Electrospinning of Bioresorbable Polymers for Tissue Engineering

Chapter 14

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Electrospinning of Bioresorbable Polymers for Tissue Engineering Scaffolds 1

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E. D. Boland , K. J. Pawlowski , C. P. Barnes , D. G. Simpson , G. E. Wnek , and G. L . Bowlin 3

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Departments of Biomedical Engineering and Anatomy, Virginia Commonwealth University, Richmond, V A 23298 Department of Chemical Engineering, Case Western Reserve University, Cleveland, OH 44106

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Since the inception of the field of tissue engineering, there has been a considerable effort to develop an "Ideal" tissue engineering scaffold. To date, investigators have developed materials such as collagen, poly(glycolic acid) (PGA), poly(lactic acid) (PLA), and polycaprolactone (PCL) for use in matrix construction with frequently unacceptable results. The construction of an "Ideal" tissue engineering scaffold requires that multiple criteria are met. The first criterion is that the material be biocompatible and function without interrupting other physiological processes. This functionality includes an ability to promote normal cell growth and differentiation while maintaining three-dimensional orientation/space for the cells. Additionally, the scaffold should not induce any adverse tissue reaction. Another criterion involves the production of the scaffold. Scaffold production must include efficient material and construction parameters. The construction must also involve a process by which the scaffold can be easily reproduced to a wide range of shapes and sizes. Once implemented in vitro or in vivo, the material should be

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© 2006 American Chemical Society

In Polymeric Nanofibers; Reneker, D., et al.; ACS Symposium Series; American Chemical Society: Washington, DC, 2006.

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completely resorbable, leaving only native tissue. The method of electrospinning provides a simple method to achieve the goals of an "Ideal" tissue engineering scaffold. Electrospinning is the deliberate application of the phenomenon of electrostatic spraying which occurs when electrical forces at the surface of the polymer solution overcome the surface tension, creating a polymer solution jet. The jet produces a fiber with diameters in the micron to nano— scale range (down to 50 nm) as the solvent evaporates. The biodegradable polymer scaffold production (flexible, quick, and simple) utilizes the jet formation of the polymer-based solution from a nozzle to a grounded rotating mandrel. We have successfully electrospun P G A and P L A as pure polymers, co-polymers, and blends with each other as well as with P C L in order to develop novel tissue engineering scaffolds. The electrospinning process allows the production of P G A scaffolds comprised of 100 nm to 1.5 micron diameter fibers. For the P L A and P L A / P C L blends, the fiber diameter produced ranges from 8-10 microns. Obviously, with various composition and fiber dimensions, we can tailor both the mechanical properties and the degradation rates of the electrospun scaffoldings. The electrospinning process also allows us to control fiber orientation in the final fibrous scaffolds, ranging from completely random to predominately aligned. Thus, with these very promising results, it may now be possible to utilize this technology to create the "Ideal" tissue engineering scaffold (mimic tissue structure on demand).

Introduction Tissue Engineering as an interdisciplinary field of engineering and life sciences is in its infancy, yet the principles are as old as interventional surgery. The desire to restore, maintain, or improve tissue function remains the end goal, but a new set of tools and techniques are emerging. To this end, the study of structure-function relationships in both normal and pathological tissues has been coupled with the development of biologically active substitutes or



190 engineered materials. One such approach is to develop either cell-containing or cell-free implants that are capable of mimicking the native tissue. In order to duplicate all the essential intercellular reactions and promote native intracellular responses, the tissue structural component that must be mimicked is the extracellular matrix (ECM). These synthetic E C M s or scaffolds should be designed to conform to a specific set of requirements. The first requirement is that the material must be biocompatible and function without interrupting other physiological processes. This functionality includes an ability to promote normal cell growth and differentiation while maintaining three-dimensional orientation/space for the cells. Secondly, the scaffold should not promote or initiate any adverse tissue reaction (/). In addition, for clinical and commercial success, scaffold production should be simple yet versatile enough to produce a wide array of configurations to accommodate the size, shape, strength and other intricacies of the target tissue (7-5). Once implemented in vitro or in vivo, the material should either be removed though degradation and absorption or incorporated through native remodeling mechanisms, leaving only native tissue. Beyond these generalized requirements, we must look at the way a single native cell interacts with its immediate surroundings. It is no longer acceptable to view a cell as a self-contained unit sitting in a passive structural network (ECM). A dynamic three-dimensional inter-relation is constantly kept in balance and influenced by both internal and external stimuli. This model requires that any scaffold material must be able to interact with cells in three dimensions and facilitate this communication. In the native tissues, the structural E C M proteins are 1 to 2 orders of magnitude smaller than the cell itself, thus, allowing the cell to be in direct contact with many E C M fibers at once, and thereby defining its three dimensional orientation. This property may be a crucial factor in determining the success or failure of a tissue engineering scaffold. We believe that electrospinning represents an advantageous processing method to meet both the general material requirements as well as the potential size issues raised above. In this overview, we will first discuss a brief history of the technique and then follow that with a simple electrospinning process that can and has been performed in the laboratory. Once the process is established, we will focus on the work done in our laboratories to merge this processing technique with the body of literature available on bioresorbable polymers to create a wide variety of scaffolds for tissue engineering.

