Essay Writing Service

1 Star2 Stars3 Stars4 Stars5 Stars (No Ratings Yet)

Electrospun Nanofibrous Polymer Composites Laminated with Magnesium Based Amorphous Alloys

do not necessarily reflect the views of

Electrospun nanofibrous polymer composites laminated with Magnesium based amorphous alloys for degradable Scaffolds


Magnesium based thin film metallic glass (MgMG) has recently pulled in huge attraction for biomedical implants because of their promising biocompatibility, bioactivity, biodegradability and mechanical properties. In this study, biodegradable poly (caprolactone) (PCL) and nano hydroxyapatite (nHA) are electrospinned at various PCL/nHA concentration and used as Scaffolds. Mg–Zn–Ca thin film metallic glasses were sputtered over the scaffold by DC magnetron sputtering process in order to obtain surface modified scaffolds of (PCL/MgMG) and (PCL/nHA/MgMG). Magnesium metallic glass coatings neither alter the chemical composition of the scaffold nor the rate of magnesium ion release nor weight loss, compared with the uncoated scaffold. The changes in the crystallinity, wettability and oxygen content of sputtered coating not only improve the mechanical adhesion strength, but also enhanced the biological response with controlled biodegradation. Moreover, the modified sample can lead to higher hemocompatibility, due to the enhanced wettability from 127 to 58° and C-O to C=O functional group ratio. These promising results reveal the clinical potential of the surface modified coating of Magnesium metallic glass on PCL and PCL/nHA scaffolds as degradable implants.

Keywords: Magnesium metallic glass, Electrospinning, Sputtering, Biodegradable, Scaffold, hemocompatibility.


1. Introduction

Coatings of Magnesium alloys have attracted increasing interest over the past decade as a superior alternative material for implants, because they address the problems of currently dominant orthopedic biomaterials, such as stress shielding issues of titanium alloys and poor mechanical stability of polymer-based devices.[1] Magnesium (Mg) based implants perform similar to natural bone with respect to the mechanical properties. In addition to this, it is biodegradable, biocompatible, osteoconductive and has anti-bacterial properties.[2] Most importantly, the human body can metabolize the degradation products of Mg in an effective manner, which eliminates the need for implant removal surgeries after the tissue has healed.[34] To realize the benefits of Mg for medical implants, it is critical to control the rate of Mg degradation in the body.[5] One approach is to engineer Mg-based alloys that not only have desired mechanical and corrosion properties, but also biocompatibility and bioactivity for specific medical applications. For the past few years, amorphous Mg-Zn-Ca ternary metallic glasses (MgMG) have been of particular interest in biomedical applications.[6-8] It has been demonstrated that Mg(OH)2 is very effective in temporarily enhancing osteoblast activity and. [9] also an effective antibacterial agent against E.coli and B.phytofirmans.. It has been developed to improve the degradation and mechanical properties of Mg alloys for high-temperature and light weight structural applications. The addition of low-cost Ca into the Mg alloy has intended to improve the precipitation hardening ability of Mg alloys. From the materials science perspective, the addition of Ca improves the room-temperature mechanical properties mainly due to the formation of fine Mg2Ca precipitates.[1011] The addition of Zn as an alloying element improved tensile and creep strength, increased age hardening response, reduced grain size, and improved alloy cast ability.[12] The maximum solubility of Zn in Mg was reported to be 2 weight (wt.) % at room temperature in an equilibrium state [13]. The addition of 7 wt. % Zn (when Ca < 1.0 wt. %) resulted in Mg-Zn-Ca alloys with corrosion (or degradation) rates promising for biomedical applications.[78] In this study, a fixed content of 5.0 wt. % Ca and a maximum of  35 wt. % Zn were selected based on numerous Mg-Zn-Ca studies favoring as-cast Mg-Zn-Ca Metallic glass with low Ca content (< 5 wt. %) and Zn/Ca wt. % ratio < 5 [14]. Mg-Zn-Ca Metallic glass coatings are considered as potential biodegradable materials for their better biological performance.[15] Despite the fact that investigations on bone cell responses to magnesium and its compounds are inadequate, a number of studies have been published for investigate the effect of enriching the surface of a biomaterial, such as hydroxyapatite, with Mg2+ ions; an important biochemical role which has been suggested for such ions in the bone system, also due to the fact that Mg2+ is the fourth most abundant cation in the human body.[9] The idea at the basis of this work is to apply sputtering technologies for the production of thin films containing Mg-Zn-Ca metallic glass (MgMG) on polymer based scaffolds for degradation implants. Reactive sputtering, electron beam deposition, and ion plating are typical deposition methods for Mg- based coatings.[16] Although it is well known that this method is characterized by a low deposition rate, DC magnetron sputtering deposition process was chosen to deposit coatings thinner than 1 μm on polymeric nanofibers commonly utilized for tissue engineering scaffolds. These polymeric nanofibrous scaffolds provides high surface to volume ratio for the nutrient transport, non–cytotoxic, and osteoconductive properties.[17] About 90 % of the materialistic properties are overcome by few polymers like Polycaprolactone (PCL), Poly-L-lactic acid (PLLA) and poly(L-lactic acid)-co -poly(ϵ -caprolactone) (PLACL) which found to be potential synthetic polymers for bone tissue regeneration [18]. Nanofibrous scaffolds for bone tissue engineering depends on various properties including scaffold composition and pore size distribution.  Also, the scaffold should be bioresorbable with a controllable degradation rate to match the replacement by new tissues.[19] Additionally, the scaffold should possess sufficient mechanical properties such as stiffness, strength and toughness. For enhancing the mechanical properties of polymers, calcium phosphate-based ceramics like nano hydroxyapatite (nHA), Ca10 (PO4)6(OH) 2, has received particular attention in recent years due to its chemical similarity to the inorganic matrix of natural bone. [20] It shows excellent biocompatibility with hard tissues and also has the tendency to bond directly to bone.[21] It is well known that nanometric thin, mechanically compliant coatings are able to modulate the interactions between cells and/or biological molecules due to their chemical composition and/or wettability,[9] on the contrary, coatings thicker than 1 μm may affect the mechanical signaling of cells through their cytoskeleton. The goal of the present study is to optimize the deposition of biodegradable amorphous coatings of Mg60-Zn35-Ca5 (wt %) Metallic glass (MgMG) over electrospinned nanofibrous scaffold (PCL/MgMG) and (PCL/nHA/MgMG) and investigated their degradation and hemocompatibility.

