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Article title: Fabrication and characterisation of biomimetic, electrospun gelatin fibre scaffolds for tunica media-equivalent, tissue engineered vascular graftsArticle reference: MSC6074Journal title: Materials Science & Engineering CCorresponding author: Dr. C. LekakouFirst author: Dr. Y. ElsayedAccepted manuscript (unedited version) available online: 30-DEC-2015DOI information: 10.1016/j.msec.2015.12.081

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FABRICATION AND CHARACTERISATION OF BIOMIMETIC, ELECTROSPUN GELATIN FIBRE SCAFFOLDS FOR TUNICA MEDIA-EQUIVALENT, TISSUE ENGINEERED VASCULAR GRAFTS

Y.Elsayed1, C.Lekakou1, F.Labeed2 and P.Tomlins3

1Advanced Materials Group, University of Surrey, Guildford, Surrey GU2 7XH, UK

2Centre of Biomedical Engineering,, University of Surrey, Guildford, Surrey GU2 7XH, UK

3National Physical Laboratory (NPL), Teddington, Middlesex, TW11 0LW, UK

Corresponding author: C.Lekakou: Email: C.Lekakou@surrey.ac.uk tel. 0044 (0)1483 689622

Abstract

It is increasingly recognised that biomimetic, natural polymers mimicking the extracellular matrix (ECM) have low thrombogenicity and functional motifs that regulate cell-matrix interactions, with these factors being critical for tissue engineered vascular grafts especially grafts of small diameter. Gelatin constitutes a low cost substitute of soluble collagen but gelatin scaffolds so far have shown generally low strength and suture retention strength. In this study, we have devised the fabrication of novel, electrospun, multilayer, gelatin fibre scaffolds, with controlled fibre layer orientation, and optimised gelatin crosslinking to achieve not only compliance equivalent to that of coronary artery but also for the first time strength of the wet tubular acellular scaffold (swollen with absorbed water) same as that of the tunica media of coronary artery in both circumferential and axial directions. Most importantly, for the first time for natural scaffolds and in particular gelatin, high suture retention strength was achieved in the range of 1.8-1.94 N for wet acellular scaffolds, same or better than that for fresh saphenous vein. The study presents the investigations to relate the electrospinning process parameters to the microstructural parameters of the scaffold, which are further related to the mechanical performance data of wet, crosslinked, electrospun scaffolds in both circumferential and axial tubular directions. The scaffolds exhibited excellent performance in human smooth muscle cell (SMC) proliferation, with SMCs seeded on the top surface adhering, elongating and aligning along the local fibres, migrating through the scaffold thickness and populating a transverse distance of 186 m and 240 m 9 days post-seeding for scaffolds of initial dry porosity of 74 and 83%, respectively.

KEYWORDS: Vascular grafts; Scaffold; Gelatin; Microstructure; Mechanical properties; Cell culture.

HIGHLIGHTS

Novel crosslinked electrospun gelatin scaffolds of specific fibre layer orientation.

These scaffolds have compliance equivalent to that of coronary artery;

and high suture retention strength, same or better than a fresh saphenous vein and

same as the tunica media of coronary artery in circumferential and axial directions.

Human smooth muscle cells had high proliferation and migration in the scaffolds.

1. Introduction

Cardiovascular disease is the cause of about 30% of deaths worldwide and in fact constitutes the leading global cause of death. Vascular grafts are needed to bypass occluded arteries, repair aneurisms [1], for haemodialysis access, and repair of cardiovascular defects in children [2]. Currently used vascular grafts fall into two categories: grafts made from synthetic materials (PTFE, Dacron®) [3,4] which have consistent mechanical properties but lack biocompatibility and are thrombogenic since they are perceived as foreign bodies by the organism triggering local platelet aggregation and thrombosis, which prevents them from being used to replace small diameter (below 6 mm) arterial vessels [5]. The second category comprises autologous grafts, such as the saphenous vein typically used as a vascular graft for the coronary artery (diameter less than 5 mm). Intensive research in tissue engineered grafts aims at manufacturing optimised scaffolds from commercially available synthetic or natural materials in which to culture own cells from the patient, so that the graft can be easily accepted and speedily integrated in the organism without any adverse effects [6].

As with synthetic vascular grafts, porous scaffolds from synthetic materials (for example: poly(l-lactic acid) (PLLA)/poly(l-lactide-co--caprolactone) (PLCL) [7], PLLA/PLA(poly-lactic acid) [8], polyethylene glycol (PEG) [9], polyurethane-graphene oxide [10]) have consistent mechanical properties and may be tailored to provide the required level of structural support and strength. Research has shown that compliance mismatch between the graft and the host vessel is another factor causing intimal hyperplasia, so efforts are made to select suitable materials so that the vascular grafts have compliance and burst pressure similar to those of native vessels [11]. For this reason, elastomeric materials have been used in the fabrication of vascular grafts such as electrospun polyurethane grafts [11,12]. However, grafts from synthetic materials still trigger platelet coagulation and, hence, are thrombogenic.

Natural polymers such as collagen, gelatin or elastin contain many functional motifs which can modulate cell-specific adhesion while they deter platelet adhesion and coagulation, as they mimic the extracellular matrix (ECM) of the arteries. Hence, many studies involve coating of the surfaces or fibres of the synthetic scaffold with natural polymers, functionalisation or grafting of the synthetic polymer surface with natural polymers such as collagen (electrospun collagen-blended poly(L-lactic acid)-co-poly(∈-caprolactone) [13]), gelatin (for example polycarbonate urethane-g-gelatin or PEGMA-g-gelatin [14-16], electrospun gelatin-blended polycaprolactone scaffolds [17], gelatin-poly(lactide-co-glycolide)(PLGA) scaffolds [18]) or functionalization with special groups (compliant poly(carbonate-urea)urethane graft (MyoLink™) with arginine-glycine-aspartate (RGD) or/and heparin (Hep) [19]; electrospun poly-e-caprolactone scaffolds functionalised with lysine groups to which heparin was anchored to improve adhesion of endothelial cells [20]).

