A hydrated phospholipid polymer gel-like layer for increased durability

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A hydrated phospholipid polymer gel-like layer for increased durability of orthopedic bearing surfaces Masayuki Kyomoto, Toru Moro, Shihori Yamane, Kenichi Watanabe, Masami Hashimoto, Sakae Tanaka, and Kazuhiko Ishihara Langmuir, Just Accepted Manuscript • DOI: 10.1021/acs.langmuir.8b01494 • Publication Date (Web): 29 Jun 2018 Downloaded from http://pubs.acs.org on July 1, 2018

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A hydrated phospholipid polymer gel-like layer for increased durability of orthopedic bearing surfaces Running title: Improved implant durability with hydrated gel-like layers

Masayuki Kyomotoa,b,c, Toru Morob,d, Shihori Yamanea,b,c, Kenichi Watanabeb,c, Masami Hashimotoe, Sakae Tanakad, Kazuhiko Ishiharaa, *

a

Department of Materials Engineering, School of Engineering, The University of Tokyo, 7-3-1 Hongo,

Bunkyo-ku, Tokyo 113-8656, Japan b

Division of Science for Joint Reconstruction, Graduate School of Medicine, The University of Tokyo, 7-3-1

Hongo, Bunkyo-ku, Tokyo 113-8655, Japan c

Medical R&D Center, Corporate R&D Group, KYOCERA Corporation, 800 Ichimiyake, Yasu 520-2362,

Japan d

Sensory & Motor System Medicine, Faculty of Medicine; The University of Tokyo, 7-3-1 Hongo,

Bunkyo-ku, Tokyo 113-8655, Japan e

Materials Research and Development Laboratory, Japan Fine Ceramics Center, 2-4-1 Mutsuno, Atsuta-ku,

Nagoya 456-8587, Japan

Keywords: Joint replacement; Polyethylene; Phosphorylcholine; Wear mechanism; Hydration; Lubrication

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ABSTRACT Recently, traditional strategies for manipulating orthopedic bearing substrates have attempted to improve their wear resistance by adjusting polyethylene substrate through cross-linking and antioxidant blending. However, further research is required on the substrate, as well as the surface focused on the structure and role of articular cartilage. We therefore develop an orthopedic bearing surface comprising a nanometer-scale hydrated gel-like layer by grafting highly hydrophilic poly(2-methacryloyloxyethyl phosphorylcholine), with the aim of mimicking the lubrication mechanism of articular cartilage, and investigate its surface characteristics, bulk characteristics, and behavior under load bearing conditions upon accelerated aging. Neither the hydrophilicity nor lubricity of the gel-like surface was influenced by accelerated aging; instead, high stability was revealed, even under strongly oxidation conditions. The characteristics of the hydrated gel-like surface potentiated the wear resistance of the cross-linked polyethylene liner, irrespective of accelerated aging. These results suggest that the hydrated gel-like surface enhances the longevity of cross-linked polyethylene bearings even under load-bearing conditions. Furthermore, the inflection point on the time series of wear can be a suitable indicator of the durability of the life-long protectant. In conclusion, the hydrated gel-like surface can positively increase orthopedic implant durability.


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INTRODUCTION The modification of surface characteristics using chemically, mechanically, or biologically suitable materials is important for preparing new multifunctional biomaterials and improving their overall performance in medical and healthcare applications. Furthermore, engineering the surface and/or interface of bioinspired and zwitterionic materials is as a highly effective and increasingly adopted strategy, particularly for cellular or tissue response at the interface of biological environments.

Specifically,

2-methacryloyloxyethyl phosphorylcholine (MPC), which is a methacrylate monomer with a phospholipid polar group, and its polymers have attracted considerable interest as surface modifiable bioinspired and zwitterionic polymers for various potential healthcare and medical applications because of their unique properties like high lubricity, good moisture retention, and good biocompatibility [1]. Several healthcare and medical products utilizing the MPC polymers have already been developed and clinically implemented; therefore, the safety and efficacy of MPC polymers as biomaterials have been well demonstrated [2]. Due to an aging population, locomotorium related diseases have increased rapidly worldwide. The repair of locomotorium related diseases remains one of the most challenging topics in clinical, biomaterial, and bioengineering fields. Total hip arthroplasty (THA) is widely recognized as successful and effective for the treatment of degenerative hip joint disease that leads to decreased daily activity and quality of life. The number of annual primary and revision THA procedures has increased significantly despite advances in surgical techniques and implant designs [3]. Aseptic loosening caused by periprosthetic osteolysis is one of the major complications of THA, limiting the clinical outcomes following revision surgery [4]. For instance, the Australian Orthopaedic Association National Joint Replacement Registry reported that the number of revision surgeries increased annually and the revision rate was shown to be 6.0% after 15 years postoperation in their latest report, even for current primary THAs, which use highly cross-linked polyethylene (CLPE) liners against large diameter femoral heads (75.9% have a diameter of 32 mm or greater) [5]. The surfaces of a natural synovial joint are covered with a specialized type of hyaline cartilage, i.e., articular cartilage. The articular cartilage is composed of a dense extracellular matrix (e.g. proteoglycans, 3 Environment ACS Paragon Plus

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collagens, and water) with sparsely distributed chondrocytes and surface-active phospholipids (e.g. phosphatidylcholine derivatives), which protect the joint interface from mechanical wear and facilitates the smooth motion of joints during daily activity [6]. Due to their charge (i.e., zwitterionic or alternating-charge), they can trap water to maintain the water-fluid and electrolyte balance in the articular cartilages, which provides hydrophilicity and works as an effective boundary lubricant [7]. Fluid thin-film lubrication by the hydrated polyelectrolyte layer of articular cartilage, known as hydration lubrication, is essential for the smooth motion of natural synovial joints. Those low friction and wear processes caused by hydration lubrication are sufficiently complex that engineered prevention of mechanical wear and facilitation of smooth motion in high load bearing interface remains an elusive goal. However, it is key to enhancing our understanding of low friction and wear processes caused by hydration lubrication at the orthopedic bearing interface and the downstream and/or spin-off benefits for the development pathways of biomaterials and related artificial joints. Biomimetic design is a successful approach for designing the surface and bulk of biomaterials; thus, the bearing surface modification for artificial joints should be investigated to mimic the structure and role of articular cartilage. To this end, we have developed an articular-cartilage-mimicked technology that allows bearing surface modification of the acetabular CLPE liners used in artificial hip joints by poly(MPC) (PMPC) grafting for the prevention of osteolysis and aseptic loosening [8–10]. Our previous study on the function and efficacy of a nanometer-scale hydrated gel-like PMPC graft layer (~100 nm thick [11]), which would be cross-linked upon gamma-ray irradiation after PMPC grafting [12], revealed that such a hydrated gel-like layer greatly enhances the wear resistance of the CLPE bearings [9]. On the other hand, the function and efficacy of the PMPC graft layer may be sometimes eradicated from the bearing surface because in vivo joint bearing is a harsh environment characterized by repetitive activity with several times higher vertical and shear loadings than the subject’s body weight [13,14]. The purpose of this study is therefore to highlight the role and/or effects of a nanometer-scaled hydrated gel-like surface (PMPC-grafted surface), even in severe environments like load bearing states undergoing 4 Environment ACS Paragon Plus

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accelerated aging. In addition to obtaining life-long orthopedic bearings, we also aim to provide evidence that the hydrated gel-like surface acts as a life-long protectant and increases the lifespan of orthopedic bearing materials. Thus, we address whether (1) the PMPC-grafted surface produces hydration lubrication characteristics despite accelerated aging, (2) the PMPC-grafted surface can enhance wear and fatigue resistance under load bearing conditions despite accelerated aging, and (3) the counter bearing material (i.e. ceramic femoral head) can compensate for the defective wear resistance of PMPC-grafted CLPE liners.

