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Electrospun fibrous polyurethane scaffolds in tissue engineering
Electrospun fibrous polyurethane scaffolds in tissue engineering
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Y. Hong* University of Texas at Arlington, Arlington, TX, USA; Joint Biomedical Engineering Program, University of Texas Southwestern Medical Center, Dallas, TX, USA *Corresponding author: [email protected]
19.1 Introduction Electrospinning of nano/submicrometer fiber fabrication is an old technique for material processing, but it is still new to tissue engineering [1]. In 1897, Rayleigh observed the electrospinning phenomenon for the first time. In 1934, Dr Formhals was the first to patent the electrospinning technique, where cellulose acetate was electrospun into filaments [2]. In recent years, electrospinning has seen improvements in apparatus design and applications, utilizing polymers suitable for biomedical applications. As a result, electrospinning has gained increasing attention in the field of biomedical engineering, due to the development of tissue engineering to repair and regenerate tissues/organs [3–5]. Classic tissue engineering is the use of a biodegradable three-dimensional scaffold combined with human cells and biological signals to form a cellularized construct to regenerate a native tissue. Because of the 3D nanofibrous structure of human tissue extracellular matrix (ECM), a scaffold with mimetic ECM structure is assigned to be an advantageous application. Electrospinning is a simple, effective technique used to process biodegradable polymers into nano/submicrometer scale fibers and has been widely utilized for tissue engineering scaffold fabrication. Various biodegradable polymers including natural polymers, synthetic polymers, and their combinations have been electrospun into fibrous scaffolds [6,7]. Biodegradable thermoplastic polyurethane having robust mechanical properties and good biocompatibility is an excellent material candidate for tissue engineering use [8], and has been processed into nanofibrous scaffolds using electrospinning (Figure 19.1). In this chapter, we will introduce the electrospinning technique, describe some of the processing parameters, and discuss applications of electrospun polyurethane fibrous scaffolds.
19.2 Electrospinning technique and apparatus Electrospinning utilizes an electrostatic force to draw fine fibers (micro-, submicrometer), and nanoscales from a polymer solution. The basic and classic electrospinning apparatus is very simple and consists of a high voltage supply, a syringe pump, a syringe loaded with polymer solution with a metal tip, and a conductive collector (Figure 19.2). Advances in Polyurethane Biomaterials. http://dx.doi.org/10.1016/B978-0-08-100614-6.00019-6 Copyright © 2016 Elsevier Ltd. All rights reserved.
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The high voltage supply can provide up to tens of kV positive voltage, which is charged to the metallic tip of the syringe. The syringe pump accurately controls the infusion rate of the polymer solution. The conductive collector is made of metal (e.g., steel or aluminum), which is grounded, and is used for fiber deposition. The tip can be either perpendicularly located on the top of the collector (Figure 19.2(a)) or horizontally placed at the side of the collector (Figure 19.2(b)). After switching on the voltage and syringe pump, polymer fibers are produced immediately and deposit on the collector. When the electrospinning is complete, the fibrous scaffold can be removed from the collector. To improve the electrospinning technique and electrospun scaffolds, the electrospinning device can be modified. For example, to increase fiber collection, an extra high voltage supply can be used to connect to the collector to provide a negative charge. This can (a)
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Figure 19.1 (a) Random and (b) aligned electrospun fibers of biodegradable polyurethane.
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Figure 19.2 The classic electrospinning setup in a perpendicular direction (a) and a horizontal direction (b).
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prevent fiber loss because the fibers with the positive charges easily deposit onto a negatively charged surface [9]. Multiple spinnerets are often used to shorten the electrospinning time to rapidly produce a large-sized sample [10,11]. The collector can also be located on an x–y axial rastering device with a rotation motor to achieve more uniform samples.
19.3 Factors that affect the electrospinning process
Electrospinning parameters
Polymer solution
The fiber diameter of an electrospun scaffold is closely related to the scaffold properties and the biological response. It is influenced by the properties of the polymer solution and the parameters of the electrospinning process (Figure 19.3). Although the electrospinning device setup is simple, it is important that these parameters be comprehensively considered to achieve an optimal scaffold. The major parameters affecting the fiber diameter and the sample morphology are described as below. Chemical structure
Concentration
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Figure 19.3 A list of electrospinning factors including properties of the polymer solution and parameters of the electrospinning process.
19.3.1 Polymer properties The chemical structure and molecular weight of the polymer significantly affects the fiber formation. The polymers must be soluble in the organic solvent, and thus crosslinked polymers and some linear polymers with high crystallinity (e.g., PTFE) cannot be electrospun. However, for tissue engineering applications, most of the biodegradable polymers are polyester, polyamide, and polyurethane based, which are feasible for electrospinning. Most importantly, the molecular weight of the polymer affects the rheological and electrical properties of the polymer solution, including viscosity, surface tension, conductivity, and dielectric strength [12]. The fiber diameter generally increases with increasing polymer molecular weight. If the molecular weight is too low, no continuous fiber is obtained, and microbeads form. Fiber formation is dependent on the interactions of the polymer chains, such as physical entanglements. Low molecular weight results in a small number of polymer chain entanglements in the solution, which makes it hard for the polymer chains to form a fiber.
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19.3.2 Polymer concentration Altering polymer concentration is an effective way to tune the fiber diameter. The polymer concentration is associated with solution viscosity, which increases with the increase of the polymer concentration and then leads to a larger fiber diameter. However, if the concentration is too low, continuous polymer fibers cannot be achieved, and polymer beads are obtained. When the concentration is too high, the polymer solution becomes very viscous and fibers are obtained along with many drops. Highly viscous solutions have difficultly flowing through the syringe tip, resulting in droplet formation. Thus, an appropriate polymer concentration is very important to the electrospinning process.
19.3.3 Solvent The solvent for polymer solution in electrospinning generally should be highly volatile. The solvent should quickly evaporate during electrospinning to solidify the polymer fiber surface. The common solvents for electrospinning include 1,1,1,6,6,6-hexafluoroisopropanol (HFIP), trifluoroacetic acid (TFA), dichloromethane (DCM), formic acid, dimethyl formamide (DMF), tetrahydrofuran (THF), etc. For biodegradable polyurethanes, HFIP and DMF are two common solvents for electrospinning use. There are little data comparing the polymer fiber morphologies spun from different solvents. The conductivity of the solvent affects electrospinning with increased conductivity reducing the fiber diameter. Polymer solution with very low conductivity may form drops during the electrospinning. Thus, in the latter case, salt may be added into the polymer solution to improve the conductivity, facilitating electrospinning [13,14].
19.3.4 Voltage Altering the voltage is another convenient way to manipulate fiber morphology; however, it is not very effective. Drops form at a relatively low voltage. At a higher voltage polymer fibers can form. Increased voltage increases the electrostatic force and can reduce the fiber diameter. However when the voltage is too high, the electrospinning process becomes unstable. Thus, an optimized voltage is critical for stable electrospinning. The voltage usually is set at around 10–40 kV.
19.3.5 Distance between the tip and the collector The distance between the tip and the collector is a factor easily neglected. The distance is related to the solvent evaporation and the electrostatic force. Thus, with the increasing distance, the solvent can evaporate completely and the fiber diameter becomes smaller. However, if the distance is too great, this causes failure of the fiber formation. If the distance is too short, the solvent may not be completely evaporated, resulting in a larger fiber diameter. Reducing the distance also can increase fiber fusion and adherence to the collector surface [15].
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19.3.6 Infusion rate The infusion rate of the polymer solution is very important in the electrospinning procedure, and can be adjusted using a syringe pump. Generally, increasing the infusion rate increases the fiber diameter. However, if the infusion rate is too slow, it takes a long time for sample fabrication, and it may induce a temporary break in the electrospinning process. Polymer drops can form when the infusion rate is too high because the fiber production rate cannot catch up with the infusion rate. The ideal situation is that a stable Taylor cone forms during electrospinning, and the fiber formation rate is equal to the infusion rate, resulting in a stable and fast electrospinning process.
19.3.7 Tip diameter The tip diameter is closely related to the formation of a stable Taylor cone, but it does not have an obvious relationship to the fiber diameter. In general, the tip inner diameter is smaller than 2 mm. The typical size of the syringe needle is 16G to 23G.
19.3.8 Collector The fiber collector is crucial for the macroscopic shape and microscopic morphology of the electrospun scaffold. The fiber collector shape and whether it is moving have no effect on fiber diameter, but it has a significant effect on fiber direction and pattern, as well as scaffold shape. The collector is made of a conductive metal, which allows it to be connected to the ground or a negative charge. For random fibrous scaffolds, a metallic disk or a rectangular sheet can be used for fiber collection and fibrous sheet formation (Figure 19.4(a)). A metal mandrel as a collector can achieve tubular fibrous scaffolds (Figure 19.4(c)). For anisotropic aligned fibrous scaffolds, a rotating disk/ cylinder (Figure 19.4(b)) or a set of two parallel blade-like collectors (Figure 19.4(d)) is required [16]. The disk diameter and its rotation speed determine the degree of fiber alignment. At a low speed, the fiber direction is random, independent of the disk diameter. At a high speed (above 1 m/s), highly aligned fibers can be achieved. However, if D
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Figure 19.4 Schematics of metallic collectors. (a) Plate, (b) rotating disk, (c) rotating mandrel, (d) two blades, (e) rotating cone.