Tissue engineered vascular grafts aim at mimicking arteries. Muscular arteries are of 3 to 10 mm diameter [21], with tunica media being their thickest layer consisting of collagen fibres with embedded elastin plates and up to 40 layers of smooth muscle cells (SMCs) which provide the contractile and relaxation response and the muscle tone [22-23]. Generally, scaffolds fully based on natural materials have demonstrated weak mechanical properties and suture strength [24]. Porous, crosslinked collagen-elastic-glycosaminoglycan scaffolds have been fabricated by gel casting and freeze-drying which reached a maximum tensile strength of 0.68 MPa [25] and a maximum yield stress of 30 kPa [26]; myoblasts were seeded [25] which after 14 days formed tubular structures generally aligned in the direction of fibres, where fibres were present.

Gelatin can be considered as a low cost alternative source of soluble collagen. In our group, we have pioneered research in gelatin scaffolds, starting with gelatin gels [27-28] which although they allowed the ingression of small, rat smooth muscle cells did not allow the migration of osteoblasts. For this reason, we have proceeded to the fabrication of porous gelatin scaffolds using electrospinning [29] that are expected to allow for easy migration of cells through the scaffold and rapid population of the scaffold with cells. Another reason is to allow for infiltration of the culture medium in vitro and the blood in vivo through the pores, providing nutrients to the cells, as it has been found that implantation of non-vascularised, multi-layer cell tissues more than 100 m thick results in necrosis [30]. An obvious hypothesis would be that the scaffold pore size should not be smaller than the dimensions of cells in suspension to allow for cell migration [31] but this is an over-simplification even for specifying a minimum pore size, as cells may crawl and deform to migrate through smaller pore sizes. Another issue is that the scaffold pores surface area is inversely proportional to the pore size, with a large surface area theoretically allowing for more cells to be attached and even migrate through the scaffold by crawling on the pore surface across and through the scaffold [32]. However, cell growth depends very much on access to nutrients which is favoured by large pore sizes [32] and pore fraction throughout the whole duration of tissue engineering until vascularisation in vivo: for example, initial high pore fraction and large pores may favour fast cell growth and migration with cells covering and blocking the pores at later stages, which means they would prevent transport of nutrients and, hence, this would lead to rapid cell death in the end. Furthermore, the mechanical properties of the scaffold are also affected by the pore fraction and size. As a result, a comprehensive investigation of the effect of pore fraction and pore size of electrospun gelatin scaffolds has been undertaken in this study in relation to smooth muscle cells in order to fabricate and test biomimetic tunica media-like vascular grafts.

Fibre and cell orientation are very important in arteries, as they affect the mechanical properties, and local contraction and dilation of the vascular vessels. In general, it has been found that SMCs have an average length of 50-100 m and a thickness of only a few micrometers, and circumferential orientation to achieve the pressures required to drive blood flow, the artery’s vasoactivity and structural integrity [33]. Investigations of collagen fibre orientation within arterial walls has reported different results: circumferential [34]. Ghazanfari et al [37] found that the fibres in porcine carotid arteries are predominantly circumferential in the outer adventitia and the tunica media and in axial orientation near the lumen. Holzapfel et al [39] measured the mechanical properties of the different layers of the coronary artery and found similar ultimate tensile strength values in the circumferential and axial direction (0.45 and 0.42 MPa, respectively, for the tunica media). This combined with the popular design of ±45o fibre composite materials for pressure vessels, led us to a targeted ±45o fibre orientation for the electrospun fibrous scaffolds in this study., helical [35], and even axial [36] orientations. While the differences may be due to the fact that different types of arteries were investigated, the clear understanding is that collagen fibre orientation is not only different between the tunica media and adventitia, but also between the different regions of the tunica media [37,38]

The current study is proposing novel, low cost, electrospun tubular, multilayer, gelatin fibre scaffolds of controlled fibre orientation, optimised in terms of mechanical properties and most importantly suture strength, as well as smooth muscle cell adhesion and migration for biomimetic, tissue engineered, tunica media-like constructs to be used as vascular grafts. Optinised fibre size and pore fraction and size, as well as fibre orientation and gelatin crosslinking has led us to achieve innovative tissue engineered constructs, based on a low cost natural scaffold material, but with equivalent mechanical properties to those of natural artery (in particular coronary artery) and a suture strength at the level required for vascular grafts, which is a first for gelatin scaffolds.

2. Materials and methods

2.1 Scaffold fabrication

Solutions of 7-15 %w/v pork gelatin type A (Sigma Aldrich) in 2, 2, 2-Trifluoroethanol (TFE) (Sigma Aldrich) were prepared for electrospinning. The electrospinning rig used is illustrated in Figure 1. A syringe pump was used to control the feed rate of the gelatin solution, coupled with a standard 5 ml plastic syringe, and polytetrafluoroethylene  (PTFE) tubing and an 18 gauge blunted stainless steel needle connected to a positively charged high voltage power supply. The syringe was held at a 45o angle to a horizontal rotating cylindrical aluminium mandrel connected to a stepper motor controlling the rotation speed. Below the collector, a stainless steel knife-edged plate, at 45o angle in the x-z plane, in line with the needle, was connected to a negatively charged high voltage power supply. Initial experiments with the syringe at 90o angle with the rotating scaffold led to fibres becoming oriented in the circumferential direction of the rotating tubular collector. Angling the syringe and electric field at 45o with respect to the rotating drum axis led to fibres arriving in the axial direction of the drum and, while being taken up by the rotating drum the fibres were rotated to helical orientation in the rotation direction (approximately 45 o), also influenced by the rotation speed of the drum collector. To produce scaffolds with multi-layer ±45 degree fibre alignment (alternating), fibres were spun for a duration of 4 minutes, before the electrospinning was stopped and the rotator collector was flipped by 180 degrees. The spinning was then resumed for 4 minutes before the steps were repeated up to16 layers in total for scaffolds to be used for cell seeding.

