Atorvastatin lipid nanocapsules and gold nanoparticles embedded in injectable thermo‐gelling hydrogel scaffold containing adipose tissue extracellular matrix for myocardial tissue regeneration

Abstract

This study aimed to prepare, optimise, and characterise the novel hybrid hydrogel scaffold containing atorvastatin lipid nanocapsules (LNCs) and gold nanoparticles (NPs) to improve cardiomyoblasts proliferation and regeneration of myocardium. A thermo‐responsive aminated guaran (AGG) hydrogel was prepared to encompass extracellular matrix (ECM) fetched from human adipose tissue. Emulsion phase‐inversion technique was used to obtain LNCs. Biocompatibility, tensile strength, conductivity, and proliferation of human myocardial cells of the optimised formulation were studied. The LNCs have a spherical shape, and the optimised formulation showed a mean particle size of 18.79 nm, the zeta potential of − 11.4 mV, drug loading of 99.99%, and release efficiency percent over 72 h was 18.73%. The injectable thermo‐sensitive hydrogel prepared using 1 w/v% of AGG, 35 w/w% of ECM, ∼0.5 mg/ml of gold NPs and atorvastatin loaded LNCs showed the best physical characteristics. The hybrid scaffold loaded with atorvastatin and gold NPs improved the proliferation of cardiomyoblasts more than sevenfold with enhanced cell attachment to the scaffold. The tensile strength and the conductivity of the scaffold were 300 kPa and 0.14 S/m, respectively. Injectable hybrid adipose tissue prepared by ECM and AGG hydrogel loaded with atorvastatin and gold NPs showed promising physical characteristics for myocardial tissue engineering.

1 Introduction

Changing lifestyles in advanced societies, including diminishing customary physical action, diabetes, hypertension, smoking, fast foods, etc. are common causes of heart attacks among people in these communities. According to the American Heart Association, cardiovascular ailments have the most elevated mortality hazard in the world. According to the same report, one out of every 34 Americans die from stroke or other cardiovascular issues [1, 2, 3, 4, 5, 6, 7].

Myocardial localised necrosis happens when a zone of the heart experiences perpetual harm, or demise because of the absence of satisfactory oxygen. Basic intricacies after myocardial dead tissue incorporate congestive heart failure, myocardial localised necrosis, arrhythmias, for example ventricular tachycardia, ventricular fibrillation and blocked coronary illness, pericarditis, irritation of the encompassing of the heart, aspiratory embolism, blood clusters in the veins, and cardiogenic stun. Until this time, a few treatments, including drug treatment and transplantation, have been utilised by analysts to help individuals with this disease. Drug‐based treatments for reducing heart rate and lowering blood pressure, although relatively effective, have helped to control the disease and avoid secondary heart attacks.

At present, the treatment choices after a stroke and consequent congestive heart failure are as yet constrained. However, medicate treatment enhances bloodstream to the heart muscle, yet its job has been restricted in heart failure because of an irrelevant impact on ventricular recovery and increments the yield of the clogged heart. In this manner, if MI advances to end arrange heart failure, the main fruitful treatment is heart transplantation. In spite of the deficiency in the donations and further medicinal clashes cause numerous constraints for this technique. Since substitution of heart muscle cells with connective tissue happened during myocardial damage, cell treatment could be a compelling procedure to remake harmed tissue. Notwithstanding, it has just had the option to accomplish a low degree of long haul stockpiling and survival of the cells [8, 9, 10, 11, 12, 13, 14, 15].

Tissue engineering has an important role in rebuilding and regenerating the heart muscle function and a significant effect in cells survival and storage, as well as cellular differentiation by using three‐dimensional designed scaffolds implanted in the body. A standout amongst the most significant highlights in regards to the determination of materials for the development of three‐dimensional platforms in treatment of myocardial dead tissue is that the chosen material can give a reasonable domain to improve the attachment of myocytes in the heart. This component is likewise fundamental for cell expansion and better ventricular capacity. Until now, numerous biomaterials, including characteristic polymers, for example fibrin, collagen, and manufactured polymers, like, polyurethanes and polyester, have been utilised in the heart tissue building. However, the occurrence of unfavourable immune responses by the body during planting and the lack of an integrated linkage of these materials with host tissue have placed restrictions on the use of these materials by Navaei et al. [16]. Therefore, the uses of natural components have paid more attention. For example, the uses of scaffolds made out of the extracellular matrix (ECM) of tissues have succeeded in tissue engineering and medical rehabilitation in the reconstruction of soft tissues due to the presence of materials and molecules needed to maintain cells and help them grow [16]. Besides, ECM parts acquired from fat tissues contain collagen, sinewy structures of elastin, neural and reticular strands, stromal vessels, and lymph hubs. Besides, the presence of peptides, cytokines, adiponectin, and other elements have a noteworthy job in procedures like vitality homeostasis, inflammatory reactions and cells development. The organisations of real ECM parts incorporate; corrosive pepsin, collagen, sulphated glycosaminoglycan, and laminin. The decellularised ECM displays attractive mechanical properties, and its arrangement has a generally safe and ease and is achievable through liposuction medical procedure [17].

One of the significant difficulties in infarcted heart revival is to give a proper electrical conductivity in the structures embedded in this tissue. The nearness of electro‐conductive nanomaterials in the structure of the platforms can build their electrical properties, encourage the dissemination of electrical sign, cell blending, and increment the mechanical parameters. The restriction of the utilisation of these materials in cross‐breed structures for tissue designing is their frail dissolvability and extreme cytotoxicity [18, 19, 20, 21, 22, 23]. Gold nanoparticles (NPs) are high electrical conductive materials with key advantages like simple assembling and modification strategies, creation in different scales and structures, biocompatibility, and low cytotoxicity. Concentrates on the utilisation of these NPs in permeable structures of collagen, alginate, polycaprolactone, and polyacrylates, have demonstrated the upgrade of both the capacity of myocytes and furthermore the differentiation of mesenchymal undeveloped cells into myocytes [16, 24].

