Journal Information
Vol. 8. Num. 1.
Pages 1-1592 (January - March 2019)
Share
Share
Download PDF
More article options
Visits
513
Vol. 8. Num. 1.
Pages 1-1592 (January - March 2019)
Original Article
DOI: 10.1016/j.jmrt.2018.06.015
Open Access
NiTi coated with oxide and polymer films in the in vivo healing processes
Visits
513
Dan Batalua, Florin Nastaseb, Manuella Militaruc, Mihaela Gherghiceanud, Petre Badicae,
Corresponding author
badica2003@yahoo.com

Corresponding author.
a University Politehnica of Bucharest, Splaiul Independentei 313, 060042 Bucharest, Romania
b National Institute for R&D in Microtechnologies, Street Erou Iancu Nicolae 126A, 077190 Bucharest, Romania
c University of Agronomical Science and Veterinary Medicine, Boulevard Marasesti 59, 011464 Bucharest, Romania
d “Victor Babes” National Institute, Splaiul Independentei 99-101, 050096 Bucharest, Romania
e National Institute of Materials Physics, Street Atomistilor 405 A, 077125 Magurele, Ilfov, Romania
This item has received
513
Visits

Under a Creative Commons license
Article information
Abstract
Full Text
Bibliography
Download PDF
Statistics
Figures (8)
Show moreShow less
Tables (4)
Table 1. Etched samples of NiTi: roughness measured by AFM. Samples notation is as in Fig. 1.
Table 2. Electro-polished samples obtained from the NiTi chemically etched sample 4: roughness measured by AFM. Samples notation is as in Fig. 2.
Table 3. Sol–gel coated samples: roughness measured by AFM.
Table 4. Implant-tissue qualitative evaluation after acute and chronic experiments (conventional notation: * – poor, ** – average, *** – best).
Show moreShow less
Abstract

Plates of NiTi chemically etched, electro-polished, and sol–gel coated with XO2 (X=Ti, Si, Zr), or coated with oxides and dip-coated polymers of Dextro-Levo-lactide-co-glycolide (DL-PLG, 0.4μm thickness), Dextro-Levo-lactic acid (DL-PLA, 1.3μm) or poly methyl methacrylate polymer (PMMA, 1.7μm) were obtained. Smooth and uniform NiTi surfaces without significant pitting, as revealed by AFM, were prepared for chemical etching of 120s in HF:HNO3:H2O=1:5:4, followed by electropolishing 120s in H2SO4:CH3OH:H2O=1:4:5 electrolyte and using a potential of 9V. Dip-coated layer of PMMA has shown cracks and large pores and was eliminated from further experiments. Samples of pristine and coated NiTi were in vivo implanted into rabbits and extracted after 10 and 60 days. Clinically, all implants are biocompatible; all rabbits survived and a recovery process was observed for all cases. NiTi covered with SiO2, DL-PLG and SiO2/DL-PLG have shown the best healing evolution. For 10 and 60 days good recovery was found also for NiTi coated with TiO2. Coatings of ZrO2 and ZrO2/DL-PLG have shown the poorest results. The oxide coating and the roughness RZJIS that contains information on the ‘deep’ large areas in the coatings show the strongest influence on the healing processes. Work indicates the possibility of space- and time- scale controlled variation of the functional properties.

Keywords:
NiTi alloy
Coating of TiO2
SiO2
ZrO2
DL-PLG
DL-PLA
PMMA
Roughness
In vivo test
Full Text
1Introduction

1The quasi-equiatomic NiTi alloy [1] has one of the highest shape-recovery-rate among materials exhibiting the shape memory effect. It has also excellent mechanical (superelasticity) and corrosion resistance properties [2]. NiTi was proved to be a biocompatible material [3]. The shape memory effect of NiTi alloy promotes designs with changing geometries depending on temperature, while superelasticity stands against large displacements. Hence, NiTi alloy is of much interest for fabrication of various medical devices [4]. For example, the hinge mechanism that it is self-assembled through the shape memory effect of a constrained or semi-constrained elbow implant simplifies and improves the design and the surgical procedure [5,6]. Other applications are [7,8]: intravascular stents, vena cava filters, orthopaedic compression staples, bone fixation clamps, patellar fixator, shape memory implant used in scoliosis, femoral cup prosthesis, dental arch wire, drug delivery system.

