Acta Metallurgica Sinica (English Letters) ›› 2024, Vol. 37 ›› Issue (2): 213-241.DOI: 10.1007/s40195-023-01631-7
J. Sharath Kumar1, Rakesh Kumar1, Rajeev Verma1()
Received:
2023-07-22
Revised:
2023-09-25
Accepted:
2023-09-27
Online:
2024-02-10
Published:
2024-02-27
Contact:
Rajeev Verma, J. Sharath Kumar, Rakesh Kumar, Rajeev Verma. Surface Modification Aspects for Improving Biomedical Properties in Implants: A Review[J]. Acta Metallurgica Sinica (English Letters), 2024, 37(2): 213-241.
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Implant material | Typical applications | Advantages | Disadvantages | Mechanical properties | ||
---|---|---|---|---|---|---|
Young modulus (GPa) | Tensile strength (MPa) | Hardness (HV) | ||||
Natural bone [ | - | - | - | 3-20 | 130-180 | 40 |
Co-Cr alloys [ | Prostheses stems, load-bearing components in total joint replacement | High wear resistance Corrosion resistance Easily castable | Hypersensitivity and inflammatory reactions are caused due to wear debris Higher modulus elasticity leads to bone resorption | 210-240 | 655-1836 | 450 |
SS [ | Fracture plates, screws, hip nails | High wear resistance Corrosion resistance Fatigue resistant | High friction coefficient, wear debris generation, aseptic loosening | 190-205 | 586-1351 | 190 |
Ti and Ti alloys [ | Hip and knee replacement prostheses, screws, and pins for bone fixation | Biocompatible Low stiffness Corrosion resistance | Poor tribological properties, low abrasive resistance, low mechanical stability, allergic reactions | 105-114 | 760 | 340 |
Zirconium alloys [ | Replacement of hip, knee, teeth tendons, ligaments, and bone filers | Biocompatible, Higher hardness | Bone resorption, higher wear rate, higher corrosion | 150-199 | 200-495 | 1000-3000 |
Mg alloys [ | Stents, Fracture plates, Screws | Biodegradable Mechanical properties equal to cortical bone | High corrosion rate, formation of H2 gas, and premature loss of mechanical integrity | 41-45 | 200 | - |
Iron [ | Fracture plates, screws, thin-walled stents | Biodegradable Good strength and formability No gas generation | Corrosion rate is too slow, and | 210 | 370 | - |
Zn alloys [ | Fracture plates, Screws | Biodegradable Low reactivity in molten state Corrosion rate less than Mg | Poor mechanical properties, age hardening | 75 | 90 | - |
UHMWPE [ | Joint arthroplasty | Biocompatible, Abrasion resistance | Submicron particles of wear debris may contribute to osteolysis and aseptic loosening of the bone | 0.9-2.7 | 53 | 62-66 |
Table 1 Comparison of traditional implant material and their properties
Implant material | Typical applications | Advantages | Disadvantages | Mechanical properties | ||
---|---|---|---|---|---|---|
Young modulus (GPa) | Tensile strength (MPa) | Hardness (HV) | ||||
Natural bone [ | - | - | - | 3-20 | 130-180 | 40 |
Co-Cr alloys [ | Prostheses stems, load-bearing components in total joint replacement | High wear resistance Corrosion resistance Easily castable | Hypersensitivity and inflammatory reactions are caused due to wear debris Higher modulus elasticity leads to bone resorption | 210-240 | 655-1836 | 450 |
SS [ | Fracture plates, screws, hip nails | High wear resistance Corrosion resistance Fatigue resistant | High friction coefficient, wear debris generation, aseptic loosening | 190-205 | 586-1351 | 190 |
Ti and Ti alloys [ | Hip and knee replacement prostheses, screws, and pins for bone fixation | Biocompatible Low stiffness Corrosion resistance | Poor tribological properties, low abrasive resistance, low mechanical stability, allergic reactions | 105-114 | 760 | 340 |
Zirconium alloys [ | Replacement of hip, knee, teeth tendons, ligaments, and bone filers | Biocompatible, Higher hardness | Bone resorption, higher wear rate, higher corrosion | 150-199 | 200-495 | 1000-3000 |
Mg alloys [ | Stents, Fracture plates, Screws | Biodegradable Mechanical properties equal to cortical bone | High corrosion rate, formation of H2 gas, and premature loss of mechanical integrity | 41-45 | 200 | - |
