Electrophoretic deposition of hydroxyapatite‐shrimp crusts nanocomposite thin films for bone implant studies

Abstract

Hydroxyapatite‐shrimp crusts nanocomposite thin films were deposited on titanium substrates by electrophoretic technique, under different preparation conditions, for bone implant applications. Fourier transform infrared spectrometer, atomic force microscope, X‐ray diffraction (XRD), optical microscope, and scanning electron microscope were employed to characterise the synthesised films. Vickers’ micro‐hardness measurements revealed a value of 502 HV for the hydroxyapatite films and 314.55 HV for the nanocomposite films. XRD results confirmed the polycrystalline nature of the hydroxyapatite and hydroxyapatite‐shrimp nanocomposite films. The in‐vitro bioactivity test of the synthesised films in simulated body fluid showed very low dissolution rate. Antibacterial activity of synthesised films was investigated against E. coli bacteria.

1 Introduction

Biomaterials are used in medical implants to repair damaged natural tissues. Hydroxyapatite (HAp) Ca₁₀ (PO₄)₆ (OH)₂ is an important inorganic bioactive [1], and a preferred biomaterial for orthopaedics use as bulk, coatings, or composites with polymers [2], due to its similarity to the bone apatite's chemical composition and crystal structure, and its favourable osteo‐conductive and bioactive properties [3]. However, it suffers from poor mechanical features, as those of bone, thus, it cannot replace bone in load‐bearing sites. One approach for solving this problem is to apply HAp as a coating on metal implants such as Ti and its alloys [4]. A coating of HAp layer on a metal can keep the mechanical properties of the metal such as load‐bearing ability, and to assist the osteo‐integration of these implants with surrounding tissues [3]. There are several ways to deposit HAp films on metal implants; these include: sputtering, pulsed laser deposition, electrochemical, electrophoretic, and plasma spray methods. Electrochemical deposition is, however, favoured over other methods because it allows homogeneous, highly crystalline, and stoichiometric coatings at low temperatures (<100°C) [5]. Besides, it can control thickness, structure [6], morphology, composition, and adhesion by the appropriate selection of the process parameters [7]. Composition and properties of the coatings deposited electrochemically depend on typical parameters and conditions, e.g. applied current/voltage, temperature, pH value, and deposition time [8]. The increase in both temperature and current density can raise the percentage of HAp in the coated film [9]. Electrodeposition of HAp on Ti‐6Al‐4V substrates showed good adhesion and uniformity. The resulted enhanced mechanical properties and biocompatibility suggests its successful use for biomedical applications [10]. TiN was pre‐deposited by DC magnetron sputtering on stainless steel substrates, and then coated with HAp to improve the steel stability and compatibility with bone [11]. The electrochemical behaviour of TiN film on the 316L stainless steel (316LSS) material was investigated in simulated body fluid (SBF) solution for implant application. Results, after 21 days in SBF, indicated higher durability of the TiN/316LSS than that of 316LSS [12]. Nano HAp was coated on Mg‐2Zn by pulse electro‐deposition method, and then post‐treated with alkaline solution to improve the biodegradation behaviour and biocompatibility for implant applications. The microstructure and composition of the samples and their degradation and cytotoxicity in SBF showed a natural bone mimic – needle‐like HAp structure perpendicular to the substrate [13]. Carbon nanotubes/carbon fibres hybrid materials were used as templates to deposit carbonate‐containing HAp by ultrasound‐assisted electro‐deposition method [14]. Cu (II) and Zn (II) co‐substituted HAp (ZnCuHAp) coating on commercially pure titanium was fabricated to improve cyto‐compatibility and antibacterial property of pure HAp. This resulted in reduced porosity, high corrosion protection for Ti, and denser HAp coating. The ZnCuHAp coating was effective against Escherichia coli [15]. HAp and HAp/ZrO₂ composites were coated on 316L stainless steel substrates in the ZrO₂ solutions. The in‐vitro test in SBF and further scanning electron microscope (SEM) observations showed 30 folds corrosion resistance improvements of the composite coating [16]. Cerium substituted HAp (Ce‐HAp) on borate passivated 316L SS in Ringer's solution showed an enhanced anti‐corrosion performance of Ce‐HAp coating. The antimicrobial performance of the obtained coatings was studied against pathogenic gram‐positive S. aureus and negative bacterial strains E. coli [17]. A manganese‐substituted porous HAp/zinc oxide duplex layer was fabricated on the AZ91‐Mg alloy by electro‐deposition to improve the corrosion resistance and biological performance of AZ91‐Mg alloy. The potentiodynamic polarisation investigation indicated an improved corrosion resistance of AZ91‐Mg alloy by the duplex‐layer coating [18]. Pulsed electro‐deposition coating of biocompatible HAp from a diluted bath of calcium and phosphate salt over SS316 was performed to assure quick bonding between implant and body tissues with minimum recovery time for the patient [19]. Carbon nanotubes‐HAp (CNTs‐HA) composite coating was synthesised by a combined electrophoretic and pulsed electro‐deposition. Potentiodynamic polarisation and electrochemical impedance spectroscopy studies showed that the CNTs‐HA composite coatings were well adhered and protected the bare carbon/carbon composites from corrosion in SBF solution [20]. Single phase, 9.6 µm, uniform and dense fluoridated HAp (FHAp) coatings were deposited on 316L stainless steel (316LSS) substrate by electrodeposition method. In‐vitro test of the apatite crystals in SBF showed a cactus‐like shape which, after 21 days, grew to form a thick block on the surface of FHAp/316LSS leading to a decrease of the corrosion current density [21]. Magnesium‐phosphate‐doped HAp and magnesium‐hydrogen‐phosphate‐doped di‐calcium‐phosphate‐dehydrate composite coatings were electrodeposited onto an AZ31 alloy. Their corrosion via electrochemical and in‐vitro degradation tests showed corrosion resistance and stability, significantly higher than those of single coatings [22]. Strontium/copper substituted HAp (SrCuHA) coating on commercially pure titanium was studied to show their effect on antibacterial and in‐vitro cyto‐compatible properties. Cu improves HA antimicrobial properties while Sr improves the biocompatibility. SrCuHA coating, in the in‐vitro SBF, exhibited high corrosion resistance and good osteoblast adhesion, which are needed for orthopaedic and dental applications [23]. In previous publication, simple novel method was demonstrated to transform chlorapatite to nanostructured HAp thin coatings on steel substrates by using pulsed Nd: YAG laser deposition [24]. Hydroxylapatite [HA, Ca₁₀ (PO4)₆ (OH)₂] is one among many biomaterials used to coat titanium (Ti) implants. HA coatings can stimulate the formation of bone‐like structure on the surface of the metal implant and help fixing the implant with less future problems. HAp films are preferred over other implants’ coatings because of their efficient integration with bones. Their bioactive nature, however, makes them vulnerable to bacterial infection. For this reason the HA films need to be modified by adding other materials to solve this problem. Silver‐ and strontium‐modified antibacterial HA layers were electrochemically deposited onto TiO₂ nanotubes (TNs) to improve the antimicrobial properties. They showed crack‐free coatings with improved corrosion resistance, adherence, cytocompatibility, and antibacterial effect. The immersion of these coatings in SBF demonstrated cell proliferation, very high corrosion resistance, and better cytocompatibility than naked HA‐Ti structure [25]. The high bonding strength of this structure suggests their possible use in orthopaedic applications. AgHA/CS layer was deposited on Ti‐TN alloy by electro‐deposition and combined with anodisation. Ag was incorporated into HA coating to improve the antimicrobial properties. The adhesion strength and corrosion resistance increased after AgHA/CS/TN coating. The antibacterial effect of AgHA/CS/TN revealed excellent inhibition efficiency towards S. aureus and shows sustainable release of Ag ions, needed to treat bone infection [26]. The AgHA/CS/TN coatings improved surface wettability and showed efficient bone cyto‐compatibility. In this study, HAp and HAp‐shrimp crust (Hap‐SC) films were deposited on titanium substrate by electrophoresis technique to study their possible use in bone plant applications. The antimicrobial activity of HAp‐SC nanocomposite showed good antibacterial activity.