The large amount of Ni content induced a reticence in using the NiTi implants without taking supplementary precautions for avoiding side effects generated by Ni-ion release. Ni induces carcinogenesis [9], and alloys with Ni release of more than 1μg/cm2/week can give a strong allergy reaction [8]. Though a protective layer of TiO2 is formed on the surface of NiTi, the Ni surface concentration ranges between 0.4 and 27% [8] and a 10 days in vitro study has shown that there is a Ni release of 5–129μg/l [8]. To obtain a “zero” side effect of NiTi-implants use, efforts are made to block the Ni/Ti ion leaching by creating a barrier between the implant and the tissue and refs [10–15] show positive results. It is also noteworthy that Liu et al.[10] reported that NiTi coated with a TiO2 film improves blood compatibility, while Lemaire et al. [13] shows that TiO2 coating reduces Ni leaching with a factor of 14. SiO2 and ZrO2 coatings as effective barriers against Ni leaching were proved by Yang et al. [14], and Sui et al. [15], respectively. A complementary approach is to deposit functional biodegradable films, such as polymers. Biodegradable polymers are themselves of much interest. Studies were made for replacement of permanent metallic stents with fully biodegradable polymeric stents [12,16,17]. Xu et al. studied the biodegradation behaviour of PLGA (poly(lactide-co-glycolide)) stents both in vitro and in vivo for the assessment of the usefulness of biodegradable polymeric stents in human common bile duct repair and reconstruction [18]. On the other hand, Xi et al. [19] investigated the degradation behaviour of a polymeric coating on a cobalt-chromium stent platform. It is also remarkable that biodegradable films can include drugs and during their decomposition they allow their gradual release, significantly improving the recovery processes. Pan et al. [20] used poly(lactide-co-glycolide) (PLGA) as a drug reservoir to develop an emodin-eluting stent by ultrasonic-atomization-spraying method. The results were promising, indicating the potential applications of emodin-eluting stents in the treatment of cardiovascular disease. Westedt et al. [11,21] studied the formulation of biodegradable nanoparticles using solvent displacement technique for catheter-based local intraluminal drug delivery. They concluded that PVA-g-PLGA comb polyesters are suitable biodegradable polymers for the nano-encapsulation of paclitaxel.

Our work proposes a comparative analysis of deposited inorganic, organic and inorganic/organic sandwich-like coatings on a NiTi substrate. Inorganic biocompatible [14,22,23] coatings consist of an oxide XO2 (X=Ti, Si, Zr) and organic ones of biodegradable polymers such as poly 2Dextro-Levo-lactide-co-glycolide (DL-PLG), poly 2Dextro-Levo-lactic acid (DL-PLA), and the bio-stable 2poly methyl methacrylate polymer (PMMA). The main role of XO2 and PMMA coatings is to block the Ni/Ti ion diffusion into tissue and the one of the biodegradable polymeric layer is to carry and deliver the appropriate drugs. As a first step for the assessment of our coatings and to understand the in vivo processes presented in this work, polymers were not loaded with drugs. Nevertheless, our experiments (to be presented elsewhere) indicated that aspirin (acetylsalicylic acid) can be easily incorporated in our polymer coatings as an antiplatelet drug. Bare and coated NiTi samples were implanted in rabbits. At the end of the short-term acute (10 days) and long-term chronic (60 days) experiments, samples were extracted and the nearby tissue was investigated. During acute experiments, the temperature of the rabbits was monitored. Coating features vs. their implantation effects are presented and discussed. All rabbits survived the in vivo experiments indicating on the biocompatibility of all materials, but the encountered differences suggest the possibility of space- and time- scale controlled variation of the functional properties.

2Materials and methods2.1NiTi substrate preparation by chemical etching and electropolishing

NiTi shape memory alloy (50.59at. % Ni) was provided by METAL PRODUCTS Research Institute of Shanghai.