Distinctive small molecule medications, proteins, and oligonucleotides of signalling pathways are very compelling in cells relocation, development, and differentiation in tissue engineering. It has been shown that statins reversed the acute inflammatory of the tissue‐engineered blood vessels by blocking the stimulation of TNF‐α [25]. Statins may influence the cardiovascular tissues by different mechanisms of their impact on the lipids profile, and it has been proposed that they influence the vascular and immunological events. A considerable helpful pleiotropic impact of statins happens because of their effects on endothelial capacity and diminished inflammatory reactions [25]. Atorvastatin has been appeared to have a critical effect on the help to mesenchymal undifferentiated cells survival in MI since responses of inflammation, oxidative and hypoxic stresses begin in such tissues after harm and the heart recovery is stimulated after nearby infusion. Numerous examinations have considered a particular portion of this medication powerful in the treatment of heart assaults, ventricular capacity alteration in myocardial damage, vascular dilatation, a decrease of fiery procedures in patients with heart failure, and lessening the rate of unexpected cardiovascular demise in patients with incessant heart failure. It has indicated the appropriate impact on vascular endothelial tissue and capillary walls in patients with ischaemic heart failure, and has a defensive impact in patients with myocardial localised necrosis [26, 27, 28, 29, 30].

Regarding these properties, the objectives of this project were to construct a biocompatible, thermo‐sensitive, injectable nanocomposite hydrogel based on a hybrid scaffold of guaran and adipose tissue for the reconstruction of myocardial cells. This scaffold was loaded with gold NPs to enhance its electro‐conductivity and also atorvastatin lipid nanocapsules (LNCs) to stimulate myocardial tissue regeneration and proliferation. To our knowledge, there is no report on the construction of such a nanocomposite in cardiac tissue engineering.

2 Materials and methods

2.1 Materials

The BASF Chemical Company (Ludwigshafen, Germany) kindly provided Labrafac and Solutol HS15. Labrafac is a mixture of caprylic and capric acid triglycerides and Solutol HS15 is the brand name of polyethylene glycol hydroxyl stearate. Atorvastatin was kindly donated by Amin Pharmaceutical Company (Iran). NaCl and lecithin were purchased from ACROS (Belgium). Gold NPs as a colloidal suspension with a particle size of 14 nm was supplied by US Research Nanomaterials Inc, (US). Guaran was from Aldrich (US), Cardiomyoblasts or H9c2(2–1) cell line was from Pasteur Institute (Iran), and foetal bovine serum (FBS) was obtained from GIBCO Laboratories (USA). (3‐(4,5‐dimethylthiazol‐2‐yl)‐5‐(3‐carboxymethoxyphenyl)‐2‐(4‐sulfophenyl)‐2H‐tetrazolium) or MTS reagent powder was from BioVision Inc. (India). Tissue culture plates and flasks were from Corning Life Sciences (Corning, NY, USA). Dulbecco's phosphate‐buffered saline (DPBS) was from Bioidea (Iran), Modified Eagles Medium (DMEM) from Gibco BRL (Grand Island, NY, USA) and fresh deionised water was utilised for setting up every single test arrangement.

2.2 Preparation of atorvastatin LNCs

A phase separation method was used for the production of atorvastatin LNCs. Lecithin, Labrafac and atorvastatin dissolved in 100 μl of methanol were mixed to form the oily phase, and the aqueous phase consisted of Solutol HS15 and NaCl. The oily phase was added to the aqueous one, and the obtained o/w emulsion was placed under three heating and cooling cycles between 85 and 60°C for phase inversion of the emulsion. In the cooling step of the last cycle, a specific volume of cold distilled water at 0°C was added to the LNCs dispersion to dilute it. The dispersion was stirred for 5 min at room temperature before characterisation [31].

Optimisation of the LNCs formulation was carried out by studying the effects of four processing variables including; surfactant percentage in the aqueous phase, the percentage of the oily phase to the total emulsion, the drug percentage and the ratio of cold water added to dilute the primary emulsion volume. The studied responses included; particle size, surface charge, drug entrapment, and release percentage and the time required to release 50% of the drug. A response surface study based on the hybrid design was used using Design Expert Software (Version 10, US). The 16 different formulations (Table 1) proposed by the software were manufactured and characterised. The code of each formulation consists of four words, each of which represents one of the variables tested and its value.

Table 1.

Physicochemical properties of the atorvastatin LNCs (n  = 3)

Formulations	Particle size, nm	Zeta potential, mV	Drug loading, %	Release efficiency, RE%	T50%, h
S46 L10 W3 D0.42	56.1 ± 2.2	−39.8	99.99 ± 0.020	83.27 ± 3.8	15.61 ± 0.13
S34.5 L20 W1.11 D0.09	17.7 ± 3.5	−29.1	99.99 ± 0.021	44.32 ± 2.6	50.41 ± 0.10
S34.5 L20 W2.25 D0.25	17.9 ± 3.0	−14.5	99.99 ± 0.012	11.87 ± 1.6	444.98 ± 0.41
S51.95 L20 W2.25 D0.09	15.4 ± 3.6	−34.9	99.99 ± 0.015	51.46 ± 3.1	67.05 ± 0.22
S46 L30 W3 D0.42	19.6 ± 2.9	−23.3	99.99 ± 0.022	39.03 ± 3.0	132.05 ± 0.32
S34.5 L35.18 W2.25 D0.09	65.9 ± 5.4	−12.8	99.99 ± 0.02	73.89 ± 2.9	17.54 ± 0.11
S23 L30 W1.5 D0.42	37.6 ± 4.1	−23.7	99.99 ± 0.01	32.69 ± 2.1	248.24 ± 0.25
S34.5 L20 W2.25 D0.65	47.8 ± 2.9	−22.4	99.99 ± 0.016	92.54 ± 3.4	7.22 ± 0.09
S23 L10 W3 D0.42	13.6 ± 2.3	−6.2	99.99 ± 0.01	38.35 ± 2.3	437.10 ± 0.33
S46 L10 W1.5 D0.42	20.0 ± 3.8	−34.5	99.99 ± 0.012	40.35 ± 1.2	222.66 ± 0.25
S23 L10 W1.5 D0.42	17.7 ± 4.4	−45.3	99.99 ± 0.011	37.50 ± 3.4	111.40 ± 0.17
a S34.5 L4.82 W2.25 D0.09	N/A	N/A	N/A	N/A	N/A
S46 L30 W1.5 D0.42	32.4 ± 4.7	−18.6	99.99 ± 0.013	41.71 ± 2.4	271.74 ± 0.31
S23 L30 W3 D0.42	38.5 ± 6.3	−31.1	99.99 ± 0.017	82.43 ± 4.1	14.25 ± 0.11
S17.05 L20 W2.25 D0.09	30.4 ± 4.6	−21.2	99.99 ± 0.025	74.04 ± 3.25	26.20 ± 0.19
S34.5 L20 W3.39 D0.09	18.8 ± 3.4	−11.4	99.99 ± 0.018	18.71 ± 1.1	543.82 ± 0.32

