Virginia Commonwealth University VCU Scholars Compass 2014 Tissue Engineering Scaffold Fabrication and Processing Techniques to Improve Cellular Infiltration Casey Grey Virginia Comm
Introduction
This chapter outlines the role of tissue engineering in medicine, explaining how engineered strategies can replace or direct the regeneration of diseased, damaged, or non-functional tissues It reviews the key fabrication methods used to create scaffolds—biomaterial frameworks designed to support and steer directed tissue regeneration By linking scaffold design, material choice, and fabrication technique to regenerative outcomes, the chapter highlights how engineered environments promote healing and functional restoration.
A Brief Review of Common Tissue Engineering Scaffold Manufacturing Techniques
Advancements in regenerative medicine are being pioneered in tissue engineering research labs around the world, driven by the urgent needs of patients who could benefit from regenerative therapies, with estimates suggesting that as many as one in three people in the United States may benefit Tissue engineering is a field dedicated to understanding biological systems with the translational goal of either replacing entire systems when they fail or guiding them back toward normal functionality when they stray.
Tissue engineering aims to create tissue analogs that direct cellular responses to regenerate damaged tissues and organs This typically involves an acellular scaffold—natural, synthetic, or hybrid—that mimics the extracellular matrix and serves as a template for cell adhesion and growth When the right cells are seeded, the scaffold guides migration, proliferation, and differentiation, evolving from an empty framework to a populated, functional tissue analog capable of performing physiological tasks, such as forming blood vessels The ultimate goal of regenerative medicine is to restore normal tissue function by implanting a biodegradable scaffold that is gradually replaced by native tissue as healthy cells invade the site In this approach, the scaffold provides a platform to transplant donor cells and to support targeted cell infiltration and organized tissue development, transitioning from scaffold to living tissue.
4 maturation of the construct into functional tissue, after which the scaffold degrades away so that normal physiological function is not impeded by non-physiological tissue components (e.g non- degradable polymer).[5-11]
1.2 FABRICATION OF TISSUE ENGINEERING SCAFFOLDS
In tissue engineering, scaffolds often fail clinical translation because they fall into two camps: mechanically robust constructs that do not provide the microenvironment needed for cell infiltration and occupation, or highly bioactive scaffolds that cannot withstand in vivo loading The field, including our lab, aims to close this gap by refining scaffold manufacturing techniques to achieve a balance between mechanical strength and the ability to elicit a targeted cellular response Chan and Leong offer an excellent framework for an ideal tissue engineering scaffold: one with architecture that promotes full tissue infiltration and vascular network formation, biocompatibility and cytocompatibility so cells attach and remain healthy, and no cytotoxic or immunogenic byproducts Ideally, such scaffolds would be positively bioactive by incorporating growth factors (e.g., FGF) to enhance scaffold–tissue integration, and they would closely match the target tissue's intrinsic structural and mechanical properties to provide immediate support and minimize mechanical stress during healing.
6 infiltrating cells after implantation as this could propagate an undesirable inflammatory wound healing response.[7]
Autologous grafts, harvested directly from the patient, offer key advantages over exogenous tissue‑engineering scaffolds, notably their native extracellular matrix composition and structure coupled with a low potential for immunogenicity, making them the gold standard for tissue repair in many cases While not all autologous grafts come from identical tissues—for example, nerve tissue used to repair a nerve defect—when the extracellular matrix characteristics are similar or adaptable, the autologous graft can support and direct a robust regenerative response This tissue plasticity underpins the tissue engineering paradigm, as highlighted by reports such as Ezzell’s use of pig intestinal grafts for vascular, ligament, and bladder applications and Pape’s demonstration of roughly a 90% success rate using autologous iliac crest bone grafts to bridge tibial gaps Together, these findings illustrate the regenerative promise of autologous grafts and the ongoing potential of tissue engineering approaches.
Although autologous grafts offer many benefits for tissue regeneration and are often regarded as the gold standard, they are not invariably the first-line treatment The core limitation is that the donor tissue must come from the same patient, which restricts the available supply and can cause donor-site morbidity from harvesting As a result, clinicians frequently consider alternative options—such as allografts or synthetic scaffolds—that reduce the demand on donor tissue while still supporting effective regeneration.