NiTi substrates were smoothened by electrochemical etching, considering that a smooth surface provides an increased corrosion resistance and, hence, the activity of Ni towards interaction with human body can be suppressed. Plates of NiTi with size of approximately 13mm in length, 4mm in width and 0.2mm in thickness were first polished with SiC abrasive paper with an ultrafine grit size of 2000, followed by diamond paste polishing up to the grit size of 8000. After polishing, samples were subject to chemical etching using different ratios of water and acid solutions, and different immersion times. Samples and chemical etching conditions are gathered in the caption of Fig. 1. Surface of the as-prepared NiTi plates was observed by optical microscopy and by tapping-mode atomic force microscopy (AFM, Ntegra Prima). Although the RRMS (root mean squared roughness), and RZJIS roughness (based on Japanese Industrial Standard – JIS, taken in 10 points) was not the lowest for sample 4 (immersed in HF:HNO3:H2O=1:5:4 for 120s, Table 1), this sample was selected for further developments since the surface shows the highest uniformity (visually appreciated over the entire AFM image, Fig. 1(4)) and it has the lowest amount of pits, while their size is reasonably small. Moreover, the peaks in the roughness profile curves (Fig. 1) are rounded, rather than being very sharp as for samples 1 and 2. Sample 3 shows the highest values of RRMS and RZJIS. RRMS was calculated by using Eq. (1):

Fig. 1.

AFM view of the NiTi after etching – surface topography (1.5μm×1.5μm). The etching parameters are: (1) –HF:HNO3:H2O=1:4:5 for 60s (Sample 1), (2) –HF:HNO3:H2O=1:5:4 for 60s (Sample 2), (3) –HF:HNO3:H2O=1:4:5 for 120s (Sample 3) and (4) –HF:HNO3:H2O=1:5:4 for 120s (Sample 4).

(0.54MB).
Table 1.

Etched samples of NiTi: roughness measured by AFM. Samples notation is as in Fig. 1.

Roughness [nm]Samples
1  2  3  4 
RRMS  1.963  1.351  6.679  4.535 
RZJIS (10 points)  14.68  9.207  25.40  22.02 

The roughness profile contains ‘n’ ordered, equally spaced points along the trace, and ‘yi’ is the vertical distance from the mean line to the ‘ith’ data point. Height is assumed to be positive in the up direction, away from the bulk material. The average distance typically based on five of the highest peak and lowest valley over the entire sampling length defines RZJIS according to formula:

where Rpi and Rvi are the ith highest peak, and lowest valley, respectively. We observe that RRMS mainly provides information of an average surface roughness (“uniformity”) while RZJIS shows the highest variation in the roughness along the measured curve profile and therefore provides information on the ‘deep’ regions from the surface.

The surface finishing of the selected sample 4 was performed by electro-polishing for different conditions of applied voltage and time. Electropolishing conditions applied on etched sample 4 leading to samples denoted 4.1, 4.2, 4.3 and 4.4 are shown in the caption of Fig. 2. The lowest level of pitting corrosion as revealed by optical microscopy (Fig. 2) was encountered for sample 4.4 (U=9V, 120s). On the other hand, this sample had the largest RRMS (Fig. 2, Table 2). The value of RZJIS was similar to those for samples 4.2 and 4.3 and lower than for sample 4.1. Taking into account the rounded shapes of the peaks in the roughness profile curves expected to promote a better corrosion behaviour, we selected sample 4.4 for subsequent coating steps.

Fig. 2.

AFM view of the etched NiTi after electro-polishing – surface topography. The electro-polishing parameters are: (4.1) U=3V for 60s (Sample 4.1), (4.2) U=9V for 60s (Sample 4.2), (4.3) U=3V for 120s (Sample 4.3), and (4.4) U=9V for 120s (Sample 4.4).

(0.61MB).
Table 2.

Electro-polished samples obtained from the NiTi chemically etched sample 4: roughness measured by AFM. Samples notation is as in Fig. 2.

Roughness [nm]Samples
4.1  4.2  4.3  4.4 
RRMS  2.372  1.947  1.526  3.281 
RZJIS (10 points)  16.61  11.31  10.97  11.72 
2.2Sol–gel coatings of XO2 (X=Ti, Si, Zr) on NiTi

Different routes were used for preparation of sol–gels. For ZrO2 we started from aqueous solutions of ZrOCl2+H2O. To remove the resulted HCl, amine-formaldehyde resin was added. The water was replaced by distillation with methoxyethanol. Finally, a stable ethanolic sol with particles of ZrO2 were obtained. For SiO2 we obtained two solutions by mixing (A) tetraethyl orthosilicate (TEOS), C2H5OH, and NH3·H2O, and by mixing (B) TEOS, C2H5OH, and HCl. The two solutions were mixed (A+B), followed by stirring and refluxing. For TiO2 we started from Ti(OC4H9) dissolved in ethyl alcohol, and acetyl acetone was added. A solution of acetic acid, H2O and ethyl alcohol was added to the previous one. The new solution was stirred, ultrasonicated and kept for 1 week at room temperature, followed by drying at 70°C. More details about concentration of solutions, temperatures, and time are given in [24–26]. The NiTi plates were immersed into the sol–gel solution, extracted with a constant speed of 1cm/min and dried at 70°C for 12h in all three cases. The thickness of the oxide layer was of about 90nm. Both optical and atomic force microscopy indicates that the highest uniformity and the lowest RRMS and RZJIS are obtained for the SiO2 film, followed by TiO2 (Table 3). The poorest surface quality is obtained for ZrO2 (Fig. 3, Table 3).