(Each digit in the formulation codes represents the quantity of the studied variable and S = Solutol%, L = Labrafac%, W = volume ratio of diluting aqueous phase to initial emulsion, and D = Drug%).

^a This formulation did not produce any emulsion due to the technical issue.

2.3 Estimation of LNCs particle size and surface charge

Particle size, polydispersity index, and zeta potential of LNCs were determined by the technique of PCS (photon correlation spectroscopy) using a Zetasizer instrument (Zetasizer 3600, Malvern Instrument Ltd., Worcestershire, UK) at 25°C. Each sample was diluted about 30 times with deionised water prior to analysis, and each measurement was done in triplicate.

2.4 Drug loading efficiency in LNCs

The strategy utilised for estimating the medication entrapment in the LNCs was an indirect technique. For this purpose, 1 ml of the LNCs suspension was centrifuged (Eppendorf 5430, Germany) at 13,000 rpm for 15 min in Amicon filter tubes (Cut‐off 10 kDa, Amicon Ultra, Ireland). The absorbance of the filtrate was determined spectrophotometrically (UV‐mini‐1240 CE‐Shimadzu, Japan) at 264 nm. The percentage of drug loading and loading efficiency were calculated by the following equations [32]:

$$\mathrm{Loading}\mathrm{efficiency}\mathrm{or}\mathrm{LE}\text{\%}\frac{\mathrm{Total}\mathrm{drug}-\mathrm{free}\mathrm{drug}\mathrm{in}\mathrm{filtrate}}{\mathrm{Total}\mathrm{drug}}$$

$$\mathrm{Loading}\mathrm{percent}\mathrm{or}L\text{\%}=\mathrm{Weight}\mathrm{of}\mathrm{entrapped}\mathrm{drug}\mathrm{within}\mathrm{LNCs}/\mathrm{total}\mathrm{weight}\mathrm{of}\mathrm{LNCs}\times 100$$

2.5 Drug release studies

The in vitro atorvastatin release from LNCs was assessed by the dialysis method in phosphate buffer solution (PBS) (pH 7.4). A suitable quantity of NLCs was set into a dialysis pack (12,000 Da) and dialysed against proper volumes of PBS to keep up the sink condition (at 37°C and 100 rpm). At foreordained time interims, 1 ml of the test medium was pulled back and was supplanted with an equivalent volume of PBS. The amount of atorvastatin in the filtrate test was calculated by the spectrophotometer at 241 nm [32]. The time required to discharge half of the medication (T _50%) and the release efficiency (RE₇₂ %) were resolved for every detailing. RE was determined utilising the accompanying condition

$$RE\text{\%}=\frac{{\int }_0^{t}y\cdot \mathrm{d}t}{y_{100}\cdot t}\times 100$$

In the aforementioned equation, y denotes the drug released percent at time t. Drug Release Efficiency percent is the area under the release curve, expressed as a percentage of the curve at maximum or 100% release (y ₁₀₀) over the same time period (here in 72 h).

2.6 Preparation of human ECM from adipose tissue

Fat tissue was provided from liposuction surgery of healthy female donors who had undergone liposuction at the Ummolbanin Plastic Surgery Clinic (Isfahan, Iran). The adipose tissue was acquired from subcutaneous belly fat with assent gotten from the ladies by the plastic surgeon specialists. An approved protocol by the Isfahan University of Medical Sciences was used to obtain the tissues by the investigators while all identifying information of the donors was removed. The Declaration of Helsinki guidelines were conducted for the study. 1000 ml of the tissue was transferred into 1500 ml of an iso‐osmolar solution of sodium chloride. Then 250 ml of it was washed for ten times by 750 ml of deionised water to get rid of any blood and other residual cells. After adding 100 ml of deionised water, the mixture was homogenised by an ultra‐homogeniser (T‐18 basic ULTRA‐TURAX, IKA; Werke, KG Staufen, Germany) at 12,000 rpm for 5 min and then centrifuged at 3000 rpm for another 5 min immediately. In the end, the adipocytes of the upper layer were thrown away, and after adding 100 ml of deionised water, the whole cycle above was repeated twice. The provided gel‐like suspension was used for the preparation of the hydrogel scaffold [17]. The amount of glycosaminoglycan in decellularised fat tissue tests was estimated utilising the Blyscan Sulphated Glycosaminoglycan examine kit (Biocolor Ltd., Carrickfergus, UK). First, 50 mg/ml of the dry weight acquired ECM was processed in the solution of 0.1 mg/ml proteinase K, 10 mM Tris cushion (pH 8.0), 50 mM NaCl and 1 mM EDTA for 24 h at 50°C. The glycosaminoglycan substance was estimated by the kit producer's convention. The absorbance was estimated in a 96‐well plate at 656 nm utilising a microplate reader (Biorad, USA). Each measure was performed in triplicate, and the outcomes were accounted for as mean ± standard deviation (SD).