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the disk diameter is too small, it is difficult to achieve aligned fibers. For two parallel blade-like collectors, the electrostatic force distribution allows highly aligned fibers to deposit between the blades (Figure 19.4(d)). The fiber alignment decreases with the increasing sample thickness for the two-blade collector since the electronic field is affected by the fiber deposition. Compared to two-blade collectors, the rotating disk method produces a sample with a lower level of fiber alignment, but the sample alignment degree is not thickness dependent. Various collectors have been designed for fabricating fibrous materials for different applications besides the above-noted common collectors (Figures 19.4(a–d)). For example, a cone collector was used to obtain a fibrous sheet with regionally different fiber alignment to mimic heart valve fiber structure. The fiber alignment gradient is due to the different linear speed at different diameters [17] (Figure 19.4(e)). Using a metallic patterned collector one can obtain a fibrous pattern sheet for cell behavior research [18,19], Thus, rationally designing unique collectors is an effective way to fabricate novel fibrous scaffolds for tissue engineering use. Besides the above-noted parameters, there exist some other factors affecting electrospinning, such as the surrounding environment (e.g., humidity [20]). However, all these parameters interact and cannot be individually considered to achieve a perfect fibrous scaffold. All parameters must be comprehensively studied and optimized for electrospinning control.
19.4 Methods to enhance cellular infiltration of electrospun scaffolds Tissue engineering requires that cells extensively infiltrate into the scaffold for new tissue formation. However, because of superfine fiber diameters, the electrospun polymer scaffolds are very dense with a pore size of ∼1 μm, while the cell size is around 7 μm in general. Hence, the cells cannot infiltrate into the scaffolds, and it is also difficult to achieve good transport into the scaffold. Improving cell infiltration of the electrospun scaffold is critical for tissue engineering. A variety of approaches have been developed for this purpose. Three mechanisms for enhancing cell infiltration include (1) accelerate scaffold degradation to provide more space for cell growth; (2) reduce electrospun fiber intersections to loosen the scaffold structure to allow more cell infiltration; and (3) directly load cells into the electrospun scaffold. Some specific methods are described below.
19.4.1 Coelectrospinning Coelectrospinning is described as blending two or more polymers for electrospinning. To accelerate the degradation, some relatively quickly degradable materials are combined with the major polymer in a single solution to achieve a composite scaffold. The material introduction can increase the scaffold degradation rate, which may allow more cell infiltration, and it also can render some extra functionality to
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the scaffold. Biodegradable polyurethane was blended with porcine dermal extracellular matrix powder in HFIP and then electrospun into a fibrous patch [21]. The ECM material can leach out of the fiber patch and also quickly degrade. The patch was implanted into the rat full-thickness abdominal wall defect model and exhibited better cell infiltration than for the polyurethane alone. However, the cellular infiltration only occurred at the patch peripheries, and poor cell infiltration was observed at the center of the scaffold.
19.4.2 Unique collector design Unique fiber collector designs have been attempted to loosen the electrospun scaffold to improve cell infiltration. The principle of the collector design is to reduce the fiber intersections. For example, a half-ball collector containing pillars was designed to achieve cotton ball-like fibrous scaffolds by changing the fiber deposition space (Figure 19.5(a)) [22]. An ethanol bath was used as a collector to obtain low-density electrospun polycaprolactone scaffolds (Figure 19.5(b)) [23]. The ethanol quickly stabilizes the polymer fiber surface to reduce the intersections between fibers. D
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Figure 19.5 (a) A half-ball collector and (b) an ethanol bath collector.
19.4.3 Porogen method Porogen leaching is a popular method for preparing porous scaffolds. This method can also be used for improving the structure of the electrospun scaffold. The porogen and the electrospun fibers can be simultaneously deposited on the collector, and then the porogen is removed, which can build macropores inside the electrospun scaffold. These large pores can allow for cell infiltration. This concept has been demonstrated by simultaneously depositing polycaprolactone polymer fiber and sodium chloride particles through a specifically designed coaxial needle [24]. The inner tube held the polymer solution, while the outer annular region held the salt particles. After the salt particles were removed by water immersion, the electrospun scaffold showed a delaminated layer structure, which allowed for cell infiltration into the scaffold. Salt particles can also be codeposited with
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electrospun fibers using a sieve along with a vibrating orbital shaker [25]. The particles were uniformly dispersed in a hyaluronic acid/collagen fibrous scaffold. After salt leaching, the scaffold contained large pores with cubic shapes, which allowed chondrocyte growth and proliferation inside the scaffold.
19.4.4 Sacrificed fiber method The sacrificed fiber method combines two kinds of fibers into a scaffold and then removes one of the fibers, which reduces the fiber intersections and loosens the scaffold (Figure 19.6). Poly(ethylene glycol) has been usually used for the sacrificed fibers because it is water soluble and easily removed by immersing the scaffold in water. For this purpose, the electrospinning device requires two spinnerets. One spinneret is for a polymer solution to produce the targeted polymer fibers, and the other one is for the polymer solution to produce the sacrificial fibers. The two spinnerets can be placed parallel or perpendicularly with two kinds of fibers depositing on a collector. For example, two different fibers of slow-degradable polycaprolactone and water-soluble, sacrificial poly(ethylene oxide) (PEO) were combined into a composite [26]. The PEO fibers were selectively removed using water to increase the pore size of the electrospun scaffold, which facilitated cell infiltration and enhanced matrix distribution.
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Figure 19.6 Schematic procedure of the sacrificed fiber method.
19.4.5 Concurrent electrospray/electrospin method The concurrent electrospray/electrospin method is an effective method for loosening the fibrous scaffold, allowing extensive cell infiltration. At least two spinnerets are required. One spinneret is used to produce the electrospun polymer fibers, while the other spinneret is used to electrospray a flowable liquid, such as PBS, cell culture medium, or even pregel solution (Figure 19.7) [27–29]. These droplets produced by electrospray and the electrospun fibers concurrently deposit on a collector. The droplets can accelerate the polymer fiber surface solidification and prevent direct overlap of fibers to reduce the fiber intersections, and their volume can also occupy the space inside the scaffold, which can loosen the electrospun scaffolds. Electrospun poly(ester urethane) urea and electrosprayed cell culture medium concurrently deposited on a rotating cylinder collector were used to achieve a looser fibrous scaffold, called a “wet-electrospun” scaffold [27], Furthermore, the pregel solution from enzymatically digested dermal extracellular matrix was electrosprayed, concurrently depositing with
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the electrospun PEUU fibers [28]. After the pregel solution was gelled at 37 °C, a hydrogel/electrospun fiber composite scaffold was obtained. The above two scaffolds exhibited extensive cell infiltration after 4 weeks of implantation in a rat full-thickness abdominal wall defect model, while the conventional polymer scaffolds showed very poor cell infiltration.
19.4.6 Cell microintegration method The cell microintegration method is a direct way to cellularize an electrospun scaffold. The electrosprayed cell suspension and the polymer fibers concurrently deposit on a collector (Figure 19.7). It is notable that the cells can survive after high voltage treatment during electrospray. The solvent residual does not obviously induce severe cell death. Vascular smooth muscle cells, cardiac progenitor cells, and mesenchymal stem cells have been electrosprayed concurrently with electrospinning biodegradable polyurethane, which resulted in cell/fiber microintegrated constructs [30–33]. The cells inside the fibrous scaffold survived and proliferated, and the stem cells differentiated into the expected primary cells. These above methods are feasible and effective for changing the architecture of the dense fibrous scaffolds for improved cell infiltration. However, the scaffold loosening inevitably induces a decrease of the mechanical strength of scaffolds. It can also result in the accelerated degradation of the submicrometer fibrous scaffold as the fluid and the cells can rapidly penetrate into the scaffold. Thus, one must consider these changes of material properties prior to applying a loosened scaffold for tissue engineering.
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19.5 Electrospun polyurethane scaffolds in tissue engineering applications Tissue engineering is the use of a biodegradable scaffold combined with cells and signal molecules to regenerate a native tissue. As a promising candidate, the electrospun polyurethane scaffold has been applied for a variety of tissue repair and regeneration applications. The biodegradable PU scaffold has robust mechanical properties with good elasticity and surgical handling, and it also exhibits good biocompatibility. Other materials and bioactive molecules can also be combined with electrospun polyurethane scaffolds to improve their biofunctionality. The polyurethane scaffolds have been used for tissue engineering blood vessels, myocardia, heart valves, and abdominal walls, as well as skeletal muscle [34], bone [35], and neural tissue [36] (Figure 19.8). 6NHOHWDOPXVFOH
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Figure 19.8 Tissue engineering applications of biodegradable electrospun polyurethane scaffolds.
