Modication of the Interface Between Degradable Magnesium Alloy and Physiological Environment for Controlled Biodegradation Using Fibrous Polymer/Metal-Organic Framework Nanocomposite Coatings

Magnesium-based implants (MBIs) have recently attracted great attention in bone regeneration due to elastic modulus similar to bone. Nevertheless, the degradation rate and hydrogen release of MBIs in the body have to be tackled for practical applications. In the present study, we present a metal organic framework (MOF) nanoplates to reduce the degradation rate of AZ91 magnesium alloy. Zeolitic imidazolate frameworks (ZIF-8) with specic surface area of 1789 m 2 .g -1 were prepared by solvothermal methods, and after dispersion in a chitosan solution (10%w/w), the suspesnsion was electrospun on the surface of AZ91 alloy. Studying of the degradation rate in simulated body uid (SBF) by electrochemical analysis including potentiodynamic polarization and electrochemical impedance spectroscopy reveals that the degradation rate of the surface modied implants is ~20 times less than the unmodied specimens. The reduced alkalization of the physiological environment and hydrogen release due to the implant degradation are shown. In vitro studies by broblasts and MG63 osteosarcoma cells exhibit improved cell adhesion and viability. The mechanisms behind the improved degradiation resistance and enhanced bioactivity are presented and discussed. Surface modication of MBIs by MOF-chitosan coatings is a promosing strategy to control the biodegradation of magnesium implants for bone regeneration.


Introduction
In spite of the self-healing ability of bones, severe injuries caused by accident or fracture require bone grafting [1]. Besides autografts, biocompatible and bioactive scaffolds are required to accelerate the healing process of damaged bone [2]. Recently, biodegradability, biocompatibility and compatible mechanical properties of magnesium-based alloys with bone have raised signi cant interest for bone implantation [3]. Magnesium itself is the fourth most abundant element in the body that participates in most of the biological functions [4], enzymatic reactions [5], tissue repair [2], protein and nucleic acid synthesis, mitochondrial activity, plasma membrane robustness, and RNA transcription processes [6], [7].
Many studies have shown that porous magnesium scaffolds can easily be degraded at endosomal sites and dramatically stimulate new bone formation and stimulate angiogenesis [5]. Of interest, the implant degradation products enhance osteoblast proliferation, differentiation, and expression of osteoporotic markers, while the excess amount does not cause cytotoxicity in the human body and is generally excreted through the urethra [7].
Albite all the aforementioned advantages of MBIs, the low corrosion resistance of magnesium alloys in physiological environments (PE) results in a rapid loss of mechanical integrity and growth of hydrogen bubbles that severely impair the bone healing process [8]. Therefore, the bioreaction rate of MBIs in PE must signi cantly be reduced to facilitate the use of light weight magnesium alloys for bone implantation. Laboratory studies have estimated the critical tolerance of H 2 to be less than 0.01 ml/cm 2 .day [9]. So far, different strategies have been employed to reduce the biodegradation rate of BMIs in PE. Alloying and surface modi cation are the most widely utilized strategies to minimise the complications of magnesium implantation into the human body [10]. It has been shown that non-toxic alloying elements such as Ca, Zn, Sr, and Si improve the mechanical stability and biodegradation rate of MBIs in PE [11]. However, the biocompatibility and degradation rate are still de cient for practical applications and may trigger a series of side effects to surrounding tissues after implantation [12].