Parametric studies of the electrospinning process were carried out in which the following process parameters were changed: the applied voltage, the feed rate and the concentration of the gelatin solution. The input parameters were varied in turn, while the rest of the parameters were fixed. The control setup was as follows: feed of 10% w/v gelatin solution at a constant flow rate of 2.5 ml/hr, a rotating mandrel revolving at a speed of 95 rpm, collecting the electrospun fibres 25 cm from the 18 inch gauge needle, with +22.5 kV dc applied voltage. The knife edged electrode with a negative voltage of 3 kV was situated 5 cm below the mandrel, in direct alignment with the needle depositing the electrospun fibres. The so produced scaffolds were dried overnight in a vacuum desiccator. The scaffolds were then crosslinked in glutaraldehyde vapour in another vacuum desiccator containing a beaker with a solution of 10 %w/v glutaraldehyde in water for two hours.

2.2 Physical characterisation of the electrospun scaffolds

Extensive image analysis was carried out of scanning electron microscopy (SEM) micrographs of the electrospun scaffolds to measure fibre diameter and orientation. Scaffolds were fixed on an aluminium stub before being sputter coated with 20 nm particles of 60%-40% gold-palladium mix (EEMS 575X). The scaffolds were then viewed under a Hitachi S3200 scanning electron microscope (SEM). To analyse the SEM micrographs, horizontal lines perpendicular to the fibres were drawn at the corners of 5 images, selected at random from across each scaffold. Ten lines that perpendicularly crossed a fibre diameter were measured, giving a total of n = 50 fibres measured in each image.

Porosity was determined using a gravimetric approach. The scaffold average thickness, H, and area, A, were measured under a light microscope. The scaffolds were then weighed and their mass, m, recorded. The apparent scaffold density, app, was then determined according to the relation:

(1)

Using the bulk density of gelatin (1340 kg m-3), the scaffold porosity, , was calculated according to the relation:

(2)

Pore size is a difficult feature to quantify using visual techniques, as what area constitutes a pore can be different between observers, thus an estimate of the pore radius, rp, was established using a relation for random fibre meshes [40] where rp is the pore radius, ε is the scaffold porosity, and df is the average fibre diameter.

(3)

Experiments of scaffold swelling in water and scaffold degradation in water as a function of time were conducted for the crosslinked gelatin scaffolds immersed in de-ionised water at 37 oC. The mass of the scaffold was measured as a function of time and was plotted as the ratio of the mass of the wet scaffold versus the mass of the dry scaffold as a function of time up to 3 weeks. The wet scaffolds were also imaged under the light microscope and an image analysis was conducted to determine their microstructural parameters. The permeability of wetted scaffolds (immersed and swollen in de-ionised water at 37 oC for 48 hours) was measured using transverse flow experiments. The wetted scaffold was clamped on a grid, below a burette. Corn oil was chosen as a viscous Newtonian fluid flowing transversely through the porous scaffold under the pressure difference, P, of the hydraulic head, h, of the corn oil in the burette over the scaffold: P = oil g h, where oil is the density of the corn oil (measured as 950.6 kg/m3) and g is the gravity. The permeability, k, of the scaffold was determined using Darcy’s law:

(4)

where U is the velocity of the corn oil measured from the change of the oil level in the burette versus time, and H are the porosity and thickness of scaffold, respectively, and is the viscosity of corn oil (measured in a Brookfield viscometer as = 65 mPa.s).

2.3 Mechanical and suture retention testing of the scaffolds

Wetted 50x10 mm2 rectangular, electrospun and crosslinked scaffolds were pulled until failure using a uniaxial tensile testing machine (Instron 6025) at crosshead speed of 5 mm/min. A pipette filled with deionised water at 37οC was allowed to drip continuously and slowly at the corner of the fixed scaffold above the gripper to maintain a wet sample throughout the experiment, in order to simulate the hydrated environment at the temperature of the human body, allowing for comparison between the mechanical properties of fabricated scaffolds and natural arteries.

Suture retention testing was also carried out, where one end of the electrospun, crosslinked and wetted tubular construct was clamped to the stage of the Instron tensile testing machine, and the other end was knotted at four quadrants using polypropylene suture (5-0, USA) 3 mm from the edge of the scaffold. The sutures were then pulled at a speed of 5 mm/min until all the sutures broke from the wet scaffolds.

2.4 Cell culture and biological characterisation of the scaffolds

Human umbilical vein smooth muscle cells (ATCC, LGC standards, UK) were cultured in medium-231 and complemented with 25 ml smooth muscle growth supplement (Cascade, USA). The final concentrations of the components in the supplemented medium were 7.5% v⁄v foetal bovine serum, 2 ng⁄ml human basic fibroblast growth factor, 0.5 ng⁄ml human epidermal growth factor, 5 ng⁄ml heparin, 5 µg⁄ml insulin, and 0.2 µg⁄ml BSA. 5ml of penicillin-streptomycin (10000 units of penicillin and 10000 µg of streptomycin per ml) were added to the 500 ml of medium, to form the final working culture medium.