19.5.1 Blood vessel tissue engineering Autologous vascular transplantation is the gold standard for treating coronary and peripheral vascular diseases. However, the resource for healthy autologous vasculatures is limited in terms of the age and health status of the patient. Thus, the tissue engineering approach for regenerating a native blood vessel has been proposed. Biodegradable polyurethane has been processed into tubular scaffolds by electrospinning, which was applied for blood vessel regeneration, because it has robust mechanical properties with good elasticity and biocompatibility. Two main barriers to blood vessel tissue engineering are material thrombosis (blood compatibility) and hyperplasia (smooth muscle cell overproliferation), which induce restenosis and failure of the implant. Hence, prior to the use of polyurethane scaffolds for blood vessel replacement, their hemocompatibility must be improved. Additionally, it is thought that the mechanical mismatch with the scaffold and the native blood vessel is a main reason to induce hyperplasia; thus, polyurethane scaffolds are required to mechanically match with the native blood vessel. A biodegradable poly(ester urethane)urea (PEUU) was blended with a phospholipid polymer of poly(methacryloyloxyethyl p hosphorylcholine-co-methacryloyloxyethyl butyl urethane) (PMBU) (molar ratio = 70/30), in HFIP solvent, and then coelectrospun
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into a small diameter conduit (inner diameter, i.d. = 1.2 mm) as a biodegradable vascular graft [37]. The PEUU provided the mechanical support, while PMBU contributed to improve blood compatibility; PMBU addition significantly reduced ovine blood platelet deposition. The generated conduits have compliances comparable to those of the human artery. The 8 week in vivo implantation in a rat aorta model showed that the PMBU addition significantly increased the patency (75%) compared to PEUU alone (25%). A thin layer of neointimal including a layer of von Willebrand factor (vWF)-positive endothelial cell-like cells and a layer of alpha-smooth muscle actin-positive smooth muscle cell-like cells was observed by immunohistological staining. In a related strategy, the nonthrombogenic polymer can be grafted on the luminal surface of an electrospun biodegradable PEUU conduit [38]. This method can provide a blood-compatible surface without adverse effects on the PU conduit properties. The phospholipid polymer (poly(methacryloyloxyethyl phosphorylcholine-co-acrylic acid), PMA, molar ratio = 70/30) was grafted onto the lumen of an electrospun PEUU conduit [38]. The PMA coating significantly reduced the ovine blood platelet deposition. After 8 weeks of implantation into a rat aorta model (end to end implantation), the phospholipid polymer-coated PEUU conduit had a higher patency of 92% than that of PEUU conduit alone (40%). A continuous endothelial cell layer formed on the luminal surface. At 12 weeks, the coated PEUU conduit had mechanical properties including compliance, stiffness, and tensile strength comparable to those of native aorta. Biodegradable electrospun PU scaffolds can also be combined with drug release to improve blood compatibility. The biodegradable PU was mixed with dipyridamole (DPA) in HFIP, and then electrospun into a drug-eluting small diameter conduit (i.d. = 1.2 mm) [39]. The DPA introduction increased the initial modulus of the PU scaffold because the DPA had strong hydrogen bonding with the polyurethane. This scaffold showed DPA released up to 90 days. The DPA-loaded scaffold showed improved blood compatibility in terms of lower red cell hemolysis and platelet deposition compared to the scaffold without DPA. The DPA-eluting scaffold also supported endothelial cell growth, while inhibiting smooth muscle cell proliferation. Although the above scaffolds are promising as biodegradable vascular grafts for in vivo tissue engineered blood vessels, the dense fibrous structure limits the cell infiltration into the scaffold. A bilayer scaffold including an outer layer of an electrospun PEUU scaffold and an inner layer of porous PEUU scaffold was developed to address the cellularization issue [15]. The electrospun layer provided the mechanical support, while the porous scaffold layer supplied space for cell loading and infiltration. Prior to implantation, the cells including primary cells or stem cells were uniformly seeded into the porous layer using a customized vacuum device [15,40–42]. Rat vascular smooth muscle cells were seeded into the bilayer conduit and cultured in vitro for 2 days, and then this cellularized scaffold was implanted into the rat aorta model [42]. After 8 weeks, the patency of the cellularized scaffolds increased by 75% compared to 38% of the acellular scaffold. The failed scaffolds were blocked due to the intimal hyperplasia. The patent scaffolds contained a neointimal layer consisting of multiple layers of the immature contractile smooth muscle
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cells and a monolayer of endothelial cells. Muscle-derived stem cells and human pericytes were also seeded into this bilayer scaffold and then implanted into the rat aorta model. Both cellularized scaffolds showed markedly improved patency after 8 weeks of implantation compared to the acellular scaffolds [40,41]. It is notable that the human pericyte-seeded scaffolds showed 100% patency without dilation [41]. Good tissue remodeling and the existence of collagen and elastin were observed for the patent pericyte-seeded scaffolds. The multiple layers of smooth muscle cells and a monolayer of endothelial cells were detected on the lumen. Although, it was observed that the smooth muscle cells were not fully contractile, these cellularized bilayer scaffolds showed great promise for tissue engineered vascular grafts, and a long-term animal study is expected. In addition, direct cellularization for electrospun tubular scaffolds can be accomplished by a cell-microintegration technique. Rat vascular smooth muscle cells were simultaneously electrosprayed while electrospinning PEUU fibers onto a rotating mandrel to form a cell-infused tubular scaffold [30]. The cells survived and proliferated inside the scaffold. Biomechanical testing of the cellularized scaffold showed that it had similar mechanical properties to those of human coronary arteries and saphenous veins. In a recent study, mesenchymal stem cells also were combined with electrospun poly(ester carbonate) urethane fibers using a microintegration method to form a vascular graft [32]. This graft also showed similar mechanical properties with the native blood vessel. Unfortunately, no animal implant result is available at present.
19.5.2 Cardiac tissue engineering In the area of cardiac tissue engineering, electrospun polyurethane scaffolds have not yet been implanted into animals. However, some reports showed promising in vitro results. Neonatal rat cardiomyocytes were seeded on either random or aligned fibrous biodegradable PU sheets [43]. The aligned fibrous scaffold yielded highly oriented cardiomyocytes. The cells on the aligned scaffold showed a low steady state level of atrial natriuretic protein, and the continuous release of the protein implied that the cell phenotype turned into a more mature status. Furthermore, murine embryonic stem cell-derived cardiomyocytes were seeded on the aligned or random fibrous polyurethane scaffolds along with the mouse embryonic fibroblasts [44]. Compared to the random scaffold, the aligned scaffold resulted in increased anisotropy of rod-shaped cells and promoted sarcomere organization. Coculturing with the mouse fibroblasts further improved the sarcomere organization. In addition, mouse cardiosphere-derived cells were microintegrated with electrospun fibers of a blend of biodegradable polyurethane and a biodegradable thermosensitive hydrogel through a concurrent electrospray/electrospin method [31]. Blending the softer hydrogel and the stiffer polyurethane modulated the global modulus, single fiber modulus, fiber density, and alignment of the construct by altering fabrication parameters. It was found that the construct with low moduli (50–60 kPa) significantly promoted the cardiosphere-derived cell differentiation into mature cardiomyocytes in vitro after a 1 week culture.
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19.5.3 Heart valve tissue engineering A low flexural stiffness is one of the most important features in the design of a biodegradable scaffold for heart valve tissue engineering. Altering the electrospinning parameters can modulate the microstructure and mechanical properties of the biodegradable polyurethane fibrous scaffold. Increasing rastering rates can significantly decrease the fiber intersections, which reduces the bending modulus [45,46]. For a wet-electrospun biodegradable polyurethane scaffold, the fiber intersections associated with the stiffness reduced with increasing the raster speed from 0.3 to 30 cm/s [45]. A cell microintegrated scaffold prepared by combining rat smooth muscle cells and biodegradable polyurethane fibers at a slow rastering rate (3 cm/s) exhibited mechanical anisotropy similar to that of the native porcine pulmonary valve [45]. Furthermore, adding a secondary fiber method can tune the material bending stiffness. When the polyurethane fibers were combined with polycaprolactone (PCL) electrospun fibers, the fiber intersections and tensile modulus increased with the increase of PCL weight ratio [46]. When the PEO fibers were sacrificed from a combined scaffold of the polyurethane and PEO fibers, the fiber intersections and bending modulus of the formed polyurethane fibrous scaffolds were significantly reduced [46]. Recently, by depositing electrospun fibers on a rotating conical mandrel, a curvilinear biodegradable polyurethane fibrous scaffold was prepared [17]. The formed scaffold had a curvilinear fiber structure similar to that of the native pulmonary valve leaflet. Under quasistatic loading, the scaffolds with the curvilinear fiber microstructures had reduced strain concentrations compared to the scaffolds fabricated using a conventional cylindrical mandrel.