Surface modi cation either through physical methods (morphological adjustment without any change in the components) or chemical techniques (surface functionalization, ion infusion, and coatings) are alternative approaches to reduce the biodegradation rate [13]. Among different methods, sol-gel [14], electrophoretic deposition (EPD) [15], and plasma electrolyte oxidation (PEO) [16] have gained a considerable attraction in recent years. Different studies have shown that surface modi cation by depositing a coating layer not only improves the biocompatibility, but also encourages the bone ingrowth, osseointegration induction, and mesenchymal stem cell proliferation [17], [18].
Polymer coatings provide several advantages for the surface modi cation of MBIs. Depositing a thin layer of polymeric materials on the surface acts as physical barrier to make the biodegradation rate of the substrate comparable with the bone regeneration rate of the damaged tissue [19]. The barrier lm ought to be mechanically stable and hydrophilic to increase cell adhesion and to stimulate bone cell proliferation [3], [20]. Sang et al. [21] have shown that surface modi cation of Mg−6%Zn −10% Ca 3 (PO 4 ) 2 by chitosan prevent severe corrosion in SBF with a signi cant reduction in the formation of hydrogen bubbles. Heakal and Bakry [22] employed spin-coating to deposit a thin chitosan lm on AXJ530 alloy. They showed that the biodegradation rate of the magnesium alloy in PBS was signi cantly retarded.
Studies of Tiyyagura et al. [10] indicated that along with the reduced corrosion rate, the chitosan lm promoted the formation of hydroxyapatite layer on the surface. Panahi et al. [23] employed electrospinning to prepare polycaprolactone-bioactive glass (BG) brous composite coatings on AZ91 alloy. Detailed electrochemical studies have determined that the composite coating signi cantly decreases the corrosion current through retarding ion and electron transport between the surface and the liquid environment. On the other hand, the BG particles enhance the bioactivity, cell adhesion, and cell proliferation.
In the present work, we used chitosan/MOF composite coatings for the surface modi cation of BMIs. To our best knowledge, the role of polymer/MOF composite coatings on the biodegradation rate of MBIs in PE has not been reported yet. We employed biocompatible ZIF-8 nanoplates with a porous crystalline structure that have high aqueous stability and exceptional chemical resistance against polar and nonpolar solvents [24]. Due to very high surface area, ZIF-8 nanoplates have high adsorption capacity, thus making this material a suitable candidate for drug delivery [24], [25], which could provide additional advantages to prepare drug-eluting coatings on MBIs to prevent bacterial infections [26] or to deliver growth factors on-site [27], [28]. Since the coating layer ought to have mechanical integrity and adherence to the surface during the degradation, brous chitosan/ZIF-8 MOF coatings were prepared. Chitosan is a well-known biocompatible biopolymer with rapid blood clotting and homeostatic effect [29], [10]. Recent studies on BMIs implants have determined that the chitosan-modi ed implants are resistant to osteomyelitis (bone infection) caused by Staphylococcus aureus along with an improved osteogenesis after implantation [30]. In vivo studies have also determined limited in ammatory response in several 8-week-old adult rats [5]. The brous structure of the composite lm provides suitable microenvironment for cell attachment and spreading. The nano bers can also be penetrated with cells to treat or replace biological targets and hinder the degradation of implants in our body through the duration of treatment [31], [32]. We studied the effect of the MOF/chitosan composite lms on the biodegradation rate by the Tofel polarization test and electrochemical impedance spectroscopy (EIS). The biocompatibility of the surface modi ed Mg alloy is demonstrated and the improved biological performance is discussed based on experimentally measured hydrogen gas evolution and changes in the pH of the cell environment. The MOF/chitosan lms show promising capacity for surface modi cation of MBIs for bone regeneration.