7 donor tissue is no longer serving its original function) and the recovery trauma from tissue harvesting surgery can outweigh the benefits of the graft.[9,23]
Allogenic grafts bypass donor-site morbidity for the recipient by using donor tissue from another person, which can be cadaveric or sourced from living donors such as kidney donors While decellularization and scaffold processing for allogenic grafts are rigorous, there remains a real risk of immunological response, inflammation, and rejection that can threaten graft survival Chemical decellularization and detergent washes effectively remove resident cells to reduce immunogenicity, but these treatments can also strip away growth factors bound to the extracellular matrix, which are often preserved in autologous grafts and support tissue regeneration A key advantage of decellularized allogenic grafts is the absence of donor-site morbidity while providing a scaffold that mimics the target tissue’s extracellular matrix Nevertheless, even with processing, there is always a risk of disease transmission associated with allogenic grafts.
A hydrogel has two component characteristics: the "gel" part of hydrogel refers to the polymer structure which consists of physically or chemically crosslinked hydrophilic polymers while the
Hydrogel gets its name from its capacity to absorb and retain water; in aqueous environments, the strongly hydrophilic polymer backbone causes the gel to swell as water is taken up, raising its water content from roughly 20% to as high as 99% by weight This high water content generally makes hydrogels very biocompatible Upon implantation, the hydrogel expands due to water influx, and soluble bioactive components embedded during design—such as those used for drug delivery—are released via hydration-driven mechanisms that can involve cell- or enzymatic-mediated processes, or, in conductive hydrogels, an external electrical signal These properties position hydrogels as prime candidates for drug delivery applications, and incorporating bioactive factors into the hydrogel matrix enables extended release of targeted pharmaceuticals.
Hydrogels' polymer networks can be based on natural polymers (for example collagen), synthetic polymers (for example poly(D,L-lactic acid-co-glycolic acid), PLGA), or combinations of natural and synthetic polymers To form a coherent structure, the separate polymer fibers are crosslinked together either physically or chemically Physical crosslinking involves chain entanglement, hydrogen bonding, ionic interactions, and strong van der Waals forces, and is typically characterized by weaker and reversible bonds Chemical crosslinking creates covalent bonds in a strong, irreversible reaction Although physically crosslinked hydrogels are structurally weaker, this approach is widely used because of their excellent biocompatibility.
Physical crosslinking avoids chemical agents that can harm cells, alter growth-factor activity, or provoke inflammation, making it favorable for systems with sensitive physiological components A typical example is injecting alginate droplets into a calcium bath, where rapid ionic gelation encapsulates growth factors or cells Although physically crosslinked hydrogels are highly biocompatible, their weak bonds often limit performance under physiological loads or during scaffold handling in tissue engineering To combine biocompatibility with mechanical robustness, researchers are pursuing dual-crosslinking strategies that begin with a physically crosslinked precursor to hold shape, followed by a secondary step that forms permanent covalent bonds through chemical crosslinking When the initial physical network is strong enough to retain structure and the chemical treatment is kept within biocompatible limits, this approach can yield strong, biocompatible hydrogels, and work in hydrogel development remains vigorously active.
Historically, hydrogel systems in tissue engineering have been favored for applications where high mechanical strength is not required and where there are no strict geometrical constraints on the implanted scaffold Notably, hydrogels have excelled in drug delivery within these contexts.
Hydrogels have numerous applications in tissue engineering, but achieving the mechanical strength needed to withstand physiological load-bearing conditions remains a major challenge Their tendency to swell in aqueous environments and the variations in physiological conditions also complicate the design of hydrogels that meet specific geometrical requirements While hydrogels typically lack strong mechanical strength, they offer excellent porosity, biocompatibility, and potential stimulus-responsive expansion—such as expansion after implantation or in response to heat or an electrical signal—keeping them a focus of active tissue engineering research.