19.5.4 Abdominal wall reconstruction An appropriate biodegradable patch is assigned to treat abdominal wall trauma and hernia through tissue regeneration. An electrospun biodegradable polyurethane scaffold having high elasticity, biocompatibility, and robust mechanical strength meets these needs. However, once again cell infiltration is a major concern due to the dense fibrous structure of the electrospun scaffolds. A wet-electrospun scaffold was fabricated by simultaneously electrospinning PEUU fibers and electrospraying cell culture medium [27]. After 8 weeks of implantation into a rat full-thickness abdominal wall defect model, extensive cellular infiltration was observed for the wet-electrospun PEUU scaffold, while no cellular penetration was seen for the conventional electrospun PEUU scaffold. Biaxial testing exhibited anisotropic tissue remodeling. To further improve the bioactivity of electrospun abdominal wall scaffold, dermal extracellular matrix (ECM) powder was blended with the PEUU, and then electrospun into a scaffold sheet [21]. After an 8 week implantation, it was shown that more positive alpha-smooth muscle actin staining was found surrounding PEUU/ECM hybrid scaffolds compared to the PEUU scaffold alone. However, extensive cell penetration was not found. Through concurrently electrospraying dermal ECM pregel solution and electrospinning the PEUU, a dermal ECM hydrogel/ PEUU fibers hybrid scaffold was fabricated [28]. After a 4 week implantation into the rat abdominal wall model, the PEUU fibers/dermal ECM hydrogel hybrid scaffolds
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exhibited extensive cell infiltration with good tissue integration. The tissue remodeling was significantly improved. To further reinforce the dermal ECM hydrogel/PEUU fibers hybrid scaffold, a sandwich scaffold consisting of two fiber-rich layers (top and bottom) of wet-electrospun PEUU scaffolds and one fiber-poor layer (middle) of the hydrogel/ fiber hybrid scaffold was developed using a concurrent electrospray/electrospin method [29]. The three layers were achieved through sequentially electrospraying phosphate buffer solution (PBS), pregel solution and PBS. This sandwich scaffold had significantly greater mechanical strength than the control of the hydrogel/fiber hybrid scaffolds without the two fiber-rich layers. In the rat full-thickness abdominal wall defect model, the control failed at 8 weeks of implantation because of the implant thinning. The sandwich sample showed similar thickness with the native rat abdominal wall at 8 weeks and 12 week, and its explant had increased M2-type macrophages and better tissue remodeling with mimetic structure and mechanical anisotropy comparable to those of the native abdominal wall.
19.6 Summary and future trends Currently, scaffolds with mimetic structures and mechanical properties of the target native tissue are desired for tissue engineering. Biodegradable polyurethanes, which are mimetic with soft tissue mechanical behavior, have robust mechanical properties with high elasticity and flexibility. The electrospinning technique can process the polymer into nanoscale fibers, which can simulate the extracellular matrix microstructure of the tissue. Thus their combination has high potential for achieving an expected tissue engineering scaffold. In the next step, new biodegradable polyurethanes with additional property optimizations and biofunctionalities are desired. Reducing the initial modulus, increasing the mechanical strength, and combining bioactivity and biofunctions will be major areas of focus in new polymer development. Next, it will be necessary to develop new techniques to facilitate cell infiltration into the electrospun scaffold. Current approaches to the cell infiltration are still complex with limitations. A simple and effective way to achieve a bimodal pore structure would be a significant advantage. Finally, further in vivo testing of novel scaffold constructs is desired. The current biodegradable polyurethane fibrous scaffolds are known to be biocompatible and able to support primary cell growth and stem cell differentiation. The in vivo study, especially for specific tissue replacement, will confirm the feasibility of using polyurethane scaffolds for these applications. In summary, combining biodegradable polyurethanes and electrospinning can achieve promising fibrous scaffolds for tissue engineering applications. By altering the properties of the polymer solution and electrospinning parameters, the characteristics of fibrous scaffolds can be modulated to simulate native tissue. Furthermore, through design evolution of the electrospinning process, it is feasible to address the major challenge of cell penetration into the electrospun scaffolds. Polyurethane fibrous scaffolds are proposed for the regeneration of a variety of tissues, especially soft tissues. Such scaffolds with attractive structural, mechanical, and biodegradable properties have the potential to advance tissue engineering into clinical applications.
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Acknowledgments I greatly thank the financial support from the University of Texas at Arlington, and the American Heart Association (No. 14BGIA20510066).
References [1] Bhardwaj N, Kundu SC. Electrospinning: a fascinating fiber fabrication technique. Biotechnol Adv 2010;28:325–47. [2] Formhals A. Apparatus for producing artificial filaments from material such as cellulose acetate. US Patent No 1975504. 1934. [3] Zhong S, Zhang Y, Lim CT. Fabrication of large pores in electrospun nanofibrous scaffolds for cellular infiltration: a review. Tissue Eng Part B Rev 2012;18:77–87. [4] Aravindan V, Sundaramurthy J, Suresh Kumar P, Lee YS, Ramakrishna S, Madhavi S. Electrospun nanofibers: a prospective electro-active material for constructing high performance Li-ion batteries. Chem Commun 2015;51:2225–34. [5] Barnes CP, Sell SA, Boland ED, Simpson DG, Bowlin GL. Nanofiber technology: designing the next generation of tissue engineering scaffolds. Adv Drug Deliv Rev 2007;59:1413–33. [6] Ingavle GC, Leach JK. Advancements in electrospinning of polymeric nanofibrous scaffolds for tissue engineering. Tissue Eng Part B Rev 2014;20:277–93. [7] Kai D, Jin G, Prabhakaran MP, Ramakrishna S. Electrospun synthetic and natural nanofibers for regenerative medicine and stem cells. Biotechnol J 2013;8:59–72. [8] Guelcher SA. Biodegradable polyurethanes: synthesis and applications in regenerative medicine. Tissue Eng Part B Rev 2008;14:3–17. [9] Hong Y, Fujimoto K, Hashizume R, Guan J, Stankus JJ, Tobita K, et al. Generating elastic, biodegradable polyurethane/poly(lactide-co-glycolide) fibrous sheets with controlled antibiotic release via two-stream electrospinning. Biomacromolecules 2008;9:1200–7. [10] Tomaszewski W, Szadkowski M. Investigation of electrospinning with the use of a multijet electrospinning head. Fibers Text East Eur 2005;13:22–6. [11] Theron SA, Yarin AL, Zussman E, Kroll E. Multiple jets in electrospinning: experiment and modeling. Polymer 2005;46:2889–99. [12] Haghi KA, Akbari M. Trends in electrospinning of natural fibers. Phys Status Solidi 2007;204:1830–4. [13] Zong XH, Kim K, Fang D, Ran SF, Hsiao BS, Chu B. Structure and process relationship of electrospun bioabsorbable nanofiber membranes. Polymer 2002;43:4403–12. [14] Sui X, Wiesel E, Wagner HD. Enhanced mechanical properties of electrospun nano-fibers through NaCl mediation. J Nanosci Nanotechnol 2011;11:7931–6. [15] Soletti L, Hong Y, Guan J, Stankus JJ, El-Kurdi MS, Wagner WR, et al. A bilayered elastomeric scaffold for tissue engineering of small diameter vascular grafts. Acta Biomater 2010;6:110–22. [16] Xie J, Li X, Lipner J, Manning CN, Schwartz AG, Thomopoulos S, et al. “Aligned-to-random” nanofiber scaffolds for mimicking the structure of the tendon-to-bone insertion site. Nanoscale 2010;2:923–6. [17] Hobson CM, Amoroso NJ, Amini R, Ungchusri E, Hong Y, D’Amore A, et al. Fabrication of elastomeric scaffolds with curvilinear fibrous structures for heart valve leaflet engineering. J Biomed Mater Res Part A 2015;103(9):3101–6.