Iron [ | Fracture plates, screws, thin-walled stents | Biodegradable Good strength and formability No gas generation | Corrosion rate is too slow, and | 210 | 370 | - |
Zn alloys [ | Fracture plates, Screws | Biodegradable Low reactivity in molten state Corrosion rate less than Mg | Poor mechanical properties, age hardening | 75 | 90 | - |
UHMWPE [ | Joint arthroplasty | Biocompatible, Abrasion resistance | Submicron particles of wear debris may contribute to osteolysis and aseptic loosening of the bone | 0.9-2.7 | 53 | 62-66 |
References | Material | Implant | Survival time | Type of failure |
---|---|---|---|---|
Shahemi et al. [ | UHWMPE | Acetabular cup of hip replacement | - | Failed due to aseptic loosening from wear generation 10 mm thickness of the cup was reduced to 2.3 mm (min) and 8.8 mm (max) |
Affatato et al. [ | Zirconium | Femoral head of joint replacement | The mean life of implants 6 years | In the study, most femoral heads failed due to aseptic loosening |
Gervais et al. [ | SS 316L | Bone-locking compression plates | - | High fatigue cycle, various stress risers in plate design, poor bone condition |
Paliwal et al. [ | Ti-6Al-4V & Co-Cr-Mo | Femoral hip prosthesis | Case i. 38 months Case ii. 28 months Case iii. 18 months | Two of the three failed catastrophically at the stem junction All implants failed due to fretting, pitting, plastic deformation, and stress-induced corrosion cracks There was evidence of metal ions released due to corrosion |
Magnissalis et al. [ | Ti6Al4V (porous coated) | Femoral stems | 24 months | Microcracks formed on the surface due to high stress Due to porous coating, 75% of fatigue strength is reduced compared to uncoated |
Shahgaldi et al. [ | SS 316L | Femoral Nail plate | 2.5 years | Wear, fatigue, stress corrosion |
Table 2 Case study reports of failed implants after surgery
References | Material | Implant | Survival time | Type of failure |
---|---|---|---|---|
Shahemi et al. [ | UHWMPE | Acetabular cup of hip replacement | - | Failed due to aseptic loosening from wear generation 10 mm thickness of the cup was reduced to 2.3 mm (min) and 8.8 mm (max) |
Affatato et al. [ | Zirconium | Femoral head of joint replacement | The mean life of implants 6 years | In the study, most femoral heads failed due to aseptic loosening |
Gervais et al. [ | SS 316L | Bone-locking compression plates | - | High fatigue cycle, various stress risers in plate design, poor bone condition |
Paliwal et al. [ | Ti-6Al-4V & Co-Cr-Mo | Femoral hip prosthesis | Case i. 38 months Case ii. 28 months Case iii. 18 months | Two of the three failed catastrophically at the stem junction All implants failed due to fretting, pitting, plastic deformation, and stress-induced corrosion cracks There was evidence of metal ions released due to corrosion |
Magnissalis et al. [ | Ti6Al4V (porous coated) | Femoral stems | 24 months | Microcracks formed on the surface due to high stress Due to porous coating, 75% of fatigue strength is reduced compared to uncoated |
Shahgaldi et al. [ | SS 316L | Femoral Nail plate | 2.5 years | Wear, fatigue, stress corrosion |
Fig. 8 Hip implant femoral stems stress distributions: a without any porous (fully dense), b femoral stem with 40% porous structure, c femoral stem with axial graded porous structure increased distally, d femoral stem with radial graded porous structure increased inwardly [61]
Fig. 9 Corrosion on a surface: (1) on the surface of an aqueous environment, the metal dissolves and cations are eliminated (oxidation); (2) remaining electrons are drawn to a differential charge at another surface location (reduction); and (3) metal-oxide or metal hydroxide form as by-products of this reaction [65]
Biomaterials | Effect |
---|---|
Nickel [ | Have an effect on skin (dermatitis) |
Cobalt [ | Anaemia B inhibits iron from being absorbed into the bloodstream |