2. Materials and Methods

2.1. Scaffolds by Electrospinning:

Electrospinning was used to fabricate PCL and PCL/nHA nanofibrous scaffolds. The process of electrospinning requires an optimization of various parameters, including voltage, distance between the needle tip and collector, as well as the concentration of polymer/blends, solution flow rate, etc to obtain the random nanofibrous scaffold. PCL polymer granules (Mw 80000), Ammonium Phosphate ((NH)3PO4), Calcium chloride (CaCl2) and Chloroform was purchased from Sigma Aldrich. Diethyl ether and Ethanol was purchased from Fisher Scientific. To fabricate scaffolds, PCL and nHA particles were homogeneously dispersed in the below mentioned suspension by ultrasonic stirring. The PCL/nHA composite scaffolds with different nHA proportions (0 %, 10 %, 30 %, and 50 % wt % of nHA) were obtained by dissolving in (3:1 v/v Chloroform: methanol) at a concentration of 10 % and 8 % w/v solution. Then the solutions  were loaded in a 5 ml syringe with a 27 G needle and a high voltage of 14 to 19 kV were applied to the needle tip via the power supply, with a flow rate of 1.0 mL/h. A positively charged jet was formed from the syringe needle and nanofibers were sprayed onto a grounded apparatus at a distance of 10 cm away from the tip of the needle. A dehumidifier is used to control the humidity of the electrospinning hood to avoid bead formation due to the significant loss of charge from the spinning head that will result on the difference in solvent evaporation. Nanofibers collected on aluminum foil and coverslips were used for chemical, mechanical and biological analysis. The electrospun nanofibrous scaffolds were dried for a week at room temperature to remove any residual reagent.

2.2. Magnesium metallic glass coating by Sputtering:

The Mg-Zn-Ca thin film metallic glass were coated on electrospinned PCL and PCL/nHA nanofibrous scaffolds by DC magnetron sputtering in an argon atmosphere. The sputtering target was made into circular disc by inductively heating the high purity elements (purity level – 99.5% Ca, 99.99% Mg and 99.99% Zn) by keeping them in a graphite crucible and made into the shape by pouring the melt into a copper mould under a high-purity argon atmosphere (99.999%). Optimization of the process parameters such as sputter gas pressure, power, substrate temperature, target to substrate distance etc for the magnetron sputtering system was done. The target used for depositing the coating is 52 mm diameter and 3 mm thick high purity Mg60Zn35Catarget. Prior to deposition, the scaffolds were subjected to UV sterilization. The base pressure of the chamber before deposition was pumped to 5×10−5mbar, and Ar pressure of 2×10−3 mbar was maintained in the sputtering chamber while the sputtering was carried out. The coatings over the scaffold were deposited at a room temperature for duration of 10 min. For all the experiments, the sputtering power was maintained at 180 W. The PCL and PCL/nHA scaffolds were plasma processed in a downstream position with respect to the discharges, in order to preserve their structural, mechanical and biological properties.