Brief History of Electrospinning: Pioneering Work and Mechanism th

Electrospinning can be documented as far back as Lord Rayleigh in the 19 Century. He theorized and subsequently determined through experiments that a sufficiently large electrical charge can cause a droplet at the tip of a



191 nozzle to overcome surface tension and eject in a stream (6,7). Specifically, this work focused on a process known as electrospraying. This phenomenon and that of electrospinning are fundamentally identical. For very low viscosity liquids, the jet that is ejected dissociates into droplets due to overwhelming surface tension, resulting in a sprayed coating on the collection plate (8). In higher viscosity liquids, the jet does not break up, and the result is a continuous stream or, if dried, a fiber. Zeleny further documented the phenomenon of electrospraying in 1917(0,7). Many years later, Taylor studied and explained the events that occur at the nozzle as the electric field deforms the fluid and it forms a stream (6,8). When the voltage is initially applied to the solution, the droplet at the nozzle forms a hemispherical surface. As the electric field is increased, the surface undergoes a shape change from hemispherical to spherical and eventually to conical. These changes are due to the balancing of the increasing solution charge with its surface tension, and the final conical shape is known as the Taylor cone (9). Countless experiments since have shown that when the liquid has a finite conductivity (even very small values), the stable configuration known as the Taylor cone is actually a cone with a thin column or jet of liquid emerging from the tip. Such "cone-jet" configurations are commonly referred to as "Taylor cones," even though the conditions under which they are formed vary greatly from the assumptions underlying Taylor's analysis. The first patent (U.S. Patent 1,975,504) in electrospinning was granted to Formhals in 1934 for a process that produced fine fibers from a cellulose acetate solution (6,10). He was later granted related patents (U.S. Patents 2,116,942; 2,160,962; and 2,187,306) in 1938, 1939, and 1940. In the last decade, there has been increasing interest in the electrospinning process. A concise review of the process and history of electrospinning can be found in an article by Reneker and Chun (6). Some examples of particularly interesting work follow. Srinivasan and Reneker dissolved Kevlar® (poly(p-phenylene terephthalamide)) in sulfuric acid and were able to obtain fibers with diameters of 1.3 jim that exhibited properties similar to that of commercial Kevlar® fibers spun from the liquid crystalline state (77). For these experiments, the fibers were collected in a grounded water bath, which allowed for removal of residual acid. Carbon nanofibers with diameters in the range of 100-500 nm have also been electrospun from polyacrylonitrile solution and molten mesophase pitch (12, 13). Doshi and Reneker demonstrated the electrospinning of poly(ethylene oxide) (PEO) in water to obtain fibers with diameters in the range of 0.5-5.0 |im (9). Other groups have also succeeded in electrospinning fibers from P E O and biodegradable materials such as poly(lactic acid) (PLA) (8, 14). Bognitzki formed P L A fibers with an average diameter of one micron from a solution containing tetraethyl benzylammonium chloride (TEBAC). T E B A C , which is an organosoluble salt, facilitated a decrease in fiber diameter by increasing solution surface tension and electrical conductivity (7). PCL, P L A , PGA, and poly(D, L -