Aluminium [ | Alzheimer’s disease |
Chromium [ | Ulcers and damage to nervous system |
Vanadium [ | Toxic in the elementary state |
Table 3 Effect on the human body due to increased metal ions
Biomaterials | Effect |
---|---|
Nickel [ | Have an effect on skin (dermatitis) |
Cobalt [ | Anaemia B inhibits iron from being absorbed into the bloodstream |
Aluminium [ | Alzheimer’s disease |
Chromium [ | Ulcers and damage to nervous system |
Vanadium [ | Toxic in the elementary state |
Category | Techniques | Features | References |
---|---|---|---|
Physical methods | Shot Penning | Simple and low cost Promote attachment the both of tissue cells and bacteria | Jemat et al. [ |
Additive manufacturing | Creating complex 3D structures Material saving | Yuan et al. [ | |
Thermal spraying | Economical and safe | Tang et al. [ | |
Laser treatment | Achieving complex and precise topography | Hindy et al. [ | |
Magnetron sputtering | Strong adhesive and homogeneous coating can produce | Liu et al. [ | |
Friction stir processing | Without melting metal, Uniformity in corrosion | Shunmugasamy et al. [ | |
Chemical methods | Anodizing | An accelerated electrochemical process | Hall et al. [ |
Sol-gel | Low-temperature technique Drugs delivery | Adams et al. [ | |
Alkali treatment | Extending uniformly Do not damage mechanical properties | Yao et al. [ |
Table 4 Categorised surface modification methods and their features [11]
Category | Techniques | Features | References |
---|---|---|---|
Physical methods | Shot Penning | Simple and low cost Promote attachment the both of tissue cells and bacteria | Jemat et al. [ |
Additive manufacturing | Creating complex 3D structures Material saving | Yuan et al. [ | |
Thermal spraying | Economical and safe | Tang et al. [ | |
Laser treatment | Achieving complex and precise topography | Hindy et al. [ | |
Magnetron sputtering | Strong adhesive and homogeneous coating can produce | Liu et al. [ | |
Friction stir processing | Without melting metal, Uniformity in corrosion | Shunmugasamy et al. [ | |
Chemical methods | Anodizing | An accelerated electrochemical process | Hall et al. [ |
Sol-gel | Low-temperature technique Drugs delivery | Adams et al. [ | |
Alkali treatment | Extending uniformly Do not damage mechanical properties | Yao et al. [ |
References | Coating/reinforcement | Substrate | Method | Mechanical properties | Observations |
---|---|---|---|---|---|
Singh et al. [ | Ha-Nb | Mg alloy (ZK60) | Plasma spray | i. Nb-650 HV ii. HA-230 HV iii. HA-10%Nb-240 HV iv. HA-20%Nb-255 HV v. HA-30%Nb-270 HV | Higher Nb thermal conductivity led to a slower cooling rate in HA zones, which helped to form more brittleness |
Abhijith et al. [ | TiNbMoMnFe | 304L SS | HVOF | i. 720 ± 3 HV0.5 (milling time 5 h) ii. 871 ± 5 HV0.5 (10 h) iii. 1021 ± 2 HV0.5 (15 h) | The wettability and surface hardness of the coating increased The corrosion rate decreased while increasing the milling time of powders |
Wang et al. [ | Tantalum | TiO2 nanotubes | Plasma spray (Thickness 100-300 nm) | - | An increase in surface roughness and hydrophilicity was observed Coating helped for enhancing osteogenic activity |
Ebrahimi et al. [ | HA and HA/Al2O3 &SiO2 | Pure Ti | Plasma spray | i. HA coating 420 HV ii. HA/Al2O3 &SiO2 coating 390 HV | Bi-layer coating improved roughness and wettability Microhardness decreased due to an increase in porosity and large grain size Bi-layer coating improved cell viability and proliferation by up to 90% in 72 h |
Stübinger et al. [ | Ti/HA bilayed coating | Polyether ether ketone (PEEK) | Plasma spray | i. Surface roughness 30 ± 3 µm ii. Porosity 56% | Higher bone-to-implant contact rates compare to a single-layer and uncoated Coated implants showed better osseointegration |
Sathish et al. [ | YSZ Al2O3-13 wt%TiO2 Bilayered | Ti-13Nb-13Z | Plasma spray | i. 617 ± 25 HV ii. 820 ± 22 HV iii. 1096 ± 10 HV | The lowest corrosion rate is shown in bilayer coating (0.0005 mm/yr) High porosity in Al2O3-13 wt%TiO2 coating due to higher melting point of alumina |
