Understanding Electrospinning

Electrospinning is a versatile and highly controllable fiber fabrication technique that has become a cornerstone of tissue engineering, particularly for creating scaffolds that mimic the nanofibrous architecture of the native extracellular matrix (ECM). In the context of cartilage repair, electrospun scaffolds offer a promising platform to support chondrocyte function, guide matrix deposition, and ultimately restore joint function. The process fundamentally relies on a high-voltage electric field (typically 10–30 kV) applied between a conductive spinneret and a collector. When the electric force overcomes the surface tension of a polymer solution or melt, a charged jet is ejected, undergoes whipping instability due to electrostatic repulsion, and stretches into continuous fibers with diameters ranging from nanometers to a few micrometers. The solvent evaporates during flight (in solution electrospinning) or the melt solidifies (in melt electrospinning), leaving randomly oriented or aligned fibers deposited on the collector. This method allows precise, reproducible tuning of fiber diameter, alignment, porosity, and surface topography—all critical parameters for cartilage-mimicking scaffolds.

Unlike traditional scaffold fabrication techniques such as salt leaching, gas foaming, or solvent casting, which produce pores in the tens to hundreds of micrometers but lack continuous fiber topography, electrospinning generates a dense network of fibers that recapitulate the physical cues chondrocytes encounter in vivo. Cartilage itself is a dense, avascular tissue with a highly organized ECM dominated by collagen type II fibers (50–200 nm diameter) interwoven with proteoglycan aggregates. Electrospun scaffolds can approximate this hierarchical structure, providing both mechanical support and biological guidance for cell attachment, proliferation, and differentiation. Moreover, the high surface-area-to-volume ratio of nanofiber scaffolds enhances adsorption of proteins from culture media or the biological environment, promoting integrin-mediated cell signaling.

Beyond structural mimicry, electrospinning offers the ability to incorporate bioactive molecules directly into fibers, either by blending them with the polymer solution, through coaxial spinnerets (core–shell fibers), or via post-spinning functionalization. This versatility makes electrospinning an attractive platform for delivering growth factors, cytokines, or drugs in a controlled, sustained manner—a feature that is especially valuable for cartilage regeneration, where inflammation and degeneration are often comorbid. The remainder of this article will examine the materials commonly used, the major electrospinning techniques, scaffold design considerations, current applications, and the challenges that remain before clinical translation becomes routine.

Materials Used in Cartilage-mimicking Scaffolds

The choice of polymer(s) is perhaps the most critical decision in electrospinning for cartilage engineering. The materials must be biocompatible, biodegradable at a rate matching neotissue formation, processable via electrospinning, and able to support chondrogenic phenotypes. Both natural and synthetic polymers have been explored, often in blends to combine desirable properties. Below is a detailed discussion of the most commonly used materials.

Synthetic Polymers

  • Polycaprolactone (PCL): PCL is a semi-crystalline, biodegradable polyester with a slow degradation rate (2–4 years in vivo), excellent mechanical properties, and FDA approval for many medical devices. Its low melting point (~60 °C) and solubility in common organic solvents (e.g., dichloromethane, chloroform) make it easy to electrospin. However, PCL is hydrophobic and lacks cell-adhesion motifs, so it is often blended with natural polymers or surface-modified (e.g., plasma treatment, collagen coating) to improve chondrocyte attachment and spreading.
  • Poly(lactic-co-glycolic acid) (PLGA): PLGA is a copolymer of lactic and glycolic acids whose degradation rate can be tailored by the monomer ratio. Fibers degrade via hydrolysis, producing acidic byproducts that are generally well-tolerated. PLGA scaffolds have been shown to support chondrocyte proliferation and ECM synthesis. Like PCL, PLGA is hydrophobic and may require blending with hydrophilic materials or growth factors to enhance bioactivity.
  • Poly(lactic acid) (PLA) and poly(glycolic acid) (PGA): PLA (especially the L isomer, PLLA) and PGA have also been electrospun for cartilage applications. PLLA is more crystalline and degrades slowly, while PGA degrades rapidly. Their mechanical properties (especially stiffness) can influence chondrocyte phenotype; very stiff substrates may promote fibrocartilage rather than hyaline-like cartilage. Therefore, blending with softer components or adjusting fiber alignment can improve outcomes.