MATERIALS AND METHODS Chemical surface modification by PMPC grafting Benzophenone (BP) and acetone were purchased from Wako Pure Chemical Industries, Ltd. (Osaka, Japan). Industrially synthesized MPC (purity ≥ 98.0%) was purchased from NOF Corp. (Tokyo, Japan). A compression-molded polyethylene (PE) bar stock (GUR1020 resin, Orthoplastics Ltd, Lancashire, UK) was irradiated with a dose (50 kGy) of gamma-rays in a N2 gas atmosphere and annealed at 120 °C for 7.5 h in N2 gas in order to facilitate cross-linking; this PE material is hereafter referred to as CLPE. CLPE samples were then machined from the bar stocks after cooling, washed with ethanol, and dried at room temperature for 1 h in a vacuum. The CLPE samples were immersed for 30 s in acetone containing 10 mg/mL BP, and then dried in the dark at room temperature to remove the acetone. MPC was dissolved in degassed pure water to a concentration of 0.5 mol/L [11]. The BP-coated CLPE samples were then immersed in the aqueous MPC solution. Photoinduced-radical graft polymerization was carried out on the CLPE surfaces using an ultraviolet (UV) irradiation (UVL-400HA ultra-high-pressure mercury lamp; Riko-Kagaku Sangyo Co., Ltd., Funabashi, Japan) with an intensity of 5 mW/cm2 at 60 °C for 90 min [15,16]. A filter (model D-35; Toshiba Corp., Tokyo, Japan) was used to permit the sole passage of UV light with a wavelength of 350 ± 50 nm. After polymerization, the PMPC-grafted CLPE samples were removed, washed with pure water and ethanol, and dried at room temperature for 1 h in a vacuum. 5 Environment ACS Paragon Plus

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The untreated CLPE and PMPC-grafted CLPE samples were sterilized with a 25-kGy dose of gamma-ray irradiation under N2 gas, then subjected to conditions of accelerated aging, i.e., exposure to 80 °C in air for 21 or 63 days (3 or 9 weeks) [17]. This practice, referring to the ASTM F2003 standard, was not intended to simulate any change that may occur in the untreated CLPE and PMPC-grafted CLPE liner following implantation, and also has not been directly correlated with the shelf life of components that have been sealed in a low-oxygen package, such as N2 gas, but is simply meaningful to experimentally generate structural changes of the CLPE substrate.

Oxidative degradation tests The oxidative degradation (oxidation index) of the CLPE and PMPC-grafted CLPE samples before and after accelerated aging was evaluated by Fourier-transform infrared (FT-IR) microspectroscopy according to the ASTM F2102 standard. A thin film (100–200 µm thick) of the cross-section was sliced from each of the samples. FT-IR spectra were obtained using a microscopic FT-IR analyzer (Spectrum BX, Perkin-Elmer Corp., MA, USA) for 100 scans over a range of 800–2000 cm–1 at a resolution of 4 cm–1. The oxidation index was defined as the ratio of the carbonyl peak area at 1720 cm–1 to the methylene peak area at 1360 cm–1. We evaluated four films for each sample surface (0–100 µm from the surface), and the mean values were calculated for the oxidation index.

Wettability and friction tests Static-water contact angles of the CLPE and PMPC-grafted CLPE samples before and after accelerated aging were measured by employing the sessile-drop method using an optical-bench-type contact-angle goniometer (Model DM300, Kyowa Interface Science Co., Ltd., Saitama, Japan). Drops of purified water (1 µL) were deposited on the surface of each sample and the contact angles were directly measured after 60 s using a microscope. For each sample, 15 areas were evaluated with the mean values reported for the static contact angles of water. 6 Environment ACS Paragon Plus

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Unidirectional friction tests were performed using a ball-on-plate machine (Tribostation 32, Shinto Scientific Co., Ltd., Tokyo, Japan). A total of 10 specimens of each of the CLPE and PMPC-grafted CLPE samples before and after accelerated aging were evaluated. A 9-mm-diameter pin made of a Co–Cr–Mo alloy was also prepared. The surface roughness (Ra) of the pin was < 0.01 µm, comparable to that of currently used femoral head products. The friction test was performed for each specimen at room temperature using a load of 0.98 N (the contact stress roughly calculated by Hertzian theory was approximately 25 MPa), a sliding distance of 25 mm, and a frequency of 1 Hz. A maximum of 100 cycles were performed, using pure water for lubrication. The mean dynamic coefficients of friction were determined by averaging the values of five data points taken from the 96–100th cycles.

Swelling tests The swelling ratio and cross-link density of the CLPE and PMPC-grafted CLPE substrates that underwent accelerated aging were evaluated according to previously reported methods [16]. We divided three sample pieces (23 × 23 × 1 mm3) from the surface (0–1 mm depth) and bulk (3–4 mm depth) of each the CLPE and PMPC-grafted CLPE substrates before and after accelerated aging. The sample pieces were allowed to swell for 72 h in p-xylene containing 0.5 mass% 2-t-butyl-4-methylphenol at 130 °C, then immersed in acetone and dried at 60 °C under vacuum. The swelling ratio (w/w) was determined from the increase in the weights of samples before and after swelling. The network chain density was calculated using the Flory-Rehner equation, and the cross-link density was defined as the mole fraction of the cross-linked units [18].

Mechanical tests The mechanical properties of the CLPE and PMPC-grafted CLPE samples that underwent accelerated aging were evaluated using a series of tests. A double-notched (notch depth = 4.57 ± 0.08 mm) Izod impact test was performed according to the ASTM F648 standard, with six samples tested for each of the CLPE and PMPC-grafted CLPE substrates before and after accelerated aging. A small punch test was performed 7 Environment ACS Paragon Plus

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according to the ASTM F2183 standard, using a disk specimen with a diameter and thickness of 6.4 mm and 0.5 mm, respectively, and a crosshead speed of 0.5 mm/min. We divided five sample pieces from the surface (0–0.5 mm depth) and bulk (1.5–2.0 mm depth) of each of the CLPE and PMPC-grafted CLPE substrates before and after accelerated aging.