One major limitation of nearly all tissue engineering scaffold fabrication methods is the lack of fine control over scaffold architecture While conventional techniques can adjust general properties such as porosity and mean fiber size, they cannot precisely place specific structural elements In contrast, three-dimensional (3D) printing builds scaffolds through layer-by-layer deposition of a curable polymer, enabling explicit control over the location of every structural component The typical workflow starts with generating a 3D computer model of the desired scaffold, which is automatically segmented into z-stacks by software and then sent to the 3D printer For example, if the cured height of the extruded polymer is 1 mm, a 1 cm tall scaffold would be digitally sliced into ten layers, with each layer corresponding to a defined z-position in the final construct.
The z-axis is partitioned into ten 1‑mm tall sections, and each layer is sequentially prepared for printing Each layer is sent to the 3D printer, which deposits the polymer onto a substrate to create that layer After completing a layer, the print head moves 1 mm in the z-direction and begins extruding the next layer When all ten layers are deposited, and if the polymer curing properties and printing timing are properly controlled, they will bond to form a coherent three-dimensional scaffold.
Figure 1.1 A commercially available 3D printer (http://3dprinterplans.info/tag/makerbot- replicator-2/)
3D printing technology has been used in bone regeneration [38], culturing cardiac cells [39,40], and for the creation of vascular and nerve grafts.[41] In addition to fabricating scaffolds, 3-D
3D printing technology also has the potential to create “copies” of tissues from medical images such as MRI scans, providing clinicians with a useful tool to assess treatment strategies, including a specific surgical approach or potential procedural risks on a patient-by-patient basis A major drawback is its inability to produce structural elements that match the size of native extracellular matrix Scaffolds printed with 3D printing technology available before 2010 were limited to components spanning a few hundred microns, far larger than native tissue matrix elements Today, the method can produce smaller structural features, but consistently printing sub-micron-sized components in a defined pattern has not been achieved across many compositional elements.
Figure 1.2 Jung et al show examples of 3D printing scaffolds A) whole scaffold, scale 2mm, B-D) SEM images of different scaffold fabrications [46] Note the precise control over scaffold architecture exhibited in 3D scaffolds
Cellular Responses in Wound Healing and Tissue Regeneration
Wound healing involves a complex cascade of physiological events that remains only partially understood, reflecting the intricate nature of tissue repair While the major theme here centers on cellular interactions with man-made extracellular matrix analogues—a focused subset of wound healing—understanding both upstream triggers and downstream consequences helps guide targeted research To equip readers with practical insight, this chapter reviews the general principles of wound healing, outlining the key phases, cellular players, and signaling pathways that constitute the repair process and set the stage for advancements in biomaterials and regenerative medicine.
Cellular Responses in Wound Healing and Tissue Regeneration
Casey P Grey 1 and David G Simpson 2
1Department of Biomedical Engineering and 2 Department of Anatomy and Neurobiology
The main objective of tissue engineering is to restore lost tissue function by clarifying what normal tissue function entails and how it diverges in various pathological states A tissue is best understood as a coordinated ensemble of cells that work together to perform a specific task, and even “healthy” tissue is dynamic, with a small, controlled amount of cell death that supports maintenance This dynamic balance is shaped by the ongoing interplay of chemical and physical signals and a cell’s ability to respond to them, which together determine tissue status and performance Clinically, pathology often manifests as reduced tissue function, pain associated with tissue activity, or both, highlighting the link between function and homeostasis Tissue dysfunction arises when the physiological environment is altered—through changes in cues, exposure to insults, or impaired cellular function—and understanding this environment is essential for guiding regenerative therapies.
Cells are the fundamental currency of living organisms.[59] In complex organisms, one or more types of cells grouped together to perform a specific function or functions are called a tissue.[60]
An organism is a coordinated collection of tissues that functions as a unified system aimed at prolonging its existence It sustains itself by maintaining homeostasis—regulating internal conditions—and, in a broader sense, by reproducing to improve the odds of long-term genetic survival.
Figure 2.1 Levels of physiological organization in a mammal.[65]
Cells in an organism act as the workforce that sustains the body's extracellular matrix, carrying out specialized tasks at a basal level—some are highly active, like immune cells surveying for threats, while others, such as fibroblasts, remain largely dormant as they monitor local signals; when environmental cues change, cells respond, for example osteoblasts respond to serum calcium and bone-promoting factors like bone morphogenetic proteins by depositing a protein and mineral matrix rich in collagen and hydroxyapatite to build and maintain the skeletal system; there are more than 200 distinct cell types in the human body, including epithelial cells, muscle myocytes, fibroblasts, neurons, stem cells, macrophages, and various circulating blood cells; wound healing is a coordinated effort that involves multiple cell types across tissues.