Xiao et al. [ | FeCoNiCrMn | 304 Stainless steel | Plasma spray | Hardness 273 ± 20 HV | Increasing H2 flow rate, there was a change in microstructure, wear resistance and microhardness Mn and Cr elements formed oxides compared to other elements |
Tüten et al. [ | TiTaHfNbZr | Ti6Al4V | Magnetron sputtering | i. Uncoated 3.46 ± 0.17 GPa ii. HEA-coated 12.51 ± 0.34 GPa | The coating is produced with a fine-grain amorphous structure Coating enhanced mechanical properties (hardness and elastic modulus) |
Meghwal et al. [ | AlCoCrFeNi | Stainless steel | HVOF | i. HEA coating 7 ± 0.6 GPa ii. SS316L 3 ± 0.1 GPa | The HEA-coated sample showed superior corrosion resistance than SS316L Due to pitting corrosion interconnected micro-pits converted into larger holes in Al-Ni-rich regions |
Ahn et al. [ | CoCrFeMnNi | Low alloy steel | CS | i. HEA coating 423 HV ii. Heat treated at 550 °C is 399 HV iii. Heat treated at 850 °C is 220 HV | CS HEA coating showed ultra-fine grains H850 °C grains fully recrystallized (2.35 µm) resulted in a decrease in hardness values |
Peighambardoust et al. [ | Ti1.5ZrTa0.5Nb0.5Hf0.5 | SS 316L Co-Cr-Mo Ti6Al4v | Magnetron sputtering | Vickers hardness values of three substrates after coating i. 11.43 GPa ii. 11.49 GPa iii. 11.45 GPa | SS316L and Co-Cr-Mo implants showed pitting corrosion while Ti alloys did not show any pitting signs. But, more corrosion resistance is shown in Co-Cr-Mo |
Liu et al. [ | Fe25Co25Ni25Al10Ti15 | Ti6Al4V | Magnetron sputtering | - | HEA coating reduced the dry friction coefficient by 37% Porous structure of T64 helped to accommodate friction particles which helped to absorb impact energy Increasing porous size helped to reduce the friction coefficient up to pd 30 μm |
Henao et al. [ | HA-TiO2 | Ti-6Al-4V | HVOF | - | Coating showed more stable electrochemical behaviour than uncoated Porosity of the coating helped to form a bone-like apatite structure after 7 days of immersion in SBF |
Sun et al. [ | Ti -Mg | Ti | CS | Porous size 70-100 μm | Porous Ti coating helped to reduce stress shielding around the bone |
Lynn et al. [ | HA | Ti-6Al-4V | Plasma spray | - | Up to 100 μm of HA coating thickness does not affect fatigue resistance of Ti-6Al-4V substrate Residual stresses did not vary while increasing coating thickness |
Sampath et al. [ | PCL/HA | AZ31 | FSP | - | CaP coating reduced the degradation rate of AZ31 alloy The controlled degradation made the surface stable and conducive to cell adhesion and proliferation |
Shunmugasamy et al. [ | - | EZ33A Mg alloy | FSP | Corrosion rate up to 1 mm/year up to 8 weeks after that 0.7 mm/year next 4 weeks | FSP grain refinement, uniformly distributed secondary phases and preferred basal texture inhibit the corrosion process FSP results showed more uniformity of corrosion in Mg alloys Steady release of Mg ions aids implant-bone interface development |
Table 5 Coated material systems and their properties of biomaterial
References | Coating/reinforcement | Substrate | Method | Mechanical properties | Observations |
---|---|---|---|---|---|
Singh et al. [ | Ha-Nb | Mg alloy (ZK60) | Plasma spray | i. Nb-650 HV ii. HA-230 HV iii. HA-10%Nb-240 HV iv. HA-20%Nb-255 HV v. HA-30%Nb-270 HV | Higher Nb thermal conductivity led to a slower cooling rate in HA zones, which helped to form more brittleness |
Abhijith et al. [ | TiNbMoMnFe | 304L SS | HVOF | i. 720 ± 3 HV0.5 (milling time 5 h) ii. 871 ± 5 HV0.5 (10 h) iii. 1021 ± 2 HV0.5 (15 h) | The wettability and surface hardness of the coating increased The corrosion rate decreased while increasing the milling time of powders |
Wang et al. [ | Tantalum | TiO2 nanotubes | Plasma spray (Thickness 100-300 nm) | - | An increase in surface roughness and hydrophilicity was observed Coating helped for enhancing osteogenic activity |