Impact-to-wear tests Impact-to-wear tests were conducted using a pin-on-disk testing machine (Ortho POD; AMTI, Watertown, MA, USA), according to the ASTM F732 standard. CLPE and PMPC-grafted CLPE disks that underwent accelerated aging were used for the wear and control soak (for correction of water-absorption increments) tests (n = 3, respectively). Disks with a thickness of 6 mm were attached to the pin-on-disk testing machine with a titanium-aluminum-vanadium alloy (Ti–6Al–4V) fixation component that had an 8-mm diameter hole to simulate an acetabular shell with a screw hole [19]. The Co–Cr–Mo alloy pins had a 30-mm surface curvature radius and surface roughness of Ra < 0.01 µm. An aqueous mixture of 27 vol% fetal bovine serum (Biowest, Nuaille, France) containing 20 mmol/L ethylenediaminetetraacetic acid (EDTA) and 0.1 mass% sodium azide was used at 37 °C as the lubricant. Testing was performed on a frequent unidirectional sliding setting with a maximum impact load of 150 N, sliding distance of 10 mm, and frequency of 1 Hz for a maximum of 2.0 × 106 cycles. Gravimetric wear was determined by weighing the disks. Soak controls were used to compensate for fluid absorption by the specimens. However, this correction was not considered to be perfect because only the tested disks were continuously moved and subjected to the load. After the impact-to-wear test, the sliding surface morphology of the disks was evaluated using a noncontact optical three-dimensional (3D) profiler (Talysurf CCI Lite; Taylor Hobson Ltd., Leicester, UK) with a green light-emitting diode as the light source at ×10 magnification. The inner fractures of the disks were scanned by micro computed tomography (µCT, InspeXio; Shimadzu Corp., Kyoto, Japan) with a resolution of 0.03 mm per pixel. The 3D images were reconstructed using 3D image processing software (TRI/3D-BON-C; RATOC system engineering co., ltd., Tokyo, Japan). 8 Environment ACS Paragon Plus

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Hip-simulator wear test A 12-station hip simulator (MTS Systems Corp., Eden Prairie, MN, USA) using the CLPE and PMPC-grafted CLPE liners before and after accelerated aging of 3 weeks was used for the wear test according to the ISO 14242-3 standard. Four test specimens of each of the CLPE, aged CLPE, PMPC-grafted CLPE, and eight test specimens of the aged PMPC-grafted CLPE liners were prepared with an inner and outer diameter of 32 and 48 mm, respectively. A Co–Cr–Mo alloy (910 metal; KYOCERA Corp., Kyoto, Japan) and zirconia-toughened alumina (ZTA, BIOCERAM AZ209; KYOCERA Corp.) femoral heads measuring 32 mm in diameter were used as the femoral head. An aqueous mixture of 25 vol% fetal bovine serum (Bio west), 20 mmol/L EDTA, and 0.1 mass% sodium azide was used as the lubricant. The lubricant was replaced every 5.0 × 105 cycles. Gait cycles were applied to simulate a physiological loading curve (Paul-type) with double peaks at 1793 and 2744 N and multidirectional (biaxial and orbital) motion with a frequency of 1 Hz. Testing was continued for 1.0 × 107 cycles. Gravimetric wear was determined by weighing the liners at intervals of every 5.0 × 105 cycles. Load-soak controls (n = 2) were used to compensate for fluid absorption by the specimens, according to the ISO 14242-2 standard. The weight loss of each of the tested liners was corrected by subtracting the weight gain resulting from the load-soak control.

Statistical analyses The mean values of the six groups (CLPE [pre], 3-weeks aged CLPE, 9-weeks aged CLPE, PMPC-grafted CLPE [pre], 3-weeks aged PMPC-grafted CLPE, and 9-weeks aged PMPC-grafted CLPE) or the five groups (CLPE [pre] against Co–Cr–Mo alloy, 3-weeks aged CLPE against Co–Cr–Mo alloy, PMPC-grafted CLPE [pre] against Co–Cr–Mo alloy, 3-weeks aged PMPC-grafted CLPE against Co–Cr–Mo alloy, and 3-weeks aged PMPC-grafted CLPE against ZTA) were compared by one-factor analysis of variance (ANOVA), and the significant differences of all comparable properties were determined by post-hoc testing using the Bonferroni method. All statistical analyses were performed using an add-on (Statcel 4, OMS Publishing Inc., 9 Environment ACS Paragon Plus

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Tokorozawa, Japan) to Microsoft Excel® 2013 (Microsoft Corp., Redmond, WA, USA).

RESULTS AND DISCUSSION Accelerated aging increased the oxidative degradation of the CLPE and PMPC-grafted CLPE surface (Fig. 1A). Surface oxidation indices of the CLPE and PMPC-grafted CLPE surfaces after 3 weeks of accelerated aging were 2.1 and 1.7, respectively, and those values were in the moderate-to-severe (1 < oxidation index ≤ 3) range of oxidation. The indices after 9-weeks accelerated aging were extremely high over the threshold of critical oxidation (oxidation index > 3). The aging also slightly affected the oxidative degradation of the CLPE and PMPC-grafted CLPE substrates (Fig. 1B). However, all bulk oxidation indices of the substrates were below the low oxidation threshold (oxidation index < 1). Accelerated aging affected the hydrophilicity of the PMPC-grafted CLPE surface, albeit only slightly. The static-water contact angle on the CLPE surface decreased markedly from ~95° to ~20° after PMPC grafting (Fig. 2A). There was no significant difference in the static-water contact angle among the non-aged, 3-weeks-aged, and 9-weeks-aged CLPE. Conversely, the contact angle of 9-weeks aged PMPC-grafted CLPE was slightly but statistically higher compared with non-aged, 3-weeks aged PMPC-grafted CLPE. On the other hand, accelerated aging did not affect the lubricity (i.e. friction responses) of the CLPE surface and PMPC-grafted CLPE surfaces. The dynamic coefficients of friction between the Co–Cr–Mo alloy and the PMPC-grafted CLPE surface decreased markedly despite accelerated aging; the interfaces exhibited an ~70% reduction in the coefficients compared to the interfaces with untreated CLPE (Fig. 2B). No significant difference was observed in the dynamic coefficients of friction among the non-aged, 3-weeks-aged, and 9-weeks-aged samples. Accelerated aging greatly affected the physical properties of both CLPE and PMPC-grafted CLPE, especially on the surface (Fig. 3). The surface cross-link densities of both CLPE and PMPC-grafted CLPE decreased significantly with accelerated aging. The bulk (substrate) cross-link density was ~30% less with an accelerated aging period of 9 weeks. There was no significant difference in these surface and bulk cross-link 10 Environment ACS Paragon Plus