2.3. Characterization

X-ray diffraction (XRD) analysis was carried to identify the crystal structure of the surface modified scaffolds using Bruker D8 Advance diffractometer with Cu Kα radiation. XRD spectra were recorded in the range of 10–60◦ 2θ in a step size of 0.02◦ which is the angle incident beam against the surface of the sample. Various functional groups present in the polymers samples were identified by Fourier transform infrared spectroscopy (FTIR; TENSOR 27, Germany).  The surface morphology of nanofibrous scaffolds was characterized by scanning electron microscope (SEM, TESCAN) with an accelerating voltage of 15 kV after coating with gold. The average diameter of the nanofibers were calculated and the composition was analyzed using energy dispersion X-Ray (EDX, equipped with SEM) analysis to confirm the presence of magnesium metallic glass coating on the scaffolds surface. Mechanical testing of the scaffolds was performed using a Universal Testing Machine.The compressive modulus was defined as the initial linear modulus and the yield strength was determined from the intersection of the two tangents on the stress-strain curve around the yield point. Five scaffolds were mechanically tested for each sample.

2.4. Biological analysis

2.4.1. In Vitro Degradation

For degradation experiments, the samples were finely cut into 20mm diameter disks, weighed, and totally immersed in 6 well plate filled with 10mL of phosphate buffer solution PBS (pH = 7.4), where they were evenly incubated at 37°C. The specimens were recovered after 1, 2 and 4 weeks under the same conditions, carefully dried at room temperature, and weighed to determine weight loss. Weight loss were evaluated by weighing, taking into account the original weight of each sample ( 0) and the residual weight, after degradation of the same specimen ( ) that had completely dried. The weight loss percentage, L%, was calculated by the following equation:[22]

Weight loss percentage ( L %) was estimated with the following equation:

WL% =W0-WrW0× 100


2.4.2. Hemocompatibility

Blood collection and preparation are the vital steps to carry out hemocompatibility testing. Blood was drawn by veinpuncture from aspirin-free healthy adult human donors and anti-coagulated with dipotassium EDTA (1.5 mg/L) or tri-sodiumcitrate (3.2%). Anticoagulant removes ionized calcium (Ca+) through a process referred to as chelation. This process forms an insoluble calcium salt that prevents blood coagulation. The amount of blood withdrawn must be within prescribed limits so as to maintain the proper ratio with the anticoagulant; otherwise, the blood cells may be damaged and/or anticoagulation may be unsatisfactory. Hemolysis was determined by the direct contact method. After fibrinogen was removed, the blood sample was adjusted by adding an appropriate quantity of normal saline (PBS) solution. The tested material was <0.5 mm thick and a total surface area, including both sides, of 2 cm2, placed in a 6 well plate.  10 mL saline solution was added to the 0.2 mL blood in a well plate containing evaluated materials. For the positive control, 10 mL of purified water and 0.2 mL of defibrinogenized blood were added to the test tube. As a negative control, 10 mL of saline and 0.2 mL of defibrinogenized blood were added to a plate and incubated for 3 h at 37°C under a static condition. After incubation, the treated samples were washed with normal saline and fixed in 2.5% (v/v) glutaraldehyde solution for 2 h. . Then the scaffolds were rinsed with distilled water and increasing the alcohol concentration (20 %, 40 %, 60 %, 80 % and 100 % of ethanol) twice with time interval. Three replicates were tested for each sample. Then the dried scaffolds are subjected SEM analysis to check the RBCs attachment over the scaffolds.