Wound healing is a complex process influenced by both tissue type and individual variation Different tissues show distinct responses to injury and treatment, and the same tissue can heal differently across patients Skin is commonly injured and thus one of the most thoroughly studied tissues in wound healing Its easy accessibility makes it a preferred model; for example, in rodents it is simpler to create a 6 mm skin lesion than a 6 mm cardiac infarct, and skin regeneration can be monitored with minimal manipulation because of its superficial location Although other tissues differ from skin in certain aspects of their wound response, they generally share essential recovery milestones such as debridement, targeted cellular infiltration, and revascularization.
Rather than trying to control the entire wound-healing cascade, many treatment strategies aim to trigger the core early healing events and then let the downstream cascade of less-understood minor processes unfold naturally The superficial nature of the skin makes it possible to observe critical events and timing points in wound healing, helping researchers determine which processes require careful monitoring and how these insights can be extrapolated to other tissues Moreover, the predictable and accessible healing response of skin enables non-invasive observation of the wound-healing process in humans.
The skin’s cellular compartment is dominated by epithelial cells, the biological interfaces between tissue and non-tissue substances throughout the body [59] These cells line the surface of the skin, blood vessel linings, and the linings of hollow internal organs, forming a barrier that both enables the transport and diffusion of beneficial materials and prevents harmful substances from penetrating the body [59, 61] In wound healing, epithelial tissue is a major research focus because severe adult injuries often replace it with fibrous connective tissue, which can cause a loss of epithelial properties and potentially lead to chronic wound‑induced tissue dysfunction [73].
Small surface-area wounds heal rapidly, especially when injuries are superficial, because skin cells stay in close contact and the loss of cell-to-cell contacts triggers migration and proliferation to close the wound With limited scar tissue, barriers to cell migration are few, allowing faster closure Most small wounds close within about a week, though the time to fully mature and restore original tissue properties can exceed a year depending on the extent and depth of tissue involved In contrast, wounds covering larger surface areas take much longer to heal because of greater tissue loss, often requiring dressings or scaffolds to support the healing process A key focus is keeping the wound bed moist, since a dry, scarred surface can hinder cell infiltration and delay healing.
Although wound healing unfolds as a continuum of events, it is useful to conceptualize it in stages, beginning with hemostasis to stop blood loss When a vessel is injured, circulating platelets rapidly adhere to the exposed sub-endothelial matrix, triggering activation of platelet integrins and promoting spreading and aggregation to form a stabilizing clot As platelets accumulate, they release soluble factors such as platelet-derived growth factor (PDGF) that support subsequent repair processes The clotting cascade also generates a fibrin mesh that embeds the platelet plug, creating a temporary matrix that both stops bleeding and provides a scaffold for the next phases of wound healing.
Following hemostasis, the wound enters a transient inflammatory phase marked by enhanced perfusion through vasodilation and local angiogenesis, along with the activation and migration of a broad range of cells into the wound bed Circulating neutrophils are among the first responders to infiltrate the injured tissue, drawn in by chemical and physical signals presented by the vascular endothelium at sites of compromised vessels Their primary role is to combat the large early influx of bacteria typical of most wounds, while they also release cytokines and growth factors that are among the mediators active in the early stages of wound healing.
Wound healing is enhanced by processes that recruit macrophages and fibroblasts to the injury site Growth factors released by various cell types at the wound site—epidermal growth factor (EGF), transforming growth factor-alpha (TGF-α), and heparin-binding EGF-like growth factor (HB-EGF)—promote improved wound closure and increased cell migration into the wound, driving more efficient tissue repair.