Ebrahimi et al. [ | HA and HA/Al2O3 &SiO2 | Pure Ti | Plasma spray | i. HA coating 420 HV ii. HA/Al2O3 &SiO2 coating 390 HV | Bi-layer coating improved roughness and wettability Microhardness decreased due to an increase in porosity and large grain size Bi-layer coating improved cell viability and proliferation by up to 90% in 72 h |
Stübinger et al. [ | Ti/HA bilayed coating | Polyether ether ketone (PEEK) | Plasma spray | i. Surface roughness 30 ± 3 µm ii. Porosity 56% | Higher bone-to-implant contact rates compare to a single-layer and uncoated Coated implants showed better osseointegration |
Sathish et al. [ | YSZ Al2O3-13 wt%TiO2 Bilayered | Ti-13Nb-13Z | Plasma spray | i. 617 ± 25 HV ii. 820 ± 22 HV iii. 1096 ± 10 HV | The lowest corrosion rate is shown in bilayer coating (0.0005 mm/yr) High porosity in Al2O3-13 wt%TiO2 coating due to higher melting point of alumina |
Xiao et al. [ | FeCoNiCrMn | 304 Stainless steel | Plasma spray | Hardness 273 ± 20 HV | Increasing H2 flow rate, there was a change in microstructure, wear resistance and microhardness Mn and Cr elements formed oxides compared to other elements |
Tüten et al. [ | TiTaHfNbZr | Ti6Al4V | Magnetron sputtering | i. Uncoated 3.46 ± 0.17 GPa ii. HEA-coated 12.51 ± 0.34 GPa | The coating is produced with a fine-grain amorphous structure Coating enhanced mechanical properties (hardness and elastic modulus) |
Meghwal et al. [ | AlCoCrFeNi | Stainless steel | HVOF | i. HEA coating 7 ± 0.6 GPa ii. SS316L 3 ± 0.1 GPa | The HEA-coated sample showed superior corrosion resistance than SS316L Due to pitting corrosion interconnected micro-pits converted into larger holes in Al-Ni-rich regions |
Ahn et al. [ | CoCrFeMnNi | Low alloy steel | CS | i. HEA coating 423 HV ii. Heat treated at 550 °C is 399 HV iii. Heat treated at 850 °C is 220 HV | CS HEA coating showed ultra-fine grains H850 °C grains fully recrystallized (2.35 µm) resulted in a decrease in hardness values |
Peighambardoust et al. [ | Ti1.5ZrTa0.5Nb0.5Hf0.5 | SS 316L Co-Cr-Mo Ti6Al4v | Magnetron sputtering | Vickers hardness values of three substrates after coating i. 11.43 GPa ii. 11.49 GPa iii. 11.45 GPa | SS316L and Co-Cr-Mo implants showed pitting corrosion while Ti alloys did not show any pitting signs. But, more corrosion resistance is shown in Co-Cr-Mo |
Liu et al. [ | Fe25Co25Ni25Al10Ti15 | Ti6Al4V | Magnetron sputtering | - | HEA coating reduced the dry friction coefficient by 37% Porous structure of T64 helped to accommodate friction particles which helped to absorb impact energy Increasing porous size helped to reduce the friction coefficient up to pd 30 μm |
Henao et al. [ | HA-TiO2 | Ti-6Al-4V | HVOF | - | Coating showed more stable electrochemical behaviour than uncoated Porosity of the coating helped to form a bone-like apatite structure after 7 days of immersion in SBF |
Sun et al. [ | Ti -Mg | Ti | CS | Porous size 70-100 μm | Porous Ti coating helped to reduce stress shielding around the bone |
Lynn et al. [ | HA | Ti-6Al-4V | Plasma spray | - | Up to 100 μm of HA coating thickness does not affect fatigue resistance of Ti-6Al-4V substrate Residual stresses did not vary while increasing coating thickness |
Sampath et al. [ | PCL/HA | AZ31 | FSP | - | CaP coating reduced the degradation rate of AZ31 alloy The controlled degradation made the surface stable and conducive to cell adhesion and proliferation |
Shunmugasamy et al. [ | - | EZ33A Mg alloy | FSP | Corrosion rate up to 1 mm/year up to 8 weeks after that 0.7 mm/year next 4 weeks | FSP grain refinement, uniformly distributed secondary phases and preferred basal texture inhibit the corrosion process FSP results showed more uniformity of corrosion in Mg alloys Steady release of Mg ions aids implant-bone interface development |
Implant surface | Wettability (WCA) | Nature of surface | Surface roughness (nm) | Bacteria | |
---|---|---|---|---|---|
S. aureus | E. coli | ||||
Untreated surface | - | - | 370 ± 40 | No reduction | No reduction |
Mirror polished surface | 31 ± 5 | 82.4% reduction | Bacteria cells doubled | ||