3. Results and Discussion

3.1. Materials properties

XRD pattern of the PCL/MgMG , PCL/nHA/MgMG scaffolds and bare magnesium metallic glass are shown in Fig. 1 There is no crystalline peaks observed in the bare magnesium metallic glass, whereas the PCL/MgMG and PCL/nHA/MgMG scaffolds shows sharp and distinct crystalline peaks. PCL, polymer with two distinct diffraction peaks at 21˚ and 23˚ and whereas PCL/nHA shows additional peaks at 25.82˚ and 31.72˚ confirming the presence of nHA (Inset of Fig. 1 (a), XRD pattern of nHA JCPDS   01-072-1243) with (002) and (211) plane arrangement of hexagonal crystal structure.[23-26] The morphological characterizations of the prepared scaffolds are given in Fig. 2. From the SEM images, the nanofibers of PCL and PCL/nHA are clearly visible with an average diameter of 650 nm. To further confirm the presence of nHA and magnesium metallic glass on PCL, PCL/nHA scaffolds, EDAX have been carried out and displayed in Fig. 2 (c, d). TEM images of PCL and PCL/nHA (Fig. 3  (a, c))  shows the nanofibers with smooth surface,[27] while after coating of magnesium metallic glass on PCL, the smoothness of nanofibers surface gets reduced (Fig. 3 (b, d)). The diameter of the uniform nanofibers is found to be: PCL (600±50 nm), PCL/nHA (650±40 nm) measured using Image J software. Moreover, the diameter of the coated PCL scaffolds gets slightly increased from ~650 nm to ~ 800 nm, which suggests that, the presence of magnesium metallic glass coating on the scaffolds surface. From EDAX, it is evident that the presence of Mg, Zn, Ca peaks in spectra indicates the formation of magnesium metallic glass on scaffolds, whereas the presence of peaks corresponds to C and P are reflected from the nHA in PCL scaffolds. Slight variation in stoichiometry is observed on elemental composition of magnesium metallic glass which may be due to the surface oxidation. A diffused ring pattern is observed in the SAED pattern (Fig.3 (e)) confirms amorphous nature of magnesium metallic glass coated PCL. While the nHA dispersed PCL scaffolds shows the crystalline spots (Fig. 3 (f)). The ring pattern in SAED observed through TEM supports the amorphous nature in the XRD results.

The electrospun and coated scaffolds were further characterized by FTIR analysis to affirm the vicinity of nHA nanoparticles, PCL and magnesium metallic glass. The vibration peak at 1727 cm-1 in Fig 4 is mainly arises from the ester group of PCL polymer. The asymmetric CH2 stretching and symmetric CH2 stretching was obtained at 2943 and 2862 respectively.  The amide I stretching at 1632 cm−1 (C=O stretching), the amide II stretching at 1525 cm−1 (N–H deformation) and the amide III stretching around 1210 cm−1 (N–H deformation) represents the presence of Mg. The phosphate ions (PO3−) are principal molecular components of nHA giving to IR absorbance in the 570–1210 cm−1 region. The characteristic vibrations of the phosphate group of nHA appeared at 1033, 603 and 560 cm−1. Stretching at 1593 cm−1 represent the C=O stretching of aliphatic and aromatic groups in Ca. These composites expressed PO43− stretch (1045 cm−1) and bending at 575 cm−1, confirm the presence of nHA in the nanocomposite of PCL/nHA.[28]  The major peak between 3800 to 2700 cm−1 represents the presence of hydroxyl groups. From this it is evident that the electrospinned nanofiber contains a phosphate group, an amino group and carboxyl groups for inducing mineralization and proliferation of cells for bone regeneration. [29]

Wettability of the scaffold is an important factor that plays the major role in biomedical applications including drug delivery, implants and tissue engineering. Hydrophilic nature allows the scaffolds to get drenched in body fluids and tends to degrade. In the present work, PCL and PCL/nHA scaffold showed hydrophobicity of 127 and 108° respectivley. During sputtering of magnesium metallic glass, PCL scaffolds were exposed to strong plasma produced by electric discharge with combination of gas for short interval. The gas becomes ionized and bombarded the surface of polymer and break the molecule chains resulting into the formation of functional groups including water compatible functional groups –OH, -NH2, -COOH and possible morphological alterations, which improves the surface energy of polymer component and will ensure the good adhesion of cells [30]. After sputtering process the contact angle reduces, the hydrophobicity of scaffolds becoming hydrophilic in nature. Fig 5 shows the PCL/MgMG and PCL/nHA/MgMG scaffolds turned into hydrophilic nature with the contact angle of 59o±5 and 48o ±5 respectively.  Increase in hydrophilicity indicates the increase in surface energy. The surface energy can be calculated from the contact angles by equation:


Where, Es is the surface energy between water and the specimen, Evl is the surface energy between water and air under ambient condition, (72.8 mJ/m2 at 20 °C) for pure water and θ is the static contact angle.[31] The calculated surface energies are shown in Table 1. The calculated surface energy of PCL and the PCL/nHA are decreasing in nature. There is an increase in surface energy for the PCL/MgMG and PCL/nHA/MgMG scaffolds. Materials with high wettability show high surface energy. Surfaces of high wettability help in adsorption of specific proteins and mineral phases which promote a strong bonding between the implant surface and the tissue.[32]