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densities between CLPE and PMPC-grafted CLPE. No differences in mechanical properties were observed between CLPE and PMPC-grafted CLPE, regardless of accelerated aging (Fig. 4). Severe accelerated aging for 9 weeks deteriorated the mechanical properties of both CLPE and PMPC-grafted CLPE. The impact strength of CLPE and PMPC-grafted CLPE did not differ significantly before and after accelerated aging of 3 weeks. The impact strength was approximately 45% lower with an accelerated aging of 9 weeks. Notably, the work to failure of CLPE and PMPC-grafted CLPE surface decreased significantly (to almost zero) with accelerated aging of 9 weeks, although the properties of CLPE and PMPC-grafted CLPE substrate did not differ during accelerated aging. The hydrated gel-like PMPC graft layer did not affect the impact fatigue resistance of CLPE until 3 weeks of accelerated aging in the impact-to-wear test with 2.0 × 106 cycles. However, the PMPC-grafted CLPE with an accelerated aging of 9 weeks statistically improved the resistance although the wear rate was increased to several hundred times (Fig. 5). In terms of the three distinctive samples of the PMPC-grafted CLPE disks with an accelerated aging of 9 weeks, the behavior of wear and/or impact fatigue becomes clearer (Fig. 6). Notably, the gravimetric wear of samples 1 and 2 started to increase after 2.5 × 105 and 1.0 × 106 cycles, respectively. In contrast, the gravimetric wear of sample 3 remained at almost zero (actual value is 0.31 mg/106 cycles). After impact loads of 2.0 × 106 cycles, we observed severe mechanical fracture and delamination in the sliding surface of samples 1 and 2; conversely, we clearly observed a crack at the sliding sub-surface (approximately 0.5 mm in depth) of sample 3, although we did not observe a delamination. The hydrated gel-like PMPC graft layer increased the durability of the CLPE liner, irrespective of the duration of accelerated aging, in the hip-simulator wear test with 1.0 × 107 cycles. As seen in Fig. 7, the accelerated aging increased the gravimetric wear of the untreated CLPE liners; its wear rate more than tripled. The aging also increased the gravimetric wear of the PMPC-grafted CLPE liners; however, the wear rate was approximately 7.6 mg/106 cycles, representing a 75% reduction compared with untreated CLPE liner. Notably, the wear rate of the PMPC-grafted CLPE liner against the ZTA femoral head was extremely low even after accelerated aging of 3 weeks, representing an 85% reduction compared with the untreated CLPE 11 Environment ACS Paragon Plus

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liner (without aging) against the Co–Cr–Mo alloy femoral head. Traditional strategies for manipulating orthopedic bearing substrates have focused on improving their wear resistance by adjusting physical properties such as molecular weight, crystallinity, and the cross-link density of PE. However, a more sophisticated understanding of the wear mechanism of a hydrophilic gel-like surface that mimics the vital role of articular cartilage could lead to the development of life-long artificial joints. Thus, biomaterials-based strategies for manipulating orthopedic bearings should target both the substrate and its surface in order to enhance durability. The results of this study indicate that a hydrated gel-like surface affects the extent of the wear resistance of the orthopedic bearing materials; in particular, the hydrated gel-like surface can ensure that CLPE bearings last a lifetime, even under load bearing conditions. The hydration and friction kinetics of the bearing surface are key to the underlying wear mechanism. The hydrated gel-like layer on the bearing surface can most effectively mimic the lubrication mechanism of articular cartilage due to its ability to manipulate orthopedic implants via PMPC grafting. The aging did not affect the hydrophilicity and lubrication characteristics of the PMPC graft layer: the PMPC-grafted CLPE surfaces were considerably more hydrophilic than the untreated CLPE surfaces, even though the CLPE substrate underwent oxidative degradation upon accelerated aging (Fig. 2A). This significant increase in hydrophilicity, due to the presence of a highly hydrophilic PMPC graft layer, is evident from the reduction in static-water contact angles on the PMPC-grafted surfaces [8,11]. Further, the PMPC-grafted surfaces with a hydrated gel-like layer influenced the friction kinetics. Fig. 2B shows that the PMPC-grafted surface had statistically lower dynamic coefficients of friction than the untreated CLPE surface despite oxidative degradation of CLPE substrates upon accelerated aging. Therefore, the hydrated gel-like surface, which achieved hydration lubrication by mimicking the lubrication mechanism of articular cartilage, is considered essential for life-long orthopedic bearings. Notably, the hydrated gel-like surface can effectively influence wear characteristics even in various controlled load bearing conditions. The majority of our previous studies only reported that the hydrated gel-like surface is more conducive to low friction behavior and/or mechanisms at the bearing interface, as 12 Environment ACS Paragon Plus

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mentioned above, although PMPC-grafted CLPE liners exhibit higher resistance to wear than untreated CLPE liners [8,9,15]. This indicates that the hydrated gel-like surface was not directly correlated with high wear resistance in previous studies. However, the results of this study (especially Fig. 6) reveal that the hydrated gel-like surface was directly correlated with not only low friction but also high wear resistance under a load bearing conditions. After accelerated aging of 9 weeks, both untreated and PMPC-grafted CLPE lost the bulk properties of the substrates owing to oxidative degradation. If it is assumed that the hydrated gel-like surface does not influence wear characteristics in load bearing conditions, the PMPC-grafted CLPE should show acute wear for the same behavior of the untreated CLPE. The acute wear behavior of 9-weeks aged CLPE met our assumption, whereas that of 9-weeks aged PMPC-grafted CLPE did not. The hydrated gel-like surface of PMPC-grafted CLPE delayed the onset of acute wear, indicating that the hydrated gel-like PMPC graft layer will act as a life-long protectant for CLPE bearing surfaces due to a difference in the surface chemical and surface physical phenomena between untreated CLPE and PMPC-grafted CLPE. Conversely, the hydrated gel-like layer would gradually be worn away from the CLPE bearing surfaces during a test duration. Furthermore, the testing cycle was detected upon removing the layer but only in impact-to-wear tests using a pin-on-disk testing machine with 9-weeks aged samples. In a previous hip simulator study [20], we found that the defective hydrated gel-like surface (i.e., low graft chain density) of PMPC-grafted CLPE delayed the onset of wear; there was a gravimetric wear relationship between untreated CLPE and PMPC-grafted CLPE when the PMPC-graft layer was removed from the surface. These results suggested that the inflection point on a time series of wear would be a suitable indicator for the durability of the PMPC graft layer. It can be assumed that cyclic impact loadings occur during walking on a daily routine for orthopedic bearings, such as heel strike and toe-off, particularly in young active patients. Therefore, we considered that the cyclic impact loadings were mainly received by the CLPE substrate in artificial hip joints, such that the quality of the CLPE substrate becomes increasingly important. Saiga et al. [21] reported previously that the impact-to-wear test was performed with a cycle composed of three types of motions: impact motion, sliding 13 Environment ACS Paragon Plus

