Improving Biocompatibility for Next Generation of Metallic Implants

Abstract

The increasing need for joint replacement surgeries, musculoskeletal repairs, and orthodontics worldwide prompts emerging technologies to evolve with healthcare's changing landscape. Metallic orthopaedic materials have a shared application history with the aerospace industry, making them only partly efficient in the biomedical domain. However, suitability of metallic materials in bone tissue replacements and regenerative therapies remains unchallenged due to their superior mechanical properties, eventhough they are not perfectly biocompatible. Therefore, exploring ways to improve biocompatibility is the most critical step toward designing the next generation of metallic biomaterials. This review discusses methods of improving biocompatibility of metals used in biomedical devices using surface modification, bulk modification, and incorporation of biologics. Our investigation spans multiple length scales, from bulk metals to the effect of microporosities, surface nanoarchitecture, and biomolecules such as DNA incorporation for enhanced biological response in metallic materials. We examine recent technologies such as 3D printing in alloy design and storing surface charge on nanoarchitecture surfaces, metal-on-metal, and ceramic-on-metal coatings to present a coherent and comprehensive understanding of the subject. Finally, we consider the advantages and challenges of metallic biomaterials and identify future directions.

Keywords: 3D Printing; Alloys; Biocompatibility; Implants; Metals; additive manufacturing.

Figure 1.

Interdependence of bulk modification and surface modification of metallic biomedical materials towards comprehensively improving biocompatibility. Representation of the current generation of some metallic implants used in the biomedical devices industry.

Figure 2.

Common bulk and surface modification approach of metallic implants to improve biocompatibility.

Figure 3.

Commercially available porous metallic orthopedic devices. A. Arcam designed hip cup using electron beam melting (EBM) technology, B. Lattice-structured implant prototype, additively manufactured by Imperial College London, C. Alphaform produced 3D printed titanium alloy bone-implant using EOSINT M 280 for the reconstruction of the hip bone for a cancer patient in 2014.

Figure 4.. a. Schematic representation of metal implant additive manufacturing techniques,

i) Powder Bed Fusion (PBF) implant fabrication involves the use of an electron beam or high energy laser to melt powder particles on a metal powder bed to fabricate the desired shape and structure of implants, ii) Directed Energy Deposition (DED) fabrication involves metal powder flow through multiple nozzles and simultaneously being exposed to a high energy laser beam which forms a metal melt pool on the substrate gradually taking shape through layer-wise deposition of molten metal and rapid cooling. b. Pore properties of fabricated metal structures, i) SEM micrograph of 80% porous Ti sponge sheet fabricated using slurry foaming showing similar pore characteristics as cancellous bone. ii) Shows the interconnected pore structure of a Ti metal cage 3D printed using electron beam melting technology with a trabecular pore diameter of 413 ± 78 μm. Pores > 300 μm are suitable for bone ingrowth and fusion with metal cage, iii) Selective laser molten (SLM) porous Ti structures with 230 μm strut size and 500 μm pore size, iv) SLM fabricated 80% porous Ta implants. c. Compressive properties of additively manufactured porous metal structures for biomedical applications i) porous Ti with 230 μm strut size and 68% porosity, ii) Ta structures with 80% porosity, iii and iv) CpTi and Ti6al4V-ELI structures with 120 μm and 500 μm pore size, respectively.

Figure 5.. a. Cell-materials interaction as a function of cellular morphology investigation,