2.7 Preparation of aminated guaran (AGG)

One gram of guaran (GG) was dissolved in 250 ml of deionised water and mixed for at least 30 min to produce a homogeneous mixture. After adding 25 ml of ethylenediamine to the GG mixture, it was mixed overnight at ambient. At that point, 50 ml of 5% solution of sodium borohydride was added to the blend and blended vivaciously for 3 h. The sediment was washed several times with acetone and then lyophilised to achieve a dry powder.

To determine the amine percent of the synthesised AGG powder, the amine groups were measured by trinitrobenzene sulfonic acid (TNBS) which reacts with the amine groups and forms a stable complex of trinitrophenol. To do it, 5 mg of AGG was dissolved in 1 ml of deionised water, and after that 1 ml of a 4 w/v% NaHCO₃ solution and 1 ml of newly made 0.5% v/v of TNBS were mixed to the blend and incubated at 60°C for 4 h. After that 1 ml of the above mixture was blended with 3 ml of 6 M HCl and kept at 40°C for another 1.5 h. The absorption of the produced trinitrophenol complex was achieved by spectroscopy method at λ _max  = 420 nm [33] and the percentage of amine groups in the AGG was calculated by the absorbance of GG as a baseline (background) at the same wavelength.

2.8 Preparation of temperature‐sensitive hydrogel scaffolds

Different ratios of adipose ECM lipid (35–50%) to AGG (0.5–2%) were mixed to get the best phase transition temperature of about 37°C. Then this ratio was kept constant and used in the preparation of different formulations of the scaffolds in which the total concentration of the hydrogel, the concentration of the optimised atorvastatin LNCs, and gold NPs were varied to optimise the scaffold formulation. The studied responses for this optimisation included; tensile strength, syringeability, and electrical conductivity.

2.9 Determination of electrical conductivity of the prepared scaffolds

The conductivity of scaffolds was measured at 25°C using a conductometer (Jenway, 3540, Germany). The average value of three determinations was obtained and reported as mean ± SD. To omit the conductivity of the sample due to the presence of water, the conductivity of the pure water was subtracted from the samples.

2.10 Determination of the tensile strength of the scaffolds

The rigidity of the frameworks was controlled by utilising the Santam® STM‐1 malleable analyser gadget (CT3‐4500, USA) with a testing rate of 10 mm/min at ambient temperature. The Young’ modulus of different formulations was calculated using the following equations:

$$\delta =\frac{F}{A}=E\times \epsilon$$

$$\epsilon =\frac{l_1-l_0}{l_0}$$

where F is the applied force per area of the gel (A) that each sample occupied. L ₀ and L ₁ are the primary and secondary distances between the upper and lower arms of the device before and after applying stretching and E is Young's modulus [34]. The Young's modulus for human myocardium is considered about 330–350 kPs [35].

2.11 Phase transition temperature of the scaffolds by stirring method and rheological measurements

Sol–gel transition temperature was measured by using a magnetic stirring bar method. 3 ml of AGG solution was placed in a 5 ml glass vial and transferred into a water bath. The temperature was increased slowly over the range of 5–40°C while stirring at 200 rpm using a 2 mm magnetic bar. The temperature, at which the magnet bar stopped motion, was considered as the gelation temperature [36].

Every single rheological estimation was performed with a rheometer (Anton Paar MCR 301, Austria). A 25 mm cone and plate apparatus was used for isothermal and non‐isothermal gelation studies. The cone angle was 1. After mixing the fluid samples were immediately moved into a Couette cell (time t  = 0), and determinations began at t  = 0 s. For time sweeping tests, a 1 Hz frequency and 2% stress–strain were used to measure storage moduli G ′ and loss moduli G ″, which were recorded as a function of time under constant temperature. For temperature sweeping tests, the advancement of G ′ and G ″ in the scope of −5 and 70°C were estimated at a warming rate of 2°C/min. Gelation was monitored during this temperature sweep test. The greatest slope of the G ′ was considered as the gelation temperature. Parallel plates (d  = 25 mm) were used in stress sweeping tests of hydrogels at 37°C.

2.12 Measuring the syringeability of the scaffolds

The injectability of the scaffold formulations was assessed by a panel test using 10 subjects who received a 23‐gauge needle‐syringe system filled with an aliquot of 2 ml of the hydrogels. The subjects were trained before starting the test and inquired about releasing 1 ml of the hydrogel from the needle and assessing the injectability in terms of the ease of the liquid formulation to flow from the syringe, utilising a subjective score from 1 to 4. In specific, the self‐assertive score for both parameters was characterised according to [37]:

• Score 1 = injection: not conceivable or exceptionally troublesome; flow: no stream or dropwise.

• Score 2 = injection: troublesome; flow: at first drop astute, after that continuous.

• Score 3 = injection: direct; stream: continuous.

• Score 4 = injection: simple; flow: continuous.

2.13 Atorvastatin release pattern from hybrid scaffolds containing LNCs

To evaluate the drug released from the LNCs loaded in the 1% AGG gel base containing 35% of ECM, the test was done in a dialysis bag sunk in PBS (pH 7.4). An appropriate amount of atorvastatin LNCs was added to 0.5 of the gel base containing ECM and gold NPs, and the whole formulation was transferred to a dialysis bag (cut‐off 12 kDa) and dialyzed at 37°C and 100 rpm, in 5 ml of PBS to keep the sink condition. At foreordained time interims, 0.5 ml of discharge medium was pulled back and was supplanted with an equal volume of the new medium. The concentration of atorvastatin within the filtrate test was measured by the spectrophotometer at 241 nm.