Materials
Biomedical grade AZ91 magnesium alloy was obtained from local market. The chemical composition of the alloy was determined by atomic absorption spectroscopy, as reported in Table 1. Medium molecular weight chitosan (Mn= 80,000 Da) and polyethilenoxide (PEO) with an average molecular weight of 4,000,000 were obtained from Sigma-Aldrich (USA). Zinc nitrate hexahydrate (Zn (NO 3 ) 2 .6H 2 O), 2methylimidazole, dimethyl sulfoxide (DMSO), and SBF were also obtained from the same company. Absolute methanol and acetic acid (50%) were purchased from Merck (Germany). Dulbecco's modi ed eagle medium (DMEM) was obtained from Gibco (Scotland). All chemicals were used without further puri cation. with a Zn 2+ /MIm molar ratio of 1:8 were dissolved in methanol by dropwise adding and gentle stirring. After mixing for 60 min, the solution was transferred in a Te on lined autoclave and treated for 24 h at 120°C. The precipitate was separated by centrifugation and washed three times by fresh methanol to remove any unreacted reagents. The resulting product was vacuum dried to attain ZIF-8 MOFs.

Surface modi cation of magnesium alloy
The AZ91 slab was cut to small cubes with dimensions 1 mm x 1 mm x 1 mm by electrical discharge machining. Before electrospinning, a surface treatment including grinding with a SiC emery paper (mesh #180) followed by chemical etching in 1M HNO 3 for 30 s was performed. As shown elsewhere [33], this pre-treatment created rougher surface and improved the adhesion of electrospun bres to the magnesium surface.
For electrospinning, at rst an aqueous solution of chitosan was prepared by dissolving the polysaccharide (4 wt.%) in 4 mL acetic acid solution (50%) with a pH=5.1. The mixture was stirred for 24 h at 250 rpm, and then ltered to remove any residuals. To facilitate electrospinning, PEO with a mass ratio of 1:4 to chitosan was added to the polymer solution and stirred overnight. For the preparation of composite suspensions containing ZIF-8 MOFs, the nanoparticles (10 wt.%) were added to the polymer solution, stirred for 2 h, and sonicated for 45 min. A single nozzle electrospinning apparatus (ES1000, FNM, Iran) was used to deposit chitosan and MOF/chitosan composite lms on the surface of the magnesium specimens. The processing parameters were determined based on trial and error efforts. The applied voltage was 21 kV. A ow rate of 1mL/h was utilized. The distance of the syringe to the aluminium rotating drum (2500 rpm) was 10 cm. To attain lms with the same thickness of 30 µm, the processing parameters were kept same for both chitosan and composite coatings.

Materials characterizations
Scanning electron microscopy (SEM, Philips, XL30) was used to study the size and morphology of MOF particles. For the phase analysis, X-ray diffraction (XRD) was carried out by a Philips instrument. The size and morphology of the bres as well as the thickness of the electrospun lms on the surface of the magnesium alloy were studied by SEM. The biodegradation products after immersion in SBF were also analysed by SEM and XRD. Besides, electron dispersive spectroscopy (EDS) was used for elemental analysis. To determine the effect of surface modi cation on the hydrophilicity, the sessile drop technique was employed according to ASTM D7334 standard. A small droplet of deionized water was placed on the surface of the samples by micropipette and photographed by a camera. The water contact angle was determined by the ImageJ software.

Degradation studies
All biodegradation evaluations were carried out in SBF at 37 °C. Potentiodynamic polarization test was carried out by a potentiostat (302N, Metrohm Autolab, Netherlands). A conventional three-electrode con guration including a working electrode (AZ91 alloy), a saturated calomel electrode (control), and a platinum counter electrode was used. Potentiodynamic polarization curves were recorded by changing the voltage in the range of -2.5 to 0 mV with a rate of 20 mV/s [34]. To establish dynamic equilibrium conditions, the specimens were immersed in SBF for 60 min prior to testing. Electrochemical impedance spectroscopy (EIS) was carried out at a 1 V DC potential (vs. the standard electrode). A sinusoidal potential amplitude (10 mV) in the frequency range of 0.1-105000 Hz was applied. The acquired EIS data were tted to an equivalent circuit by employing the ZView® software (Scribner Associates Inc., USA).
To determine the rate of hydrogen evolution during biodegradation, the specimens were inserted in a transparent plastic syringe containing 60 mL SBF. The container was incubated at 37°C for several days.
The amount of hydrogen volume was then determined by measuring the volume change according to the procedure explained in Ref. [35].

In vitro studies
L929 broblast cells were obtained from the National Cell Bank of Iran (Pasteur Institute of Iran, Tehran) All the specimens were sterilized by 70% ethanol for 30 min in a 24 well-plate and dried under a laminar ow biosafety cabinet (KG-A100, Kimia Gene Co., Iran). The specimens were then placed in 24-multi well plate and 5×10 4 cells were seeded on each sample. The control was the well plate without AZ91 sample. After incubation for selected times, 20 μL of 1 mg/mL MTT (3-[4,5-dimethythiazol-2-yl]-2,5-diphenyltetrazolium bromide) solution was added to each well and incubated for 4 h at 37 °C. To dissolve the insoluble purple formazan crystals, the medium was replaced by (DMSO). The absorbance in each well was then recorded at 540 nm through an ELISA reader [36].
For evaluation of cellular adhesion, human osteoblast-like MG-63 cells (the National Cell Bank of Iran, Pasteur Institute of Iran, Tehran) were used. The cells with a density of 5×10 4 were incubated on the specimens in a 5% CO 2 humidi ed incubator at 37 °C. After 3 days, the nonadherent cells were washed out by PBS rinsing. The attached cells were then xed by a xative solution containing glutaraldehyde and stained in an alcoholic solution for 5 min and dried. Changes in the pH of the medium was measured by a pH meter (pH7110, Inolab, Germany).