Table 3.

Sol–gel coated samples: roughness measured by AFM.

Roughness [nm]Samples
TiO2  SiO2  ZrO2 
RRMS  3.103  1.234  3.339 
RZJIS (10 points)  8.484  4.108  16.21 
Fig. 3.

AFM images of the sol–gel dip-coated TiO2, SiO2, and ZrO2.

(0.16MB).
2.3Coatings of biodegradable or biostable polymers on Si wafer

A Si wafer (10mm×10mm) for electronics applications was used as a substrate for the experiments of coating with the polymers. This substrate was selected in order to avoid uncertainties regarding the use of a substrate with a different uniformity and roughness (as in the case of oxide coatings obtained in Section 2.2) and of the use of different solvents. The solvents for DL-PLG, DL-PLA and PMMA were ethyl acetate, acetonitrile and dichloromethane [27], respectively. We used DL-PLG biodegradable copolymer ([C3H4O2]x[C2H2O2]y) in lactide:glycolide proportion of 85:15, with molecular weight (Mw) of 50,000–75,000g/mol. The rate of hydrolytic degradation of DL-PLG is influenced by copolymer ratio; i.e., 85:15 DL-PLG can be degraded within 5–6 months [28,29]. DL-PLA, with chemical formula (C3H6O3)n, can be degraded within 12–16 months [28,29]. The role of the biostable PMMA coating ([CH2C(CH3)(CO2CH3)]n, Mw=15,000g/mol) is of a supplementary organic barrier against Ni dissolution or as a replacement for the oxide coatings [30,31]. The dissolved polymers were ultrasonically stirred (300W) at room temperature. The samples were immersed in the solutions and dried by evaporating the solvent.

The thickness of DL-PLG, DL-PLA and PMMA polymer coatings on Si was different. Namely, it was about 0.4, 1.3 and 1.7μm, respectively. Fourier transformed infrared spectroscopy (Jasco FT-IR-6200 Spectrometer coupled with FT-IR Jasco IRT 3000 microscope) measurements (not shown) confirmed formation of the polymer coatings. Surface quality of the polymer coatings was checked by AFM (Fig. 4). PMMA surface has shown cracks and in some regions films were not uniform or continuous. From these reasons PMMA was abandoned and not used in further experiments. By using the same procedure, DL-PLG and DL-PLA polymers were coated on NiTi and NiTi/XO2 samples.

Fig. 4.

Polymeric dip-coated films of DL-PLG (a), DL-PLA (b), and PMMA (c).

(0.17MB).
2.4In vivo assessments of the implanted samples into rabbits

Healthy 18 rabbits were used for implantation. Their weight was between 1.5 and 1.8kg. Implants were extracted after 10 and 60 days. Time of the acute and chronic in vivo tests was selected based on the results reported in ref. [3]. Authors of [3] investigated the thickness of the tissue-reaction-capsule surrounding the implant vs. time. They found that thickness of the inflammatory capsule for NiTi implanted in rabbits, decreases reaching saturation for a time longer than approximately 60 days. It was considered that 10 days is appropriate to observe the tissue-implant reaction, while 60 days is the time to reveal the recovery process. All in vivo tests were carried out according with national and European legislation for protection of animals. Surgical preparatory, implantation and explantation steps are presented in Fig. 5a–d, f.

Fig. 5.

(a) Selection of the incision spots for implantations. (b) Preparation of the surgical area. (c) Insertion of the implants. (d) Suturing. (e) Radiography. (f) Explantation of NiTi implants.

(0.35MB).

Radiography (Fig. 5e) shows the position of the implants. Before implantation, the implants have been sterilized by UV. After explantation, small samples of muscular tissue taken from the adjacent area to the implants were fixed in glutaraldehyde. Optical microscopy images (Fig. 6, Nikon E 600 CCD camera) were taken on the as-prepared samples coloured using the Masson's trichrome staining protocol. Tissue samples of 60–80nm thickness were prepared for TEM (transmission electron microscopy) observations using an ultramicrotom and embedding in Epon, according to the standard procedure of Ultrastructural Pathology Laboratory. Contrast enhancement was made with Uranyl Acetate 1% (20min) and Reynold's solution (10min).