2.14 Determination of the enzymatic biodegradability of the scaffolds

11.9 mg (W ₀) of the scaffold was transferred to small glass vials containing 2.5 ml of PBS. The pH was adjusted on 2.5 with acetic acid, and 0.06 w/v% of egg white lysosome was added to the sample. The vials were incubated for 7 and 14 days at 37°C, and at each time point, the remaining sample was removed, washed with deionised water, dried and weighed again (W ₁). The percentage of degradation was obtained by the following equation [38]:

$$\mathrm{DR}\text{\%}=\left(W_0-W_1\right)/W_0\times 100$$

2.15 Study the proliferation of cardiomyoblast cells on scaffolds

The MTS assay was utilised to evaluate the cells proliferation on the scaffolds. The cells were kept at 37°C under 95% air and 5% CO₂. The sol hydrogels were presented into plates at 25°C, and the medium comprising of Dulbecco's Adjusted Eagle's medium (DMEM) supplemented with 1% antibiotics/antimycotics (last concentration: penicillin 100 units/ml, streptomycin 100 mg/ml and amphotericin B, 0.25 mg/ml), L‐glutamine 2 mM and 10% FBS was included. At that point, the plates were hatched at 37°C to form the hydrogels strong and 2 × 10⁴ myocardium cells were included per each well. The medium included to the hydrogel was utilised as a control. At 7 days, the number of surviving cells was calculated utilising the MTS test.

For this work, the MTS solution containing phenazine methosulphate (PMS) in DPBS was obtained by dissolving 2 mg/ml MTS powder in DPBS until a clear solution with golden‐yellow colour was prepared. At that point, the pH was balanced to 6.0–6.5 utilising 1 N HCl. Then the solution was filtered to be sterilised by passing it through a 0.2 μm filter while keeping the filtrate into a sterile, lightly protected container at 4°C for use but for long periods of storage it was kept at −20°C.

To perform the MTS assay at the moment needed, 100 µl of PMS solution (0.92 mg/ml PMS in DPBS) was added to 2.0 ml of MTS solution and 20 µl of the obtained solution was included to each well for each 100 µl of volume (last concentration of MTS will be 0.33 mg/ml and 0.25 mM of PMS). At that point, the plates were hatched 4 h at 37°C. The absorbance of the plates was read by ELIZA reader at 490 nm, and the percentage of live cells was calculated in comparison with the control [39]. The test was rehashed in three diverse days.

2.16 Adhesion of cardiomyoblast cells on scaffolds

The scaffolds were sterilised by UV radiation for 24 h and soaked with 50 µl of media. The cardiomyoblasts were seeded as much as 2 × 10⁴ cells/cm² on it. After 7 days, the cells adhered to the scaffolds, and the medium was removed. Then the scaffold was fixed for 30 min with 4% formaldehyde and washed for three times with PBS [40]. After drying at room temperature, the scaffold was gold coated and observed using scanning electron microscopy (SEM).

2.17 SEM studies of scaffolds

The morphologies of scaffolds were observed with a scanning electron microscope (Yimei, model 2300 C, Beijing Yimei Science Co., Ltd. China). The hydrogel samples were allowed to swell in water for 2 days. It was then quickly frozen using liquid nitrogen and then freeze‐dried for 1 day. The freeze‐dried samples were then coated with gold.

2.18 Statistical analysis

Design Expert (version 10, Ease Stat, USA) and SPSS software packages were utilised for statistical analysis. The results of at least three replicates were reported as mean ± SD. The one‐way analysis of variance (ANOVA) and the post‐hoc test of LSD were used for pairwise comparisons after ANOVA to correct for multiple testing. The significant level was chosen at p  < 0.05.

3 Results

3.1 Atorvastatin LNCs

To optimise the process condition of LNCs by phase inversion technique, several variables including; Labrafac, Solutol HS15 and drug concentration and also the volume ratio of the diluting aqueous phase to the initial emulsion weight were assessed. LNCs of atorvastatin were prepared by a hybrid design and their particle size, surface charge, drug loading, RE₇₂ %, and the T _50% were determined as responses. Analysis of the results was done using ANOVA, and the best model fitted to each response was suggested by Design Expert Software. The results of the studied responses of the atorvastatin LNCs are depicted in the table.

As the results of Table 1 show the mean particle size of atorvastatin loaded LNCs was between 17.‏7 and 65.9 nm.

Zeta potential of NPs determines their stability. Table 1 shows that the electric potential at the surface of LNCs and its magnitude ranged from −39.8 to −11.4 mv. Statistical analysis of the obtained data showed that not the drug content nor Labrafac concentration did not have any significant effect of the zeta potential of the LNCs (like S_34.5 L_35.18 W_2.25 D_0.09 and S_34.5 L₂₀ W_2.25 D_0.25). This may be related to the encapsulation of both these variables in the core of LNCs without any superficial effect on the surface charge of LNCs. The absolute value of zeta potential of LNCs decreased when the ratio of the diluting water to the primary emulsion weight was increased (like in S_34.5 L₂₀ W_1.11 D_0.09 and S_34.5 L₂₀ W_3.39 D_0.09), although not significantly (p  > 0.05). The effect of Solutol on the absolute amount of the zeta potential was not significant (p  > 0.05) and although in some formulations like (S₄₆ L₁₀ W_1.5 D_0.42 and S₂₃ L₁₀ W_1.5 D_0.42) or (S₂₃ L₃₀ W₃ D_0.42 and S₄₆ L₃₀ W₃ D_0.42) enhancement in Solutol amount decreased the absolute value of the zeta potential, but in other cases like (S₄₆ L₁₀ W₃ D_0.42 and S₂₃ L₁₀ W₃ D_0.42) it had the reverse effect which was not significant (p  > 0.05).

The results of Table 1 also show that all formulations have a perfect drug loading all above 99%.

Atorvastatin release through the LNCs showed to be in a sustained manner (Fig. 1). RE₇₂ % and T _50% were parameters employed for assessing the drug release behaviour of LNCs (Table 1). The greater the RE₇₂ %, and the shorter T _50%, the faster drug release rate. As shown in Table 1, the drug release efficiency and its related T _50% have a wide range of changes between 18.71 and 92.57%, and 7.22 and 543.82 h, respectively. This table shows that in both formulations which had the fastest (S_34.5 L₂₀ W_2.25 D_0.65) and the slowest release rate (S_34.5 L₂₀ W_2.25 D_0.25), the amounts of Solutol, Labrafac and diluting cold water were similar but included different contents of drug and the higher amounts of the drug caused faster drug release rate.

Fig. 1.