Tissue engineering systems are highly dependent on the chemical stability and mechanical properties of the scaffolds which shows a huge impact on the hemocompatabiliy and growth of the tissues. However nHA has the major role in boosting the mineral secretion and its low tensile strength and brittleness hampers its application to hard tissue implants. Fig. 6 shows the stress-strain curves of PCL, PCL/nHA, PCL/MgMG and PCL/nHA/MgMG scaffolds under normal room temperature. The PCL fibers show the maximum elastic strain of 114 % among the other scaffolds, whereas PCL/MgMG scaffolds showing less elastic strain of 40.7% with an ultimate stress of 2.50±0.15MPa. The MgMG coating decreases the yield strength by breaking the bond between the nHA particles with the PCL polymer by showing reduction in strain. The higher tensile strain of PCL might be reflected from the interface between the crystals, semi-crystalline region, shear alignment of the molecular chain and its unique mechanical properties of crystalline magnesium.[16] Comparing with PCL, the PCL/MgMG and PCL/nHA/MgMG scaffolds are lagging in tensile strength with respect to the other fibers and showing the elastic breaking of 40.7 and 24.4 % respectively.

3.2. Biological in-vitro characterization

3.2.1. Degradation studies

The degradation behavior of prepared scaffolds are studied in phosphate buffered solution (PBS, pH =7.4), at 37°C, for a period of 28 days. After 28 days, it is observed that the PCL/MgMG and PCL/nHA/MgMG scaffolds accelerated the weight loss and increased their capacity to absorb water during the initial degradation process. However, the magnesium metallic glass coating increases the degradation behavior of the PCL and PCL/nHA scaffolds. which might be due to the hydrolysis of ester is faster in comparison to PCL in this study over the period of degradation.[33] The effect of weight loss of the scaffolds, after degradation studies has been analyzed and given in Fig.7. The in-vitro degradation of magnesium metallic glass coated nanofibrous scaffolds are found to be safe as an implantable material. Zreiqat et al. [34] found that magnesium ions could enhance the adhesion of human bone-derived cells. It has been reported that zinc can also promote bone formation [35] and therefore, it is postulated that the magnesium metallic glass has good biocompatibility. In addition, the PCL/nHA scaffolds contained Ca and P, which could reinforce the bone cell activity. Hence, we suggested that rapid degradation of the magnesium metallic glass is not harmful to adjacent bone tissues. The excess magnesium produced by degradation can be excreted by the kidneys during degradation and no adverse effects of bubbles were observed [36]. Zinc is an essential element for humans, which is absolutely necessary and is non-toxic except at an extreme exposure. The human requirement for zinc is estimated to be 15 mg day-1.[37] The zinc content of the magnesium metallic glass in this study is found to be very less quantity of 25 wt.%, hence the zinc release rate will be 0.0965 mg cm-2 day-1. For an implant with the dimensions 4.5 x 10 mm (cylindrical, with a surface area of 1.732 cm2), the zinc release rate would be about 0.11 mg day-1, which is much lower than the suggested intake of 15 mg day-1 for adults. Besides, the released zinc could be absorbed by the surrounding tissues [38] and excreted through the gastrointestinal route and the kidney.[37] Compared to previous reports, the present work shows the very less amount of zinc release rate and therefore, it is proposed that the zinc release during degradation is safe and non-toxic for implants.

3.2.2 Apatite formation

Much has been written about apatite layer formation on materials in simulated body fluid. However, in many papers, there is no qualitative assessment of basic properties of the so-called apatite, such as atomic composition, Ca/P ratio, crystal morphology and crystallinity, amorphous calcium phosphate have been shown to provoke cytotoxicity in vitro and in vivo, likely because of high concentrations of both calcium and phosphates. In order to analyze the bioactivity of the scaffold, the calcium phosphate (CaP) has been deposited onto the substrates as a ceramic control using an immersion method. This process leads to the formation of a low-crystalline apatite (CaP) on the Scaffolds. In vitro growth of bone-like apatite particles on scaffolds is evaluated using SEM and mass increase as a function of incubation time. Fig. 8 shows the SEM micrographs of apatite formation on nanofibrous scaffolds after varying times of incubation in SBF. Substantial amounts of bone-like apatite crystals are formed on nanofibrous PCL scaffolds after 15 days of incubation in SBF Fig. 8 (a). The deposited apatite particles grow into a few hundred nanometers in size and have nanostructured surface features. The underlying nanofibers become hardly observable after 15 days of incubation for magnesium metallic glass coated scaffolds. Further incubation in the SBF leads to a continuous apatite layer formation, covering the entire fibrous surfaces Fig. 8 (c). Even after the apatite formation, the interconnected structure of the scaffolds remains same, which is important for cell migration and mass transport when used for implants. Compared to metallic glass coated scaffolds, there is no appreciable apatite formation in the PCL nanofiber. Preincorporation of nHA particles in polymer scaffolds (even at a low content of 10 w/w%) induces greater amounts of apatite formation in SBF as compared to pure PCL scaffolds. Substantial amounts of apatite crystals are grown on PCL/nHA composite scaffolds after 15 days of incubation in SBF Fig. 8(b). However, the Ca and P elements are not detectable for pure PCL scaffolds after 6 days of incubation in the SBF. The size and morphology of apatite particles on PCL/MgMG, PCL/nHA/MgMG nanocomposite scaffolds (Fig.8) are different from those on PCL scaffolds.