Spikes (pulse energy 19.1 μJ) | 160° | Superhydrophobic | 8600 ± 100 | 69.8% reduction | Bacteria cells tripled |
Laser-induced periodic surface structures (LIPSS) (pulse energy 1.01 μJ) | i. 119° ii. 26° (Treated with hot water at 90 °C for 48 h) | i. Hydrophobic ii. Hydrophilic | 90 ± 5 | i. 84.7% reduction ii. - | i. 99.8% reduction ii. 98.5% reduction |
Nanopillars (pulse energy 1.46 μJ) | 130° | Hydrophobic | 60 ± 5 | 79.9% reduction | 99.2% reduction |
Table 6 Impact of implant surface properties on antibacterial properties [137]
Implant surface | Wettability (WCA) | Nature of surface | Surface roughness (nm) | Bacteria | |
---|---|---|---|---|---|
S. aureus | E. coli | ||||
Untreated surface | - | - | 370 ± 40 | No reduction | No reduction |
Mirror polished surface | 31 ± 5 | 82.4% reduction | Bacteria cells doubled | ||
Spikes (pulse energy 19.1 μJ) | 160° | Superhydrophobic | 8600 ± 100 | 69.8% reduction | Bacteria cells tripled |
Laser-induced periodic surface structures (LIPSS) (pulse energy 1.01 μJ) | i. 119° ii. 26° (Treated with hot water at 90 °C for 48 h) | i. Hydrophobic ii. Hydrophilic | 90 ± 5 | i. 84.7% reduction ii. - | i. 99.8% reduction ii. 98.5% reduction |
Nanopillars (pulse energy 1.46 μJ) | 130° | Hydrophobic | 60 ± 5 | 79.9% reduction | 99.2% reduction |
References | Laser parameters | Dimple size | Tissue cells | Observations |
---|---|---|---|---|
Mirhosseini et al. [ | Pulse energy-100 MJ Frequency-10 HZ Wavelength-1064 nm | Diameter 156, 214, 127 μm | 2T3 osteoblast cells | More cell growth percentage was observed in small-size patterns (127 μm) It found that optimum surface roughness for cell growth, and roughness closer to cell size improved cell growth Cells were uniform and evenly spread across patterns compared to untreated surfaces Untreated samples showed higher wettability than patterned surfaces |
Mukherjee et al. [ | Frequency 1000 HZ Wavelength 50-85 μm Scanning speed 50-80 mm/s Duty cycle 60-80% | Depth 60-65 μm | MG 63 | At the same wavelength by reducing duty cycle, wider peaks were formed and at same duty cycle by reducing wavelength sharp peaks were formed It was observed that samples with similar avg. roughness, contact angle, and protein adsorption also showed different cell viability |
Mesquita-Guimaraes et al. [ | Power-25 W Velocity-200 mm/s frequency-550 Hz | 50 × 50 μm2 width 100 μm Depth 0, 40, 80 μm | MC3T3-E1 osteoblast cells | Laser micro-textured surface showed a 40% increase in cell viability compared to polished surface Optimal pattern size should be 100 to 200 μm, showing 60% more cell viability The most extensive cell layer formation was found on the laser-sintered HAp and 45S5 BG coatings, also by the cell viability values |
Purnama et al. [ | Pulse energy (E)-30, 40, 50 μJ Scan speed (v)-39, 59, 78, 97 mm s−1 Wavelength-1064 nm | Distance between patterns 25, 75, 150 μm | Endothelial cells (HUVECs) avg. size 25 μm | The adhesion of HUVEC cells is influenced by chain-like structures, which were introduced by laser structuring Due to laser treatment, some amount of iron hydroxide and impurities showed in the chemical composition analysis Adhesion, cell proliferation, and alignment of cells were significantly more in d25 surface compared to the remaining surfaces The optimal design condition was a pattern size closer to the size of a cell |
Table 7 Effect of laser surface texturing on different tissue cells (in vitro & in vivo analysis)
References | Laser parameters | Dimple size | Tissue cells | Observations |
---|---|---|---|---|
Mirhosseini et al. [ | Pulse energy-100 MJ Frequency-10 HZ Wavelength-1064 nm | Diameter 156, 214, 127 μm | 2T3 osteoblast cells | More cell growth percentage was observed in small-size patterns (127 μm) It found that optimum surface roughness for cell growth, and roughness closer to cell size improved cell growth Cells were uniform and evenly spread across patterns compared to untreated surfaces Untreated samples showed higher wettability than patterned surfaces |