After electrospinning and crosslinking of scaffolds, square scaffolds of 15 mm sides were punched from the porous fibrous mat and secured in 6-well plates. The scaffolds (300-400µm thick) were immersed in warm 25% w/v glycine solution for a period of 3 hours to remove any unreacted glutaraldehyde groups. The scaffolds were then sterilised in 6% v/v hydrogen peroxide overnight, followed by three washes in sterile phosphate buffered saline solution and incubated in culture medium containing 5% penicillin-streptomycin for a period of 24 hours. 200 µl of the cell suspension of 2.5x105 cells/ml were uniformly seeded onto one side, only the top surface of each scaffold, and incubated for 3 hours in 6-well plates. The scaffolds were then transferred to the bottom of 6-well plates where only one surface was in direct contact with the culture medium.

For histo-morphology the scaffolds were collected on days 1, 3, and 9 after seeding. The scaffolds were then impregnated and moulded in epoxy resin for sectioning. Thin slices of thickness 3-6 µm of the scaffold were cut in the transverse plane of the scaffold using a microtome. The resin was softened and cleared in xylene, then rehydrated. To prepare for SEM viewing the slices were coated and viewed as previously described in section 2.2.

Cellular proliferation throughout the whole scaffold was measured using MTS assay at 12 hr, 3, 6, and 9 days after seeding. The CellTiter 96 Aqueous One solution cell proliferation assay (MTS) (Promega) was thawed in a water bath for 10 minutes. The two reagents (1 ml of PMS solution to 20 ml of MTS solution) were mixed in a laminar flow hood following standard aseptic techniques. The scaffold was washed in PBS before being incubated in 500 μl of culture medium mixed with 100μl of the working MTS solution and incubated at 37 ºC for four hours. The 600μl of the solution was transferred into 4 or 5 wells (120 μl in each well) of a 96 well plate and the absorbance of light at 490 nm was measured using an ELx800 96 well micro-plate reader. Three measurements were taken for the same plate and the average absorbance for each well was calculated. 5 standard wells holding a combination of the culture medium and the working MTS assay, which had undergone the same steps but without a scaffold, were used to calculate the average absorbance under 490 nm and the value was subtracted from the measured absorbance for the acellular scaffolds to obtain the corrected average absorbance.

2.5 Statistical analysis

Results in tables and the data points in graphs represent average values of measurements of different samples or sample areas. Results in tables are expressed as average values ±standard deviation. Data points in graphs are presented as the average values with error bars representing the range of obtained measurements from all measured samples for the corresponding data point in the graph.

3. Results

The alignment of the opposite polarity electric field at 45 degrees to the revolving mandrel led to the fabrication of tubular scaffolds generally with 45o fibre orientation, even after the scaffolds were chemically crosslinked as seen in Figure 2.

The physical properties (fibre diameter, pore size and porosity) of the electrospun dry scaffolds, fabricated using different values of the electrospinning process parameters, are summarised in Figure 3: the average fibre diameter was measured from SEM micrographs, the average porosity was measured using the gravimetric method and the average pore radius was determined using equation (3). In general, all scaffolds had electrospun microfibres of less than 10 m diameter. Figure 3(a) shows that as the gelatin concentration of the feed solution was raised from 7.5 to 15 % w/v, the fibre diameter increased in approximately linear fashion from 1.5 to 3.8 µm. This brought a very small change of scaffold porosity (within the experimental error) while the calculated pore radius increased from 3.6 m to 10.3 m, respectively. Figure 3(b) shows that as the feed flow rate was raised from 0.5 to 7.5 ml/hr, the fibre diameter displays a gradual increase from 1.5 to 2.5 µm. Similarly, the porosity exhibits an initial fast increase from 59 to 76.7% as the flow rate was raised from 1 to 2.5 ml/hr and a slower increase up to 80% thereafter to a flow rate of 7.5 ml/hr. Consequently, the calculated pore radius varies from 4.2 to 17.2 m for a corresponding increase of feed flow rate from 1 to 7.5 ml/hr. The effect of the applied voltage was more drastic in comparison, as shown in Figure 3(c). While operating at high gelatin concentration and flow rate of feed solution, 15 %w/v and 7.5 ml/hr respectively, as the positive applied voltage was raised gradually from 15 to 30 kV, the average fibre diameter increases approximately linearly from 2.6 to 8.1 µm. This trend seems to be in direct contrast to the accepted norm of an inverse relationship between the applied voltage and the resultant fibre diameter in electrospinning [41]. On the other hand, other works have concluded the opposite effect, where an increase of applied voltage results in an increase in fibre diameter [42] and Pham et al [43] considered these conflicting results “ambiguous”. In our opinion the contrasting relationships of fibre diameter against applied voltage are related to the polarity of gelatin in the feed solution and the direction of the applied electric field. By connecting the positive voltage pole to the feed solution nozzle, it seems that an attractive force is generated that increases the deposited mass of each fibre. Tong et al [44] found that electrospinning of poly(3-hydroxybutyrate-co-3-hydroxyvalerate) (PHBV) fibres showed an increase in fibre diameter with an increase in the positively charged applied voltage, and a decrease in diameter with an increase of the negatively charged bias voltage. They argued that the existence of negatively charged molecules in their electrospinning solution would partly neutralise the positive charge, while superimposing on the negatively charged jet. Furthermore, Figure 3(c) shows that a rise of the applied voltage from 15 to 30 kV has brought a corresponding scaffold porosity increase from 63 to 76.3% and a large, corresponding, pore size (radius) increase from 7.9 to 44.7 m.