During wound healing, inflammatory signals activate macrophages, transforming them into phagocytic, pro-inflammatory cells that clear bacteria and debris and promote resolution by removing residual neutrophils As healing progresses, the inflammatory stage gives way to the proliferative phase, marked by fibroblast infiltration into the wound bed These fibroblasts express receptors for fibrin, fibrinogen, and vitronectin to facilitate migration and synthesize new extracellular matrix components—collagen, proteoglycans, and fibronectin—establishing the interstitial framework and supporting the ingrowth of blood vessels Through the secretion of fibroblast growth factor (FGF) and vascular endothelial growth factor (VEGF), fibroblasts and other infiltrating cells drive angiogenesis to meet the tissue’s metabolic needs Fibroblasts also differentiate into myofibroblasts, whose contractile activity promotes wound closure via wound contracture In the later stages of tissue recovery, the wound’s mechanical forces contribute to remodeling and maturation of the repaired tissue.
Environmental signals can revert activated fibroblasts to their quiescent state, and differentiated myofibroblasts typically undergo apoptosis since fibroblast-to-myofibroblast differentiation is considered terminal In both embryonic and adult wound healing, the wound contracts, but embryonic contraction does not appear to require fibroblast-to-myofibroblast differentiation Embryonic tissues express high levels of TGF-β3 and minimal TGF-β1, while adult tissues show the opposite pattern, with strong TGF-β1 and little TGF-β3 TGF-β1 promotes fibroblast-to-myofibroblast differentiation and drives increased collagen expression.
In adult tissues, wound healing after a serious injury typically results in scarring—a nonfunctional dermis dominated by dense, organized collagen bundles driven in part by increased collagen expression in response to pro-inflammatory signals such as TGF-β1 By contrast, embryonic tissues can achieve near-complete regeneration of the lost skin even after substantial injury, indicating that the healing outcome is not determined solely by the profile of TGF-β isoforms These signaling differences appear to prevent fibroblast-to-myofibroblast differentiation during embryonic repair, leading to less wound contraction and fewer aberrant collagen deposits The greater involvement of embryonic stem cell populations at the wound site likely also supports reconstitution of normal skin architecture in developing tissues.
Evidence indicates that the source of the final cells that infiltrate wounds and drive regeneration helps determine healing outcomes An active area of research asks how healing might change if a larger or smaller proportion of infiltrating fibroblasts originates from circulating fibrocytes—precursors to fibroblasts—versus resident fibroblasts that already exist in adjacent tissue and migrate into the injury Fibrocytes are less differentiated than resident fibroblasts and may confer benefits to regeneration by being more stem-cell–like Generally, more source cells with the capacity to adopt regenerative phenotypes correlate with better recovery For example, when a cutaneous injury does not devastate existing hair stumps, those residual stumps contribute a significant portion of the regenerated cell population Conversely, injuries that remove hair bulbs from the dermis produce wounds that do not regenerate hair, likely because the wound is infiltrated by deep dermal fibroblasts rather than those in superficial layers; deep dermal fibroblasts may lose the ability to regenerate superficial dermal layers, or the injury severity may create an environment that prioritizes wound closure over restoring normal superficial tissue function.
Creating Scaffolds Exhibiting Smooth Mechanical Gradients
Published in Biomaterials, this manuscript describes multi-layered electrospun scaffolds with controllable, gradual interlayer transitions rather than abrupt laminations, enabling precise direction of cellular responses to regenerate complex, multi-layered tissues while reducing the risk of delamination and scaffold failure The transition layer not only steers infiltrating cells toward a specific phenotype as they penetrate the scaffold but also functions as a mechanical primer to enhance the overall regenerative outcome By tackling the well-known weakness of delamination in multi-layered scaffolds, this work represents the first step toward clinically relevant tissue engineering scaffolds.
Gradient Fiber Electrospinning of Layered Scaffolds using Controlled Transitions in Fiber
Casey P Grey 1 , Scott Newton 2 , Gary L Bowlin 1 , Thomas Haas 1 , and David G Simpson 2
1Department of Biomedical Engineering and 2 Department of Anatomy and Neurobiology
Department of Anatomy and Neurobiology Virginia Commonwealth University Richmond, VA 23298 dgsimpso@vcu.edu
We characterize layered, delamination-resistant tissue engineering scaffolds produced by gradient electrospinning through an integrated workflow that combines computational fluid dynamics, fiber diameter measurements under dynamic polymer concentration changes, SEM analysis, and materials testing This approach links processing conditions to microstructural features and mechanical performance, showing that gradient electrospinning enables robust multilayer architectures with controlled fiber diameter distributions, validated by SEM and materials testing and illuminated by CFD insights into fluid dynamics during fabrication Our findings highlight how polymer concentration gradients influence fiber diameter across layers, providing design guidance for delamination resistance in scaffold systems.