Natural Polymers

  • Gelatin: Gelatin is derived from collagen and retains many cell-binding motifs (e.g., RGD sequences). It is water-soluble, biodegradable, and relatively inexpensive. Gelatin can be electrospun from aqueous solutions or co-solvent systems, but its poor mechanical strength and high solubility in water at physiological temperatures often require crosslinking (e.g., glutaraldehyde, genipin, EDC/NHS) to maintain structural integrity. Crosslinked gelatin nanofiber scaffolds have been shown to support chondrocyte attachment, proliferation, and glycosaminoglycan (GAG) deposition.
  • Collagen type I and II: Collagen is the major ECM protein in cartilage (type II) and is an ideal scaffold material due to its natural bioactivity, biocompatibility, and weak immunogenicity. Electrospinning pure collagen requires specialized solvents (e.g., 1,1,1,3,3,3-hexafluoroisopropanol, HFIP) and careful control of pH and ionic strength. Crosslinking is also necessary to prevent rapid dissolution. Collagen scaffolds closely mimic the native ECM and promote chondrogenesis, but their processing is more challenging and expensive than synthetic polymers.
  • Chitosan: Chitosan is a cationic polysaccharide derived from chitin, with structural similarity to GAGs found in cartilage ECM. It is biodegradable, non-toxic, and has antimicrobial properties. Chitosan solutions are electrospinnable from aqueous acetic acid solutions, but its high viscosity and tendency to form polyelectrolyte complexes can make spinning difficult. Blending chitosan with PCL, gelatin, or poly(vinyl alcohol) improves processability. Chitosan-based scaffolds promote chondrocyte proliferation and maintain the chondrogenic phenotype, partly due to the positive charge that facilitates binding of negatively charged GAGs.
  • Hyaluronic acid (HA): HA is a non-sulfated GAG abundant in cartilage and synovial fluid, playing key roles in joint lubrication and cell signaling. HA is highly hydrophilic and anionic, making it challenging to electrospin as a pure polymer; it is usually blended with carrier polymers (e.g., PCL, gelatin) or chemically modified (e.g., methacrylation) to enable crosslinking and fiber formation. HA-containing scaffolds enhance chondrocyte migration and ECM deposition but require careful handling to avoid rapid dissolution.

Blends of synthetic and natural polymers are increasingly used to combine the mechanical robustness of synthetics with the biological cues of naturals. Common combinations include PCL/gelatin, PLGA/collagen, and PCL/chitosan. Such blends can achieve fiber diameters in the range of 200–1000 nm, porosities >80%, and tensile moduli comparable to native cartilage (0.5–2 MPa).

Electrospinning Techniques

The basic electrospinning setup has been refined into several variants, each offering unique advantages for cartilage scaffold engineering.

Solution Electrospinning

This is the most widely used method. The polymer is dissolved in a volatile organic solvent (or a water-based system for natural polymers) to form a solution with optimal viscosity, conductivity, and surface tension. Parameters such as flow rate (0.5–5 mL/h), applied voltage (10–25 kV), needle-to-collector distance (10–20 cm), and ambient conditions (temperature, humidity) must be optimized for each polymer–solvent system. Solution electrospinning can produce continuous fibers with diameters as small as tens of nanometers. However, the use of organic solvents (e.g., HFIP, chloroform, DMF) raises concerns about residual toxicity, requiring thorough vacuum drying or post-treatment to remove trace solvents before cell seeding. Additionally, volatile solvent evaporation can create a porous surface morphology that further increases surface area—a benefit for cell adhesion but a potential source of mechanical weakness if pores become interconnecting.

Melt Electrospinning

Melt electrospinning uses a polymer melt instead of a solution, eliminating the need for organic solvents and avoiding toxicity issues. This method is particularly attractive for polymers like PCL and PLA that have well-defined melting points. Melt electrospinning typically produces larger fiber diameters (usually >1 µm) compared to solution electrospinning, because the melt viscosity is higher and the jet does not undergo the same degree of whipping instability. Nevertheless, melt electrospinning offers advantages for scaling up production (no solvent recovery needed) and for creating thicker, more mechanically robust scaffolds. Recent advances in melt electrospinning writing (MEW) allow direct writing of fibers in a controlled pattern, enabling the fabrication of ordered scaffolds with defined pore geometry—a feature that is highly desirable for cartilage tissue engineering to facilitate nutrient diffusion and cell infiltration.

Coaxial Electrospinning

Coaxial electrospinning employs a concentric dual-needle setup: an inner needle carries a core solution (often containing growth factors, drugs, or a mechanically supportive polymer) while an outer needle carries a shell polymer solution. The high voltage causes both solutions to be ejected simultaneously, forming a core–shell fiber that encapsulates the core material. In cartilage engineering, coaxial electrospinning has been used to encapsulate transforming growth factor-β1 (TGF-β1) or insulin-like growth factor-1 (IGF-1) within a PCL or PLGA shell, allowing sustained release over days to weeks. This technique reduces the initial burst release compared to simple blending and protects bioactive molecules from denaturation during electrospinning. Core–shell fibers can also be designed to have a mechanically stiff core (e.g., PCL) and a softer, bioactive shell (e.g., gelatin or collagen), mimicking the composite structure of native ECM.