Characteristics of the metal organic frameworks
Different analytical techniques were employed to determine the morphology, porosity, and structure of the synthesized ZIF-8 MOFs. Table 2. summarizes the results. Figure 1 shows the characteristics of the ZIF-8 MOFs. The XRD pattern indicates well de ned crystal planes with slightly broadened peaks (Fig. 1a). In agreement with the simulated pattern of the ZIF-8 structure [37], the particles are crystalized in hexagonal structure [38]. The SEM image indicates that the particles have thin plate-like morphology with a lateral size of 70 to 80 nm and thickness of about 22 nm, yielding an aspect ratio of 0.3 (Fig. 1b). BET analysis at 77 K determines a type I adsorption isotherm with a small hysteresis loop (Fig. 1c). Analysis of the pore size distribution (Fig. 1d) reveals the nanoporous structure of MOF with the characteristics reported in Table 2. As seen, the processed MOF nanoplates have very high speci c surface area (1778 m 2 .g -1 ) and large pore volume (1.28 cm 3 .g -1 ).
FTIR spectroscopy exhibits characteristic peaks at 993 cm -1 , 1142 cm -1 and 1421 cm -1 , which are associated with the C-N stretching vibration [39]. Two small peaks at 3130 cm -1 and 2927 cm -1 are ascribed to aromatic C-H stretching vibration and aliphatic C-H stretching vibration of the imidazole ring and the methyl group, respectively [40]. The peak at 1583 cm -1 is attributed to C=N stretching vibration.
Other peaks that appeared below 1250 cm -1 correspond to the in-plane and out-of-plane bending of the imidazole ring [41]. Zn-N stretching vibration is detected at 422 cm -1 . The peaks in the range of 1350 to 900 cm -1 could be attributed to in-plane bending of the ring while the peaks at 754 and 688 cm -1 are associated with aromatic sp 2 C-H bending [39].  The absorption bands at 1149 cm −1 and 1090 cm −1 can be attributed to asymmetric stretching of the C-O-C bridge and C-O stretching, respectively. The small bands at 608 cm −1 correspond to the wagging of the CS saccharide structure [40]. No major difference or band shifts is noticeable in the FTIR spectrum of MOF/chitosan lm (Fig. 2c), indicating that chemical reactions or new band formation have not occurred upon processing.

Biodegradation of magnesium alloy
To study the biodegradation of AZ91 alloy with and without surface modi cation, the electrochemical responses of the samples in SBF were measured. The Tofel plots of the examined specimens are shown in Fig. 3a. The extracted data from the polarization curves are summarized in Table 3. The results indicate that the magnesium alloy exhibits a relatively high corrosion current (i corr =383.6 µA cm -2 ) in SBF, which is in agreement with the literature [43]. Surface modi cation by the chitosan lm signi cantly reduces the degradation rate and corrosion current to about one third (i corr =102.6 µA cm -2 ). The potential (E corr ) also slightly increases. Interestingly, the incorporation of ZIF-8 MOFs in chitosan further decreases the degradation rate and corrosion current by two folds without changing E corr. As compared to unmodi ed magnesium alloy, i corr drops to 18.3 µA cm -2 , which is only about 5% of that of AZ91. It is important to note that unlike the cathodic slope (β c ), no signi cant change is observed in the anodic slope (β a ); hence, the protecting lms should reduce the degradation rate through controlling the cathodic reactions [42]. The amount of hydrogen evolution measured at different time intervals (up to 15 days) after immersion in SBF is shown in Fig. 3b. Evidently, the amount of gas evolution increases with increasing the immersion time for all specimens. The rate of gas evolution exhibits a declining trend with time, indicating the effect of degradation products on the corrosion rate. It seems that the products of degradation act as a physical barrier to prevent rapid degradation in PE. It is known that magnesium hydroxide (Mg (OH) 2 ) and magnesium oxide (MgO) are commonly formed during degradation and form a passive layer on the metal surface [44]. To study the degradation products, SEM and XRD were employed (Fig. 4). SEM studies show the formation of rod-shape or ower-like precipitates on the surface.