Mukherjee et al. [ | Frequency 1000 HZ Wavelength 50-85 μm Scanning speed 50-80 mm/s Duty cycle 60-80% | Depth 60-65 μm | MG 63 | At the same wavelength by reducing duty cycle, wider peaks were formed and at same duty cycle by reducing wavelength sharp peaks were formed It was observed that samples with similar avg. roughness, contact angle, and protein adsorption also showed different cell viability |
Mesquita-Guimaraes et al. [ | Power-25 W Velocity-200 mm/s frequency-550 Hz | 50 × 50 μm2 width 100 μm Depth 0, 40, 80 μm | MC3T3-E1 osteoblast cells | Laser micro-textured surface showed a 40% increase in cell viability compared to polished surface Optimal pattern size should be 100 to 200 μm, showing 60% more cell viability The most extensive cell layer formation was found on the laser-sintered HAp and 45S5 BG coatings, also by the cell viability values |
Purnama et al. [ | Pulse energy (E)-30, 40, 50 μJ Scan speed (v)-39, 59, 78, 97 mm s−1 Wavelength-1064 nm | Distance between patterns 25, 75, 150 μm | Endothelial cells (HUVECs) avg. size 25 μm | The adhesion of HUVEC cells is influenced by chain-like structures, which were introduced by laser structuring Due to laser treatment, some amount of iron hydroxide and impurities showed in the chemical composition analysis Adhesion, cell proliferation, and alignment of cells were significantly more in d25 surface compared to the remaining surfaces The optimal design condition was a pattern size closer to the size of a cell |
References | Name of microorganism | Time taken to kill microorganisms (bacteria) |
---|---|---|
Wilks et al. [ | Escherichia coli O157 (bacteria) | 90-270 min in 20 and 4 °C |
Wilks et al. [ | Listeria monocytes (bacteria) | No live bacteria were found after 60 min at room temperature |
Noyce et al. [ | Methicillin-resistant Staphylococcus aureus (MRSA) (bacteria) | A complete kill of bacteria in 45-90 min at room temperature at 4 °C within 6 h |
Noyce et al. [ | Influenza A virus | On Cu sample 2 × 106 virus particles were incubated after 24 h, 5 × 105 virus particles were observed |
Table 8 In vitro experimental study about Cu antimicrobial property examined under Nikon Eclipse ME600 microscope
References | Name of microorganism | Time taken to kill microorganisms (bacteria) |
---|---|---|
Wilks et al. [ | Escherichia coli O157 (bacteria) | 90-270 min in 20 and 4 °C |
Wilks et al. [ | Listeria monocytes (bacteria) | No live bacteria were found after 60 min at room temperature |
Noyce et al. [ | Methicillin-resistant Staphylococcus aureus (MRSA) (bacteria) | A complete kill of bacteria in 45-90 min at room temperature at 4 °C within 6 h |
Noyce et al. [ | Influenza A virus | On Cu sample 2 × 106 virus particles were incubated after 24 h, 5 × 105 virus particles were observed |
References | Coating | Substrate | Method of coating | Purpose of coating | Observations |
---|---|---|---|---|---|
Hassan et al. [ | Graphite | SS 316L | Physical vapour deposition (PVD) Coating thickness 15 ± 2-1009 ± 22 nm | Anti-corrosive, biocompatibility | The corrosion rate improved from 22 m/year to 1.4 m/year Surface roughness 0.004-0.001 μm. Microhardness 350 HV Due to graphite coating biocompatibility improved |
Kumar et al. [ | Ti-Baghdadite composite | SS 316L | CS Coating thickness 236 ± 11 μm | Anticorrosive and biocompatibility | Surface roughness and porosity of coating improved with increase in BAG composition A large improvement in corrosion rate was observed, uncoated substance (268 μm/year), and composite coating (0.154 μm/year) At 25% BAG composition coating observed cell viability at 120% |
Singh et al. [ | HA-TiO2 | Ti-35Nb-7Ta-5Zr | Plasma spray Coating thickness 185-200 μm | Anticorrosive and biocompatibility | Reinforcement of TiO2 in HA improved microstructure and mechanical bonding of coating Ecorr values of HA-TiO2 (− 420 mV) less than HA coating (− 330 mV) improvement in corrosion resistance was observed After 7 days, cell density of reinforcement coating (53 cells/cm2) was better than HA coating (35 cells/cm2) |