For the set up of the electrospinning rig presented in Figure 1, the following findings are reported regarding the effect of the collector rotation speed: when there was no rotation of the collector, the fibres were aligned parallel to the longitudinal axis of the tubular collector, due to the electric field imposed by the 45 o knife-edge plate which directed fibre deposition; as the collector rotation speed was increased the fibres were oriented obliquely, reaching an average 45 o fibre orientation at a collector rotation speed of 95 rpm; raising the collector speed to 200 rpm modified the average fibre alignment towards the circumferential direction of the collector. Hence, a standard collector rotation speed of 95 rpm was maintained. Fibre alignment was also affected by changing the distance between the feed nozzle and the collector. Figure 4 shows that as this distance was increased from 15 to 35 cm, the fibre diameter decreases from 7 to 4 m whereas the fibre fraction at 45 o alignment increases from 55 to 85%.

Crosslinking of gelatin scaffolds brought an average increase of 55% in fibre diameter (Figure 2(b) and (c)) which may be attributed to the shrinking of gelatin material and also according to Sisson et al [45] some fusion of fibres. Subsequent immersion of the crosslinked, electrospun scaffolds in de-ionised water at 37 oC, brought further swelling of the fibres to an average mass increase of 100% within the first day as shown in Figure 5. Figure 5 demonstrates that the so wetted scaffolds still immersed in de-ionised water maintained this swelling (mass ratio of wet to dry scaffold around 2) without any degradation or mass loss for a tested duration of 10 days, and some degradation thereafter reaching a mass ratio of wet to dry scaffold of 1.6 after being immersed in the water for 3 weeks. Figure 6(a) presents a summary of the microstructures and microstructural parameters of the different crosslinked, electrospun gelatin fibre scaffolds of different porosity and fibre diameter. As with the uncrosslinked, electrospun scaffolds, the general trend is a parabolic increase of porosity with fibre diameter. Furthermore, a narrow fibre orientation distribution around 45 o is observed in scaffolds of an average fibre diameter of 1 m, whereas a wider fibre orientation distribution is observed in scaffolds of large fibre diameter. Given that all microstructural parameters of scaffolds (crosslinked or uncroslinked) were measured from SEM images and only dry scaffolds were examined under SEM, the parameters of wet scaffolds have been calculated as follows:

df,wet = 2 df,dry

(5)

wet = dry (df,dry/df,wet)2

(6)

and the pore radius is calculated according to equation (3) using the data of wet fibre diameter and porosity.

Permeability tests were carried out for several wet (swollen with water), crosslinked, electrospun scaffolds of different fibre diameter and porosity. Figure 6(b) presents a Carman-Kozeny plot of this data (wet values for df and ) according to the Carman-Kozeny equation [46,47]:

(7)

where Ko is the Kozeny constant determined from the best fit in Figure 6(b) as Ko = 0.032 for the tested scaffolds.

First of all, mechanical testing of wet, crosslinked, electrospun scaffolds was carried out in both parallel and transverse to fibre directions, and as is shown in Figure 6(c) it is clear that the scaffold is more rigid and stronger in the fibre direction but it reaches higher strain in the direction transverse to the fibres. In particular, scaffolds exhibit two-fold higher ultimate tensile strength and four times higher Young’s modulus parallel to the fibre direction compared to the direction perpendicular to the fibres.

The electrospinning set up was designed to wind up the fibres at a 45 degree angle to achieve maximum strength in both circumferential and axial directions in order for the scaffolds to withstand the peristaltic blood flow. However, depending on the size of fibre diameter and porosity (see Figure 6(a)), a very narrow or wide distribution of fibre orientation may be achieved. Scaffold specimens were cut from the tubular scaffold and were subjected to tensile testing both in the axial and circumferential direction of the original tubular scaffold. Figure 7 presents the effects of the porosity, pore size and fibre diameter on the Young’s modulus and ultimate tensile strength of the tested scaffolds. It can be seen that the Young’s modulus is generally the same in the axial and circumferential directions but the tensile strength may become twice as high in the circumferential direction compared to the axial direction for scaffolds of low porosity, small pore size and small fibre diameter. As the scaffold porosity increases from 62 to 80%, Young’s modulus decreases from 1.2 to 0.2 MPa, respectively, and the ultimate tensile strength decreases from 2.1 to 0.1 MPa, respectively. Similar trends are followed at increasing pore size and fibre diameter.

The results of the suture retention strength for three types of wet, crosslinked, electrospun scaffolds are presented in Table 1. While the results show an increased suture strength with smaller, more packed fibres, the differences are not regarded as significant to affect the choice of the optimum scaffold, as other properties could be affected in a more pronounced manner.

Table 1: Test data of suture retention strength of different types of wet, crosslinked, electrospun scaffolds (5 repeat experiments for each type)

Measured parameters and properties

Scaffold1

Scaffold2

Scaffold3

Average fibre diameter (μm)

Dry

6.8 ± 0.7

4.7 ± 0.5

2.3 ± 0.2

Wet

13.6

9.4

4.6

Porosity (%)

Dry

75 ± 4

69 ± 3

65 ± 3

Wet

18.8

17.3

16.3

Suture retention strength (N)

Wet

1.81 ± 0.04

1.9 ± 0.2

1.94 ± 0.09

In order to optimise the mechanical properties of the scaffold, bidirectional ±45 o multilayer scaffolds were created with alternating layers of +45 and -45 o fibre orientation electrospun sequentially. Such scaffolds were crosslinked and seeded with human umbilical vein smooth muscle cells (SMCs) according to the procedures described in section 3. Figure 8 presents SEM micrographs of both seeded (top) surface and unseeded (bottom surface) of scaffold 3 days, 6 days and 9 days after seeding (with the cell-seeded scaffold in the culture medium at 37 oC). The ±45 o fibre orientation of the multilayer scaffold is clearly visible. Three days post-seeding, Figure 8(a) shows only a fraction of the seeded (top) scaffold surface been covered with SMCs, which generally seem well adhered onto the scaffold fibres, and also elongated and oriented along the 45 o fibre direction of the top layer as each SMC seems to be elongated and adhered along the length of a fibre. Hence, in general SMCs seem to follow the local fibre orientation. Six days post-seeding, Figure 8(b) shows growth of SMCs on the top surface and, additionally, Figure 8(c) of the bottom, unseeded surface demonstrates that the SMCs have migrated from the top surface through the porous scaffold and have reached the bottom surface on which they continue to grow in elongated structures, adhered on the scaffold fibres and oriented in the local fibre direction. Nine days post-seeding, Figure 8(d) shows that the SMCs have populated well the scaffold and a large fraction of the bottom, unseeded surface of the scaffold (although there are still open pores through which the culture medium still impregnates and provides nutrients to the cells). Furthermore, tissue has started forming locally through local cell integration at the bottom, unseeded surface.