Fig. 6.

Tissue response (B) around tested materials (A) after 10 and 60 days of implantation (200×).

(1.23MB).
3Results and discussion

All rabbits survived the in vivo tests. We monitored temperature daily, measuring it in the morning and in the evening during the first critical 10 days after implantation. The average temperature and the temperature variation interval for each rabbit are presented in Fig. 7. Temperatures for all rabbits were within the limits for healthy ones [32].

Fig. 7.

Monitored rabbits’ body temperatures during the acute experiment.

(0.1MB).

Tissue reaction fibrous capsules (see region B in Fig. 6(h) as an example) can be visualized for all samples in Fig. 6 for a constant magnification. In contact with the living tissue, the biological response is different, depending on the materials type and features at the interface. Three qualitative criteria were considered for tissue evaluation: thickness, smoothness, and the compactness of the contact surface.

The muscular tissue (blue-grey areas in Fig. 6) located in the implant vicinity reacted through a granuloma inflammatory process (red and especially dark red areas in Fig. 6) without a specific association to necrosis. Initial phases of acute/severe inflammation generally developed in the first days after implant insertion. Comparison of the collected implants at 10 and 60 days reveals a chronic inflammation, through the persistence of the inflammatory agent. This is the usual process when stimuli (related to the implanted material which associates with the necrotic tissue occurrence in the surgical injury) persist or are recurrent. Monocytes-macrophages, lymphocytes and plasma cells were the main cells involved in the inflammation process (Fig. 8). Monocytes that derived from peripheral blood transformed into macrophages cells. The intensity of cells formation was dependent on the implanted materials and the indicated cells were more often observed in the samples extracted 10 days after surgery. Specific for granuloma tissue is also formation of neo-capillary cells and fibroblast proliferation with collagen depositions. In the samples extracted after 60 days from surgery, granulomatous inflammation elements were found in most cases to contain few macrophage-epithelial cells (isolated or aggregated), lymphocytes, plasma platelets, mastocyte cells, but without multinuclear giant cells. This behaviour is a typical immune response of the host organism. The main resulting cells of the immune response are epitheliotic-type macrophages that may also have a secretory function. The immune response intensity depends on the antigenicity of the introduced materials. Materials in this work are aseptic and neutral so that a necrotic activity was not observed. Worthy to note is also the evolution with the time from surgery of the chronic granuloma elements that is reflected in their fibrous organization with observation of numerous fibroblasts, miofibroblasts and with a massive synthesis of collagen and of glucose-aminoglycans as components of the extracellular matrix. The evolution denotes a healing process, but, again, its intensity depends on implanted samples. The differences induced by different samples on contact tissue are addressed in the next paragraphs considering the already mentioned criteria for the reaction fibrous capsules.

Fig. 8.

Selected TEM images showing healing traces. (a) Bare NiTi sample. (b, c) NiTi/TiO2 sample, (d-f) NiTi/SiO2 sample, (g) NiTi/PLA sample (PG – proteoglycans, # – collagen, FB – fibroblasts, * – degranulation, PC – plasmocytes, LS – lysosomes, MP – macrophages, PMN – polymorphonuclear, GC – granulocyte, AEM – amorphous extracellular matrix, MC – monocytes, 9100×).

(0.26MB).

For inorganic coatings, the SiO2 film provides a thin, very smooth, and compact regenerated tissue after 60 days (Fig. 6f, NiTi/SiO2 60 days). The thinnest and smoothest contact tissue is observed for the biodegradable DL-PLA layer in the acute experiment (Fig. 6i, NiTi/DL-PLA 10 days). For the chronic experiment the DL-PLG layer (Fig. 6l, NiTi/DL-PLG 60 days) proved to induce a better regeneration than PLA, where the tissue does not show a good compactness (Fig. 6j, NiTi/DL-PLA 60 days). Finally, the combination SiO2/DL-PLG shows the best evolution, both for short and long-term experiment (Fig. 6o, p, NiTi/SiO2/DL-PLG, 10/60 days), hence a good acceptance by the living tissue. A recovery process is detected and, in general, it continues for 60 days (Fig. 6r, NiTi/ZrO2/DL-PLG 60 days). The encapsulation tissue thickness is stationary (e.g. NiTi, NiTi/DL-PLG, Fig. 6a, b, and k, l) or decreases after 60 days (Fig. 6q, r, NiTi/ZrO2/DL-PLG). TEM observations (Fig. 8) show that after 60 days, regeneration/healing processes are still visible. Low number of inflammatory cells such as monocytes and macrophages (follow the arrows in Fig. 8) indicate a normal tissue healing activity [33]. At the same time, they show that healing processes are not over after 60 days of implantation.

