that the corrosion products wprimarily consisted of Zn, O, and P, with significant variations depending on the composition of the samples. Notably, the addition of hydroxyapatite and iron influenced the corrosion rates and the nature of the corrosion products, suggesting a complex interaction that could enhance the material’s suitability for orthopedic applications.
4.1.3 Fe and its alloys
As an additional element, Fe demonstrates potential as a biodegradable bioimplant material. Implants composed of pure Fe and its alloys deteriorate without releasing hydrogen, and their mechanical qualities, which are comparable in strength to stainless steel, surpass those of pure Mg and its alloys. Nevertheless, the degradation rate is comparatively slower than that of magnesium and its alloys. This characteristic renders the implants unsuitable for numerous applications and may give rise to complications akin to those encountered with permanent implants combined with their ferromagnetic characteristics. Consequently, elements such as C, Mn, Si, and so forth are alloyed with pure Fe to circumvent such complications. (Gorejová et al. 2019; Zheng et al. 2014).
Biodegradable Fe and its alloys have found application in the fabrication of bone implants, cardiovascular stents, and other devices. When non-biodegradable stents are utilized, “restenosis” (narrowing of the artery) has been documented to occur. Nevertheless, this issue is rendered moot when biodegradable stents composed of Fe are utilized; hence, an additional surgical procedure to extract the stent is unnecessary. As previously stated, non-biodegradable bone implants are poisonous and necessitate an additional surgical procedure for removal during bone-related applications. Hence, biodegradable Fe material can be employed for these specific applications due to its exceptional biocompatibility and the obviation of the need for implant removal, which is a costly endeavor (He et al. 2016). Wang et al. (2017) investigated the corrosion characteristics of pure Fe and Fe-Ga alloys in SBF. The authors deduced from the OCP plot that the potential (V/SCE) value of pure Fe experiences an early decline of −0.36 to −0.47 over the initial 2000 s, followed by stabilization at 0.49 (V/SCE) after 3,600 s, as illustrated in Figure 24. They hypothesized that this decrease in OCP as immersion duration increased showed that the passive coating on the surface had deteriorated. The OCP plots of Fe-Ga alloys exhibit similar values to those of pure Fe but are marginally higher, which suggests that the surface passivation layers are comparatively more efficient. Figure 25 shows the PDP curve of pure Fe and Fe-Ga alloys submerged in SBF at 37 °C.
The OCP graph of pure Fe and Fe-Ga alloys submerged in 37 °C SBF (Wang et al. 2017) (reproduced with permission from Elsevier).
PDP graph of pure Fe and Fe-Ga alloys submerged in SBF at 37 °C temperature (Wang et al. 2017) (reproduced with permission from Elsevier).
Recently, the development of iron scaffolding via additive manufacturing has gained attractions as potential biomaterial in bone tissue engineering. A study by Sharma et al. (Sharma et al. 2020) investigated the development and characterization of topologically ordered porous iron scaffolds that were designed for bone tissue engineering applications. The key findings include the successful fabrication of scaffolds with two distinct unit cell structures of truncated octahedron and cubic structures which has interconnected porosity ranging from 50.70 % to 80 %. Their research highlights that the topology significantly influences the degradation properties, with increased weight loss and corrosion rates observed in scaffolds with larger strut sizes. Specifically, the truncated octahedron scaffolds demonstrated a maximum corrosion rate of 1.64 mm/y and a weight loss of 6.4 %. In terms of cytocompatibility, reduced cell viability was noted, particularly with increased strut sizes and longer culture times, indicating some cytotoxicity. However, all scaffold samples exhibited excellent hemocompatibility with hemolysis ratios below the acceptable limit and strong anti-platelet adhesion properties.