Atorvastatin release profiles from different formulations of LNCs (n = 3)

3.2 Optimisation of the LNCs formulation

The effect of different levels of the independent variables on the studied responses was analysed by the computer optimisation process using Design Expert Software, and the desirability function was calculated. The range of particle size was 13.6 ≤ Y ₁ ≤ 65.9 nm, with the target set at the smallest result. The constraints of zeta potential, EE, and T _50% were − 45.3 ≤ Y ₂ ≤ − 11.4 mV, 99.9%≤ Y _3, and 7.22≤ Y ₄  ≤ 543.82%, respectively. The optimum LNCs were considered those with the least particle size and the maximum of the absolute value of zeta potential, drug loading, and RE%. Considering these optimisation criteria, the optimised formulation proposed by the software was S_35.75 L_21.6 W₃ D_0.12 which contained 21.6% of Labrafac, 35.75% of Solutol HS15, 0.12% of atorvastatin and three‐fold of the diluting aqueous phase to the initial emulsion with 95% desirability. This formulation was manufactured, and all responses were measured for it. The particle size was 19 nm, zeta potential of − 11.5 mV, loading efficiency of about 100% and the T _50% of 550 h were obtained. These values were in quite an agreement with the predicted values by the software. Figs. 2 a and b show the SEM and AFM morphologies of the optimised LNCs, which are spherical and smooth.

Fig. 2.

Micrographs of optimised formulation of atorvastatin loaded LNC

(a) AFM, (b) SEM

3.3 AGG production

To determine the amination degree of the AGG, the amine groups were measured by TNBS addition, which reacts with the amine groups and forms the stable complex of trinitrophenol. The results showed the amine percentage of AGG to be 0.2%. This was obtained from the absorbance of AGG and an equal amount of GG by spectrophotometry method at the wavelength of 420 nm. These values were 1.6 and 0.4 for AGG and GG, respectively.

3.4 Preparation and characterisation of hybrid hydrogels of AGG‐lipid ECM

The importance of the sulphated glycosaminoglycan (GAGs) is in its sequestering capability of growth factors and subsequent, presenting them to cells [17]. Consequently, their presence within the matrix provides the chance for the bioactive molecules to be delivered in vitro and in vivo. Analysis of the GAG showed an average of 37.98 ± 4.28 µg sulphated GAG per each g of the dry adipose ECM, and 40.79 ± 3.55 µg sulphated GAG/g normal adipose tissue which shows no significant difference between the normal and dried adipose ECM. Fig. 3 shows the rheological characteristics of the optimised hydrogels of AGG‐lipid ECM and its phase temperature transition temperature, which is sharply at 35°C.

Fig. 3.

Viscosity changes of the optimised hybrid hydrogel of AGG‐lipid ECM with temperature and time

To each ml of the optimised hydrogel containing 1% of AGG and 35% of ECM, 0.12 μl of the LNCs dispersion and 0.5–0.75 ml of colloidal gold NPs (∼0.5 mg of gold NPs) was added to achieve the conductivity of at least 0.1–0.16 S/m which, is similar to the conductivity of the cardiac muscle [41].

The tensile test was repeated with the new composition of the scaffold, i.e. 1% of AGG, 35% of ECM, atorvastatin LNCs dispersion and 0.5 ml of colloidal gold NPs (containing 0.5 mg/ml). The results of this test indicated that incorporation of gold NPs and also the LNCs did not change the tensile strength of the scaffold and it was almost the same as Table 1 (i.e. 295 kPa for the gel base and 300 kPa for the gel containing atorvastatin LNCs and the gold NPs). Table 1 shows the best transition temperature, which was around 35°C and appropriate syringeability were obtained from the concentration of 1 w/w% of AGG and 35% of lipid ECM. This combination also showed the optimum of Young's modulus (Table 1).

3.5 Drug release profile from hybrid scaffolds loaded with atorvastatin LNCs

Fig. 4 shows the release profile of atorvastatin from hybrid scaffolds loaded with LNCs. As this figure indicates, increasing the AGG percentage in the scaffolds reduced the drug release rate from the hydrogels.

Fig. 4.

Atorvastatin release profile from different hybrid hydrogels of AGG‐lipid ECM (n = 3)

3.6 In vitro biodegradation of the scaffold

It was observed that the scaffold imbibed a lot of water at first and swelled rapidly. The AGG fraction of the scaffold degraded rapidly, but the ECM degradation was slower. At last, the results of the biodegradation rate study showed that the scaffold prepared by 1% of AGG and 35% of ECM had 24.3% weight loss after 7 days and 41.17% after 14 days.

3.7 Morphology and cardiomyoblast adhesion to the scaffold

Figs. 5 a and b show the designed scaffold had a porous structure, and the presence of lipid ECM caused cell adhesion.

Fig. 5.

Morphology and cardiomyoblast adhesion to the scaffold

(a) Porous structure of hybrid hydrogels of AGG‐lipid ECM loaded with LNCs of atorvastatin, (b) Cardiomyoblast cell adhesion to the scaffold

3.8 Proliferation of cardiomyoblast cells

Before starting the cell proliferation test on the scaffolds, it was necessary to find the IC₅₀ of atorvastatin on the cardiomyoblast cells to work with the drug concentrations, which were well less than its toxic dose. The IC₅₀ was calculated by exposing the cells to various concentrations of atorvastatin for 24 and 48 h. Fig. 6 shows the cardiomyoblasts cell viability in different concentrations of atorvastatin. As this figure indicates the IC₅₀ of atorvastatin after 24 and 48 h was calculated to be 620 and 371.0 µM, respectively.

Fig. 6.

Cardiomyoblast cells viability after 24 and 48 h exposure to different concentrations of atorvastatin to determine the IC₅₀ of the drug (n = 3) [lines show the significant difference (p < 0.05) between cell viability in different concentrations]

The scaffold loaded with LNCs was then used on the cardiomyoblast cells. However, the presence of Solutol in LNCs prevented us from using atorvastatin loaded LNCs even at lower concentrations of IC₅₀ of the drug. Therefore, the formulation was diluted to an appropriate concentration so that, the equivalent amount of 8 nM for atorvastatin was applied to the cell culture. The results of cell viability after 7 days of exposure of cells with this concentration are shown in Fig. 7. As this figure shows, the scaffold loaded by the LNCs caused twofold more proliferation than those treated with blank hydrogel.