i) SEM micrographs of hFOB morphology on DED processed 27% porous Ti and Ta and 45% porous Ta at day 11, showing Ta-27 has the most optimized surface for cell-material interaction with cellular confluence as well as excellent proliferation ii) shows highly efficient focal adhesion of human adipose-derived stem cells (hADCs) on trabecular tantalum (TT) surface. b) Fluorescence assay of dead and living cells on biocompatible metal surfaces. i) (Top left and bottom left) show fluorescence micrographs of Ti64 and Mg alloy (WE43) at 4h after seeding MG-63 osteoblast-like cells as compared to 24h fluorescence images (Top right and bottom right) ii and iii) ALP protein expression for hFOB cells on 27% porous Ti and Ta and 45% porous Ta has shown through fluorescence micrographs. In contrast to SEM micrographs, Ta-27 did not show strong ALP expression at day 11. c) In vivo bone formation and osseointegration analyses, i) X-ray and histological images of open porous SLM-processed Ta revealing a fair amount of osseous growth for both the porous Ta specimens with almost 100% bridging of the defect, ii) H&E (left) and Masson-Goldner (right) stained bone sections after 12 weeks of implantation of Ti64 implants with 178 μm pore size showing new bone formation around the porous metal implant, iii and iv) Photomicrograph showing the histology images after 4 weeks (top row) and 10 weeks (bottom row) for Ti, porous Ti (Ti-P) and porous Ti with nanotube surface (TNT-P) where signs of osteoid like new bone formation could be seen in orange/red color. Modified Masson Goldner’s trichrome staining method was used,v) In vivo biological response from tantalum parts fabricated using direct energy deposition (DED) showing early-stage osseointegration at 5 weeks (left column) as a function of designed porosities and extended new bone formation at 12 weeks (right column).

Figure 6.. a. Different mechanisms of storing electrical charge on implant surfaces to enhance biocompatibility,

i) through the natural DC electrical potential of bone, which results as a trigger response from stresses applied on the bone, ii) through external electrothermal polarization, which induces capacitive electrical charge storage on the surface of the implant, iii) through external electrical triggers post-implantation with the electrical charge applied in situ at the surgery site. b. surface characteristics of the implant favorable towards electrical charge storage, i) nanoarchitecture Ti surface with TiO₂ nanotubes which show capacitive potential for charge storage ~ 40mC/cm² . ii and iii) Micro arc oxidized (MaO) Ti surfaces with TiO₂ nanoarchitecture for electrical charge storage c. In vivo bone formation and osseointegration in rat femur model, i) Endothelial nitric oxide synthase (eNOS) and inducible nitric oxide synthase (iNOS) in ovariectomized (OVX) rats exposed to low-intensity electrical stimulation (LIES) showed similar eNOS and iNOS expressions for both in the OVX rats, ii) optical micrographs for polarized nanotube (TNT-P) showing higher osteoids like bone as well as mineralized bone formations at 5-week post implantation, d. In vitro cell-materials interaction, i) SaO2 cells cultured on MaO-TiO₂ surface showing well-attached cells with flattened, spread out morphologies with numerous filopodial extensions, ii) SEM micrographs of osteoblasts cultured on implants exposed to external 15 V stimulation on anodized nanotubular titanium, iii) well-flattened and proliferated osteoblasts on polarized TNT surfaces at 7 days post-culture.

Figure 7.. a. Morphological analyses of β-Ti alloys,

i) Low (left) and high (right) magnification SEM micrographs of TiO₂-Nb₂O₅-ZrO₂-nanotube (NT) surfaces revealing non-uniform size distribution of nanotubes (inset), with a diameter ranging between 33 and 76 nm. ii) Electron backscattered micrograph of SLM printed TNTZ (Ti-Nb-Ta-Zr) alloy (left) and as-cast TNTZ (right), mainly revealing β-Ti phase with a minority of α’ phase and grain size almost nine times smaller than as-cast TNTZ. iii) Optical micrograph (left) of EBM printed Ti2448 (Ti-24Nb-4Zr-8Sn) implants showing pore (800–900 μm) and strut (350 μm) characteristics, (right) shows the surface morphology of type I collagen immobilized Ti2448 implants with much smaller pores (10–100nm) due to partial coverage due to collagen immobilization. b. In vitro cell-material interactions of β-Ti alloys, i) SEM micrographs of spreading, attachment and proliferation of osteoblast-like Saos-2 cells on TiO₂-Nb₂O₅-ZrO₂-NT indicating a positive enhancement in the biocompatibility of these surfaces compared to Ti35Zr28Nb surfaces , ii) fluorescence images of 72h cultured mouse fibroblast cells (L929) on rhombic dodecahedron (RD) printed TNTZ (left), and body diagonal (BD) printed TNTZ (right) showing more uniform cell spreading on BD structures iii) human bone marrow-derived mesenchymal stem cells (hBMSC) adhesion and spreading with long filopodial extensions on the type I collagen immobilized Ti2448 surface (left), corroborating focal adhesion observed from immunofluorescence image of the same surface (right), iv) nitride Ti27Nb alloy shows actin cytoskeleton, and focal adhesion of endothelial progenitor cells (left) 2h post culture, (right) shows immunofluorescence vascular endothelial cell cadherin marker on nTi27Nb surface after 5days of culture.