4.2 Degradation rate of biodegradable implants (*in vivo tests)
Walker et al. (2012) conducted a study involving the implantation of pure magnesium in rats to determine its degradation rate using the weight loss technique at different evaluation intervals (7, 14, and 21 days). The findings indicated that the corrosion rate for the sample removed after seven days was consistent with that of 14 days (0.390 mm/year). However, after 21 days, the removed sample revealed a decline in corrosion rate of 0.221 (mm/y). In their research, Qin et al. (2015) evaluated the degradation rate of an MgNdZnZr alloy by inserting it into a rat. Adding Nd to the MgNdZnZr alloy enhances its corrosion resistance and strength. Additionally, it has been reported that including Zn and Zr enhances the alloy’s strength. Using the weight loss method, researchers discovered that the JDBM alloy degrades at an average rate of 0.092 (mm/year). In their own investigation, Xiao et al. (2018) examined how rapidly degradation occurs techniques in Zn-0.05 Mg alloy using PDP and weight loss techniques in Zn-0.05 Mg alloy using PDP and weight loss techniques, as illustrated in Figures 26 and 27, respectively. Furthermore, in vivo tests were conducted by placing the implants inside a rabbit. The method revealed a uniform degradation rate of 0.15 (mm/y). Razavi et al. (2015) conducted a study to determine the degradation rate of the AZ91 alloy through the weight loss method. An AZ91 alloy was implanted inside a rabbit and removed after two months. The researchers found that the weight loss was 25 (mg/cm2). Li et al. (2017b) examined the degradation rate of the MgZnZr alloy employing the method of volume loss. Both studies provide valuable insights into the degradation properties of different alloys. They implanted the MZZ alloy inside a rabbit and observed the corrosion rates. The implant demonstrated a degradation rate of 0.826 mm per year in the first month. The rate decreased to 0.697 mm/y in the second month but increased again to 0.929 mm per year by the third month (Li et al. 2017b). Huehnerschulte et al. (2011) researched the corrosion rate of ZEK100 alloy by using the volume loss method. The ZEK100 implants were inserted into a rabbit for three and six months. The corrosion rate was determined to be 0.067 (mm/year) for the 3-month group and 0.154 (mm/y) for the 6-month group. Aghion et al. (2012) investigated the corrosion rates of two magnesium alloys, namely Mg-Nd-Y-Zr (EW10) and Mg-Nd-Y-Zr-Ca (EW10X04). The study results do not present subjective evaluations and adhere to conventional academic structure, language, and style guides. The study employed the volume loss method to compute the degradation rate, and the two alloys were implanted in rats for two periods (6 and 12 weeks). After six weeks, the corrosion rate of EW10 was calculated to be 0.23 mm/year, while after 12 weeks, it was found to be 0.16 mm/year. The corrosion rate (CR) of EW10X04 was measured to be 0.16 (mm/year) after 6 weeks and 0.14 (mm/year) after 12 weeks (Aghion et al. 2012). Kraus et al. (2012) investigated the corrosion performance of the Mg-Zn-Ca-Mn alloy using the volume loss method in a rat over 24 weeks. The average degradation rate was determined to be approximately 1.58 (mm/y). Liu et al. (2019) performed research on the corrosion rate of Mg-30Sc alloy implanted in rats for 24 weeks. The weight loss method was utilized to determine the corrosion rate, which was calculated to be 0.06 mm/y. Meanwhile, Li et al. (2015) the degradation rate of three alloys – Zn-1Mg, Zn-1Sr, and Zn-1Ca – was examined using the weight loss technique, with pure Zn serving as the reference alloy. The Zn-1Mg alloy was implanted in a rat to evaluate its corrosion rate. Results showed that Zn-1Mg had a 0.17 mm/year corrosion rate, whereas Zn-1Sr and Zn-1Ca had corrosion rates of 0.22 mm/year and 0.19 mm/year, respectively.
Potentiodynamic graph of Zn and Zn-0.05 Mg (Xiao et al. 2018) (reproduced with permission from Elsevier).
Graph depicting the corrosion rate of the immersion test (Xiao et al. 2018) (reproduced with permission from Elsevier).
4.3 Degradation rate of biodegradable implants (*in vitro tests)
In a study conducted by Wang et al. (2012), they employed pure Mg alloy to observe its degradation rate in SBF. The results showed that after three days, the corrosion rate was 2.5 (mm/y), which reduced to 1.33 (mm/y) after seven days and continued to decrease to 0.4 (mm/y) after 30 days. Similarly, Witte et al. (2007) researched the corrosion performance of Mg-HA alloy through immersion and electrochemical tests. Artificial seawater and cellular solutions, with and without protein, were used for testing. Results demonstrated that the corrosion rate was 2.0–3.2 mm/y during immersion testing and 1.25 ± 0.16 mm/y during electrochemical testing. A study by Liu et al. (2019) delved into the corrosion performance of Mg-30Sc alloy via immersion in Hank’s solution for 10 days. After a ten-day immersion period, a 2.9 ± 0.1 mm/y corrosion rate was determined. Gui et al. (2019) under as-cast and as-extruded conditions, investigated to assess the corrosion performance of GZKM-1, an Mg-3.0Gd-2.7Zn-0.4Zr-0.1Mn alloy in Hank’s solution. The OCP and PDP curves are illustrated in Figure 28. The alloys underwent a 240-h immersion, during which it was observed that the corrosion rate for as-extruded GZKM-1 was 0.34 ± 0.09 (mm/y), while for as-cast GZKM-1, it was 0.38 ± 0.23 (mm/y). Shangguan et al. (2016) investigated the corrosion performance of different MgSr (cast and extruded) alloys using Hank’s solution for 24 h. In contrast to the as-extruded alloy, which exhibited a corrosion rate of 0.04 mm/y, the as-cast alloy exhibited a corrosion rate of 0.15 mm/y. Yao et al. (2018) researched the corrosion rate of Mg2.0Zn0.5Zr3.0Gd alloy in SBF by conducting immersion and electrochemical tests.