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motion with a lower pressure, and sliding motion with a higher pressure. For the aged CLPE or PMPC-grafted CLPE bearings with different bulk properties (i.e., with and without degradation) of the surface and substrate (or sub-surface) in this study, sliding motion with a higher pressure, which mimics toe-off wear behavior, was found to be more aggressive than other motions owing to the wear phenomena observed as cracks at the sliding sub-surface of sample 3. The steady wear rate (17.21 and 17.97 mg/106 cycles, respectively) during 1.0 × 106–2.0 × 106 cycles of the impact-to-wear test did not differ significantly (p > 0.05) between the CLPE and PMPC-grafted CLPE bearing. This indicates that the worst in vivo scenario is that only the new surface of the CLPE substrate is exposed when the PMPC graft layer and oxidized CLPE layer disappear from the bearing surface. The advantage of photoinduced graft polymerization becomes apparent from the fact that PMPC grafting generated high hydrophilicity and lubricity only on the surface and had no effect on the bulk characteristics of the CLPE substrate [22]. Retention of the bulk substrate properties is important in clinical applications because the orthopedic bearings must fulfill the requirements of not only surface-functional materials but also structural materials in vivo. Generally, methacrylate polymers have high resistance to thermal- and oxidative-degradations, which may precede hydrolysis [23]. Furthermore, we previously reported that aging did not affect the hydrophilicity and lubricity of the PMPC graft layer (the melting point of MPC is 138–141 °C) on the CLPE substrate blended with the antioxidant vitamin E (α-tocopherol) [24]. In this study, the PMPC also showed very high stability against thermal- and oxidative-degradation and even cross-linking with a behavior similar to that of other methacrylate polymers, and the thick graft layer formed on the CLPE surface delayed surface oxidation (Fig. 1). On the other hand, accelerated aging remarkably deteriorated the CLPE substrate in this study, although the PMPC layer prevented surface oxidation; however, the oxidative degradation of the CLPE substrate was not directly correlated with shelf life aging and in vivo aging. Oxidation of PE comprises a free radical-initiated chemical reaction and is expected to result in molecular chain scission [25]. A sequential decrease in molecular weight and cross-link density (Fig. 3) or an increase in crystallinity would be facilitated by oxidative degradation of the PE or CLPE [26]. Such changes in the chemical and physical structures 14 Environment ACS Paragon Plus

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compromise their mechanical properties (Fig. 4). In fact, the clinical impact of oxidative degradation of a CLPE bearing surface is not clear because the oxidative degradation does not govern clinical failures such as wear or osteolysis; however, in vivo oxidative degradation is generally regarded as undesirable. Previous studies have reported similar results in which the femoral head limited the access of body fluids acting as oxygen and/or reactive oxygen species carriers into the bearing interface and prevented oxidative degradation of the CLPE bearing surface [27,28]. In comparison, the wear rate (7.57 mg/106 cycles) of the aged PMPC-grafted CLPE liners against Co–Cr–Mo alloy femoral heads used in this study was slightly higher than that (3.63 mg/106 cycles) of the non-aged PMPC-grafted CLPE liners, although thermal accelerated aged CLPE and PMPC-grafted CLPE liners (according to the ASTM F2003 standard) of this hip simulator test are not directly correlated with both shelf life aging and in vivo aging (or in vivo degradation), as mentioned above. Ceramic femoral heads have generally been considered an alternative bearing surface to reduce wear. Similarly, the results of this study (Fig. 7) substantiate that ceramic femoral heads are an appropriate clinical approach for the PMPC-grafted CLPE liner. Although ceramic femoral heads may increase the risk for implant failure (although the rate of fracture of ZTA ceramic femoral heads was merely 0.001–0.003% [29]), we believe that this imperfection or defect is partially offset by the bearing interface with ceramics due to the stable lower wear rate than that obtained using Co–Cr–Mo alloy. Further, ceramic femoral heads have shown markedly less fretting-initiated crevice corrosion compared to Co–Cr–Mo alloy femoral heads, with a lower potential for metal debris and ion release leading to adverse local tissue reactions [30]. Therefore, as for further efforts, ceramic femoral heads will be used to improve wear resistance, as an alternative to Co–Cr–Mo alloy femoral heads, due to the reduction in the unexpected increase of wear caused by in vivo degradation. Notwithstanding these promising results, the present study has several limitations. First, preclinical findings do not always correspond to clinical results; we could not completely obtain the clinical efficacy, e.g. wear and oxidative degradation resistances, of PMPC-grafted CLPE liners. The clinical trials of PMPC-grafted CLPE liners were conducted at multiple medical centers between 2007 and 2009 in Japan [10]. Based on the 15 Environment ACS Paragon Plus

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preclinical evidence and the results of these clinical trials, the Ministry of Health, Labor, and Welfare, a government organization in Japan, approved the clinical use of PMPC-grafted CLPE acetabular liners (Aquala; KYOCERA Corp.) in artificial hip joints in 2011. We observed neither osteolysis nor implant fracture, which resulted in revision surgery, during 10 years of follow-up in the clinical trials; the mean steady-state wear rate was 0.002 mm/year [10]. On the other hand, Tone et al. [13] reported the oxidative degradation of PMPC-grafted CLPE acetabular liners in their retrieval study. In fact, the clinical impact of oxidation degradation remains unclear because (1) only few retrieved acetabular liners were available for the study, (2) we could not estimate the amount of difference in some properties that would be effective in an actual clinical situation, and (3) several parts of that retrieval study are inaccurate and unreliable owing to data contamination. Although its clinical significance is still the subject of scientific debate, in vivo oxidation is regarded as undesirable. To gain a better understanding of the retrieved PMPC-grafted CLPE acetabular liner, a large number of samples need to be assessed [14]. Further long-term follow-up study (UMIN000003681) for the clinical trials and the analysis of retrospective data gathered for over 45,000 clinical applications (at April, 2018) are required to confirm the validity of the in vivo performance of the PMPC-grafted CLPE acetabular liner. Second, although we prepared the thermal accelerated aged CLPE and PMPC-grafted CLPE samples according to the ASTM F2003 standard, the practices were not based on autoxidation (free radical chain process) of the in vivo environment with mechanical stress, heat, and lipid absorption [31]. Those thermal-accelerated aged CLPE and PMPC-grafted CLPE samples are meaningful CLPE substrates that have experimentally undergone structural changes, even though they do not simulate in vivo stability/degradation during implantation. Therefore, we will appeal to the field of orthopedic biomaterials that new standardized testing methods are necessary to simulate in vivo stability/degradation more accurately by controlling the complex factors. Third, we could not completely duplicate the motion and loading conditions during the impact-to-wear test and the hip-simulator wear test in terms of the variety of positions and the daily routine of the subjects, or the direction and range of loading. Since the specific motion and loading conditions were in accordance with the ASTM F732 and ISO 14242-3 standards, respectively, these results probably play an 16 Environment ACS Paragon Plus

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important role in the design of orthopedic bearing materials and in the continual assessment of current orthopedic implants. On the other hand, several findings have shown that for various reasons (e.g., differences in bearing designs and materials, unintended implant positioning, soft tissue laxity, additional patient range of motion, increased loads), unintended conditions such as edge loading can occur and the consequences can be severe, possibly leading to implant failure [4,5]. Many factors contribute to the occurrence of edge loading conditions. In regards to this issue, the ISO technical committee will finish working and discussing issues regarding severe-wear test methods with direct edge loading, according to a new standard (ISO 14242-4). Further related studies will be required in the near future.