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[18] Song W, An D, Kao DI, Lu YC, Dai G, Chen S, et al. Nanofibrous microposts and microwells of controlled shapes and their hybridization with hydrogels for cell e ncapsulation. ACS Appl Mater Interfaces 2014;6:7038–44. [19] Zhang DM, Chang J. Patterning of electrospun fibers using electroconductive templates. Adv Mater 2007;19:3664–7. [20] Nezarati RM, Eifert MB, Cosgriff-Hernandez E. Effects of humidity and solution viscosity on electrospun fiber morphology. Tissue Eng Part C Methods 2013;19:810–9. [21] Hong Y, Takanari K, Amoroso NJ, Hashizume R, Brennan-Pierce EP, Freund JM, et al. An elastomeric patch electrospun from a blended solution of dermal extracellular matrix and biodegradable polyurethane for rat abdominal wall repair. Tissue Eng Part C Methods 2012;18:122–32. [22] Blakeney BA, Tambralli A, Anderson JM, Andukuri A, Lim DJ, Dean DR, et al. Cell infiltration and growth in a low density, uncompressed three-dimensional electrospun nanofibrous scaffold. Biomaterials 2011;32:1583–90. [23] Yang W, Yang F, Wang Y, Both SK, Jansen JA. In vivo bone generation via the endochondral pathway on three-dimensional electrospun fibers. Acta Biomater 2013;9:4505–12. [24] Nam J, Huang Y, Agarwal S, Lannutti J. Improved cellular infiltration in electrospun fiber via engineered porosity. Tissue Eng 2007;13:2249–57. [25] Kim TG, Chung HJ, Park TG. Macroporous and nanofibrous hyaluronic acid/collagen hybrid scaffold fabricated by concurrent electrospinning and deposition/leaching of salt particles. Acta Biomater 2008;4:1611–9. [26] Baker BM, Shah RP, Silverstein AM, Esterhai JL, Burdick JA, Mauck RL. Sacrificial nanofibrous composites provide instruction without impediment and enable functional tissue formation. Proc Natl Acad Sci USA 2012;109:14176–81. [27] Hashizume R, Fujimoto KL, Hong Y, Amoroso NJ, Tobita K, Miki T, et al. Morphological and mechanical characteristics of the reconstructed rat abdominal wall following use of a wet electrospun biodegradable polyurethane elastomer scaffold. Biomaterials 2010;31:3253–65. [28] Hong Y, Huber A, Takanari K, Amoroso NJ, Hashizume R, Badylak SF, et al. Mechanical properties and in vivo behavior of a biodegradable synthetic polymer microfiber-extracellular matrix hydrogel biohybrid scaffold. Biomaterials 2011;32:3387–94. [29] Takanari K, Hong Y, Hashizume R, Huber A, Amoroso NJ, D’Amore A, et al. Abdominal wall reconstruction by a regionally distinct biocomposite of extracellular matrix digest and a biodegradable elastomer. J Tissue Eng Regen Med 2013. http://dx.doi.org/10.1002/ term.1834. [30] Stankus JJ, Soletti L, Fujimoto K, Hong Y, Vorp DA, Wagner WR. Fabrication of cell microintegrated blood vessel constructs through electrohydrodynamic atomization. Biomaterials 2007;28:2738–46. [31] Xu Y, Patnaik S, Guo X, Li Z, Lo W, Butler R, et al. Cardiac differentiation of cardiosphere-derived cells in scaffolds mimicking morphology of the cardiac extracellular matrix. Acta Biomater 2014;10:3449–62. [32] Wang F, Li Z, Guan J. Fabrication of mesenchymal stem cells-integrated vascular constructs mimicking multiple properties of the native blood vessels. J Biomater Sci Polym Ed 2013;24:769–83. [33] Stankus JJ, Guan J, Fujimoto K, Wagner WR. Microintegrating smooth muscle cells into a biodegradable, elastomeric fiber matrix. Biomaterials 2006;27:735–44. [34] Riboldi SA, Sadr N, Pigini L, Neuenschwander P, Simonet M, Mognol P, et al. Skeletal myogenesis on highly orientated microfibrous polyesterurethane scaffolds. J Biomed Mater Res Part A 2008;84:1094–101.
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[35] Mi HY, Palumbo S, Jing X, Turng LS, Li WJ, Peng XF. Thermoplastic polyurethane/ hydroxyapatite electrospun scaffolds for bone tissue engineering: effects of polymer properties and particle size. J Biomed Mater Res Part B Appl Biomater 2014;102:1434–44. [36] Puschmann TB, de Pablo Y, Zanden C, Liu J, Pekny M. A novel method for threedimensional culture of central nervous system neurons. Tissue Eng Part C Methods 2014;20:485–92. [37] Hong Y, Ye SH, Nieponice A, Soletti L, Vorp DA, Wagner WR. A small diameter, fibrous vascular conduit generated from a poly(ester urethane)urea and phospholipid polymer blend. Biomaterials 2009;30:2457–67. [38] Soletti L, Nieponice A, Hong Y, Ye SH, Stankus JJ, Wagner WR, et al. In vivo performance of a phospholipid-coated bioerodable elastomeric graft for small-diameter vascular applications. J Biomed Mater Res Part A 2011;96:436–48. [39] Punnakitikashem P, Truong D, Menon JU, Nguyen KT, Hong Y. Electrospun biodegradable elastic polyurethane scaffolds with dipyridamole release for small diameter vascular grafts. Acta Biomater 2014;10:4618–28. [40] Nieponice A, Soletti L, Guan J, Hong Y, Gharaibeh B, Maul TM, et al. In vivo assessment of a tissue-engineered vascular graft combining a biodegradable elastomeric scaffold and muscle-derived stem cells in a rat model. Tissue Eng Part A 2010;16:1215–23. [41] He W, Nieponice A, Soletti L, Hong Y, Gharaibeh B, Crisan M, et al. Pericyte-based human tissue engineered vascular grafts. Biomaterials 2010;31:8235–44. [42] He W, Nieponice A, Hong Y, Wagner WR, Vorp DA. Rapid engineered small diameter vascular grafts from smooth muscle cells. Cardiovasc Eng Technol 2011;2:149–59. [43] Rockwood DN, Akins Jr RE, Parrag IC, Woodhouse KA, Rabolt JF. Culture on electrospun polyurethane scaffolds decreases atrial natriuretic peptide expression by cardiomyocytes in vitro. Biomaterials 2008;29:4783–91. [44] Parrag IC, Zandstra PW, Woodhouse KA. Fiber alignment and coculture with fibroblasts improves the differentiated phenotype of murine embryonic stem cell-derived cardiomyocytes for cardiac tissue engineering. Biotechnol Bioeng 2012;109:813–22. [45] Amoroso NJ, D’Amore A, Hong Y, Wagner WR, Sacks MS. Elastomeric electrospun polyurethane scaffolds: the interrelationship between fabrication conditions, fiber topology, and mechanical properties. Adv Mater 2011;23:106–11. [46] Amoroso NJ, D’Amore A, Hong Y, Rivera CP, Sacks MS, Wagner WR. Microstructural manipulation of electrospun scaffolds for specific bending stiffness for heart valve tissue engineering. Acta Biomater 2012;8:4268–77.
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Y. Hong* University of Texas at Arlington, Arlington, TX, USA; Joint Biomedical Engineering Program, University of Texas Southwestern Medical Center, Dallas, TX, USA *Corresponding author: [email protected]
19.1 Introduction Electrospinning of nano/submicrometer fiber fabrication is an old technique for material processing, but it is still new to tissue engineering [1]. In 1897, Rayleigh observed the electrospinning phenomenon for the first time. In 1934, Dr Formhals was the first to patent the electrospinning technique, where cellulose acetate was electrospun into filaments [2]. In recent years, electrospinning has seen improvements in apparatus design and applications, utilizing polymers suitable for biomedical applications. As a result, electrospinning has gained increasing attention in the field of biomedical engineering, due to the development of tissue engineering to repair and regenerate tissues/organs [3–5]. Classic tissue engineering is the use of a biodegradable three-dimensional scaffold combined with human cells and biological signals to form a cellularized construct to regenerate a native tissue. Because of the 3D nanofibrous structure of human tissue extracellular matrix (ECM), a scaffold with mimetic ECM structure is assigned to be an advantageous application. Electrospinning is a simple, effective technique used to process biodegradable polymers into nano/submicrometer scale fibers and has been widely utilized for tissue engineering scaffold fabrication. Various biodegradable polymers including natural polymers, synthetic polymers, and their combinations have been electrospun into fibrous scaffolds [6,7]. Biodegradable thermoplastic polyurethane having robust mechanical properties and good biocompatibility is an excellent material candidate for tissue engineering use [8], and has been processed into nanofibrous scaffolds using electrospinning (Figure 19.1). In this chapter, we will introduce the electrospinning technique, describe some of the processing parameters, and discuss applications of electrospun polyurethane fibrous scaffolds.
19.2 Electrospinning technique and apparatus Electrospinning utilizes an electrostatic force to draw fine fibers (micro-, submicrometer), and nanoscales from a polymer solution. The basic and classic electrospinning apparatus is very simple and consists of a high voltage supply, a syringe pump, a syringe loaded with polymer solution with a metal tip, and a conductive collector (Figure 19.2). Advances in Polyurethane Biomaterials. http://dx.doi.org/10.1016/B978-0-08-100614-6.00019-6 Copyright © 2016 Elsevier Ltd. All rights reserved.
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The high voltage supply can provide up to tens of kV positive voltage, which is charged to the metallic tip of the syringe. The syringe pump accurately controls the infusion rate of the polymer solution. The conductive collector is made of metal (e.g., steel or aluminum), which is grounded, and is used for fiber deposition. The tip can be either perpendicularly located on the top of the collector (Figure 19.2(a)) or horizontally placed at the side of the collector (Figure 19.2(b)). After switching on the voltage and syringe pump, polymer fibers are produced immediately and deposit on the collector. When the electrospinning is complete, the fibrous scaffold can be removed from the collector. To improve the electrospinning technique and electrospun scaffolds, the electrospinning device can be modified. For example, to increase fiber collection, an extra high voltage supply can be used to connect to the collector to provide a negative charge. This can (a)
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prevent fiber loss because the fibers with the positive charges easily deposit onto a negatively charged surface [9]. Multiple spinnerets are often used to shorten the electrospinning time to rapidly produce a large-sized sample [10,11]. The collector can also be located on an x–y axial rastering device with a rotation motor to achieve more uniform samples.