3.2.3. Hemocompatibility

The prepared scaffolds samples were tested for the evaluation of hemocompatibility studies and it is provided in Fig. 9. The results shows that the of magnesium metallic glass coated scaffold are found to be highly hemocompatible when made to contact with the blood cells. It is noted that there is no blood coagulation as well as the biconcave disc shaped morphology. The blood cells remains the same even after coming in contact with the coated scaffolds as noticed in Fig. 9(d), when in contact with PCL scaffold blood coagulation is observed due to hydrophobic nature as shown in Fig. 9(a)

The increase in surface energy of PCL/MgMG is high as compared to bare PCL which may be  due to the reduction in contact angle after MgMG coated PCL scaffolds. A similar effect is observed in PCL/nHA and PCL/nHA/MgMG scaffolds. A higher surface energy helps in good cellular attachment leading to better cell spreading and their proliferation is observed from SEM image of hemolysis study (Fig. 9). From the micrograph it is seen that the red blood cells (RBCs) are fully distorted on the surface of PCL scaffolds. A higher number of echinocytes is observed on PCL surface. The presence of echinocytes indicates a stressed environment for the RBCs which lead to their transformation into echinocytes. Although no visible thrombus formation is observed on the surface of PCL scaffolds, activation of thrombogenesis cycle indicated by the thick fibrin mesh as observed from the micrograph.

The attachment of RBCs on PCL/MgMG shows the biconcave disc like morphology for the attached RBCs which can be attributed due to higher surface energy of the scaffold. The fibrous scaffold are visibly intact unlike in the case of bare PCL scaffolds, the entire surface has covered by unseparable fibrin meshes Fig 9c. In the case of PCL/nHA no fibrin meshes are observed from the SEM images. A mixture of biconcave RBCs and fewer echinocytes are observed on the surface of the scaffolds indicating the hemocompatibile nature of the scaffold. The increased contact angle of PCL/nHA have attributed to the formation of RBCs into echinocytes. A partially distorted RBCs along with a higher number of echinocytes as compare to PCL/nHA is observed on the surface of PCL/nHA/MgMG scaffolds. This echinocytes transformation is mainly attributed due to the higher water permeability induced to the scaffold during plasma deposition, creating a tensed environment for the blood cells thereby hindering their better attachment and spreading. [39]


In summary, we have fabricated the PCL based nanofibrous scaffolds using electrospinning method for biodegradable implants. The mechanical strength, chemical composition wettability and degradation of the scaffolds, has improved by the modification of PCL scaffolds by dispersing nHA and magnesium metallic glass coating. However, bone-like apatite layer has grown uniformly throughout the surface modified scaffolds after immersion in SBF solution, which has not only demonstrated the bioactivity of the scaffold but also provided a method to generate biomimetic surfaces. The mechanical properties of the metallic glass coated scaffolds with required tensile strength and elongation were suitable for implant applications. Degradation of the PCL/MgMG and PCL/nHA/MgMG resulted in deterioration in the mechanical properties and mechanical integrity was rapidly lost in the early stages of hydrolysis, but more slowly in the later stages of in vitro degradation. Magnesium and Zinc elevated the degradation rate of the magnesium metallic glass which was slower than that of pure PCL in SBF. Magnesium alone is not suitable to use as the bone grafting material. Hence, magnesium alloy with Zn and Ca are used as to form the metallic glass thin films over PCL matrix so as to make the scaffolds implantable as well. PCL/MgMG and PCL/nHA/MgMG hybrids have controlled release of magnesium. This suggested that the resulting scaffolds should have better hemocompatibility and improved mechanical properties for orthopedic implants.


Authors thank the Department of Science and Technology New Delhi, for the research grant No. SB/S3/ME/042/2014 to carry out this work.

Conflicts of interest

The authors declare no conflict of interests that could potentially influence or bias the submitted work.


[1] H Wang, Z Shi (2011) In vitro biodegradation behavior of magnesium and magnesium alloy Journal of Biomedical Materials Research Part B: Applied Biomaterials 98: 203.

[2] MP Staiger, AM Pietak, J Huadmai, G Dias (2006) Magnesium and its alloys as orthopedic biomaterials: a review Biomaterials 27: 1728.