2 Experimental work

In this study, two electrolyte solutions were prepared in order to deposit HAp thin films on titanium substrates. In the first electrolyte, diammonium hydrogen phosphate (NH₄)₂ HPO₄ and calcium nitrite Ca(NO₃)₂ were dissolved in deionised water at a volume molar ratio of 3.861 mM:6.435 mM, respectively. In the second electrolyte, calcium nitrite Ca(NO₃)₂ and sodium phosphate Na₃ PO₄ were dissolved in deionised water at a volume molar ratio of 6.435 mM:3.851 mM, respectively. The phosphate source was added slowly to the calcium source giving Ca/P = 1.67 and pH = 6.1. The shrimp crusts (SCs) were milled by an electrical miller for 15 min, dissolved in 2.25% H₂ O₂, and then mixed with the electrolyte. The final mixture was then stirred for 10 min to get homogenous solution. DC potential difference of 65 V was used between the 6 mm separated platinum sheet anode and titanium sheet cathode with different deposition conditions. The deposition temperature 25–70°C was controlled by applying 65 V for certain periods of time – up to 2 h. Hydrogen peroxide was added to the electrolyte to occupy 2.25% of the total electrolyte volume. The details of deposition conditions are summarised in Table 1.

Table 1.

Electrophoresis deposition conditions

Samples	Deposition materials		Voltage, V	Time, min	Electrolyte temp., °C	RMS surface roughness, mm	Additives vol.%
E1	Ca(NO3)2	(NH4)2 HPO4	65	120	25	0.964	—
E2	Ca(NO3)2	(NH4)2 HPO4	65	120	45	0.964	2.25%H2 O + 0.07 g CS
E3	Ca(NO3)2	(NH4)2 HPO4	65	120	70	2.33	2.25%H2 O + 0.07 g CS

Finally, the samples were heat treated for 2 h at 400°C using a temperature‐controlled tube furnace. The films’ surface morphology and root mean square roughness were investigated by using atomic force microscope (SPM‐model AA3000, Angstrom Advanced Inc., USA). X‐ray diffractrometer (PHILIPS PW, CuKα, λ  = 0.154 nm) was used to search structural properties. Vickers’ hardness study was conducted by using digital micro‐hardness tester (HVS‐1000, LARYEE). Finally, films’ thickness was measured by weighing method using a four digits balance. SEM (Tescan VEGA3) was used to investigate the morphology and structure of the synthesised films. The in‐vitro test was performed by immersing the samples in SBF for 1 month (720 h). SBF was prepared by adding the compounds listed in Table 2 to a 750 ml deionized water (DIW), then stirred under fixed temperature of 37°C and PH 7, until all compounds were dissolved.

Table 2.

SBF preparation compounds [27]

Compound	Amount
NaCl	7.99 g
NaHCO3	0.35 g
KCl	0.22 g
K2 PH4. 3H2 O	0.22 g
MgCl2. 6H2 O	0.3 g
1 kmol/m3 HCl	40 cm3
CaCl2	0.27 g
Na2 SO4	0.07 g
(CH2 OH)3 CNH2	appropriate amount to adjust pH

Gram‐negative E. coli bacterial stain (obtained from the Medical Microbiology Laboratory, Biotechnology Branch, Department of Applied Science, University of Technology, Baghdad‐Iraq) was used in the present study. The bacterial suspension was prepared by making a saline suspension of isolated colonies selected from nutrient agar plate, grown for 18–24 h. Approximately 20 ml of nutrient agar was poured into sterilised petri dishes. The plates were left overnight at room temperature to check for any contamination to appear. Agar wells of 6 mm diameter were prepared with the help of a sterilised stainless steel cork‐bore.

3 Results and discussion

Fig. 1 illustrates the SEM image of SC particles showing different morphologies with size in the range of 1–6 μm.

Fig. 1.

SEM image of SC particles

Fig. 2 shows a cross‐sectional view of SC‐HAp films prepared at different deposition temperatures. It is clear that the film thickness of sample E3 is smaller than that of E1 and E2 samples. Fig. 3 a shows an optical micrograph of E1 sample. It illustrates the formation of compact and homogeneous HAp film of high density of pores but with some micro‐cracks. The HAp film prepared at 45°C temperature (Fig. 3 b) shows smaller microstructural cracks but with relatively unstable film thickness [28]. Fig. 3 c illustrates an optical micrograph of SC‐HAp film deposited at 70C°, showing non‐uniform film thickness. Deposition at high temperature disperses the particles more than at low temperature. The formation of pores can be ascribed to the gas evolvement from cathode when using high deposition voltage [29].