OCP and PDP graphs of GZKM-1 as-cast and as-extruded (Gui et al. 2019) (reproduced with permission from Elsevier).
After immersion testing, they discovered a corrosion rate of 0.417 (mm/y), and after electrochemical testing, a corrosion rate of 0.116 ± 0.002 (mm/y) was found (Yao et al. 2018). Chen et al. (2017) conducted a comprehensive study on the ZK60 magnesium alloy, analyzing various properties, including its corrosion rate. The alloy underwent corrosion testing in Hank’s solution, with results indicating a corrosion rate of 0.247 ± 0.0019 mm/y for the as-cast ZK60 alloy and 0.494 ± 0.0025 mm/y for the as-extruded ZK60 alloy. Ibrahim et al. (2017) studied Mg-1.2Zn-0.5Ca alloy and investigated the effects of different aging durations on the alloy’s corrosion rate. The results showed that after being heat treated for 2 h, the alloy exhibited the minimum corrosion rate value of 4.4 mm/y in the immersion test, as well as the lowest corrosion rate value of 8.9 mm/y in the electrochemical test. The results showed that after being heat treated for 2 h, the alloy exhibited the minimum corrosion rate value of 4.4 mm/year in the immersion test, as well as the lowest corrosion rate value of 8.9 mm/y in the electrochemical test. Li et al. (2017a) investigated various properties, such as corrosion performance, of Mg-Zn-Sr, Mg-Zn-Zr, and Mg-Zn-Zr-Sr alloys by immersing them in SBF for 15 days. It was observed that the corrosion rate of Mg-Zn-Sr alloy was 4.07 ± 0.71 (mm/y), while it was 5.357 ± 0.75 (mm/y) for Mg-Zn-Zr-Sr alloy and 6.92 ± 0.3 (mm/year) for Mg-Zn-Sr alloy. Figure 29 shows the weight loss of the alloys with time, while Figure 30 shows the corresponding corrosion rate with time.
Weight loss graph of all the three alloys (Li et al. 2017a) (reproduced with permission from Elsevier).
Corrosion rate graph of all the three alloys (Li et al. 2017a) (reproduced with permission from Elsevier).
5 Recommendations for future research
The field of implant technologies, which includes both biodegradable and non-biodegradable types, is on the verge of making substantial progress via focused research and development. This review provides detailed suggestions to advance the next phase of implant innovation, focusing on improving effectiveness, patient-centered outcomes, and safety across all areas. Enhancing the corrosion resistance of materials is vital to mitigate the substantial effects of corrosion on the longevity of implants. This entails the development of alloys, coatings, or surface treatments that can endure the corrosive conditions present in the human body. Research should be conducted to explore methods for reducing fretting corrosion, especially in regions with reduced oxygen levels, such as the vicinity of screw heads. This may include the development of implants with enhanced geometric characteristics or coatings to mitigate the effects of fretting corrosion. Investigation into sophisticated surface coatings that might improve the durability of metallic implants is vital, particularly for materials such as Titanium alloys and stainless steel. The sustained effectiveness of implants must provide coatings that reduce surface roughness and the generation of wear particles.
There is a need to investigate the incorporation of intelligent technologies into implants for continuous monitoring. Timely intervention and prevention of premature failures may be facilitated by real-time data about corrosion, wear, and structural integrity, which enables early identification of possible concerns. The disparity between the modulus of elasticity of human jawbones and metallic implants, especially in dental contexts, requires attention. Implant designs or innovative materials that minimize stress shielding effects may effectively avoid loosening and failure over time. It is essential to explore materials and designs that facilitate the repair and restoration of implants, particularly those made of hard titanium. Enhancing the repairability of implants may extend their lifespan and decrease the need for intrusive operations during reimplantation. It is advisable to investigate using antibacterial coatings to reduce the likelihood of bacterial infections occurring near implants. This is especially important in dental applications since prolonged contact with the oral environment increases the risk of infection. It is necessary to carry out extensive clinical studies to evaluate the long-term efficacy of both biodegradable and non-biodegradable implants. This study will provide significant data on practical efficacy, potential problems, and patient results linked to various implant materials. It is necessary to examine the practicality and advantages of implant designs tailored to individual patients. Personalized implants can reduce problems associated with differences in stiffness and improve overall compatibility, resulting in improved implant success rates. Optimal degradation rates for biodegradable implants should be determined according to the unique requirements of medical applications. The degradation profiles should be customized to correspond with the desired implantation period, balancing the need for surgical removal and the implant’s life.