19.3 Factors that affect the electrospinning process
Electrospinning parameters
Polymer solution
The fiber diameter of an electrospun scaffold is closely related to the scaffold properties and the biological response. It is influenced by the properties of the polymer solution and the parameters of the electrospinning process (Figure 19.3). Although the electrospinning device setup is simple, it is important that these parameters be comprehensively considered to achieve an optimal scaffold. The major parameters affecting the fiber diameter and the sample morphology are described as below. Chemical structure
Concentration
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Figure 19.3 A list of electrospinning factors including properties of the polymer solution and parameters of the electrospinning process.
19.3.1 Polymer properties The chemical structure and molecular weight of the polymer significantly affects the fiber formation. The polymers must be soluble in the organic solvent, and thus crosslinked polymers and some linear polymers with high crystallinity (e.g., PTFE) cannot be electrospun. However, for tissue engineering applications, most of the biodegradable polymers are polyester, polyamide, and polyurethane based, which are feasible for electrospinning. Most importantly, the molecular weight of the polymer affects the rheological and electrical properties of the polymer solution, including viscosity, surface tension, conductivity, and dielectric strength [12]. The fiber diameter generally increases with increasing polymer molecular weight. If the molecular weight is too low, no continuous fiber is obtained, and microbeads form. Fiber formation is dependent on the interactions of the polymer chains, such as physical entanglements. Low molecular weight results in a small number of polymer chain entanglements in the solution, which makes it hard for the polymer chains to form a fiber.
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19.3.2 Polymer concentration Altering polymer concentration is an effective way to tune the fiber diameter. The polymer concentration is associated with solution viscosity, which increases with the increase of the polymer concentration and then leads to a larger fiber diameter. However, if the concentration is too low, continuous polymer fibers cannot be achieved, and polymer beads are obtained. When the concentration is too high, the polymer solution becomes very viscous and fibers are obtained along with many drops. Highly viscous solutions have difficultly flowing through the syringe tip, resulting in droplet formation. Thus, an appropriate polymer concentration is very important to the electrospinning process.
19.3.3 Solvent The solvent for polymer solution in electrospinning generally should be highly volatile. The solvent should quickly evaporate during electrospinning to solidify the polymer fiber surface. The common solvents for electrospinning include 1,1,1,6,6,6-hexafluoroisopropanol (HFIP), trifluoroacetic acid (TFA), dichloromethane (DCM), formic acid, dimethyl formamide (DMF), tetrahydrofuran (THF), etc. For biodegradable polyurethanes, HFIP and DMF are two common solvents for electrospinning use. There are little data comparing the polymer fiber morphologies spun from different solvents. The conductivity of the solvent affects electrospinning with increased conductivity reducing the fiber diameter. Polymer solution with very low conductivity may form drops during the electrospinning. Thus, in the latter case, salt may be added into the polymer solution to improve the conductivity, facilitating electrospinning [13,14].
19.3.4 Voltage Altering the voltage is another convenient way to manipulate fiber morphology; however, it is not very effective. Drops form at a relatively low voltage. At a higher voltage polymer fibers can form. Increased voltage increases the electrostatic force and can reduce the fiber diameter. However when the voltage is too high, the electrospinning process becomes unstable. Thus, an optimized voltage is critical for stable electrospinning. The voltage usually is set at around 10–40 kV.
19.3.5 Distance between the tip and the collector The distance between the tip and the collector is a factor easily neglected. The distance is related to the solvent evaporation and the electrostatic force. Thus, with the increasing distance, the solvent can evaporate completely and the fiber diameter becomes smaller. However, if the distance is too great, this causes failure of the fiber formation. If the distance is too short, the solvent may not be completely evaporated, resulting in a larger fiber diameter. Reducing the distance also can increase fiber fusion and adherence to the collector surface [15].
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19.3.6 Infusion rate The infusion rate of the polymer solution is very important in the electrospinning procedure, and can be adjusted using a syringe pump. Generally, increasing the infusion rate increases the fiber diameter. However, if the infusion rate is too slow, it takes a long time for sample fabrication, and it may induce a temporary break in the electrospinning process. Polymer drops can form when the infusion rate is too high because the fiber production rate cannot catch up with the infusion rate. The ideal situation is that a stable Taylor cone forms during electrospinning, and the fiber formation rate is equal to the infusion rate, resulting in a stable and fast electrospinning process.
19.3.7 Tip diameter The tip diameter is closely related to the formation of a stable Taylor cone, but it does not have an obvious relationship to the fiber diameter. In general, the tip inner diameter is smaller than 2 mm. The typical size of the syringe needle is 16G to 23G.
19.3.8 Collector The fiber collector is crucial for the macroscopic shape and microscopic morphology of the electrospun scaffold. The fiber collector shape and whether it is moving have no effect on fiber diameter, but it has a significant effect on fiber direction and pattern, as well as scaffold shape. The collector is made of a conductive metal, which allows it to be connected to the ground or a negative charge. For random fibrous scaffolds, a metallic disk or a rectangular sheet can be used for fiber collection and fibrous sheet formation (Figure 19.4(a)). A metal mandrel as a collector can achieve tubular fibrous scaffolds (Figure 19.4(c)). For anisotropic aligned fibrous scaffolds, a rotating disk/ cylinder (Figure 19.4(b)) or a set of two parallel blade-like collectors (Figure 19.4(d)) is required [16]. The disk diameter and its rotation speed determine the degree of fiber alignment. At a low speed, the fiber direction is random, independent of the disk diameter. At a high speed (above 1 m/s), highly aligned fibers can be achieved. However, if D
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Figure 19.4 Schematics of metallic collectors. (a) Plate, (b) rotating disk, (c) rotating mandrel, (d) two blades, (e) rotating cone.
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the disk diameter is too small, it is difficult to achieve aligned fibers. For two parallel blade-like collectors, the electrostatic force distribution allows highly aligned fibers to deposit between the blades (Figure 19.4(d)). The fiber alignment decreases with the increasing sample thickness for the two-blade collector since the electronic field is affected by the fiber deposition. Compared to two-blade collectors, the rotating disk method produces a sample with a lower level of fiber alignment, but the sample alignment degree is not thickness dependent. Various collectors have been designed for fabricating fibrous materials for different applications besides the above-noted common collectors (Figures 19.4(a–d)). For example, a cone collector was used to obtain a fibrous sheet with regionally different fiber alignment to mimic heart valve fiber structure. The fiber alignment gradient is due to the different linear speed at different diameters [17] (Figure 19.4(e)). Using a metallic patterned collector one can obtain a fibrous pattern sheet for cell behavior research [18,19], Thus, rationally designing unique collectors is an effective way to fabricate novel fibrous scaffolds for tissue engineering use. Besides the above-noted parameters, there exist some other factors affecting electrospinning, such as the surrounding environment (e.g., humidity [20]). However, all these parameters interact and cannot be individually considered to achieve a perfect fibrous scaffold. All parameters must be comprehensively studied and optimized for electrospinning control.
19.4 Methods to enhance cellular infiltration of electrospun scaffolds Tissue engineering requires that cells extensively infiltrate into the scaffold for new tissue formation. However, because of superfine fiber diameters, the electrospun polymer scaffolds are very dense with a pore size of ∼1 μm, while the cell size is around 7 μm in general. Hence, the cells cannot infiltrate into the scaffolds, and it is also difficult to achieve good transport into the scaffold. Improving cell infiltration of the electrospun scaffold is critical for tissue engineering. A variety of approaches have been developed for this purpose. Three mechanisms for enhancing cell infiltration include (1) accelerate scaffold degradation to provide more space for cell growth; (2) reduce electrospun fiber intersections to loosen the scaffold structure to allow more cell infiltration; and (3) directly load cells into the electrospun scaffold. Some specific methods are described below.
19.4.1 Coelectrospinning Coelectrospinning is described as blending two or more polymers for electrospinning. To accelerate the degradation, some relatively quickly degradable materials are combined with the major polymer in a single solution to achieve a composite scaffold. The material introduction can increase the scaffold degradation rate, which may allow more cell infiltration, and it also can render some extra functionality to
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the scaffold. Biodegradable polyurethane was blended with porcine dermal extracellular matrix powder in HFIP and then electrospun into a fibrous patch [21]. The ECM material can leach out of the fiber patch and also quickly degrade. The patch was implanted into the rat full-thickness abdominal wall defect model and exhibited better cell infiltration than for the polyurethane alone. However, the cellular infiltration only occurred at the patch peripheries, and poor cell infiltration was observed at the center of the scaffold.
19.4.2 Unique collector design Unique fiber collector designs have been attempted to loosen the electrospun scaffold to improve cell infiltration. The principle of the collector design is to reduce the fiber intersections. For example, a half-ball collector containing pillars was designed to achieve cotton ball-like fibrous scaffolds by changing the fiber deposition space (Figure 19.5(a)) [22]. An ethanol bath was used as a collector to obtain low-density electrospun polycaprolactone scaffolds (Figure 19.5(b)) [23]. The ethanol quickly stabilizes the polymer fiber surface to reduce the intersections between fibers. D
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Figure 19.5 (a) A half-ball collector and (b) an ethanol bath collector.