CONCLUSIONS In this study, we developed an orthopedic bearing surface with a nanometer-scale gel-like layer by grafting highly hydrophilic PMPC, and subjecting the bearing material to conditions of accelerated aging. The surface and bulk characteristics were then investigated under load bearing conditions. Neither the static-water contact angle nor the dynamic coefficient of friction of the PMPC-grafted surface were influenced by accelerated aging; therefore, it exhibited high oxidative stability. The wear resistance of the CLPE surface was improved by the hydrated gel-like layer regardless of accelerated aging although the hydrated gel-like layer did not appear to affect fatigue resistance. Notably, the hydrated gel-like surface was found to enhance the lifetime of CLPE bearings even in load bearing conditions, and the inflection point on the time series of wear can indicate its durability. The hydrated gel-like surface will therefore positively affect the durability of orthopedic implants.

AUTHOR INFORMATION Corresponding Author* Kazuhiko Ishihara Department of Materials Engineering, School of Engineering, The University of Tokyo 17 Environment ACS Paragon Plus

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Hongo 7-3-1, Bunkyo-ku, Tokyo 113-8656, Japan Phone: +81-3-5841-7124, Fax: +81-3-5841-8647 E-mail address: [email protected]

Author Contributions The manuscript was written through contributions of all authors. All authors have given approval to the final version of the manuscript. †These authors contributed equally. Masayuki Kyomoto† Toru Moro† Shihori Yamane† Kenichi Watanabe† Masami Hashimoto† Sakae Tanaka† Kazuhiko Ishihara†

Acknowledgments We thank Dr. Yoshio Takatori of The University of Tokyo for valuable discussions and suggestions. We also thank Dr. Nobuaki Moriguchi, Mr. Kenichi Saiga, and Mr. Takashi Sasaki of KYOCERA Corporation for their technical assistance.


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References [1] Ishihara, K.; Ueda, T.; Nakabayashi, N. Preparation of Phospholipid Polymers and Their Properties as Polymer Hydrogel Membranes. Polym. J. 1990, 22, 355–360. [2] Kyomoto, M.; Moro, T.; Ishihara, K. In Phospholipid polymer grafted highly cross-linked UHMWPE, UHMWPE Biomaterials Handbook Third Edition, Kurtz, S. M. (Ed.),William Andrew: Oxford, 2015; pp 352–368. [3] Bozic, K. J.; Kamath, A. F.; Ong, K.; Lau, E.; Kurtz, S.; Chan, V.; Vail, T. P.; Rubash, H.; Berry, D. J. Comparative Epidemiology of Revision Arthroplasty: Failed THA Poses Greater Clinical and Economic Burdens Than Failed TKA. Clin. Orthop. Relat. Res. 2015, 473, 2131–2138. [4] Bozic, K. J.; Kurtz, S. M.; Lau, E.; Ong, K.; Vail, T. P.; Berry, D. J. The Epidemiology of Revision Total Hip Arthroplasty in the United States. J. Bone Joint Surg. Am. 2009, 91, 128–133. [5] Graves, S.; Turner, C. In Annual Report 2017, Australian Orthopaedic Association National Joint Replacement Registry, Australian Orthopaedic Association: Adelaide, 2017. [6] Kirk, T. B.; Wilson, A. S.; Stachowiak, G. W. The Morphology and Composition of the Superficial Zone of Mammalian Articular Cartilage. J. Orthop. Rheumatol. 1993, 6, 21–28. [7] Goldberg, R.; Schroeder, A.; Silbert, G.; Turjeman, K.; Barenholz, Y.; Klein, J. Boundary Lubricants with Exceptionally Low Friction Coefficients Based on 2D Close-Packed Phosphatidylcholine Liposomes. Adv. Mater. 2011, 23, 3517–3521. [8] Moro, T.; Takatori, Y.; Ishihara, K.; Konno, T.; Takigawa, Y.; Matsushita, T.; Chung, U. I.; Nakamura, K.; Kawaguchi, H. Surface Grafting of Artificial Joints with a Biocompatible Polymer for Preventing Periprosthetic Osteolysis. Nat. Mater. 2004, 3, 829–837. [9] Moro, T.; Takatori, Y.; Kyomoto, M.; Ishihara, K.; Hashimoto, M.; Ito, H.; Tanaka, T.; Oshima, H.; Tanaka, S.; Kawaguchi, H. Long-Term Hip Simulator Testing of the Artificial Hip Joint Bearing Surface Grafted with Biocompatible Phospholipid Polymer. J. Orthop. Res. 2014, 32, 369–376. [10] Moro, T.; Takatori, Y.; Tanaka, S.; Ishihara, K.; Oda, H.; Kim, Y. T.; Umeyama, T.; Fukatani, E.; Ito, H.; 19 Environment ACS Paragon Plus