Figure 8.. a. Surface morphologies of Ti-Ta alloys,

i, ii and iii) show SEM micrographs of Ti-25Ta surfaces with acicular α’ and α” platelet microstructure for SLM fabricated (i) and DED fabricated (iii) Ti-25Ta alloys, while (ii) shows more β-phase stabilization for Ti-25Ta alloys, iv) exhibits equiaxed microstructure containing both α” and ω phases,v) shows Ti-35Ta microstructure, which was subjected to heat treatment to achieve homogenized distribution of Ta. b. In vitro cell-material interaction of Ti-Ta alloy surfaces, i) human adipose-derived stem cells show well-attached morphology and proliferation on Ti-30Ta surfaces nanotube architecture, ii) well proliferated osteoblast cells on Ti-25Ta surface with uniform surface area coverage and layering shows enhanced biocompatibility of this alloys. c, In vivo implantation of Ti-Ta alloys, i and ii) show implant-bone apposition of Ti-Ta alloys from a radiograph of rabbit femurs and CT-scan of rat femurs respectively. d, Biological response in a dynamic in vivo environment, i) showing optical micrographs of Ti-10Ta and Ti-25Ta bone-implant sections from rat femur implantation revealing enhanced osteoid formation (red areas) in NT modified alloy surfaces while (ii) corroborates the results exhibiting higher trabecular bone formation at 7 weeks in rabbit femur implantation.

Figure 9.. a. mechanism of stent placement

for percutaneous coronary intervention showing Co-Cr multi-vision link stents (commonly used permanent device) compared to bioresorbable Mg alloy and pure Zn devices (degraded in 12 months)^–. b. In vivo osteogenesis and osseointegration profile for binary Zn alloy materials at 8 weeks post-implantation. Zn-Li, Zn-Mg, and Zn-Ca showed the most efficient performance in terms of Zn degradation in load-bearing sites. c. surface characteristics of Mg alloys for bone implants, (top left) optical micrograph of the expandable tubular Mg alloy stent showing characteristic cross-linked circumferential noose- shaped structures. (Bottom and right): Electron microscopy at low and high magnification, respectively, Optical micrographs of Mg alloys surfaces consisting of Mg-0.6Zr-0.5Sr-2Sc (left) and Mg-0.6Zr-0.5Sr-3Sc (right), which showed enhanced biocompatibility due to the addition of Sc>1. d. antibacterial performance of SS, Mg, and Zn implants showing highest contact killing for Zn. e. in vitro cytocompatibility of Mg alloys, (top) and (bottom) shows confocal immunofluorescence images of MC3T3-E1 cells on untreated Mg alloy ZK60 surface (top) as compared to higher and enhanced cell viability on Ta suboxide functionalized ZK60 surface (TO-ZK60) in (bottom). f. The mechanical properties of binary Zn alloy materials compared to pure Zn show that Zn-Li alloys exhibit better stability. g. Permanent and bioresorbable (BRS)-based drug-eluting stents (DES). (a) in vivo SEM images of permanent (Co-Cr)-DES compared highlighting everolimus-eluting stent (EES) over its competitors, (b) late stenosis as a result of the molecular level effect of strut malapposition, (c) 28-day CT images showing high embedding and low protrusion and area of uncovered struts of Magnitude-BRS as compared to the Absorb-BVS (bioresorbable vascular scaffolds) (d) a nitrided-Fe DES and (e) a PDLLA (poly DL-lactic acid)-carrying Rapamycin coated Mg-JDBM stent.