The next set of results focus on the migration of SMCs though the scaffold for two types of scaffolds, for which details are given in the caption of Figure 9. The graph of cell migration data derived from SEM images in this study agrees with similar histology data derived from the cross-sections of stained cell seeded scaffolds in our parallel study reported in [48]. Scaffold-S2 has greater porosity and pore size (given its larger fibre diameter, as dictated by equation (3), than scaffold-S1, which seems to allow for a higher rate of cell migration through the scaffold (Figure 9(b)) compared to scaffold-S1 which has lower porosity and smaller pore size. MTS assay results in Figure 9(c) demonstrate overall cell proliferation in the whole scaffold: both scaffolds seem to have a similar number of adhered cells 12 hr post-seeding; scaffold-S1 exhibits a faster initial growth from day 0.5 to 3 post-seeding, but due to its lower porosity and pore size cell growth slows down after day 3, presumably due to poor infiltration of the culture medium through the scaffold. Scaffold-S2, which has higher porosity and larger pores, seems to provide good infiltration to the culture medium and cell migration throughout the whole duration and exhibits accelerated cell growth between days 6 and 9 post-seeding. In general, both types of scaffolds are suitable for tissue engineering of vascular grafts using SMCs and they survive very well for the duration of 9 days in culture medium, as tested in this study and shown in Figure 9.

4. Discussion

It is clear from Figures 3 and 4 that by varying the process parameters of electrospinning it is possible to change the microstructural parameters of the electrospun scaffold, including fibre diameter, pore size, porosity and fibre alignment. Taking into account that the optimised crosslinking procedure brings an average increase of 55% in fibre diameter and wetting in de-ionised water at 37 oC swells the fibres by a further 100%, the following ranges of wet, crosslinked, electrospun scaffold microstructures may be obtained by selective changes in the electrospinning parameters from the standard conditions (feed solution of 10 w/v gelatin at constant flow rate of 2.5 ml/hr, collector rotating at 95 rpm at 25 cm distance from the feed nozzle, feed nozzle at +22.5 kV, knife edge electrode at -3 kV below the collector): by changing the feed solution gelatin concentration from 7.5 to 15 w/v, the wet scaffold fibre and pore size can be changed within the corresponding range of (df,wet = 4.65 m, rp,wet = 1.97 m) to (df,wet = 11.8 m, rp,wet = 4.8 m) for an approximately constant porosity wet = 15%; furthermore, by changing the feed flowrate from 0.5 to 7.5 ml/hr, the scaffold microstructure can be changed within the corresponding range of (df,wet = 4.65 m, wet = 15%, rp,wet = 2 m) to (df,wet = 7.8 m, wet = 20%, rp,wet = 4.3 m); while operating at the high limit of gelatin concentration 15 %w/v and flow rate of the feed solution 7.5 ml/hr, changing the voltage of the positive electrode from 15 to 30 kV led to the corresponding change in scaffold microstructure of ((df,wet = 8.1 m, wet = 16%, rp,wet = 3.4 m) to (df,wet = 25.1 m, wet = 19%, rp,wet = 13.1 m). Overall, wet, crosslinked, electrospun microfibrous scaffolds could be made with wet fibre diameter ranging from 4.6 to 25 m, wet porosity ranging from 15 to 20% and wet pore radius ranging from 2 to 13 m. Tissue engineered grafts were targeted to replace small calliper muscular arteries (such as coronary artery) in this study, hence the scaffolds were designed to be tunica media-equivalent for their in vitro population with SMCs. Pore sizes of the produced wet scaffolds may be compared to the dimensions of the seeded adhered SMCs in this study, which were estimated to be of an average cell thickness of 1.5 m and average cell lateral area of 12x103 m2 [48].