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Kyomoto, M.; Oshima, H.; Tanaka, T.; Kawaguchi, H.; Nakamura, K. Clinical Safety and Wear Resistance of the Phospholipid Polymer-Grafted Highly Cross-Linked Polyethylene Liner. J. Orthop. Res. 2017, 35, 2007–2016. [11] Kyomoto, M.; Moro, T.; Miyaji, F.; Hashimoto, M.; Kawaguchi, H.; Takatori, Y.; Nakamura, K.; Ishihara, K. Effect of 2-Methacryloyloxyethyl Phosphorylcholine Concentration on Photo-induced Graft Polymerization of Polyethylene in Reducing the Wear of Orthopaedic Bearing Surface. J. Biomed. Mater. Res. A 2008, 86, 439–447. [12] Kyomoto, M.; Moro, T.; Miyaji, F.; Konno, T.; Hashimoto, M.; Kawaguchi, H.; Takatori, Y.; Nakamura, K.; Ishihara, K. Enhanced wear resistance of orthopaedic bearing due to the cross-linking of poly(MPC) graft chains induced by gamma-ray irradiation. J. Biomed. Mater. Res. B Appl. Biomater. 2008, 84, 320– 327. Erratum in: J. Biomed. Mater. Res. B Appl. Biomater. 2008, 85, 301. [13] Tone, S.; Hasegawa, M.; Puppulin, L.; Pezzotti, G.; Sudo, A. Surface Modifications and Oxidative Degradation in MPC-Grafted Highly Cross-Linked Polyethylene Liners Retrieved from Short-Term Total Hip Arthroplasty. Acta Biomater. 2018, 66, 157–165. [14] Sakoda, H.; Okamoto, Y.; Haishima, Y.; Sugano, N. Methods to Evaluate the Presence of Hydrophilic Modification Layer on the Surface of Retrieved Acetabular Liner. Proc. Orthop. Res. Soc. U.S.A. 2018, 1800. [15] Kyomoto, M.; Moro, T.; Konno, T.; Takadama, H.; Yamawaki, N.; Kawaguchi, H.; Takatori, Y.; Nakamura, K.; Ishihara, K. Enhanced Wear Resistance of Modified Cross-Linked Polyethylene by Grafting with Poly(2-methacryloyloxyethyl phosphorylcholine). J. Biomed. Mater. Res. A 2007, 82, 10– 17. [16] Kyomoto, M.; Moro, T.; Yamane, S.; Hashimoto, M.; Takatori, Y.; Ishihara, K. Effect of UV-Irradiation Intensity on Graft Polymerization of 2-Methacryloyloxyethyl Phosphorylcholine on Orthopedic Bearing Substrate. J. Biomed. Mater. Res. A 2014, 102, 3012–3023. [17] Kurtz, S. M.; Muratoglu, O. K.; Buchanan, F.; Currier, B.; Gsell, R.; Greer, K.; Gualtieri, G.; Johnson, 20 Environment ACS Paragon Plus

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R.; Schaffner, S.; Sevo, K.; Spiegelberg, S.; Shen, F. W.; Yau, S. S. Interlaboratory Reproducibility of Standard Accelerated Aging Methods for Oxidation of UHMWPE. Biomaterials 2001, 22, 1731–1737. [18] Shen, F. W.; McKellop, H. A.; Salovey, R. Irradiation of Chemically Crosslinked Ultrahigh Molecular Weight Polyethylene. J. Polym. Sci. B: Polym. Phys. 1996, 34, 1063–1077. [19] Kyomoto, M.; Moro, T.; Takatori, Y.; Tanaka, S.; Ishihara, K. Multidirectional Wear and Impact-to-wear Tests of Phospholipid-polymer-grafted and Vitamin E-Blended Crosslinked Polyethylene: A Pilot Study. Clin. Orthop. Relat. Res. 2015, 473, 942–951. [20] Kyomoto, M.; Moro, T.; Takatori, Y.; Kawaguchi, H.; Ishihara, K. Cartilage-Mimicking, High-Density Brush Structure Improves Wear Resistance of Crosslinked Polyethylene: A Pilot Study. Clin. Orthop. Relat. Res. 2011, 469, 2327–2336. [21] Saiga, K.; Kyomoto, M.; Watanabe, K.; Taketomi, S.; Kadono, Y.; Takatori, Y.; Tanaka, S.; Ishihara, K.; Moro, T. Effects of Material Thickness and Surface Modification of Cross-linked Polyethylene with Poly(2-Methacryloyloxyethyl Phosphorylcholine) on Its Deformation Behavior, Wear Resistance, and Durability Under Repetitive Impact-to-sliding Motion. Biotribology 2017, 10, 35–41. [22] Kyomoto, M.; Moro, T.; Konno, T.; Takadama, H.; Kawaguchi, H.; Takatori, Y.; Nakamura, K.; Yamawaki, N.; Ishihara, K. Effects of Photo-induced Graft Polymerization of 2-Methacryloyloxyethyl Phosphorylcholine on Physical Properties of Cross-linked Polyethylene in Artificial Hip Joints. J. Mater. Sci. Mater. Med. 2007, 18, 1809–1815. [23] Peterson, J. D.; Vyazovkin, S.; Wight, C. A. Stabilizing Effect of Oxygen on Thermal Degradation of Poly (methyl methacrylate). Macromol. Rapid. Commun. 1999, 20, 480–483. [24] Kyomoto, M.; Moro, T.; Yamane, S.; Watanabe, K.; Hashimoto, M.; Tanaka, S.; Ishihara, K. A Phospholipid Polymer Graft Layer Affords High Resistance for Wear and Oxidation Under Load Bearing Conditions. J. Mech. Behav. Biomed. Mater. 2018, 79, 203–212. [25] Bracco, P.; Brunella, V.; Luda, M. P.; Zanetti, M.; Costa, L. Radiation-Induced Crosslinking of UHMWPE in the Presence of Co-agents: Chemical and Mechanical Characterisation. Polymer 2005, 46, 21 Environment ACS Paragon Plus

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10648–10657. [26] Kyomoto, M.; Miwa, Y.; Pezzotti, G. Strain in UHMWPE for Orthopaedic Use Studied by Raman Microprobe Spectroscopy. J. Biomater. Sci. Polym. Ed. 2007, 18, 165–178. [27] Kurtz, S. M.; MacDonald, D. W.; Mont, M. A.; Parvizi, J.; Malkani, A. L.; Hozack, W. Retrieval Analysis of Sequentially Annealed Highly Crosslinked Polyethylene Used in Total Hip Arthroplasty. Clin. Orthop. Relat. Res. 2015, 473, 962–971. [28] Kurtz, S. M.; Hozack, W. J.; Purtill, J. J.; Marcolongo, M.; Kraay, M. J.; Goldberg, V. M.; Sharkey, P. F.; Parvizi, J.; Rimnac, C. M.; Edidin, A. A. 2006 Otto Aufranc Award Paper: Significance of In Vivo Degradation for Polyethylene in Total Hip Arthroplasty. Clin. Orthop. Relat. Res. 2006, 453, 47–57. [29] Massin, P.; Lopes, R.; Masson, B.; Mainard, D.; French Hip & Knee Society (SFHG). Does Biolox Delta Ceramic Reduce the Rate of Component Fractures in Total Hip Replacement? Orthop. Traumatol. Surg. Res. 2014, 100 (6 Suppl), S317–321. [30] Kyomoto, M.; Shoyama, Y.; Saiga, K.; Moro, T.; Ishihara, K. Reducing Fretting-Initiated Crevice Corrosion in Hip Simulator Tests Using a Zirconia-Toughened Alumina Femoral Head. J. Biomed. Mater. Res. B Appl. Biomater. 2018, in press. [doi: 10.1002/jbm.b.34062] [31] Costa, L.; Bracco, P. In Mechanisms of cross-linking, oxidative degradation, and stabilization of UHMWPE, UHMWPE Biomaterials Handbook Third Edition, Kurtz, S. M. (Ed.),William Andrew: Oxford, 2015; pp 467–487.


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Figure captions Fig. 1.