Gradient electrospinning enables a continuously variable polymer concentration within the electrospinning jet, yielding scaffolds with controlled transitions in fiber diameter along the Z-axis This approach allows fabrication of scaffolds that present markedly different fiber sizes and material properties on opposing surfaces while eliminating boundary layers that lead to delamination failures In materials testing, bi-layered laminated electrospun scaffolds demonstrate how gradient control can tailor interfacial properties and produce robust, delamination-resistant constructs.
Two-layer scaffolds composed of sub-250 nm and 1000 nm diameter polycaprolactone (PCL) fibers exhibit ductile behavior and multiphasic failure When fabricated by gradient electrospinning with such fibers on opposing surfaces, the scaffolds fracture and fail as a unified, mechanically integrated structure Gradient electrospinning also eliminates the anisotropic strain properties observed in scaffolds made from highly aligned fibers In burst tests, scaffolds composed of aligned fibers produced via gradient electrospinning demonstrate superior material properties compared with scaffolds made from random or aligned fibers produced from a single polymer concentration or as bi-layered laminated structures.
Modern understanding describes the extracellular matrix (ECM) and cellular components of hollow organs—most notably blood vessels, cartilage, and skin—as organized into layered structures that are mechanically integrated to withstand routine physiological stresses In arteries, this is classically characterized by three layers—the tunica intima, tunica media, and tunica adventitia—though their exact structure and function are more intricate than the simple labels imply The tunica intima comprises a single layer of endothelial cells supported by a basement membrane, with an internal elastic lamina often present The tunica media contains alternating layers of smooth muscle cells surrounded by an ECM rich in proteoglycans, reticular fibers, and fibrils of type I collagen, with smooth muscle arranged radially around the lumen and also in a longitudinal spiral along the vessel Surrounding them, the tunica adventitia features thick collagen and elastic fibers, scattered fibroblasts, and a network of nerves and lymphatics These integrated components collectively determine the vessel's mechanical properties and functional responses.
Although each layer of a prototypical blood vessel has distinct structural and material properties, the layers are mechanically integrated and function as a unified tissue The transmission of mechanical stresses across the boundaries between layers with different properties is critical to maintaining the vessel’s structural integrity and coordinated function.
Without true integration, mechanical stresses accumulate at the interfaces between layers, increasing the risk of delamination where the layers separate at the boundary Proper integration allows these stresses to cross the boundary and dissipate into the adjacent material, reducing boundary failure A key driver for the future success of tissue‑engineered materials is the development of scaffolds that replicate both the architecture and function of native target tissue Although this approach is conceptually sound, efforts to fully recapitulate an organ’s specific structural elements and functional properties with tissue engineering have largely fallen short.
Electrospinning remains a versatile processing strategy for producing physiologically relevant tissue engineering scaffolds This adaptable technique can selectively process native, synthetic, and blended polymers into nano- to micron-scale fibers that mimic native ECM dimensions Early vascular constructs produced by electrospinning used uniform fibers engineered to withstand physiological mechanical loads, while later designs feature multiple fiber types deposited in discrete layers to better replicate the structure of native vessels By enabling a tunica intima–like inner layer of very small diameter fibers, overcoated with larger diameter fibers, electrospinning facilitates the creation of composite tunica media and tunica adventitia The finer inner fibers provide adhesion sites along the luminal surface, whereas the larger fibers contribute mechanical stability to the engineered construct.
Unfortunately, this direct approach, and related methods designed to entangle fibers of different
39 compositions [17,50,107] produce a laminated structure with boundary layers that can be fabricated to resemble the architecture of the native extracellular matrix (ECM) From a structural perspective, this configuration mirrors native ECM, while functionally, under mechanical loading, stress concentrates at the interfaces between layers, making the device susceptible to multiphasic delamination failure [17].