Electrochemical impedance spectroscopy
To explore the mechanism of biodegradation retardant effect of chitosan and MOF nanoplates, electrochemical impedance spectroscopy (EIS) in SBF at 37°C was employed. The resulting Nyquist and Bode plots are shown in Fig. 5. In the Nyquist plots, two capacitive loops in both high and low frequency regions are observed (Fig. 5a). In concurrence, the phase angle Bode plots exhibit two-time constants (Fig. 5b). After immersing the specimens in the PE solution, the biodegradation products are formed on the surface and two capacitive loops (at high-and low-frequency regions) develop due to changes in the charge transport, surface chemistry, and mass transfer [45]. The high-frequency region of the EIS relayed on the electrolyte penetration process including water uptake and the electrolyte interference. The diameter of the loops is directly proportional to the surface lm resistance for mass and charge transport [46], [47]. The unmodi ed magnesium alloy exhibits relatively a small loop that indicates a rapid biodegradation rate. The surface modi cation by chitosan lm increases the diameter of the circle with double humps. The behaviour of the specimen can be described based on capacitive behaviour originated from the degradation of the magnesium alloy in the medium and inductive behaviour resulted from continuous adsorption and desorption chemicals elements or ions and formation of corrosion products [35]. The incorporation of ZIF-8 MOFs further enlarges the circles with little effects of the inductive behaviour. Therefore, the chitosan-MOF lm has further increased the degradation resistance of the magnesium substrate.
To get better insight about the mechanism of biodegradation, the equivalent circuit for the Nyquist plot was constructed (Figs. 5c,d). The values obtained by data tting are reported in Table 4. As so the low frequency region for the bare and chitosan-MOF specimens are too much noisy, an equivalent circuit with one time constant was used to t the EIS plots. However, a circuit with two-time constants was used to t the chitosan-modi ed AZ91 sample. R s is representative of the solution resistance which is in series with the unit of the oxide layer system. For the bare electrode, R1 and R 2 in parallel with constant phase elements (CPE 1 and CPE 2 ) are the resistance of naturally formed oxide lm on the substrate and charge transfer resistance of the faradaic process on the metal surface, respectively [46], [47]. For the chitosan and chitosan-MOF specimens, the resistances (R 1 and R 2 ) and constant phase elements (CPE 1 and CPE 2 ) represent the resistance of the porous outer layer and the inner coating/substrate interface, respectively. This part of circuits was not quantitatively evaluated for the bare and chitosan-MOF coated AZ91 electrodes as the corresponding error was relatively high. Noted that the value of tting errors should be less than 1% [48], [49]. Besides, CPE which models the behaviour of a double layer (as an imperfect capacitor) was considered instead of a capacitive element because of the surface inhomogeneity and possible diffusional factors affecting the shape of the semi-circles [50]. Table 4 indicates that the R 2 value for the chitosan-MOF specimen (1256 Ω) is much higher than that of the chitosan-modi ed AZ91 (651.9 Ω) and the bare sample (291.3 Ω). This nding a rms corrosion resistance enhancement of AZ91 through surface modi cation by the chitosan-MOF lm via inserting a physical barrier against diffusion/penetration of the corrosive ions from the solution toward the metal surface. For the chitosanmodi ed AZ91 specimen, the value of R2 (651.9 Ω) is far higher than that of R 1 (10.6 Ω). This observation could indicate the vital role of the magnesium/coating interface on the biodegradation process.

In vitro biocompatibility and cell adhesion
To assay the in vitro biocompatibility of the specimens, L929 cells were used for MTT assay (Fig. 6a). The unmodi ed AZ91 alloy exhibits poor biocompatibility with a cell viability of less than 50% even after a short incubation time (24 h). A longer incubation time further decreases the cell viability, indicating the cytotoxic effect of Mg degradation, as reported in literature [51]. The magnesium alloy modi ed with chitosan lm exhibits higher cell viability (>90%) at the early stage but reduced to about 70% at prolonger