Oxidative degradation of the untreated CLPE and PMPC-grafted CLPE before and after accelerated aging. Oxidation indices of surface (0 mm depth) and bulk (3 mm depth) of the CLPE and PMPC-grafted CLPE are expressed as means (n = 4) ± 95% confidence intervals. Significant differences were observed among all six groups of samples. *p < 0.05 and **p < 0.01 for one-factor ANOVA and post-hoc test.

Fig. 2.

(A) Static-water contact angle and (B) dynamic coefficient of friction of untreated CLPE and PMPC-grafted CLPE before and after accelerated aging. Data are expressed as the means (n = 15 and 10, respectively) ± 95% confidence intervals. **One-factor ANOVA and post-hoc test; significant differences (p < 0.01) were observed among all six sample groups.

Fig. 3.

Cross-link density of the untreated CLPE and PMPC-grafted CLPE before and after accelerated aging. Cross-link densities of surface (0–1 mm depth) and bulk (3–4 mm depth) of the CLPE and PMPC-grafted CLPE are expressed as means (n = 3) ± 95% confidence intervals. Significant differences were observed among all six groups of samples. *p < 0.05 and **p < 0.01 for one-factor ANOVA and post-hoc test.

Fig. 4.

Mechanical properties of the untreated CLPE and PMPC-grafted CLPE before and after accelerated aging. (A) Impact strength are expressed as the means (n = 6) ± 95% confidence intervals. (B) Work to failure of the surface (disk from 0–0.5 mm depth) and bulk (disk from 1.5–2.0 mm depth) are expressed as means (n = 5) ± 95% confidence intervals in the small punch test. Significant differences were observed among all six groups. *p < 0.05 and **p < 0.01, one-factor ANOVA and post-hoc test.


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Fig. 5.

Wear of the untreated CLPE and PMPC-grafted CLPE disks before and after accelerated aging in the impact-to-wear test. Time course of gravimetric wear of (A) the pre-aged, 3-weeks aged, and (B) 9-weeks aged disks during the test. (C) Wear rate (5.0 × 104–2.0 × 106 cycles) of all six groups of disks. Data are expressed as means (n = 6) ± 95% confidence intervals. Significant differences were observed among all six groups. *p < 0.05 and **p < 0.01 for one-factor ANOVA and post-hoc test.

Fig. 6.

Wear and/or impact fatigue behaviors of the distinctive three samples among the 9-weeks aged PMPC-grafted CLPE disks in the impact-to-wear test. (A) Time course of gravimetric wear of the distinctive samples during the test. (B) 3D profile and (C) µCT images of the disks after the test. Dotted line on 3D profile image of sample 3 indicates the cross-section for µCT observation.

Fig. 7.

Wear of the untreated CLPE and PMPC-grafted CLPE liners in the hip-simulator wear test. (A) Time series of gravimetric wear of the untreated CLPE and PMPC-grafted CLPE liners before and after accelerated aging of 3 weeks during the test. (B) Time series of gravimetric wear of the 3-weeks aged PMPC-grafted CLPE liners against the Co–Cr–Mo alloy or ZTA femoral head. (C) Wear rate (5.0 × 105–1.0 × 107 cycles) of all five groups of liners. Data are expressed as means (n = 4) ± 95% confidence intervals, respectively. *p < 0.05 and **p < 0.01, one-factor ANOVA and post-hoc test.


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Graphical abstract


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Figure 1. Oxidative degradation of the untreated CLPE and PMPC-grafted CLPE before and after accelerated aging. Oxidation indices of surface (0 mm depth) and bulk (3 mm depth) of the CLPE and PMPC-grafted CLPE are expressed as means (n = 4) ± 95% confidence intervals. Significant differences were observed among all six groups of samples. *p < 0.05 and **p < 0.01 for one-factor ANOVA and post-hoc test. 91x46mm (600 x 600 DPI)

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Figure 2. (A) Static-water contact angle and (B) dynamic coefficient of friction of untreated CLPE and PMPCgrafted CLPE before and after accelerated aging. Data are expressed as the means (n = 15 and 10, respectively) ± 95% confidence intervals. **One-factor ANOVA and post-hoc test; significant differences (p < 0.01) were observed among all six sample groups. 102x58mm (600 x 600 DPI)


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Figure 3. Cross-link density of the untreated CLPE and PMPC-grafted CLPE before and after accelerated aging. Cross-link densities of surface (0–1 mm depth) and bulk (3–4 mm depth) of the CLPE and PMPCgrafted CLPE are expressed as means (n = 3) ± 95% confidence intervals. Significant differences were observed among all six groups of samples. *p < 0.05 and **p < 0.01 for one-factor ANOVA and post-hoc test. 105x62mm (600 x 600 DPI)


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Figure 4. Mechanical properties of the untreated CLPE and PMPC-grafted CLPE before and after accelerated aging. (A) Impact strength are expressed as the means (n = 6) ± 95% confidence intervals. (B) Work to failure of the surface (disk from 0–0.5 mm depth) and bulk (disk from 1.5–2.0 mm depth) are expressed as means (n = 5) ± 95% confidence intervals in the small punch test. Significant differences were observed among all six groups. *p < 0.05 and **p < 0.01, one-factor ANOVA and post-hoc test. 215x565mm (600 x 600 DPI)


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Figure 5. Wear of the untreated CLPE and PMPC-grafted CLPE disks before and after accelerated aging in the impact-to-wear test. Time course of gravimetric wear of (A) the pre-aged, 3-weeks aged, and (B) 9-weeks aged disks during the test. (C) Wear rate (5.0 × 104–2.0 × 106 cycles) of all six groups of disks. Data are expressed as means (n = 6) ± 95% confidence intervals. Significant differences were observed among all six groups. *p < 0.05 and **p < 0.01 for one-factor ANOVA and post-hoc test. 223x605mm (600 x 600 DPI)


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Langmuir

Figure 6. Wear and/or impact fatigue behaviors of the distinctive three samples among the 9-weeks aged PMPC-grafted CLPE disks in the impact-to-wear test. (A) Time course of gravimetric wear of the distinctive samples during the test. (B) 3D profile and (C) µCT images of the disks after the test. Dotted line on 3D profile image of sample 3 indicates the cross-section for µCT observation. 148x268mm (600 x 600 DPI)


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Figure 7. Wear of the untreated CLPE and PMPC-grafted CLPE liners in the hip-simulator wear test. (A) Time series of gravimetric wear of the untreated CLPE and PMPC-grafted CLPE liners before and after accelerated aging of 3 weeks during the test. (B) Time series of gravimetric wear of the 3-weeks aged PMPC-grafted CLPE liners against the Co–Cr–Mo alloy or ZTA femoral head. (C) Wear rate (5.0 × 105–1.0 × 107 cycles) of all five groups of liners. Data are expressed as means (n = 4) ± 95% confidence intervals, respectively. *p < 0.05 and **p < 0.01, one-factor ANOVA and post-hoc test. 214x556mm (600 x 600 DPI)


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A hydrated phospholipid polymer gel-like layer for increased durability

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