In summary, there are notable differences among the behaviour of the samples. Considering the thickness/smoothness/compactness qualitative criteria for the inflammatory capsule, our results are gathered in Table 4.

Table 4.

Implant-tissue qualitative evaluation after acute and chronic experiments (conventional notation: * – poor, ** – average, *** – best).

No.  Implant  Acute (10 days)  Evolution  Chronic (60 days) 
NiTi  * *  →  ** 
NiTi/TiO2  ***  →  *** 
NiTi/SiO2  ↗  *** 
NiTi/ZrO2  **  ↘ 
NiTi/DL-PLG  ↗  *** 
NiTi/DL-PLA  ***  ↘ 
NiTi/TiO2/DL-PLG  ***  ↘  ** 
NiTi/SiO2/DL-PLG  **  ↗  *** 
NiTi/ZrO2/DL-PLG  → 

There are situations that are worse at 60 days than at 10 days (Fig. 6, NiTi/ZrO2/DL-PLG 60 days). Therefore, the implant interface can either accelerate the healing response, or delay it. This observation shows that certain films on the implant can play an important role in tissue regeneration. The factors influencing regeneration processes are known to be multiple. Due to complexity it is not possible to establish confidently the key factors and observed correlations can be local fitting only a certain case. Bearing in mind this important observation, we shall remind that Figs. 3, 4 and Table 3 revealed different patterns and qualities of coated films. Results suggest that materials and surface details play an important role in the implant-tissue interaction. The effects of the material type and of the surface patterns cannot be separated. Nevertheless, our data may give some hints concerning the major influence for our specific samples. In our assessment analysis we started from the assumption that corrosion processes can be suppressed for a lower roughness of the surface. The possibility of a controlled corrosion on a space and time scale is expected to provide the background for a better healing process (e.g. faster, easier and more predictable without complications). In Sections 2.2 and 2.3 we have noted that one has to discuss the surface patterns of our films from the viewpoint of a macro and micro roughness somehow revealed by RZJIS and RRMS. Moreover, we also pointed out that peaks in the roughness profile curves may show sharp or rounded shapes. All these details may contribute to corrosion and bio processes at interface. For the oxide coatings and to a less extent for the oxide/polymer coatings the influence of the oxide coating seems to be essential (Table 4). A strong influence is given by the oxide coating surface quality: the lowest RZJIS and RRMS values for SiO2 and the highest for ZrO2 are in agreement with the best healing results for the first oxide and with the poorest one for the second oxide, respectively. We have seen that a coated film is composed of smooth areas mainly described by RRMS and large ‘deep’ circular areas stronger involved in estimation of RZJIS. One observes from Table 4 that, while the variation of RRMS for TiO2 and ZrO2 (Table 3) is small (3.103 and 3.339nm), a better healing process is for sample TiO2 with a lower RZJIS (4.108nm for TiO2 and 16.21nm for ZrO2) and less ‘deep’ type areas. The result evidences the strong influence of the ‘deep’ areas on recovery processes and further research is necessary. As regarding the polymer layer, it also shows a morphology (Fig. 4) resembling the one for the oxide coatings. The ‘deep’ areas in the polymer case are often pores. While they can also influence the healing processes at interface on space and time scale depending on their distribution and size, they can be useful for designing of the drug-eluting films.

4Conclusion

From a clinical viewpoint, all implants show good in vivo biocompatibility. All rabbits survived to the experiments and a good recovery process was observed for all cases. NiTi samples coated with SiO2, DL-PLG, SiO2/DL-PLG have shown the best results for the recovery progress. Good results for the short and long experiments were also observed for NiTi/TiO2. Coatings of ZrO2, DL-PLA and ZrO2/DL-PLG have shown the poorest results in the long-term experiment. The possibility of space and time scale controlled healing processes by using suitable coatings is envisioned. Materials and surface patterns of the samples at interface show a strong influence on the recovery processes. Oxides and especially ‘deep’ areas reflected by roughness RZJIS are the strongest factors influencing healing processes.