Figure 10.. a. popular mechanisms of effective ceramic coatings on metal surfaces,

i) Plasma-spray coating molten ceramic powder particles onto the surface of a heated material to-be-coated. The ceramic powder is injected into the high-temperature plasma, ii) Electrophoretic deposition technique involves the deposition of charged ceramic particles in a colloidal solution onto a conducting metal surface, iii) infiltration process involves the use of nanoscale ceramic particles into porous metal network resulting in entrapment of the ceramic nanoparticles and coating the metal, iv) DED additive manufacturing technique involves premixing ceramic and metal powders to coat metal surfaces under high-temperature laser melting and rapid solidification. b, Morphologies of coating surfaces and coating-substrate interfaces, i) (left) interfacial SEM micrograph and (right) micrograph of the particle size distribution of 45S5 bioglass coated on stainless steel (AISI 304) with the fine dry ground (DR) particle size less than 63μm. DR63 powder coating revealed fewer cracks and pores at the interface due to decreased temperature gradient during cooling, ii) Interfacial microstructure of Ag-oxide and Mg-oxide doped HA-coated on DED fabricated Ti+3%HA substrate showed well-bonded coating with the substrate with no cracks and delamination, iii) CaP deposition using immersion in simulated body fluid on polycaprolactone (PCL) coated metal substrate reveals nanotextured HCA like layer formation on PCL strands, iv) Cross-section microstructure of 45S5 and silver particles embedded in PEEK coated on SS316L, shows uniform co-deposition of 45S5 and Ag particles and homogenous microstructure from SEM (left) as well as a backscattered image (right). c, in vitro cellular activity on coated metal surfaces, i) 45S5 coating on AISI 304 substrate shows the formation of a uniform layer of HA after 7-day immersion in SBF, indicating enhanced bioactivity, ii) Doped HA coating on DED processed Ti64 surfaces shows not only well-attached osteoblasts but also differentiation of osteoblast at day 11 on these increased biocompatible coated metals, iii) At week-8, HCA coated PCL strands shows cell sheet formation around each strand of PCL when exposed to a basal media, iv) PEEK and bioglass composite on SS316L exhibited well-spread MG63 osteoblast-like cell on the surface at 4 days after incubation, with no evident detrimental effect from Ag addition in the coating, d, in vivo dynamic biological response, i) Tissue integration within the doped HA coating on Ti64 can be seen in week 10 with ZnSiAg HA-coated Ti64 indicating better osseointegration. The higher osteoid formation is seen with HA; however, ZnSiAg-HA has a higher total bone formation, ii) H&E staining of rat explant section reveals new bone formation and angiogenesis for HA-coated Ti64 with natural medicinal compound curcumin doping in the HA matrix, iii) Mineralized bone filling pores of HA+ PCL coated scaffolds at 8 weeks implantation in athymic rats, iv) High-magnification image of Ti discs coated with HCA and BMP-2, showing bone tissue in direct contact with the coating and within the surrounding connective tissue.

Figure 11.. a. Mechanisms of immobilization of biomolecules on metal implant surfaces,

i) Physical adsorption of proteins where the proteins get adsorbed on a hydrophobic surface eventually modifying the surface characteristic to hydrophilic and favorable for cellular attachment resulting in cascades of cell-material interactions, ii) physical entrapment of biomolecule on porous metal surfaces where the biomolecule gets entrapped in the pores of the metal structure and gets released over time, iii) covalent binding of biomolecules on the surface of metals leading to specific functionalization of the metal surface to enhance biocompatibility b, surface characteristics of metal functionalized with biomolecules, i) Rat tail type I collagen immobilized on Mg surface(AZ31) showing collagen self-assembly structure varying with collagen concentration. At 200μg/ml, the collagen coating showed longer thin fibrils forming a multilayer network structure, ii) Polycationic electrolytes (PAH) coupled with polyanionic salmon DNA were immobilized on Ti surfaces showed spheroid-like deposition of biomolecules on the surface of these materials aiding in cellular interactions c, In vitro biological response of biomolecule coated metal surfaces, i) (left and right) shows MC3T3 cells growing and proliferating on AZ31 and Mg alloys respectively with an increased number of living cells on AZ31 surface at day 7, ii) osteoblast-like cell growth on PAH/DNA immobilized Ti surface at day 4 (left) and day 16 (right) respectively. On day 16, a large formation of collagen fibrils, mineralized globules, and extracellular matrix on the treated surfaces was observed.