19.4.3 Porogen method Porogen leaching is a popular method for preparing porous scaffolds. This method can also be used for improving the structure of the electrospun scaffold. The porogen and the electrospun fibers can be simultaneously deposited on the collector, and then the porogen is removed, which can build macropores inside the electrospun scaffold. These large pores can allow for cell infiltration. This concept has been demonstrated by simultaneously depositing polycaprolactone polymer fiber and sodium chloride particles through a specifically designed coaxial needle [24]. The inner tube held the polymer solution, while the outer annular region held the salt particles. After the salt particles were removed by water immersion, the electrospun scaffold showed a delaminated layer structure, which allowed for cell infiltration into the scaffold. Salt particles can also be codeposited with
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electrospun fibers using a sieve along with a vibrating orbital shaker [25]. The particles were uniformly dispersed in a hyaluronic acid/collagen fibrous scaffold. After salt leaching, the scaffold contained large pores with cubic shapes, which allowed chondrocyte growth and proliferation inside the scaffold.
19.4.4 Sacrificed fiber method The sacrificed fiber method combines two kinds of fibers into a scaffold and then removes one of the fibers, which reduces the fiber intersections and loosens the scaffold (Figure 19.6). Poly(ethylene glycol) has been usually used for the sacrificed fibers because it is water soluble and easily removed by immersing the scaffold in water. For this purpose, the electrospinning device requires two spinnerets. One spinneret is for a polymer solution to produce the targeted polymer fibers, and the other one is for the polymer solution to produce the sacrificial fibers. The two spinnerets can be placed parallel or perpendicularly with two kinds of fibers depositing on a collector. For example, two different fibers of slow-degradable polycaprolactone and water-soluble, sacrificial poly(ethylene oxide) (PEO) were combined into a composite [26]. The PEO fibers were selectively removed using water to increase the pore size of the electrospun scaffold, which facilitated cell infiltration and enhanced matrix distribution.
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Figure 19.6 Schematic procedure of the sacrificed fiber method.
19.4.5 Concurrent electrospray/electrospin method The concurrent electrospray/electrospin method is an effective method for loosening the fibrous scaffold, allowing extensive cell infiltration. At least two spinnerets are required. One spinneret is used to produce the electrospun polymer fibers, while the other spinneret is used to electrospray a flowable liquid, such as PBS, cell culture medium, or even pregel solution (Figure 19.7) [27–29]. These droplets produced by electrospray and the electrospun fibers concurrently deposit on a collector. The droplets can accelerate the polymer fiber surface solidification and prevent direct overlap of fibers to reduce the fiber intersections, and their volume can also occupy the space inside the scaffold, which can loosen the electrospun scaffolds. Electrospun poly(ester urethane) urea and electrosprayed cell culture medium concurrently deposited on a rotating cylinder collector were used to achieve a looser fibrous scaffold, called a “wet-electrospun” scaffold [27], Furthermore, the pregel solution from enzymatically digested dermal extracellular matrix was electrosprayed, concurrently depositing with
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the electrospun PEUU fibers [28]. After the pregel solution was gelled at 37 °C, a hydrogel/electrospun fiber composite scaffold was obtained. The above two scaffolds exhibited extensive cell infiltration after 4 weeks of implantation in a rat full-thickness abdominal wall defect model, while the conventional polymer scaffolds showed very poor cell infiltration.
19.4.6 Cell microintegration method The cell microintegration method is a direct way to cellularize an electrospun scaffold. The electrosprayed cell suspension and the polymer fibers concurrently deposit on a collector (Figure 19.7). It is notable that the cells can survive after high voltage treatment during electrospray. The solvent residual does not obviously induce severe cell death. Vascular smooth muscle cells, cardiac progenitor cells, and mesenchymal stem cells have been electrosprayed concurrently with electrospinning biodegradable polyurethane, which resulted in cell/fiber microintegrated constructs [30–33]. The cells inside the fibrous scaffold survived and proliferated, and the stem cells differentiated into the expected primary cells. These above methods are feasible and effective for changing the architecture of the dense fibrous scaffolds for improved cell infiltration. However, the scaffold loosening inevitably induces a decrease of the mechanical strength of scaffolds. It can also result in the accelerated degradation of the submicrometer fibrous scaffold as the fluid and the cells can rapidly penetrate into the scaffold. Thus, one must consider these changes of material properties prior to applying a loosened scaffold for tissue engineering.
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19.5 Electrospun polyurethane scaffolds in tissue engineering applications Tissue engineering is the use of a biodegradable scaffold combined with cells and signal molecules to regenerate a native tissue. As a promising candidate, the electrospun polyurethane scaffold has been applied for a variety of tissue repair and regeneration applications. The biodegradable PU scaffold has robust mechanical properties with good elasticity and surgical handling, and it also exhibits good biocompatibility. Other materials and bioactive molecules can also be combined with electrospun polyurethane scaffolds to improve their biofunctionality. The polyurethane scaffolds have been used for tissue engineering blood vessels, myocardia, heart valves, and abdominal walls, as well as skeletal muscle [34], bone [35], and neural tissue [36] (Figure 19.8). 6NHOHWDOPXVFOH
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19.5.1 Blood vessel tissue engineering Autologous vascular transplantation is the gold standard for treating coronary and peripheral vascular diseases. However, the resource for healthy autologous vasculatures is limited in terms of the age and health status of the patient. Thus, the tissue engineering approach for regenerating a native blood vessel has been proposed. Biodegradable polyurethane has been processed into tubular scaffolds by electrospinning, which was applied for blood vessel regeneration, because it has robust mechanical properties with good elasticity and biocompatibility. Two main barriers to blood vessel tissue engineering are material thrombosis (blood compatibility) and hyperplasia (smooth muscle cell overproliferation), which induce restenosis and failure of the implant. Hence, prior to the use of polyurethane scaffolds for blood vessel replacement, their hemocompatibility must be improved. Additionally, it is thought that the mechanical mismatch with the scaffold and the native blood vessel is a main reason to induce hyperplasia; thus, polyurethane scaffolds are required to mechanically match with the native blood vessel. A biodegradable poly(ester urethane)urea (PEUU) was blended with a phospholipid polymer of poly(methacryloyloxyethyl p hosphorylcholine-co-methacryloyloxyethyl butyl urethane) (PMBU) (molar ratio = 70/30), in HFIP solvent, and then coelectrospun
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into a small diameter conduit (inner diameter, i.d. = 1.2 mm) as a biodegradable vascular graft [37]. The PEUU provided the mechanical support, while PMBU contributed to improve blood compatibility; PMBU addition significantly reduced ovine blood platelet deposition. The generated conduits have compliances comparable to those of the human artery. The 8 week in vivo implantation in a rat aorta model showed that the PMBU addition significantly increased the patency (75%) compared to PEUU alone (25%). A thin layer of neointimal including a layer of von Willebrand factor (vWF)-positive endothelial cell-like cells and a layer of alpha-smooth muscle actin-positive smooth muscle cell-like cells was observed by immunohistological staining. In a related strategy, the nonthrombogenic polymer can be grafted on the luminal surface of an electrospun biodegradable PEUU conduit [38]. This method can provide a blood-compatible surface without adverse effects on the PU conduit properties. The phospholipid polymer (poly(methacryloyloxyethyl phosphorylcholine-co-acrylic acid), PMA, molar ratio = 70/30) was grafted onto the lumen of an electrospun PEUU conduit [38]. The PMA coating significantly reduced the ovine blood platelet deposition. After 8 weeks of implantation into a rat aorta model (end to end implantation), the phospholipid polymer-coated PEUU conduit had a higher patency of 92% than that of PEUU conduit alone (40%). A continuous endothelial cell layer formed on the luminal surface. At 12 weeks, the coated PEUU conduit had mechanical properties including compliance, stiffness, and tensile strength comparable to those of native aorta. Biodegradable electrospun PU scaffolds can also be combined with drug release to improve blood compatibility. The biodegradable PU was mixed with dipyridamole (DPA) in HFIP, and then electrospun into a drug-eluting small diameter conduit (i.d. = 1.2 mm) [39]. The DPA introduction increased the initial modulus of the PU scaffold because the DPA had strong hydrogen bonding with the polyurethane. This scaffold showed DPA released up to 90 days. The DPA-loaded scaffold showed improved blood compatibility in terms of lower red cell hemolysis and platelet deposition compared to the scaffold without DPA. The DPA-eluting scaffold also supported endothelial cell growth, while inhibiting smooth muscle cell proliferation. Although the above scaffolds are promising as biodegradable vascular grafts for in vivo tissue engineered blood vessels, the dense fibrous structure limits the cell infiltration into the scaffold. A bilayer scaffold including an outer layer of an electrospun PEUU scaffold and an inner layer of porous PEUU scaffold was developed to address the cellularization issue [15]. The electrospun layer provided the mechanical support, while the porous scaffold layer supplied space for cell loading and infiltration. Prior to implantation, the cells including primary cells or stem cells were uniformly seeded into the porous layer using a customized vacuum device [15,40–42]. Rat vascular smooth muscle cells were seeded into the bilayer conduit and cultured in vitro for 2 days, and then this cellularized scaffold was implanted into the rat aorta model [42]. After 8 weeks, the patency of the cellularized scaffolds increased by 75% compared to 38% of the acellular scaffold. The failed scaffolds were blocked due to the intimal hyperplasia. The patent scaffolds contained a neointimal layer consisting of multiple layers of the immature contractile smooth muscle
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cells and a monolayer of endothelial cells. Muscle-derived stem cells and human pericytes were also seeded into this bilayer scaffold and then implanted into the rat aorta model. Both cellularized scaffolds showed markedly improved patency after 8 weeks of implantation compared to the acellular scaffolds [40,41]. It is notable that the human pericyte-seeded scaffolds showed 100% patency without dilation [41]. Good tissue remodeling and the existence of collagen and elastin were observed for the patent pericyte-seeded scaffolds. The multiple layers of smooth muscle cells and a monolayer of endothelial cells were detected on the lumen. Although, it was observed that the smooth muscle cells were not fully contractile, these cellularized bilayer scaffolds showed great promise for tissue engineered vascular grafts, and a long-term animal study is expected. In addition, direct cellularization for electrospun tubular scaffolds can be accomplished by a cell-microintegration technique. Rat vascular smooth muscle cells were simultaneously electrosprayed while electrospinning PEUU fibers onto a rotating mandrel to form a cell-infused tubular scaffold [30]. The cells survived and proliferated inside the scaffold. Biomechanical testing of the cellularized scaffold showed that it had similar mechanical properties to those of human coronary arteries and saphenous veins. In a recent study, mesenchymal stem cells also were combined with electrospun poly(ester carbonate) urethane fibers using a microintegration method to form a vascular graft [32]. This graft also showed similar mechanical properties with the native blood vessel. Unfortunately, no animal implant result is available at present.