Conflicts of interest

The authors declare no conflicts of interest.

Acknowledgements

Authors acknowledge Romanian National Authority for Scientific Research and Innovation, CCCDI-UEFISCDI, grants AMCSIT-CEEX/194 – ANGIOMAT and 74-COFUND-M-ERA.NET II – BIOMB (within PNCDI III).

References
[1]
W.J. Buehler, J.V. Gilfrich, R.C. Wiley.
Effect of low-temperature phase changes on the mechanical properties of alloys near composition NiTi.
J Appl Phys, 34 (1963), pp. 1475-1477
[2]
K. Otsuka, X. Ren.
Physical metallurgy of Ti-Ni-based shape memory alloys.
Progr Mater Sci, 50 (2005), pp. 511-678
[3]
J. Ryhanen, M. Kallioinen, J. Tuukkanen, J. Junila, E. Niemela, P. Sandvik, W. Serlo.
In vivo biocompatibility evaluation of nickel-titanium shape memory metal alloy: muscle and perineural tissue responses and encapsule membrane thickness.
J Biomed Mater Res, 41 (1998), pp. 481-488
[4]
L.L. Meisner, V.P. Sivokha.
Physical and biochemical principles of the application of NiTi-based alloys as shape-memory implants.
Shape memory implants, pp. 61-72
[5]
Batalu ND, et al. Total constricted elbow prosthesis made of shape-memory alloy with hinge-like fixation and coupling system based on shape-memory effect. Patent No. RO131261-A0.
[6]
Batalu ND, et al. Semiconstrained total elbow prosthesis made of shape-memory alloys, with coupling system based on shape-memory effect. Patent No. RO131379-A0.
[7]
I.P. Lipscomb, L.D.M. Nokes.
The applications of shape memory alloys in medicine.
Paston Press Ltd, (1996), pp. 153
[8]
Y. L’Hocine.
Shape memory implants.
Springer, (2000), pp. 349
[9]
B. Zambelli, V.N. Uversky, S. Ciurli.
Nickel impact on human health: an intrinsic disorder perspective.
Biochim Biophys Acta, 1864 (2016), pp. 1714-1731
[10]
J.X. Liu, D.Z. Yang, F. Shi, Y.J. Cai.
Sol–gel deposited TiO2 film on NiTi surgical alloy for biocompatibility improvement.
Thin Solid Films, 429 (2003), pp. 225-230
[11]
U. Westedt, M. Wittmar, M. Hellwig, P. Hanefeld, A. Greiner, A.K. Schaper, T. Kissel.
Paclitaxel releasing films consisting of poly(vinyl alcohol)-graft-poly(lactide-co-glycolide) and their potential as biodegradable stent coatings.
J Contr Release, 111 (2006), pp. 235-246
[12]
X. Wang, S.S. Venkatraman, F.Y.C. Boey, J.S.C. Loo, L.P. Tan.
Controlled release of sirolimus from a multilayered PLGA stent matrix.
Biomaterials, 27 (2006), pp. 5588-5595
[13]
V. Lemaire, B. Sicotte, S. Allard.
Surface modification treatments to reduce Ni leaching from porous nitinol Porous metals and metallic foams.
METFOAM, 2007 (2008), pp. 291-294
[14]
S. Yang, F. Zhou, T. Xiao, D.B. Xu, Z. Li, Z. Xiao, Z.A. Xiao.
Surface modification with SiO2 coating on biomedical TiNi shape memory alloy by sol–gel method.
Trans Nonferrous Met Soc China, 25 (2015), pp. 3723-3728
[15]
J.H. Sui, W Cai.
Formation of ZrO2 coating on the NiTi alloys for improving their surface properties.
Nucl Instrum Meth Phys Res Sect B-Beam Interact Mater Atoms, 251 (2006), pp. 402-406
[16]
S.S. Venkatraman, L.P. Tan, J.F.D. Joso, Y.C.F. Boey, X.T. Wang.
Biodegradable stents with elastic memory.
Biomaterials, 27 (2006), pp. 1573-1578
[17]
S. Venkatraman, L.P. Tan, T. Vinalia, K.H. Mak, F. Boey.
Collapse pressures of biodegradable stents.
Biomaterials, 24 (2003), pp. 2105-2111
[18]
X. Xu, T. Liu, K. Zhang, S. Liu, Z. Shen, Y. Li, X. Jing.
Biodegradation of poly(l-lactide-co-glycolide) tube stents in bile.