[3] F Witte, N Hort, C Vogt, et al. (2008) Degradable biomaterials based on magnesium corrosion Current opinion in solid state and materials science 12: 63.

[4] N Hort, Y Huang, D Fechner, et al. (2010) Magnesium alloys as implant materials–principles of property design for Mg–RE alloys Acta biomaterialia 6: 1714.

[5] AF Cipriano, A Sallee, R-G Guan, et al. (2015) Investigation of magnesium–zinc–calcium alloys and bone marrow derived mesenchymal stem cell response in direct culture Acta biomaterialia 12: 298.

[6] Z Xu, C Smith, S Chen, J Sankar (2011) Development and microstructural characterizations of Mg–Zn–Ca alloys for biomedical applications Materials Science and Engineering: B 176: 1660.

[7] B Zhang, Y Wang, L Geng (2011) Biomaterials-Physics and ChemistryInTech,

[8] B Zhang, Y Hou, X Wang, Y Wang, L Geng (2011) Mechanical properties, degradation performance and cytotoxicity of Mg–Zn–Ca biomedical alloys with different compositions Materials Science and Engineering: C 31: 1667.

[9] E Sardella, V Armenise, R Gristina, P Favia Plasma sputter-deposition of Mg-containing coatings for the regeneration of bone tissue.

[10] S Farahany, HR Bakhsheshi-Rad, MH Idris, MRA Kadir, AF Lotfabadi, A Ourdjini (2012) In-situ thermal analysis and macroscopical characterization of Mg–xCa and Mg–0.5 Ca–xZn alloy systems Thermochimica acta 527: 180.

[11] Y Ortega, MA Monge, R Pareja (2008) The precipitation process in Mg–Ca–(Zn) alloys investigated by positron annihilation spectroscopy Journal of Alloys and Compounds 463: 62.

[12] C Boehlert, K Knittel (2006) The microstructure, tensile properties, and creep behavior of Mg–Zn alloys containing 0–4.4 wt.% Zn Materials Science and Engineering: A 417: 315.

[13] J Nie, B Muddle (1997) Precipitation hardening of Mg-Ca (-Zn) alloys Scripta Materialia 37: 1475.

[14] AF Cipriano, C Miller, HN Liu (2014) Advanced Materials ResearchTrans Tech Publ,

[15] Y Zhao, G Wu, Q Lu, et al. (2013) Improved surface corrosion resistance of WE43 magnesium alloy by dual titanium and oxygen ion implantation Thin Solid Films 529: 407.

[16] TS Narayanan, I-S Park, M-H Lee (2015) Surface Modification of Magnesium and Its Alloys for Biomedical Applications: Modification and Coating Techniques. Elsevier,

[17] WS Khan, F Rayan, BS Dhinsa, D Marsh (2011) An osteoconductive, osteoinductive, and osteogenic tissue-engineered product for trauma and orthopaedic surgery: how far are we? Stem cells international 2012.

[18] K Rezwan, Q Chen, J Blaker, AR Boccaccini (2006) Biodegradable and bioactive porous polymer/inorganic composite scaffolds for bone tissue engineering Biomaterials 27: 3413.

[19] MI Sabir, X Xu, L Li (2009) A review on biodegradable polymeric materials for bone tissue engineering applications Journal of materials science 44: 5713.

[20] P Turon, LJ del Valle, C Alemán, J Puiggalí (2017) Biodegradable and Biocompatible Systems Based on Hydroxyapatite Nanoparticles Applied Sciences 7: 60.

[21] W Suchanek, M Yoshimura (1998) Processing and properties of hydroxyapatite-based biomaterials for use as hard tissue replacement implants Journal of Materials Research 13: 94.

[22] E Díaz, I Sandonis, MB Valle (2014) In vitro degradation of poly (caprolactone)/nHA composites Journal of Nanomaterials 2014: 185.

[23] JE Oliveira, LH Mattoso, WJ Orts, ES Medeiros (2013) Structural and morphological characterization of micro and nanofibers produced by electrospinning and solution blow spinning: a comparative study Advances in Materials Science and Engineering 2013.

[24] A Haider, KC Gupta, I-K Kang (2014) PLGA/nHA hybrid nanofiber scaffold as a nanocargo carrier of insulin for accelerating bone tissue regeneration Nanoscale research letters 9: 314.

[25] M Sadjadi, M Meskinfam, B Sadeghi, H Jazdarreh, K Zare (2010) In situ biomimetic synthesis, characterization and in vitro investigation of bone-like nanohydroxyapatite in starch matrix Materials Chemistry and Physics 124: 217.

[26] V Thomas, S Jagani, K Johnson, et al. (2006) Electrospun bioactive nanocomposite scaffolds of polycaprolactone and nanohydroxyapatite for bone tissue engineering Journal of nanoscience and nanotechnology 6: 487.