19.5.2 Cardiac tissue engineering In the area of cardiac tissue engineering, electrospun polyurethane scaffolds have not yet been implanted into animals. However, some reports showed promising in vitro results. Neonatal rat cardiomyocytes were seeded on either random or aligned fibrous biodegradable PU sheets [43]. The aligned fibrous scaffold yielded highly oriented cardiomyocytes. The cells on the aligned scaffold showed a low steady state level of atrial natriuretic protein, and the continuous release of the protein implied that the cell phenotype turned into a more mature status. Furthermore, murine embryonic stem cell-derived cardiomyocytes were seeded on the aligned or random fibrous polyurethane scaffolds along with the mouse embryonic fibroblasts [44]. Compared to the random scaffold, the aligned scaffold resulted in increased anisotropy of rod-shaped cells and promoted sarcomere organization. Coculturing with the mouse fibroblasts further improved the sarcomere organization. In addition, mouse cardiosphere-derived cells were microintegrated with electrospun fibers of a blend of biodegradable polyurethane and a biodegradable thermosensitive hydrogel through a concurrent electrospray/electrospin method [31]. Blending the softer hydrogel and the stiffer polyurethane modulated the global modulus, single fiber modulus, fiber density, and alignment of the construct by altering fabrication parameters. It was found that the construct with low moduli (50–60 kPa) significantly promoted the cardiosphere-derived cell differentiation into mature cardiomyocytes in vitro after a 1 week culture.
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19.5.3 Heart valve tissue engineering A low flexural stiffness is one of the most important features in the design of a biodegradable scaffold for heart valve tissue engineering. Altering the electrospinning parameters can modulate the microstructure and mechanical properties of the biodegradable polyurethane fibrous scaffold. Increasing rastering rates can significantly decrease the fiber intersections, which reduces the bending modulus [45,46]. For a wet-electrospun biodegradable polyurethane scaffold, the fiber intersections associated with the stiffness reduced with increasing the raster speed from 0.3 to 30 cm/s [45]. A cell microintegrated scaffold prepared by combining rat smooth muscle cells and biodegradable polyurethane fibers at a slow rastering rate (3 cm/s) exhibited mechanical anisotropy similar to that of the native porcine pulmonary valve [45]. Furthermore, adding a secondary fiber method can tune the material bending stiffness. When the polyurethane fibers were combined with polycaprolactone (PCL) electrospun fibers, the fiber intersections and tensile modulus increased with the increase of PCL weight ratio [46]. When the PEO fibers were sacrificed from a combined scaffold of the polyurethane and PEO fibers, the fiber intersections and bending modulus of the formed polyurethane fibrous scaffolds were significantly reduced [46]. Recently, by depositing electrospun fibers on a rotating conical mandrel, a curvilinear biodegradable polyurethane fibrous scaffold was prepared [17]. The formed scaffold had a curvilinear fiber structure similar to that of the native pulmonary valve leaflet. Under quasistatic loading, the scaffolds with the curvilinear fiber microstructures had reduced strain concentrations compared to the scaffolds fabricated using a conventional cylindrical mandrel.
19.5.4 Abdominal wall reconstruction An appropriate biodegradable patch is assigned to treat abdominal wall trauma and hernia through tissue regeneration. An electrospun biodegradable polyurethane scaffold having high elasticity, biocompatibility, and robust mechanical strength meets these needs. However, once again cell infiltration is a major concern due to the dense fibrous structure of the electrospun scaffolds. A wet-electrospun scaffold was fabricated by simultaneously electrospinning PEUU fibers and electrospraying cell culture medium [27]. After 8 weeks of implantation into a rat full-thickness abdominal wall defect model, extensive cellular infiltration was observed for the wet-electrospun PEUU scaffold, while no cellular penetration was seen for the conventional electrospun PEUU scaffold. Biaxial testing exhibited anisotropic tissue remodeling. To further improve the bioactivity of electrospun abdominal wall scaffold, dermal extracellular matrix (ECM) powder was blended with the PEUU, and then electrospun into a scaffold sheet [21]. After an 8 week implantation, it was shown that more positive alpha-smooth muscle actin staining was found surrounding PEUU/ECM hybrid scaffolds compared to the PEUU scaffold alone. However, extensive cell penetration was not found. Through concurrently electrospraying dermal ECM pregel solution and electrospinning the PEUU, a dermal ECM hydrogel/ PEUU fibers hybrid scaffold was fabricated [28]. After a 4 week implantation into the rat abdominal wall model, the PEUU fibers/dermal ECM hydrogel hybrid scaffolds
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exhibited extensive cell infiltration with good tissue integration. The tissue remodeling was significantly improved. To further reinforce the dermal ECM hydrogel/PEUU fibers hybrid scaffold, a sandwich scaffold consisting of two fiber-rich layers (top and bottom) of wet-electrospun PEUU scaffolds and one fiber-poor layer (middle) of the hydrogel/ fiber hybrid scaffold was developed using a concurrent electrospray/electrospin method [29]. The three layers were achieved through sequentially electrospraying phosphate buffer solution (PBS), pregel solution and PBS. This sandwich scaffold had significantly greater mechanical strength than the control of the hydrogel/fiber hybrid scaffolds without the two fiber-rich layers. In the rat full-thickness abdominal wall defect model, the control failed at 8 weeks of implantation because of the implant thinning. The sandwich sample showed similar thickness with the native rat abdominal wall at 8 weeks and 12 week, and its explant had increased M2-type macrophages and better tissue remodeling with mimetic structure and mechanical anisotropy comparable to those of the native abdominal wall.
19.6 Summary and future trends Currently, scaffolds with mimetic structures and mechanical properties of the target native tissue are desired for tissue engineering. Biodegradable polyurethanes, which are mimetic with soft tissue mechanical behavior, have robust mechanical properties with high elasticity and flexibility. The electrospinning technique can process the polymer into nanoscale fibers, which can simulate the extracellular matrix microstructure of the tissue. Thus their combination has high potential for achieving an expected tissue engineering scaffold. In the next step, new biodegradable polyurethanes with additional property optimizations and biofunctionalities are desired. Reducing the initial modulus, increasing the mechanical strength, and combining bioactivity and biofunctions will be major areas of focus in new polymer development. Next, it will be necessary to develop new techniques to facilitate cell infiltration into the electrospun scaffold. Current approaches to the cell infiltration are still complex with limitations. A simple and effective way to achieve a bimodal pore structure would be a significant advantage. Finally, further in vivo testing of novel scaffold constructs is desired. The current biodegradable polyurethane fibrous scaffolds are known to be biocompatible and able to support primary cell growth and stem cell differentiation. The in vivo study, especially for specific tissue replacement, will confirm the feasibility of using polyurethane scaffolds for these applications. In summary, combining biodegradable polyurethanes and electrospinning can achieve promising fibrous scaffolds for tissue engineering applications. By altering the properties of the polymer solution and electrospinning parameters, the characteristics of fibrous scaffolds can be modulated to simulate native tissue. Furthermore, through design evolution of the electrospinning process, it is feasible to address the major challenge of cell penetration into the electrospun scaffolds. Polyurethane fibrous scaffolds are proposed for the regeneration of a variety of tissues, especially soft tissues. Such scaffolds with attractive structural, mechanical, and biodegradable properties have the potential to advance tissue engineering into clinical applications.
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Acknowledgments I greatly thank the financial support from the University of Texas at Arlington, and the American Heart Association (No. 14BGIA20510066).
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