It is clear from Figure 8 that SMCs tend to adhere, elongate and orientate along the fibres of scaffold. Given the similar values of ultimate tensile strength in the axial and circumferential directions of tunica media of the human coronary artery [39] and the popular ±45o fibre orientation of continuous fibre composites used in the walls of pressure vessels, a ±45o fibre orientation was targeted in the electrospinning of microfibrous scaffolds (with shifts to circumferential orientation as seen in Figure 6(a)), which was maintained after crosslinking of gelatin. This resulted in similar Young’s modulus (or compliance) in the axial and circumferential direction of the produced wet tubular scaffolds, as seen in Figure 7(a),(c),(e), in the range of 0.1-1.2 MPa for a corresponding wet porosity varying between 20 and 15%, respectively. This can be compared to the Young’s modulus of the coronary artery as a whole (including cells) which has been reported to be in the range of 1.06-4.1 MPa [49] depending on age and sex. Furthermore, Ozolanta et al [49] reported that the coronary artery walls consist of adventitia 200 m thickness relatively constant with age, tunica media 200 m relatively constant with age, and tunica intima of thickness increasing with age from 100 to 400 m on average which causes the undesirable artery stiffening. In general, hyperelastic constitutive models have been fitted for the different parts of coronary artery, with the Young’s modulus at large final deformation on the basis of ultimate stress-strain, E, of the different layers of the human coronary artery found to be [39]: Eadventitia,circ = 2.2 MPa, Eadventitia,axial = 1.5 MPa, Emedia,circ = 0.55 MPa, Emedia,axial = 0.57 MPa, Eintima,circ = 0.66 MPa, Eintima,axial = 0.71 MPa. Hence, from the point of view of Young’s modulus all scaffolds of this study may be suitable as vascular constructs for the coronary artery as they are expected to be combined with an internal layer, intima-equivalent for the growth of endothelial cells. The proposed vascular graft will be based on a tubular, multi-layer, fibrous scaffold with an internal layer of axial fibre orientation (intima-equivalent) and further electrospun layers of ±45 o fibre orientation (tunica media-equivalent, as in this study). The intima-equivalent internal layer will have axial fibre orientation which will be achieved by angling the line of feed nozzle needle to knife-edged electrode in the x-y plane at an angle opposite to the direction of collector rotation, in order to compensate for the effect of the collector rotation on fibre alignment. Hence, it may be assumed that tunica-media scaffolds with a wet porosity of 15-19% (corresponding dry porosity of 62-75%) may be suitable to be used in vascular grafts (combined with an internal, intima-equivalent layer) to match the compliance of coronary artery. It must be mentioned that the saphenous vein mostly used as autologous vascular graft of the coronary artery is very anisotropic and much stiffer than the coronary artery, with a Young’s modulus of 2.3 MPa (small deformation) to 43 MPa (large deformation) in the circumferential direction and 23.7 MPa (small deformation) to 130 MPa (large deformation) in the axial direction [50]: this mismatch in compliance between the autologous graft and the native artery may cause intimal hyperplasia [51], which could be eliminated using the tissue engineered vascular grafts proposed in this study.

The tensile strength of the tunica media-equivalent, tubular, wet, crosslinked, electrospun scaffolds of this study varied from 0.1 MPa (in both circumferential and axial directions) to 2.1 MPa (circumferential direction) and 0.8 MPa (axial direction) for the wet porosity changing from 20 to 15% respectively, compared to 0.45 MPa (circumferential direction) and 0.42 MPa (axial direction) tensile strength of the tunica media of coronary artery [39]. Hence, it may be assumed that tunica-media scaffolds with a wet porosity of 15-18.8% (corresponding dry porosity of 62-75%) could be used in vascular grafts. This is the first report where crosslinked gelatin fibre scaffolds have achieved such high strength, higher than the maximum tensile strength of 0.68 MPa reported for porous crosslinked collagen-elastin-glycosaminoglycan scaffolds [25].

The suture retention strength of wet crosslinked electrospun scaffolds of wet porosity in the range of 16-18.8% was found to vary from 1.94 to 1.81 MPa respectively (Table 1) which is very close to the typically accepted minimal limit of 2 N for a tissue engineered graft for implantation [52] and the same or better than the suture retention strength of fresh saphenous vein (1.81 MPa [53]) and even synthetic fibre scaffolds (max 1.2 N for elastomeric polyester urethane fibre scaffolds[12]. While these values for the wet acellular scaffolds of this study are slightly lower than the required limit, cell seeding has shown to increase the suture retention force, meaning that after cell seeding the scaffolds will most likely cross the 2 N threshold.

Figure 6 indicates that the wet, crosslinked electrospun scaffolds in this study have a suitable permeability range of 10-12 – 10-11 m2 for the impregnation of the low viscosity culture medium. However, scaffold porosity and permeability are expected to be reduced as cells grow through the scaffold and block the pores. So while an increase in the porosity and pore size of scaffolds would facilitate cellular migration and good access to nutrients, this comes at the expense of mechanical properties. In particular, Figure 9 shows that a wet, crosslinked, electrospun scaffold of wet porosity wet = 20.8%, wet fibre diameter df,wet = 5 m and wet pore diameter of dp,wet = 5.7 m exhibits excellent cell growth and migration throughout a thickness of 240 m from the seeded top surface of the scaffold after 9 days, whereas a wet, crosslinked, electrospun scaffold of lower porosity wet = 18.5%, df,wet = 3.6 m and dp,wet = 3.6 m exhibits cell migration throughout a thickness of 186 m from the seeded top surface after 9 days and cell growth that has slowed down overall after day 3. The microstructural parameters of the latter scaffold seem to be at the border of the accepted mechanical properties and biological properties for cell growth. However, it must be added that cellular scaffolds of higher initial porosity are expected to exhibit higher mechanical properties than when tested in acellular form, which would extend the range of mechanically suitable scaffolds to higher porosities.

5. Conclusions

Novel crosslinked, tubular, electrospun, multi-layer, gelatin microfibre scaffolds have been fabricated which by aligning the gelatin fibres at ±45 o (alternating by layer) and optimising gelatin crosslinking achieved compliance and strength values similar to those of the tunica media of coronary artery, when tested as acellular and wetted (swollen in de-ionised water at 37 oC): a Young’s modulus of 0.5 to 1.2 MPa in both circumferential and axial directions and an ultimate tensile strength of 0.6 to 2.1 MPa in the circumferential direction and 0.45 to 0.8 MPa in the axial direction for a corresponding change of the wet scaffold porosity from 18.8 to 15% (equivalent range of dry scaffold porosity of 75 to 62%). The Young’s modulus values of our grafts create a compliance match better than that for saphenous vein whereas the strength values of these gelatin grafts are higher than that for natural vascular grafts in the literature [25]. Furthermore, these wet, gelatin fibre scaffolds reached for the first time in the literature high suture retention strength (1.8-1.94 MPa) still in acellular form, very close to the typically accepted lower limit of 2 N for a vascular graft and same or better than the suture retention strength of fresh saphenous vein (1.81 MPa [53]. The scaffolds allowed adherence and high rate of cell migration and proliferation for seeded human umbilical vein smooth muscle cells (SMCs) to a migration depth of 186 m and 240 m from the seeded surface through the scaffold thickness 9 days post-seeding for scaffolds of wet porosity of 18.5% and 20.8%, respectively (equivalent dry scaffold porosity of 74% and 83%) in static cell culture medium.