Fig. 7.

Cardiomyoblast cells proliferation after 7 days exposure to the scaffold of hybrid hydrogels of AGG‐lipid ECM loaded with LNCs of atorvastatin and gold nanoparticles, drug‐free scaffold, free drug and blank scaffold (n = 3). The difference between all studied groups was significant (p < 0.05)

4 Discussion

The physicochemical properties of atorvastatin LNCs can influence on the LNCs physical stability, the drug release profile, and their biocompatibility. The statistical analysis of the obtained data by Design Expert Software showed that Labrafac and drug concentrations have a significant effect on the particle size of LNCs (p  < 0.05). In some cases, when these two variables were similar (like S₄₆ L₃₀ W₃ D_0.42 and S₂₃ L₃₀ W₃ D_0.42) (or S₄₆ L₃₀ W₃ D_0.42 and S₄₆ L₃₀ W_1.5 D_0.42) other variables like the Solutol and the ratio of the diluting water were effective on the particle size of the LNCs although non‐significant (p  > 0.05). By increasing in the Solutol HS15 content of NLCs, their particle size decreased (Table 1). This result is by the findings of previous studies [42]. Thanks to the stabilisation effect of Solutol HS15 at higher concentrations, which reduces the particle size, the LNCs will be protected from aggregation. Statistical analysis of data by the software showed that decreasing in the Labrafac concentration enhanced the particle size of the LNCs significantly (p  < 0.05) (like in formulations of S₂₃ L₁₀ W₃ D_0.42 and S₂₃ L₃₀ W₃ D_0.42). This may be due to the growing effect of the lipid phase on the size of LNCs. Heurtault et al. [43] also reported that increasing the Labrafac concentration caused grow of the LNCs particle size possibly due to the accumulation of oil in the centre of the particles and therefore, increments of LNCs particle size. Besides, the particle size of LNCs increased with the enhanced ratio of the diluting water to the primary emulsion weight (like in S_34.5 L₂₀ W_1.11 D_0.09 and S_34.5 L₂₀ W_3.39 D_0.09), but this change was not significant (p  > 0.05). The results of Table 1 also show that increasing the drug content of LNCs caused increments in the particle size significantly (p  < 0.05) (like S_34.5 L₂₀ W_2.25 D_0.65 and S_34.5 L₂₀ W_2.25 D_0.25).

None of the contemplated factors had a huge impact on the LNCs’ zeta potential (Table 1). Even though the impact of Solutol on zeta potential of LNCs had been unimportant yet increment of its content reduced the supreme estimation of zeta potential. Adsorption of more surfactant on the shell of LNCs and reduction of the concealing properties of polyethylene glycol of Solutol on steric obstruction may diminish the absolute value of zeta potential [42]. Su et al. [44] also reported that incorporation of PEG‐monostearate in NLCs decreased their surface charge. Although the zeta potential of NLCs was not affected significantly by the volume ratio of the diluting aqueous phase to the initial emulsion but its negative value enhanced with increments of this variable (like S₄₆ L₃₀ W₃ D_0.42 and S₄₆ L₃₀ W_1.5 D_0.42). Diminishing the last temperature of the emulsion with expanding the volume of cold water can prompt crystalline reorientation of lipid (lecithin) and could be a conceivable reason of augmentation in total estimations of the zeta potential of the LNCs. Electric dipoles are provided by the phospholipids of lecithin in water. Initially, if phosphoric gatherings of phospholipids were available in the external part of the shell, the nearness of phosphoric gatherings would incite negative zeta potential. Despite what might be expected, if terminal gatherings of phosphatidylcholine were available in the external part of LNCs shell, the presence of N‐functionalities brought about positive surface charge [42].

As the results of Table 1 indicate all formulated LNCs have a high drug loading efficiency (<99%) which is related to the physicochemical properties of atorvastatin with its rather low partition coefficient [log P (octanol/water)] of 6.36 and low solubility in water (0.835 mg/ml) [45]. This causes the rapid change of its location during the phase inversion of the emulsion and to be loaded successfully in LNCs.

The results of Table 1 show higher contents of Solutol HS15 decreased the drug release rate, although with no significant effect (p  > 0.05). Abdel‐Mottaleb et al. [46] and Safwat et al. [47] also concluded that the drug release rate from NLCs decreased at higher Solutol HS15 content but with no significant effect (p  > 0.05). The observations also showed that increasing in a volume ratio of the diluting aqueous phase to the initial emulsion increased RE₇₂ % (Table 1). Varshosaz et al. [42] also reported increment in the RE with increasing the cold water volume during the dilution step. The possible reason is that by increasing the volume of diluting phase, the viscosity of the dispersion decreased and consequently the resistance to diffusion of the drug also decreased, which resulted in an increase in drug release rate. The enhancement in Labrafac percent increased the RE₇₂ % too (like S₂₃ L₁₀ W₃ D_0.42 and S₂₃ L₃₀ W₃ D_0.42) but its effect was not significant (p  > 0.05) (Table 1).

After optimisation of the atorvastatin LNCs, the thermo‐responsive hybrid scaffold was designed. AGG was synthetised and characterised for its degree of amination by TNBS reagent. The results showed 75% amine groups substitution on GG. Jana et al. [48] also synthesised carboxymethyl guaran (CMGG) from the native guaran (GG) and afterward arranged aminated CMGG by conjugating it to ethylenediamine. At that point, fish scale collagen and aminated CMGG were cross‐linked by ceftazidime through non‐covalent ionic interaction and utilised it for wound mending. TNBS assay confirmed that 45% of amino groups of ethylenediamine interacted with CMGG.