[27] Q Wei (2012) Functional nanofibers and their applications. Elsevier,

[28] MP Prabhakaran, J Venugopal, S Ramakrishna (2009) Electrospun nanostructured scaffolds for bone tissue engineering Acta biomaterialia 5: 2884.

[29] JR Venugopal, S Low, AT Choon, AB Kumar, S Ramakrishna (2008) Nanobioengineered electrospun composite nanofibers and osteoblasts for bone regeneration Artificial organs 32: 388.

[30] D Yan, J Jones, X Yuan, et al. (2013) Plasma treatment of electrospun PCL random nanofiber meshes (NFMs) for biological property improvement Journal of Biomedical Materials Research Part A 101: 963.

[31] RN Wenzel (1936) Resistance of solid surfaces to wetting by water Industrial & Engineering Chemistry 28: 988.

[32] P Jojibabu, BR Sunil, TS Kumar, U Chakkingal, V Nandakumar, M Doble (2013) Wettability and in vitro bioactivity studies on titanium rods processed by equal channel angular pressing Transactions of the Indian Institute of Metals 66: 299.

[33] D Milovac, TC Gamboa-Martínez, GG Ferrer, M Ivanković, H Ivanković (2014) 16th European Conference on Composite Materials, ECCM16,

[34] H Zreiqat, C Howlett, A Zannettino, et al. (2002) Mechanisms of magnesium‐stimulated adhesion of osteoblastic cells to commonly used orthopaedic implants Journal of Biomedical Materials Research Part A 62: 175.

[35] M Hashizume, M Yamaguchi (1993) Stimulatory effect of β-alanyl-L-histidinato zinc on cell proliferation is dependent on protein synthesis in osteoblastic MC3T3-E1 cells Molecular and Cellular Biochemistry 122: 59.

[36] S Zhang, X Zhang, C Zhao, et al. (2010) Research on an Mg–Zn alloy as a degradable biomaterial Acta biomaterialia 6: 626.

[37] H Tapiero, KD Tew (2003) Trace elements in human physiology and pathology: zinc and metallothioneins Biomedicine & Pharmacotherapy 57: 399.

[38] L Xu, G Yu, E Zhang, F Pan, K Yang (2007) In vivo corrosion behavior of Mg‐Mn‐Zn alloy for bone implant application Journal of Biomedical Materials Research Part A 83: 703.

[39] G Benga, T Borza (1995) Diffusional water permeability of mammalian red blood cells Comparative Biochemistry and Physiology Part B: Biochemistry and Molecular Biology 112: 653.

Figure legends

Figure 1:  X-ray diffraction pattern of PCL/MgMG, PCL/nHA/MgMG and MgMG

Figure 2: SEM Micrographs of (a) PCL (b) PCL/HA and EDAX of (c) PCL/MgMG (d) PCL/nHA/MgMG

Figure 3: TEM images of (a) PCL, (b) PCL/MgMG, (c) PCL/HA, (d) PCL/HA/MgMG and SAED pattern of (e) PCL/MgMG and (f) PCL/HA/MgMG.

Figure 4: FTIR Spectra of pure (a) PCL, (b) PCL/HA, (c) PCL/MgMG and (d) PCL/HA/MgMG

Figure 5: Water contact angle for (a) PCL, (b) PCL/MgMG, (c) PCL/HA, (d) PCL/HA/MgMG

Figure 6: Stress-strain curve of pure PCL and Surface modified scaffolds

Figure7: The dynamic weight loss degradation of the uncoated and coated Scaffolds in PBS at different stage

Figure 8: SEM micrograph of apatite layer formation over of (a) PCL, (b) PCL/MgMG, (c) PCL/HA, (d) PCL/HA/MgMG scaffolds

Figure 9: Scanning electron micrograph of erythrocytes, representing changes of RBCs morphology after contacting with (a) PCL, (b) PCL/MgMG, (c) PCL/HA, (d) PCL/HA/MgMG.


Table 1:  Contact angles of the scaffolds and their surface energy


Specimen Contact angle Surface energy 


PCL 127° -43.812
PCL/nHA 108° -22.496
PCL/MgMG 59o 37.494

EssayHub’s Community of Professional Tutors & Editors
Tutoring Service, EssayHub
Professional Essay Writers for Hire
Essay Writing Service, EssayPro
Professional Custom
Professional Custom Essay Writing Services
In need of qualified essay help online or professional assistance with your research paper?
Browsing the web for a reliable custom writing service to give you a hand with college assignment?
Out of time and require quick and moreover effective support with your term paper or dissertation?