Multijet and Needleless Electrospinning

Traditional single-needle electrospinning has low throughput (≈0.1–1 mL/h), limiting its use for large-scale scaffold fabrication. Multijet systems using multiple needles or a rotating drum with multiple spinnerets can increase production rates. Needleless electrospinning, in which an electrode (e.g., a rotating wire or cylinder) picks up polymer solution from a bath and initiates multiple jets simultaneously, can achieve throughput orders of magnitude higher. Although these methods produce less uniform fiber diameters than single-needle spinning, they are being researched for commercial-scale production of cartilage scaffolds. For cartilage repair, where defects are often small (1–5 cm²) and patient-specific, the throughput may not be a limiting factor, but for future off-the-shelf products, scalability will become essential.

Aligned Fiber Collectors

The orientation of electrospun fibers significantly affects cell behavior. Random fibers produce isotropic mechanical properties and promote cell spreading in all directions, whereas aligned fibers induce contact guidance and anisotropic mechanical properties—important for mimicking the zonal organization of articular cartilage, where collagen fibers in the superficial zone are oriented parallel to the joint surface, in the middle zone are random, and in the deep zone are perpendicular to the subchondral bone. To produce aligned fibers, collectors can be designed as rapidly rotating drums (up to several thousand rpm), parallel electrode arrays, or rotating discs. Chondrocytes cultured on aligned fibers show increased collagen type II and aggrecan gene expression compared to those on random fibers, and the resulting scaffold can withstand higher tensile loads along the fiber direction. For full-thickness cartilage defects, multilayered scaffolds with different fiber orientations in each zone are being explored to reproduce the native anisotropy.

Designing Cartilage-mimicking Scaffolds

Designing an electrospun scaffold for cartilage requires a balance of structural, mechanical, and biological properties. The scaffold must support cell spreading, maintain a round morphology (typical of chondrocytes), promote ECM deposition, and resist mechanical loads from joint movement. Key design parameters include fiber diameter, alignment, porosity, pore size, degradation rate, and surface chemistry.

Mechanical Properties

Articular cartilage is a viscoelastic material with a compressive modulus of 0.5–2 MPa and a tensile modulus of 5–15 MPa. Electrospun scaffolds typically have a lower compressive modulus (0.05–0.5 MPa) due to their high porosity and small fiber diameter, but they can be strengthened by increasing fiber diameter, reducing porosity, or blending with stiffer polymers (e.g., adding nanocrystalline cellulose or hydroxyapatite). For cartilage, an ideal scaffold should be stiff enough to withstand joint loading but not so stiff that it prevents cell-mediated matrix remodeling. Dynamic compressive loading in bioreactors has been shown to upregulate chondrogenic markers, so scaffold mechanical properties also influence how cells respond to mechanical signals.

Porosity and Pore Size

High porosity (>80%) is necessary for nutrient diffusion and waste removal in an avascular tissue like cartilage. However, electrospun scaffolds often have small pores (a few micrometers) due to the dense packing of fibers, which can impede cell infiltration. Various strategies have been developed to increase pore size: incorporating sacrificial fibers (e.g., water-soluble PEO that can be leached out), cryogenic electrospinning (collecting fibers on a cold mandrel to form ice crystals that create macropores), or combining electrospinning with particulate leaching. Pore sizes in the range of 100–300 µm are considered optimal for chondrocyte ingrowth. Cell seeding efficiency can be improved by dynamic seeding methods (e.g., centrifugation or orbital shaker) or by creating macroporous architectures using 3D printing combined with electrospinning.

Degradation Rate

The scaffold should degrade at a rate that matches new ECM deposition—typically over several months for cartilage repair. Fast-degrading polymers like PGA (weeks to months) may lose mechanical integrity before sufficient matrix is produced. Slow-degrading polymers like PCL (years) may persist too long, hindering complete integration. Blends and copolymers (e.g., PLGA with tailored monomer ratio) allow fine-tuning of degradation kinetics. Enzyme-degradable crosslinks (e.g., matrix metalloproteinase-sensitive peptides) can also be incorporated to make degradation cell-responsive.

Surface Functionalization

To improve bioactivity, electrospun fibers can be coated with ECM proteins (collagen, fibronectin, laminin) or loaded with growth factors. Immobilization of TGF-β1 or bone morphogenetic protein-2 (BMP-2) on the fiber surface has been shown to enhance chondrogenesis of mesenchymal stem cells (MSCs). Additionally, incorporating adhesive peptides (RGD, YIGSR) can promote chondrocyte attachment without the risks associated with animal-derived proteins.