Different percentages of AGG were also tested for their mechanical properties (Table 2), but the formulation containing 1 w/w% of AGG, which showed the best transition temperature showed only 200 kPa Young's modulus while the minimum required Young's modulus for human myocardia is 300–350 kPa [35]. Therefore, different concentrations of lipid ECM were added to 1% of AGG (which showed the best phase transition temperature at 35°C) and the syringeability of different ratios of AGG and lipid ECM used for the preparation of the injectable hydrogels were tested. The possible mechanism of gel formation between the synthetic AGG hydrogel and adipose ECM may be the production of layered porous structure with the fibrous texture of ECM and physical entanglement of these fibres with polymer chains in the microstructure of the hydrogels. It seems that the junction zone of AGG hydrogels is tightly formed by the incorporation of ECM fibres.

Table 2.

Phase transition temperature, syringeability and Young's modulus of different ratios of AGG and lipid ECM used for the preparation of hybrid hydrogels

Hydrogel type	AGG concentration, w/w%	Lipid ECM, w/w%	Phase transition temperature, °C	Syringeability (overall scorea)	F, N	L 0, mm	L 1, mm	Young's modulus, kPa
AGG0.5 ECM0	0.5	0	41	39	34.9	0.1	0.49	89.5
AGG1 ECM0	1	0	35	37	85.74	0.1	0.54	192
AGG2 ECM0	2	0	27	34	201.3	0.1	0.63	381
AGG1 ECM35	1	35	36	27	62.8	0.1	0.31	295
AGG1 ECM40	1	40	37	24	150.1	0.1	0.39	510

(^a The overall score is calculated by adding each person's score for each formulation).

As the results of Table 2 show the best transition temperature, which was around 35°C and appropriate syringeability obtained from the concentration of 1 w/w% of AGG and 35% of lipid ECM. This formulation was quite liquid at 4°C and at room temperature while gelation occurred at 35°C sharply (Fig. 3). This combination also showed the optimum of Young's modulus. It seems AGG and lipid ECM produce an interpenetrating polymer network that traps the NPs without any need for the toxic cross‐linking agents. The conceivable explanation behind the lot higher mechanical strength of the present engineered hydrogel than synthetic hydrogels might be because of its layered permeable and porous structure comprising of nanoorganised layers and filaments of ECM.

Murali et al. [33] likewise integrated an injectable hydrogel framework including biocompatible AGG, Fe₃ O₄ –ZnS core–shell NPs, and doxorubicin hydrochloride. They demonstrated that amination of guaran brought about an influx of water molecules and in this way produced the hydrogel without utilising harmful cross‐linking substances and hydrogel arrangement was seen at 37°C.

Tensile strength or rigidity is the capacity of a material to withstand a pulling (pliable) power. It is usually estimated in units of force per cross‐sectional area. The capacity to oppose breaking under pliable pressure is a standout amongst the most significant and broadly estimated properties of materials utilised in basic applications. The relative solidness or tractable modulus of the hydrogels is gotten from the stress–strain graph. Tseng et al. [49] studied the flexural behaviour of quasi‐laminated and single‐layer hydrogels of poly(ethylene glycol) diacrylate scaffold to be 118.25 ± 14.05 kPa to 223.40 ± 23.69 kPa. In another study by Yu et al. [50] the carbon nanotube (CNT)–collagen hydrogel with a CNTs loading level of 2 wt% exhibited Young's modulus of 1.85 ± 0.44 kPa which was supposed to be used in neonate rats.

Schmedlen et al. [51] found Young's modulus of photo cross‐linkable polyvinyl alcohol hydrogels from 125 ± 13 kPa for 15 wt% PVA to 838 ± 194 kPa for 30 wt% PVA and 820 ± 201 kPa for densely cross‐linked hydrogels. The optimised scaffold (AGG₁ ECM₃₅) showed about 45% of the drug after about 21 days while the NLCs themselves release about 70% of the drug at the same period (Fig. 4). This shows the sustained release of the drug in the scaffold, which can stimulate slowly the regeneration of the cardiac tissue. Also, the optimised scaffold had 24.3% weight loss after 7 days and 41.17% after 14 days. Tiwari et al. [52] used the required time for the disappearing of the hydrogels to assess the degradability. They found GG naturally degraded by endo‐1,4‐β‐mannanase and generated hydroxyl radicals. They studied the behaviour of GG–methacrylate in Tris–HCl buffer digestion solution (pH 8.0) at 37°C with 5% CO₂. They found the level of in vitro enzymatic biodegradation of the hydrogels diminished straightly with the gel content and the level of methacrylation of the particular macromonomers. For example, while 32% gel substance corrupted in only 44 h, test with 86% of gel substance did not totally break down, even following 35 days.

In tissue engineering, a polymeric framework with a well‐characterised design develops as a promising scaffold for cells attachment and aides their multiplication and differentiation into the ideal tissue or organ. A perfect model for the recovery should imitate clinical states of tissue damage, make a lenient microenvironment for the dissemination of supplements, gases and development factors and allow angiogenesis. These conditions are provided by a 3D support scaffold. The most important factors that influence cell adhesion to the scaffold are its porosity degree and surface properties that influence the composition and orientation of proteins that will be adsorbed onto material surfaces [53]. Fig. 5 shows the designed scaffold had microchannel‐like structure and well cell adhesion to the scaffold. Moreover, the scaffold loaded by the LNCs caused twofold more proliferation of cardiomyoblasts compared to those treated with blank hydrogel (Fig. 7).

5 Conclusion

The hybrid scaffold of AGG and lipid ECM containing LNCs of atorvastatin and gold NPs produced a suitable hydrogel scaffold with promising syringeability, tensile strength, and conductivity results. The LNCs containing atorvastatin could control drug release for more than one month. This hybrid hydrogel was produced just by the interpenetration of the polymer and ECM without using any toxic cross‐linking agent used in other hydrogels. The presence of atorvastatin LNCs significantly enhanced cardiomyoblasts proliferation, and the lipid ECM caused cells adhesion on the scaffold. Thereafter, this combination may have interesting implications for the development of injectable thermo‐sensitive hydrogels for cardiac tissue engineering as well as regeneration of other soft tissues.