192 lactide-co-glycolide) (PLGA) have been successfully electrospun into matrices with submicron fiber diameters and known degradation times for tissue engineering applications (15-18). Until recently, few researchers have concentrated on critical analysis of the electrospinning process variables. The effects of polymer solution concentration and spinning voltage (electric field) on fiber morphology have been detailed by Deitzel et al. (8) and Sukigara et al. (19). Gibson et al. have studied transport properties of the fibers and fiber mats (15,16,20% and mathematical models describing the travel of the fiber jet have been recently detailed by Reneker et al. (17) and Fridrikh et al. (21). From these bodies of work, we know that diameter, morphology, and orientation of the resultant fibers are a function of the many variables involved. The properties of the polymer solution, including its molecular weight, chemistry, and nature can greatly alter the ability of a fiber to be spun. The polymer to solvent ratio in solution affects viscosity, surface tension, and conductivity, all of which can influence fiber diameter. Voltage appears to play a role in determining fiber uniformity through jet stability. As voltage increases, the number of "beads on a string" defects increases. This is because the rate at which the solution is removed from the tip of the nozzle exceeds the rate at which the solution is delivered to the tip. Observationally, this can be seen as a jet initiating from the inside surface of the nozzle rather than from the Taylor cone (8, 14). These are not the only variables but are thought to be the most changeable (7, 9). Additionally, if fibers are still wet when they contact the collecting plate, they will fuse at junctions through a process commonly referred to as solvent welding. The higher the polymer concentrations with respect to solvent volumes, the drier the fiber produced (8,14). Solution infusion rates into the system and distances between the initiation tip and collecting plate also affect fiber dryness (7,8,14). The principal variable in fiber dryness is solvent volatility, the computations of which must account for the temperature, environmental conditions, and distance to the target if trying to maximize fiber dryness and minimize the formation of fused fibers (9).

A Simple Laboratory Apparatus for Electrospinning A simple apparatus (Figure 1) includes: a narrow conduit through which polymer solution is introduced to the system by a suitable pump or pressurized reservoir, a power supply capable of D C voltages in the 5 - 3 0 kilovolt range, and a conductive target. Because it is convenient to use stainless steel tubing



193 from which hypodermic needles are produced, the injection conduit is frequently referred to as a "needle." Opposite the needle at some distance away is a grounded target that can take on virtually any configuration as long as it is conductive. For this illustration, consider a simple, static metal plate. Therefore, any potential applied to the needle (or to the entering liquid by means of a contact electrode) produces an electric field at the needle tip. The intensity of this field depends upon the potential difference between the needle and the target, the diameter of the needle, and the distance to the target. As previously described, the field induces flow of the emerging liquid toward the target in the form of a Taylor cone. When the applied voltage is sufficient to overcome surface tension, a tiny stream of fluid is ejected from the tip of the needle. If sufficiently high polymer concentrations are present, continuous fiber formation will occur rather than the droplets predicted by Rayleigh instability. The charges on the fibers eventually dissipate to the surrounding environment, and the final product of the process is a non-woven mat that is composed of fibers with diameters on the order of nanometers to microns (//). Polymer melts can also be electrospun (12). In actuality, the technique is even more versatile, as we have found that electrospun fibers can also be collected on a dielectric material interposed between the grounded target and the spinning solution. Thus, a wide variety of substrates other than metals can be coated.

Figure 1. A schematic of a simple electrospinning system capable ofproducing micro- to nano-scale fibers which collect to become a biodegradable polymer tissue engineering scaffold.


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Electrospinning of Bioresorbable Polymers


Rationale for Electrospinning Bioresorbable Polymers The bioresorbable polymers under investigation in our laboratory are aliphatic polyesters of the poly(a-hydroxyl acids). These are some of the most commonly used biodegradable polymers and have the most clinical experience and initial acceptance. Their generalized formula is - [ ~ 0 - C H ( R ) - C O - ] - where R equals H for glycolic acid, C H side chain for lactic acid (both L and D isomers are possible), and C H C H along the backbone for caprolactone. Degradation rates vary from as short as two weeks for some P G A constructs to as long as a few years for PCL. This degradation is believed to be a function of hydrolytic attack initially in amorphous regions and then more slowly in crystalline regions. This degradation is characterized by a sharp decrease in local pH and a decrease in crystallinity (22,23). As will be shown later in this text, careful combinations of polymers combined with the fiber diameters possible with electrospinning can create the flexibility needed to address the vast array of tissue engineering scaffold requirements. 3

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Poly(glycolic acid) Poly(glycolic acid) (PGA) possesses moderate crystallinity (30% - 50%), a high melting point (225°C), and low solubility in organic solvents. Through clinical experience with suture, P G A was found to have better than average tissue compatibility and reproducible mechanical properties such as strength, elongation, and knot retention (23). P G A can be processed into fibers through the traditional extrusion methods. Due to the mechanics of the systems, fibers are restricted to diameters above 10 jam. Since electrospinning does not involve mechanical drawing, we have demonstrated the ability to produce scaffolds with individual fiber diameters from approximately 1.5 \im to less than 0,2 \xm (24). This is accomplished by dissolving P G A in l,l,l,3,3,3-hexafluoro-2-propanol (HFP) at various weight to volume ratios, the effect of which can be seen in the ability to control fiber diameter as illustrated in Figure 2. Diametric control, while controlling orientation, is the key to tailoring the mechanical properties of a P G A scaffold, including tangential modulus of the construct, stress and strain to failure, and strength retention; smaller fibers have larger surface area to volume ratios and lose their strength faster during degradation. Testing has shown that P G A is a good material when an initially tough (high strength and elasticity) material is needed and when it is desirable to have the construct degrade as fast as possible. However, a possible negative