Fig. 2.

Cross‐sectional view of HAp‐SC composite prepared at different temperatures

(a) 25°C, (b) 45°C, (c) 70°C

Fig. 3.

Optical micrographs of HAp‐SC films prepared at different temperatures

(a) 25°C, (b) 45°C, (c) 70°C

The key requirement for bone implant is the preparation of carbonated apatite in order to mimic the bone's chemical composition. The tested commercial HAp powder shows the same peaks of standard HAp given in [30]. Fig. 4 shows Fourier transform infrared spectrometer (FTIR) spectra of samples: E1, E2, and E3 with similar peaks, but of higher or lower intensities. These peaks are located at 573–604 cm⁻¹, assigned to antisymmetric v ₄ bending mode of PO₄ ⁻³, 962 and 1040 cm⁻¹ for symmetric v ₁ stretching PO₄ ⁻³ mode in addition to the antisymmetric v ₃ stretching PO₄ ⁻³ mode. The peaks at 3572, 1635, 3750 cm⁻¹ and the broad 3200–3400 cm⁻¹ indicate stretching vibration O–H, bending vibration of O–H, O–H stretching, and absorbed H₂ O, respectively. The peaks at 873, 1410, 2374, and 2925 cm⁻¹ indicate the presence of v ₂ stretching CO₃ ⁻² mode, the stretching CO₃ ⁻² vibration v₂ mode, the carbonyl group, and the C–H stretching bond of the CH₂ group, respectively. The peaks at 1517 and 1745 cm⁻¹ are indexed to C=C, and carboxyl stretching mode, respectively. The obtained carbonated apatite is attributed to the 872 cm⁻¹ peak [31]. The SC composition is similar to the polymeric material that exists in the composition of real bone. The only difference is its higher polymer percentage in SC than in real bone. The FTIR transmission spectrum (Fig. 4 a) of tested shrimp‐HAp composite powder shows peaks at 557–602, 875, 963, 1040, 1423, 1630, 2925, 3200–3400, and 3564 cm⁻¹ which belong to HAp. The remaining peaks are located at 1076, 1159, 1380, 1410, 1540, 1660, 1747, and 2852 cm⁻¹ indicating C‒O stretching, C‒O‒C stretching, asymmetric stretch vibration v ₃ of CO₃, amide II (NH₂), amide I (C=O), C=O, and asymmetric mode of CH₂, respectively. The main peaks of SCs appeared in these films spectra shown in Fig. 4 a, and are located at 1540 and 1660 cm⁻¹ referring to amide I and amide II, respectively.

Fig. 4.

FTIR spectra of HAp‐SC films

(a) E1, (b) E2, (c) E3

These two peaks prove the deposition of polymeric SCs particles with the HAp particles on titanium substrates, usually utilised in bone implant applications. Films containing valleys between relatively large accumulated particles give the film a greater surface roughness, a critical characteristic for bone cells film growth and penetration, to enhance osteo‐integration and increase the film surface area [32]. Pore size on the human bone's surface is around 100 nm, and therefore, a comparable value to this size, between the grains of the film will facilitate the penetration and hooking of the growing tissue to the implant. Grain size and voids may have close values; therefore, grain size between 80–200 nm was one of the targets in this work. To study the effect of morphology and surface roughness, atomic force microscope (AFM) studies were carried out and the results are presented in Fig. 5. The E1 film is uniform and porous, but covers a small area of the substrate due to the formation of H₂ bubbles which prevent the deposition of HAp. The addition of H₂ O₂ has reduced porosity and H₂ bubbles formation, and increased the film covering area and film deposited weight. The highest average grains size and its percentage from AFM studies were 118.22 nm and 86.55%, respectively, with grains size increased with deposition temperature. The results obtained from E1, E2, and E3 films show uniform and porous surfaces. The XRD patterns of E1, E2, and E3 films, treated at 400°C for 2 h, show titanium peaks in addition to others, resulted from the deposited films. The peaks in Fig. 6 indicate polycrystalline HAp structure according to the standard data (JCPDS No. 09‐0432). Peaks belonging to the α‐Ca₃ (PO₄)₂ phase located at 36.4° and 37.8° were detected from all the films according to the standard data (JCPDS No. 29‐0359). E2 film's patterns show higher diffraction peak at 37.8° than in E1 and E3 films, originated from SC. The addition of polymeric SCs powder has broadened the peaks, which agrees with the AFM findings. The increase in peak intensity of the titanium substrate in E2 film refers to the crystallinity decrease after polymeric SCs addition. The grains size was measured using Scherer equation and the results are presented in Table 3. The increase in (002) and (004) peaks intensity of E1 and E3 films patterns as compared to the standard peaks intensity is attributed to the HAp crystals elongation. This is a characteristic of the bone appetite. These results are in good agreement with some published results [33]. The increase in peak intensity from one condition to another indicates a crystallinity degree increase. The high crystallinity reduces the film dissolution inside the body and enhances the film growth [32]. The formation of OH ions close to the substrate surface, during the electrochemical deposition process, results in pH increase at this position, which in turn increases the HAp formation at pH ≥ 6 [30]. The addition of H₂ O₂ increases the formation of OH ions, which in turn, increases the pH. In addition, the increase in electrolyte temperature allows the formation of more OH ions through the dissociation reaction of H₂ O₂. The best crystallinity was obtained when treating at 400°C, but reduced after adding the SC powder. Vickers’ hardness (HV) of the composite films was measured and listed in Table 4.