Discussion
In this study, ZIF-8 nanoplates were employed and incorporated in the chitosan nano ber to modify the biodegradation rate of AZ91 Mg alloy in PE. It was shown that the MOF particles decorated the surface of the nano bers and slightly decreased their diameters. The reduced sizes can be attributed to the increase in the size of the Taylor cone and decrease in the jet velocity due to a change in the viscosity of the solution [52], [53]. Since ZIF MOFs are hydrophilic [37], the particles tend to be pushed toward the outer layer and collected on the surface. Meanwhile, no remarkable agglomeration was noticed.
It was also interesting to observe that the chitosan coating containing MOFs exhibited better protection against biodegradation in PE. Electrochemical studies revealed that the coatings were mainly operated as a physical barrier. EIS studies determined that the impedance of the surface-modi ed specimens was higher than that of the bare metal. It seemed that the coating layer reduced the ion diffusion and thus decreased the corrosion rate [35]. The Bode plot determined low-frequency and high-frequency resistor zones and a capacitive behaviour in the mid-frequency range due to the maximum phase changes (θ max ) vs. logf [54]. Based on the shape and values of θ at high frequencies, it appeared that the electrochemical response of the surface modi ed samples had a higher tendency toward capacitive behaviour. At low frequencies, due to absorption and reduction of corrosion products such as Mg (OH) and Mg (OH) 2 at the interface between the electrolyte solution and the metal surface, an induction loop appears that is equivalent in shape to the inductor and offers resistance [55]. The Nyquist plots and the capacitive loops also indicated the existence of a distinct layer on the surface of the specimens. The diameter of the capacitive ring indicated the higher polarization resistance of the chitosan-MOF composite specimen [56].
Here, the role of the precipitates on the corrosion rate should also be considered, It was shown that corrosion of magnesium in PE was accompanied with the formation of precipitates on the surface (Fig.  4). Degradation and dissolution of magnesium increased the pH of the medoium that promoted phosphate formation due to saturating of SBF at highly alkaline conditions [50].
The cell viability assay indicated improved biocompatibility of chitosan-modi ed specimens over the period of 72 hours. Measurements of the pH value of the culture medium (Fig. 6b) indicated that for the bare metal, the pH exceeded beyond the physiological level and gained a high value (8.6) that cells can tolerate. The raise of the pH is attributed to biodegradation of the metal surface and the release of calcium ion in the cell culture medium as well as the formation of hydrogen bubbles [10]. The surface modi cation reduced the amount of pH raise as the coating layer acted as a barrier against direct contact of the medium and the metal surface. Not only the fast release of Mg in the culture medium affects the viability due to alkalization, degradation of MOFs may in uence the cell response. Hook et al. [24] showed that at concentrations beyond 30µg.mg-1, ZIF-8 exhibited cytotoxicity due to the release of Zn2+ in mitochondrial ROS products. This negative effect suppresses the cell cycle in G2/M phase, which is irreversible and nally caused DNA damage and ultimately activates cellular apoptosis pathways [34]. We used low amount of ZIF-8 in our mats, therefore, the effect of Zn ions should be marginal. It was also shown that cell adhesion on the surface of the brous coatings was improved. Cells usually interact better with surfaces that have more pores and roughness [57]. Hydrophilicity and mechanical durability of the surface also affect the interaction of cells with the material at the interface [51], [30]. For instance, it has been shown that MG63 osteoblast-like cells signi cantly proliferate at rough surfaces such as acidetched or plasma [57], [58]. Hydrophilic surfaces have much more integration and adhesion with osteoblastic cells with implants compared to hydrophobic or less hydrophilic surfaces [51].

Conclusions
To control the biodegradation rate of AZ91 magnesium alloy in physiological environments, composite lms of ZIF-8 MOF/chitosan were deposited on the surface by electrospinning. Plate-like MOF particles with lateral sizes of 70 to 80 nm with an aspect ratio of 0.3 and speci c surface area of 1789 m 2 .g -1 were prepared by solvothermal methods. Continuous, uniform and bead-free bers with an average diameter of 30 µm were deposited on the surface of the magnesium alloy. The incorporation of ZIF-8 particles in the chitosan coating did not signi cantly changed the size and uniformity of the bers, but the particles tended to decorate the surface of the polymer. In vitro evaluations of biodegradation in PE by electrochemical methods indicated that the corrosion current of AZ92 alloy decreased to one third after surface modi cation by the brous chitosan coating. The incorporation of MOF particles further reduced the corrosion current by two orders as well. The rate of hydrogen evolution was also reduced signi cantly. EIS determined that the polymer and composite lms operated as a physical barrier and increased the charge transfer resistance while changing the capacitive behaviour of the biodegradation process. The degradation products were changed from magnesium compounds for the bare metal to hydroxyapatite precipitates with more complex Ca-Zn-Si compounds. In vitro cytotoxicity examinations also revealed that surface modi cations improved the biocompatibility due to reduced degradation rate of Mg in the physiological medium. In contrast to the bare metal, cell adhesion with lopodia formation and cell spreading were also improved. Therefore, the MOF-chitosan composite coating showed promising potential for the surface modi cation of magnesium alloys to limit their biodegradation in vitro. In vivo examinations, however, are required to evaluate the biological response of the surface modi ed BMIs to reveal their potential applications for bone regeneration.