As a central challenge in tissue engineering, researchers strive to develop scaffolds that replicate the mechanical and functional properties of the target tissue A clear strategy to reduce delamination in multi-layered constructs is to modulate the transitional properties of boundary domains The production of a continuous fiber gradient within an electrospun scaffold as a means to diminish the impact of boundary layers inherent to any composite has not been examined in depth Only limited experiments have targeted functionally graded tissue-engineering scaffolds with nanoparticle gradients, focusing on how nanoparticle distribution across the Z-axis influences the overall tensile strength under various processing conditions.
This study describes and characterizes gradient fiber electrospinning, a technique that enables the fabrication of layered electrospun scaffolds with a gradual, continuous transition in average fiber diameter along the Z-axis These diameter gradients diminish stress concentrations at the boundaries between different fiber types, improving scaffold mechanical integrity We prototype gradient electrospinning with two scaffold designs, and the first prototype demonstrates the feasibility of achieving smooth diameter transitions across the scaffold height.
Gradient-fiber electrospun polycaprolactone (PCL) scaffolds were produced in two prototypes, designated 22-gauge and 18-gauge The 22-gauge scaffold has one face composed of electrospun PCL fibers with a mean cross-sectional diameter of 0.17 ± 0.09 μm, increasing across the Z-axis to 0.78 ± 0.89 μm on the opposite face, while the 18-gauge scaffold shows a similar gradient, ranging from 0.24 ± 0.12 μm on one surface to 0.89 ± 0.96 μm on the other, with the total average fiber diameter across the 18-gauge scaffold being larger and exhibiting a steeper Z-axis gradient than the 22-gauge prototype These architectural features impart unique material properties to the constructs, and the gradient-fiber electrospun scaffolds were evaluated against pure fiber controls and laminated-structure scaffolds using computational fluid dynamics (fluid modeling), output polymer concentration versus time (experimental mixing characteristics), output fiber diameter versus time (experimental electrospinning characteristics), mechanical testing (tensile/burst, overall scaffold failure properties), and scanning electron microscopy (SEM) analysis of scaffolds before, during, and after mechanical testing to characterize failure modes.
All computer drawings and meshes were developed in Gambit (Version 2.4) Fluid models were analyzed in Fluent (Version 12.0) using 1,000 iterations or until convergence, with graphical representations prepared in Tecplot Gradient electrospinning was modeled using a 3 mL plastic BD syringe (ID = 0.876 cm) as a reservoir, with the intermediate disk containing a central port sized to the equivalent of a 22-gauge (ID = 0.413 mm) or an 18-gauge (ID = 0.838 mm) needle segment The high-concentration polycaprolactone (65,000 MW) solution (top reservoir) was modeled at 200 mg/mL with a viscosity of 1.11×10^7 kg/m·s and a density of 958 kg/m^3, while the low-concentration PCL solution (bottom reservoir) was modeled at 100 mg/mL with a viscosity of 4.6×10^6 kg/m·s and a density of 951 kg/m^3 Electrospinning outlet was an 18-gauge needle, and the model incorporated a mass flow rate of 8 mL/hr False colors represent the magnitude of fluid velocity, measured in mm/s.
All reagents were obtained from Sigma unless noted Polycaprolactone (PCL, 65,000 Da) was suspended in trifluoroethanol (TFE) at 100–200 mg/mL and electrospun Electrospinning syringes capped with an 18-gauge blunt-tipped needle were installed into a syringe pump, delivering solution at 8 mL/h into a static electric field of 18 kV (Spellmen) Spinning occurred across a 20 cm gap onto a grounded cylindrical metal target (11.75 cm long, 6.33 mm diameter) designed to rotate and translate laterally (4 cm/s over a 12 cm distance) to promote an even coating of polymer Scaffolds were collected at 700 rpm.
(random) or 7,000 rpm (aligned) Laminated scaffolds were produced by sequentially spinning
100 mg/mL PCL onto the target, this layer was overcoated with a second, separate layer of fibers spun from 200 mg/mL PCL
In conventional electrospinning, fiber diameter increases with polymer concentration By engineering a polymer concentration gradient within the electrospinning reservoir, a continuously varying polymer concentration can be delivered to the jet, enabling a gradient in fiber diameters along the produced material Traditional approaches to realize this gradient have relied on entangled or multiple electrospinning jets to create spatial variation in fiber size.