In conclusion, the optimised tissue engineered constructs could be considered tunica media-equivalent, tissue engineered vascular grafts. The proposed scaffolds could be electrospun sequentially on an initially electrospun gelatin fibre layer of axial orientation and seeded with SMCs on the outer surface of the tubular scaffolds (tunica media-equivalent) and endothelial cells on the inner surface (intima-equivalent) of the overall scaffold to obtain a fully biomimetic vascular graft.

Acknowledgments

The study was funded by the University of Surrey-NPL Partnership.

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-+zyx

Figure 1. A schematic of the electrospinning rig to produce 45o oriented fibres around a rotating cylindrical mandrel controlled using a stepper motor.

(a)

(b)

(c)

Figure 2. Dry electrospun scaffolds fabricated with gelatin solution feed concentration of 10 %w/v, feed rate of 2.5 ml/hr, applied positive voltage of 22.5 kV, at 25 cm distance between the feed and the collector, rotating at 95 rpm; (a) tubular uncrosslinked scaffolds, (b) SEM micrograph of uncrosslinked electrospun scaffold-single layer, (c) SEM micrograph of crosslinked electrospun scaffold-single layer. The black arrows in the images indicate the circumferential direction of electrospinning.

Figure 3. Graphs of fibre diameter, pore size (radius) and porosity of dry electrospun scaffolds as a function of electrospinning parameters: (a) gelatin concentration of feed solution, (b) feed solution flow rate and (c) applied positive voltage at feed nozzle.

50556065707580859095100234567891010203040Fibre fraction aligned at 45 degrees (%)Diameter (µm)Distance of nozzle from collectorFibre diameterFibre alignment

Figure 4. Graph of the effect of the distance between the feed nozzle and the collector on the scaffold fibre diameter and fibre alignment.

Figure 5. Graph of the mass ratio of the wet versus dry scaffold as a function of time of the scaffold being immersed in de-ionised water at 37 oC.

Figure 6. Crosslinked, electrospun scaffolds: (a) Graph of porosity against fibre diameter and fibre orientation distribution (for corresponding experimental points of the graph) of the dry, crosslinked scaffolds; SEM micrographs for fibre orientation measurements show a single scaffold layer. (b) Permeability data for the wetted, crosslinked, electrospun scaffolds plotted in a Carman-Kozeny relationship plot. (c) Stress-strain data from tensile testing of the wetted, crosslinked, electrospun scaffolds in the directions parallel (higher curve) and transverse (lower curve) to the fibres.

00.20.40.60.811.21.41.6606570758085Young's Modulus (MPa)Porosity (%)Axial LoadingCircumferential LoadingAxial loading fitCircumferencial loading fit00.511.522.53606570758085Ultimate tensile strength (MPa)Porosity (%)Axial LoadingCircumferential LoadingAxial loading fitCircumferential loading fit

(a)(b)

00.20.40.60.811.21.41.6051015Young's Modulus (MPa)Pore radius (µm)Axial LoadingCircumferential LoadingAxial loading fitCircumferencial loading fit

(c)(d)

00.511.522.53051015Ultimate tensile strength (MPa)Pore radius (µm)Axial LoadingCircumferential LoadingAxial loading fitCircumferencial loading fit

(e)(f)

00.10.20.30.40.50.60.70.80.911.11.21.31.41.50123456Young's Modulus (MPa)Fibre diameter (µm)Axial LoadingCircumferential LoadingAxial loading fitCircumferential loading fit00.511.522.530123456Ultimate tensile strength (MPa)Fibre diameter (µm)Axial LoadingCircumferential LoadingAxial loading fitCircumferential loading fit

Figure 7. Young’s modulus (a,c,e) and ultimate tensile strength (b,d,f) for wet, crosslinked, electrospun gelatin scaffolds as a function of the structural parameters of the dry scaffold: (a,b) scaffold porosity, (c,d) pore radius, and (e,f) fibre diameter.

(a) (b)

(c)

(d)

Figure 8. SEM micrographs of a bi-directional, multilayer scaffold with ±45 fibre orientation, fibre diameter of 2.1±0.3µm (dry) /4.2µm (wet) and porosity of 74±8% (dry) /18.5% (wet): (a) 3 days post seeding -seeded surface; no cells have reached the unseeded surface yet. (b) 6 days post seeding -seeded surface; (c) 6 days post seeding- unseeded surface; (d) 9 days post seeding- unseeded surface.

Figure 9. (a) Histological studies of cell migration: graph of distances of cellular migration from the seeded surface as a function of cell culture time in the scaffold, acquired from the histology of SEM micrographs of slices of the scaffold-S1 and scaffold-S2. On the left, examples of SEM micrographs showing cell migration across an image of a section of the scaffold-S1 (from left). (b) Overall cell proliferation MTS assay results for multilayer scaffolds S1 and S2, where S1: df = 1.8 µm (dry)/ 3.6 µm (wet) and = 74% (dry)/ 18.5 (wet); and S2: df = 2.5 µm (dry)/ 5 µm and = 83% (dry)/ 20.8% (wet).

(a)

(b)

(c)

(i)

(ii)

(iii)

(a)

(b)

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4050607080900510152002468Porosity (%)Average fibre or pore size (µm)Feed flow rate (ml/hr)

Fibre diameterPore sizePorosity

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