Fig. 5.

Three‐dimensional AFM images and grains size distribution of HAp‐SC films

(a) E1, (b) E2, (c) E3

Fig. 6.

XRD patterns of E1, E2, and E3 samples

Table 3.

XRD data of HAp‐SC films prepared at various deposition temperatures

$2\theta$, deg.	Orientation	Type	FWHM, deg.			Crystallite size, nm
			E1	E2	E3	E1	E2	E3
218	(200)	HAp	0.1276	—	0.1176	63.382	—	68.771
22.9	(111)	HAp	0.1511	—	0.1411	53.628	—	57.429
25.35	(201)	HAp	0.1814	—	0.1714	44.872	—	47.490
25.9	(002)	HAp	0.1528	—	0.1428	53.331	—	57.065
28.15	(102)	HAp	0.2683	—	0.2583	30.516	—	31.698
28.95	(210)	HAp	0.2877	—	0.2777	28.508	—	29.535
31.75	(211)	HAp	0.1016	0.1648	0.0916	81.265	50.100	90.137
32.2	(112)	HAp	0.4044	0.2014	0.3944	20.438	41.039	20.957
32.9	(300)	HAp	0.21	0.2399	0.2	39.431	34.517	41.403
34.05	(202)	HAp	0.1566	—	0.1466	53.038	32.393	56.656
36.4	(206)	$\alpha -\mathrm{C}{\mathrm{a}}_3(\mathrm{P}{\mathrm{O}}_4){}_2$	0.1333	0.2476	0.1311	62.711	33.761	63.763
37.8	(436)	$\alpha -\mathrm{C}{\mathrm{a}}_3(\mathrm{P}{\mathrm{O}}_4){}_2$	0.18604	0.2930	0.27906	45.116	28.647	30.77
39.2	(212)	HAp	0.0864	—	0.0764	97.565	—	110.336
39.8	(310)	HAp	0.1503	—	0.1403	56.190	—	60.195
42.05	(311)	HAp	0.19	—	0.18	44.777	—	47.264
42.3	(302)	HAp	0.086	—	0.076	99.012	—	112.040
43.8	(113)	HAp	0.106	—	0.096	80.744	—	89.155
44.35	(400)	HAp	0.086	—	0.076	99.715	—	112.835
45.3	(203)	HAp	0.1047	—	0.0947	82.186	—	90.864
46.7	(222)	HAp	.	—	0.1444	.	—	59.900
48.1	(312)	HAp	0.0869	—	0.0769	100.070	—	113.083
49.45	(213)	HAp	0.21	—	0.2	41.632	—	43.713
50.5	(321)	HAp	0.1074	—	0.0974	81.751	—	90.144
51.3	(410)	HAp	0.20487	—	0.19487	42.999	—	45.206
52.1	(420)	HAp	0.21	—	0.2	42.091	—	44.195
53.15	(004)	HAp	0.16891	—	0.15897	52.550	—	55.855
54.45	(104)	HAp	0.0766	—	0.0666	116.589	—	134.094
55.9	(322)	HAp	0.07153	—	0.06153	125.678	—	146.104
58.05	(501)	HAp	0.092	—	0.082	98.715	—	110.753
average crystallite size						65.660	36.742	71.083

Table 4.