Polym Degrad Stabil, 93 (2008), pp. 811-817
[19]
T. Xi, R. Gao, B. Xu, L. Chen, T. Luo, J. Liu, Y. Wei, S. Zhong.
In vitro and in vivo changes to PLGA/sirolimus coating on drug eluting stents.
Biomaterials, 31 (2010), pp. 5151-5158
[20]
C.J. Pan, J. Wang, N. Huang.
Preparation, characterization and in vitro anticoagulation of emodin-eluting controlled biodegradable stent coatings.
Colloids Surf B: Biointerfaces, 77 (2010), pp. 155-160
[21]
U. Westedt, M. Kalinowski, M. Wittmar, T. Merdan, F. Unger, J. Fuchs, S. Schäller, U. Bakowsky, T. Kissel.
Poly(vinyl alcohol)-graft-poly(lactide-co-glycolide) nanoparticles for local delivery of paclitaxel for restenosis treatment.
J Contr Release, 119 (2007), pp. 41-51
[22]
J. Huang, P. Dong, W. Hao, T. Wang, Y. Xia, G. Da, Y. Fan.
Biocompatibility of TiO2 and TiO2/heparin coatings on NiTi alloy.
Appl Surf Sci, 313 (2014), pp. 172-182
[23]
Y. Li, L. Li, B. Li.
Direct write printing of three-dimensional ZrO2 biological scaffolds.
Mater Des, 72 (2015), pp. 16-20
[24]
S.M. Attia, J. Wang, G. Wu, J. Shen, J. Ma.
Nanostructure study of TiO2 films prepared by dip coating process.
J Mater Sci Technol, 18 (2002), pp. 31-33
[25]
S.M. Attia, J. Wang, G. Wu, J. Shen, J. Ma.
Review on sol–gel derived coatings: process, techniques and optical applications.
J Mater Sci Technol, 18 (2002), pp. 211-217
[26]
Q. Zhang, J. Shen, J. Wang, G. Wu, L. Chen.
Sol–gel derived ZrO2-SiO2 highly reflective coatings.
Int J Inorg Mater, 2 (2000), pp. 319-323
[27]
A. Delgado, C. Evora, M. Llabras.
Degradation of DL-PLA-methadone microspheres during in vitro release.
Int J Pharm, 140 (1996), pp. 219-227
[28]
A. Tschakaloff, H.W. Losken, R. von Oepen, W. Michaeli, O. Moritz, M.P. Mooney, A. Losken.
Degradation kinetics of biodegradable dl-polylactic acid biodegradable implants depending on the site of implantation.
Int J Oral Maxillofac Surg, 23 (1994), pp. 443-445
[29]
M. Zilberman, K.D. Nelson, R.C. Eberhart.
Mechanical properties and in vitro degradation of bioresorbable fibers and expandable fiber-based stents.
J Biomed Mater Res B: Appl Biomater, 74 (2005), pp. 792-799
[30]
C. Nastase, A. Dumitru, F. Nastase, A. Morozan, S. Vulpe, D. Batalu.
Comparative study of deep-coating and plasma processing PMMA thin films.
J Optoelectron Adv Mater, 12 (2010), pp. 944-947
[31]
R. Premraj, D. Mukesh.
Biodegradation of polymers.
Indian J Biotechnol, 4 (2005), pp. 186-193
[32]
D. Robertshaw.
Temperature regulation and thermal environment.
Dukes’ physiology of domestic animals, 12th ed., pp. 962-975
[33]
E. Ciobotaru, M. Militaru, T. Soare, I. Ionascu, M. Gherghiceanu, C. Vlagioiu, G. Dinescu, D. Batalu, R. Bololoi.
Histological and electron microscopy features in local muscular biocompatibility of NiTi alloys coated with oxides and polymers. Acute and chronic experimental model in rabbit.
Sci Works C: Vet Med, LIII (2008), pp. 69-79
Copyright © 2018. Brazilian Metallurgical, Materials and Mining Association
Journal of Materials Research and Technology

Subscribe to our newsletter

Article options
Tools
Cookies policy
To improve our services and products, we use cookies (own or third parties authorized) to show advertising related to client preferences through the analyses of navigation customer behavior. Continuing navigation will be considered as acceptance of this use. You can change the settings or obtain more information by clicking here.