JunRong et al. [ | Tantalum | Ti6Al4V | CS Max thickness 380 μm Min thickness 24 μm | Biocompatibility, and wear resistance | The rough and porous coating formed The Microhardness of the outer part of the coating is less than the inner part due to porosity Good bioactivity showed in the form of cell proliferation |
Henao et al. [ | HA | Ti6Al4V | HVOF Four coating layers mean thickness of coating 350 ± 7 μm | Anticorrosive | Graded porosity was observed in the coating Corrosion resistance improved from (0.00164-0.000975 mA/cm2) after coating The bioactivity of the implant increased after coating After 28 days, the bone-like apatite phase consolidated and sealed the graded coating's porosity |
Zhao et al. [ | Reductive graphene oxide-Ag nanoparticle-Al | Mild Steel | CS | Antibacterial | The coating showed high antibacterial properties against Escherichia coli Homogeneity of the coating was observed |
Vilardell et al. [ | Pure Ti | - | CS | Biocompatibility | High roughness (30-50 μm) and high wettability (WCA 28°) Higher cell viability and good cell proliferation |
Al-Mangour et al. [ | SS67%-Co33.3% | Mild steel | CS Coating thickness 0.5-3.4 mm | Anti-corrosive | SS-33%Co coating showed a better corrosion rate than SS coating Annealing (1100 °C) after coating improved densification and porosity reduction |
Table 9 Different purposes of coating on implants
References | Coating | Substrate | Method of coating | Purpose of coating | Observations |
---|---|---|---|---|---|
Hassan et al. [ | Graphite | SS 316L | Physical vapour deposition (PVD) Coating thickness 15 ± 2-1009 ± 22 nm | Anti-corrosive, biocompatibility | The corrosion rate improved from 22 m/year to 1.4 m/year Surface roughness 0.004-0.001 μm. Microhardness 350 HV Due to graphite coating biocompatibility improved |
Kumar et al. [ | Ti-Baghdadite composite | SS 316L | CS Coating thickness 236 ± 11 μm | Anticorrosive and biocompatibility | Surface roughness and porosity of coating improved with increase in BAG composition A large improvement in corrosion rate was observed, uncoated substance (268 μm/year), and composite coating (0.154 μm/year) At 25% BAG composition coating observed cell viability at 120% |
Singh et al. [ | HA-TiO2 | Ti-35Nb-7Ta-5Zr | Plasma spray Coating thickness 185-200 μm | Anticorrosive and biocompatibility | Reinforcement of TiO2 in HA improved microstructure and mechanical bonding of coating Ecorr values of HA-TiO2 (− 420 mV) less than HA coating (− 330 mV) improvement in corrosion resistance was observed After 7 days, cell density of reinforcement coating (53 cells/cm2) was better than HA coating (35 cells/cm2) |
JunRong et al. [ | Tantalum | Ti6Al4V | CS Max thickness 380 μm Min thickness 24 μm | Biocompatibility, and wear resistance | The rough and porous coating formed The Microhardness of the outer part of the coating is less than the inner part due to porosity Good bioactivity showed in the form of cell proliferation |
Henao et al. [ | HA | Ti6Al4V | HVOF Four coating layers mean thickness of coating 350 ± 7 μm | Anticorrosive | Graded porosity was observed in the coating Corrosion resistance improved from (0.00164-0.000975 mA/cm2) after coating The bioactivity of the implant increased after coating After 28 days, the bone-like apatite phase consolidated and sealed the graded coating's porosity |
Zhao et al. [ | Reductive graphene oxide-Ag nanoparticle-Al | Mild Steel | CS | Antibacterial | The coating showed high antibacterial properties against Escherichia coli Homogeneity of the coating was observed |
Vilardell et al. [ | Pure Ti | - | CS | Biocompatibility | High roughness (30-50 μm) and high wettability (WCA 28°) Higher cell viability and good cell proliferation |
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