195 attribute is also related to the degradation rate. Sharp increases in localized pH may cause unwanted tissue responses if the region does not have a high buffering capacity or sufficient mechanisms for the rapid removal of metabolic waste (23). In terms of the in vitro and in vivo biocompatilbilty testing of electrospun PGA, the authors have recently published studies which evaluated P G A and acid treated P G A in cell culture and in an animal model (25). In general the cellular and in vivo responses were a function of fiber diameter and acid pretreatment. The acid pretreatment improved biocompatibility via a hypothesized mechanism of fiber polymer surface hydrolysis of ester bonds, exposing carboxylic acid and alcohol groups (26), which may improve vitronectin binding thereby improving the ability of cells to adhere to the surface.

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Figure 2. (Top) Electrospun PGA (0.143 g/ml in HFP) showing the random fiber orientation (1,600X magnification). (Bottom) This graph illustrates the ability to linearly controlfiberdiameter as a function of initial concentration of PGA in HFP. Fibers ranging from less than 0.2 microns to over 1.2 microns were produced (24).


As has been shown with P G A , electrospinning is an attractive approach for the production of sub-micron diameter fibers, which are of interest as tissue engineering scaffolds. The fiber diameter, as well as the fiber orientation, is important in controlling material properties of the scaffolds produced. In a published study (24), the authors have evaluated the mechanical properties of the different P G A electrospinning concentrations in both aligned and random fiber orientation scaffolds. Figure 3 summarizes the results of this study in terms of the elastic modulus and strain to failure of P G A in a uniaxial model. The overall results exhibit a correlation between the fiber diameter and orientation and the elastic modulus and strain to failure.

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Figure 3. (Top) Modulus of elasticity and (Bottom) strain at failure measurements presented as a function of PGA concentration tested along the longitudinal and orthogonal axes (alignedfibers) and in a random mat (24).


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Poly(lactic acid) Poly(lactic acid) (PLA) is a very popular tissue engineering scaffold material and is available in either the L or D isomer. The L isomer is more commonly used, since it is derived from the only naturally occurring stereoisomer of lactic acid. The presence of the methyl group makes P L A more soluble in organic solvents and more hydrophobic than P G A . Despite a similar or higher degree of crystallinity compared to PGA, P L A commonly degrades in 30 to 50 weeks due to steric hindrance of the methyl group (23). Our laboratory testing has shown fiber formation by electrospinning P L A (27) in a range of 1/7 to 1/10 weight/volume ratios in both chloroform and methylene chloride. The decrease in concentration produces a similar decrease in fiber diameter, from about 10.5 |im to near 1 \im. If electrospun in HFP at similar concentrations, then one can create fibers down to 100 nm in diameter with P L A . Mechanically, crystalline P L A exhibits relatively high strength and modulus and low strain to failure. Given that the mechanical properties of P L A present strong and non-compliant scaffolds, it becomes obvious why P L A has been extensively used in the tissue engineering of bone and cartilage.

Polydioxanone Polydioxanone (PDS) is a highly crystalline degradable polyester originally developed for wound closure. The generalized formula of PDS is - [ 0 - C H - C H - 0 - C H - C = 0 ] - and it has a degradation rate falling between that of P G A and P L A . PDS exhibits excellent flexibility due to the incorporation of an ester oxygen in the monomer backbone. PDS has been electrospun in HFP at concentrations between 42 and 167 mg/ml, producing fibers that ranged linearly from 0.18 |xm to 1.4 jam, respectively, as seen in Figure 4. As a suture material, PDS is often criticized for shape memory. These sutures recoil to their original spooled shape which leads to kinking as fine sutures are tightened. While this is a negative trait for a suture material, it may provide rebound and kink resistance if formed into a vascular conduit. This unique trait coupled with excellent strength and elasticity may prove valuable for soft tissue engineering (28). The authors have evaluated the mechanical properties of the different PDS electrospinning concentrations in both aligned and random fiber orientation scaffolds (28). Figure 5 summarizes the results of this study in terms of the elastic modulus and peak stress of PDS in a uniaxial model. The overall results demonstrate a correlation between the fiber diameter and orientation and the elastic modulus and strain to failure. 2

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Electrospinning of Bioresorbable Polymers for Tissue Engineering

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