Efforts should be directed toward developing regulatory solid toward developing solid regulatory standards and guidance for implant materials. Implementing thorough assessment criteria would improve the safety and effectiveness of both biodegradable and non-biodegradable implants. It is advisable to allocate funds towards educational programs that aim to enhance knowledge and understanding among patients, healthcare professionals, and industry stakeholders on the intricacies, developments, and safety precautions related to implant technology.
To summarize, future research should prioritize tackling material obstacles, boosting implant monitoring, and optimizing the safety and efficacy of both biodegradable and non-biodegradable implants. This interdisciplinary methodology will enhance the advancement of implants that provide extended longevity, diminished problems, and enhanced patient results.
6 Conclusions
The human body possesses a remarkable ability to self-heal in certain instances; however, many conditions, whether arising from disease, injury, or accidents, exceed the body’s natural healing capacity and necessitate external medical intervention. In such cases, implants – both biodegradable and non-biodegradable – serve as critical tools for restoring lost function by closely mimicking the structure and mechanical properties of the affected tissues or body parts. Metallic implants, specifically, have been widely used in clinical settings for decades due to their mechanical strength, durability, and versatility.
This review has provided a comprehensive exploration of the role of metallic implants, with a particular focus on the distinction between biodegradable and non-biodegradable materials. Non-biodegradable metallic implants, such as stainless steel and titanium alloys, have been indispensable in orthopedic and dental applications due to their high strength, corrosion resistance, and long-term stability. However, the permanence of these materials often leads to complications, such as the need for secondary surgeries for removal, long-term foreign body reactions, and potential implant failure over time. In contrast, biodegradable metallic implants, primarily based on magnesium, iron, and zinc alloys, offer a promising alternative. These materials are designed to gradually degrade and be absorbed by the body, eliminating the need for removal and reducing the risk of long-term complications. This review has highlighted the corrosion and degradation mechanisms of these materials, emphasizing their potential to revolutionize the field of medical implants. Biodegradable implants not only provide mechanical support during the critical healing phase but also offer a more patient-friendly approach by naturally disappearing once their function is fulfilled.
Despite these advantages, the widespread adoption of biodegradable implants is still hindered by challenges, particularly the need to finely tune their degradation rates. Implants that degrade too quickly may fail to provide sufficient support during healing, while those that degrade too slowly may persist longer than necessary, defeating the purpose of biodegradability. The ability to control the degradation behavior of these materials in various physiological environments remains a key area of ongoing research.
In summary, biodegradable metallic implants represent the future of implant technology, offering solutions to the drawbacks associated with traditional non-biodegradable metals. However, for biodegradable materials to reach their full potential, further research is needed to address the complexities of degradation kinetics, material performance, and biocompatibility. The insights from this review contribute to bridging the gap between current clinical practices and the future of implantable medical devices, advancing both research and practical applications in the field. With continued advancements, biodegradable implants hold the promise of becoming a cornerstone in improving patient’s experience with implants and reducing the burdens associated with long-term implant use.
Acknowledgments
The authors appreciate and acknowledge the support provided by the Mechanical Engineering Department and the Interdisciplinary Research Center for Advanced Materials, King Fahd University of Petroleum & Minerals (KFUPM) for conducting this research.
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Research ethics: Not applicable.
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Informed consent: Not applicable.
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Author contributions: All authors have accepted responsibility for the entire content of this manuscript and approved its submission.
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Use of Large Language Models, AI and Machine Learning Tools: None declared.
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Conflict of interest: The authors state no conflict of interest.
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Research funding: None declared.
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Data availability: Not applicable.
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Articles in the same Issue
- Frontmatter
- Review
- Exploring the future of metallic implants: a review of biodegradable and non-biodegradable solutions
- Original Articles
- Exploring the inhibitive performance of amino-dicarbonitrile derivative for Q235 steel protection in Cl3CCOOH solution: effect of temperature and synergism with iodide ions
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