Applications and Current Research

Electrospun nanofiber scaffolds have been investigated for in vitro cartilage model systems, in vivo repair of osteochondral defects in animal models, and early-stage clinical trials. In in vitro studies, human chondrocytes have been seeded on electrospun PCL/gelatin scaffolds and maintained chondrogenic phenotype for up to 28 days, with increasing GAG and collagen type II content. Coaxial fibers releasing TGF-β1 have been shown to induce differentiation of human MSCs into chondrocyte-like cells without the need for external growth factor supplementation.

In vivo studies in rabbit, goat, and sheep models have shown promising results. For example, electrospun PCL/chitosan scaffolds implanted into rabbit knee osteochondral defects promoted fibrocartilage repair with better integration than untreated controls. In a goat model, bilayered scaffolds combining an electrospun PLGA layer (cartilage side) with a microporous bone graft substitute (bone side) achieved near-complete healing at 6 months. A few human pilot studies have used electrospun collagen or PLGA scaffolds for cartilage repair in combination with autologous chondrocytes or MSCs, with improvements in pain and function scores. However, long-term follow-up and randomized controlled trials are still lacking.

Beyond scaffold‐based repair, electrospun fibers are used in drug delivery for inflammation control (e.g., incorporating anti-inflammatory drugs like diclofenac) and as barrier membranes guided tissue regeneration or for joint lubrication (electrospun hyaluronic acid mats).

Challenges and Future Directions

Despite the tremendous progress, several challenges remain before electrospun cartilage scaffolds become a standard clinical option.

Cell Infiltration

The limited pore size of electrospun mats—often below 20 µm—restricts cell migration into the scaffold’s interior, leading to a rim of cells on the surface and a necrotic core. While techniques like salt leaching, cryogenic spinning, and hybrid 3D printing/electrospinning have improved infiltration, achieving uniform cell distribution throughout thick scaffolds (>2 mm) remains difficult. Future work may focus on dynamic culture systems (perfusion bioreactors) to enhance nutrient transport and on creating scaffolds with graded porosity that directs cellular colonization.

Vascularization and Anisotropy

Cartilage is avascular, so scaffolds do not require blood vessel ingrowth. However, for osteochondral defects (involving both cartilage and bone), the subchondral bone portion must undergo vascularization. Electrospun scaffolds for bilayer or gradient constructs need to be designed with pore sizes and growth factors (e.g., VEGF for the bone side) that support both chondrogenesis and osteogenesis. Additionally, mimicking the zonal structure of cartilage (superficial, middle, deep zones) with different fiber orientations and biochemical compositions is a current research focus: some groups are using electrospinning to produce layered mats or 3D-printed templates that guide fiber deposition.

Scaling and Sterilization

Moving from lab-scale to industrial production while maintaining quality and reproducibility is a major engineering challenge. Electrospinning is sensitive to ambient conditions (humidity, temperature), polymer batch variability, and needle clogging. Needleless and multijet systems are being refined, but uniformity issues persist. Sterilization methods (ethylene oxide, gamma irradiation, e-beam) must not degrade the scaffold or residual solvents; supercritical CO₂ sterilization is emerging as a gentle alternative. Regulatory approval requires demonstration of biocompatibility, mechanical consistency, and lot-to-lot reproducibility.

Long-Term Stability and Integration

After implantation, the scaffold must degrade over time while the new tissue replaces it. However, degradation byproducts (e.g., lactic acid from PLGA) can lower local pH, potentially causing inflammation or calcification. Controlling degradation through polymer design and using basic salts to buffer acidity are strategies under investigation. Integration with host cartilage—a notoriously difficult problem—is enhanced by forming a seamless interface through in situ crosslinking or bioadhesive coatings.

Future Directions

Advanced electrospinning technologies such as near-field electrospinning (NFES) allow direct writing of fibers with micrometer precision, enabling patient-specific scaffold shapes that match a cartilage defect’s geometry (obtained from MRI or CT scans). Hybrid systems combining electrospinning with 3D bioprinting can produce scaffolds with both micro- and macro-scale hierarchies: a 3D-printed frame provides structural integrity and large pores, while electrospun nanofibers fill the interstices to provide cell-instructive topography. Smart scaffolds incorporating conductive polymers (e.g., polyaniline) for electrical stimulation—which promotes chondrogenic differentiation—and biodegradable electronics for real-time monitoring of tissue maturation are future possibilities. Clinical trials with electrospun cartilage scaffolds are anticipated to expand within the next five years, particularly for small focal defects in young patients. With continued material innovation and process control, electrospinning will undoubtedly play a key role in the next generation of cartilage regeneration therapies.

For further reading, consult recent reviews on electrospinning for cartilage tissue engineering (PubMed link), coaxial electrospinning for growth factor delivery (Biomedical Materials), and the role of fiber alignment in chondrogenesis (Acta Biomaterialia).