Effect of deposition temperatures on hardness of and grain size of HAp‐SC films

Sample	Vickers’ hardness, HV	Film thickness, μm	Grain size, nm
El	437	0.396	76
E2	314.55	3.138	82
E3	502	2.526	93

HV values were inversely proportional to the grain size [2], and decreased after the addition of SCs (E1). The decrease in hardness after adding the SC is due to the formation of nanocomposite, weak interfacial bonding of HAp and SC, and defects [34]. The H₂ O₂ addition has increased the film's thickness by reducing the H₂ bubbles on the substrate surface and permitting larger number of HAp particles to reach the substrate [35] (see Table 4). The addition of SC did not affect the deposition mechanism. The polymer was ionised and acquired positive charge during the application of electric field. Fig. 7 shows the SEM images of HAp film deposited on a titanium substrate before and after being immersed in SBF.

Fig. 7.

SEM images of HAp‐SC films deposited at various temperatures

(a) 25°C, (b) 45°C, (c) 70°C

It is clearly seen that no significant variation was noticed in the microstructure of HAp after 15 days (Fig. 7 b) and 30 days (Fig. 7 c) of immersion in SBF, indicating the good durability of these films for biomedical applications. The SEM images confirmed the formation of flower like‐lamellar structures of HAp coatings deposited on titanium substrates. Agglomeration of HAp particles was noticed for all coated films and the particles size were ranged between 50 and 400 nm. SEM images confirmed the formation of porous network structure, needed for the bone formation to provide rich sites for osseous tissue [10]. The grain size obtained from SEM is in good agreement with AFM results and indicates a good result for mimicking bone implant surface. The slight difference of grain size determined from XRD and AFM measurements originates from the fact that XRD is based on the size of defects‐free volume, whereas the AFM visualises directly the grains without considering the degree of structural imperfections [36]. The bioactivity test, based on dissolution rate with error bars, was investigated for the E3 sample and presented in Fig. 8 and found to be 0.85 nm h⁻¹. The antimicrobial activity of SC‐HAp composite nanostructure film (sample E3) was tested on E. coli bacteria and found to have good and higher antibacterial activity (Fig. 9) than the HAp ceramic nanostructure film, for all the concentrations studied (0.03–0.12 mg/ml). This is an encouraging result when compared to dissolution rate of HAp published elsewhere 1.71–3.42 nm h⁻¹ [37]. The effect of SC, SC‐HAp, surface morphology on cellular growth depends on concentration and biological activity of adsorbed proteins [38]. Inset of Fig. 9 a and Fig. 9 b is the inhibition zones induced by HAp and HAp‐SC samples against gram‐negative E. coli bacteria. The presence of an inhibition zone indicates a biocidal action of nanostructured HAp surface that may involve disrupting the bacterial membrane [39]. It is obvious that bioactivity strongly depends on concentration.

Fig. 8.

Dissolution time versus HAp film thickness for E3 sample

Fig. 9.

Histogram of antibacterial activity of

(a) HAp composite, (b) HAp‐SC composite

Inset is the inhibition zone

4 Conclusions

Electrophoresis deposition was found to be a successful technique to prepare crystalline HAp thin film coatings on titanium substrates. HAp‐SC nanocomposite films produced fairly hard and good (80–200 nm) grain size range, low dissolution rate but relatively low crystallinity. The HAp films revealed high crystallinity, high hardness, porous surface, and excellent (80–200 nm) grain size range. The in‐vitro test of sample E3, based on SEM and dissolution investigations, showed porous, stable, flower like‐lamellar structures of low dissolution rate films. Based on the results obtained, it is possible to use these films for orthopaedic applications and achieve a good control of the crystalline size, namely, the range ∼80–200 nm is usually required for bone implant applications. The antimicrobial activity of SC‐HAp composite nanostructure film was tested on E. coli bacteria and found to have good and high antibacterial activity.