Three-Dimensional Printed Biomimetic Polymer Scaffolds for Wound Healing

2.1 Synthetic polymer

The superior biodegradability of synthetic polymers like PCL, PLA, and PGA makes them popular for use in tissue engineering and medication delivery, offering little toxicity and the capacity to create permeable scaffolds. PLA is FDA-approved, highly processable, and well-suited for bone tissue applications, particularly when combined with bioactive materials such as gelatin or hydroxyapatite. Because of its higher tensile strength as well as faster rate of degradation, PGA is frequently combined with PLA to customize its characteristics. Their copolymer, PLGA, provides tunable degradation profiles and high biocompatibility, making it highly suitable for biomedical applications. However, materials like pure PCL exhibit limited cell adhesion, necessitating surface modifications or blending strategies to improve cellular interactions [9].

2.2 Natural polymer

Natural polymers such as collagen, gelatin, chitosan, alginate, and plant-derived polysaccharides exhibit excellent biocompatibility and effectively simulate the ECM environment [17].

3.1 Biocompatibility

In the context of tissue engineering, a polymeric scaffold’s primary and most critical property is its biocompatibility. Cells must be able to adhere, migrate, and proliferate on the scaffold before depositing a new ECM. Following implantation, the scaffold or tissue-engineered construct should elicit only a minimal immune response to avoid severe inflammation that could impair healing or lead to rejection [28].

3.2 Biodegradability

The ultimate goal of tissue engineering is for the body’s own cells to gradually replace the implanted scaffold or tissue-engineered construct. Scaffolds are not intended to serve as permanent implants; instead, they must degrade naturally to allow cells to synthesize their own ECM [28]. The degradation byproducts should be biocompatible and safely eliminated from the body without causing harm to other organs. To ensure that scaffold degradation occurs in parallel with new tissue formation, a controlled immune response and the regulated recruitment of cells, such as macrophages, are required. With the increasing application of tissue engineering techniques in clinical practice, immunological considerations have become an essential focus of research [29, 30].

3.3 Mechanical property

It is essential that a scaffold is mechanically compatible with the tissue it is meant to repair, providing structural stability during implantation [28]. When repairing load-bearing structures like bone, the scaffold must have sufficient initial integrity to endure the full duration of tissue remodeling. Critically, this strength must be balanced with porosity, as an overly dense structure can impede vital vascularization and cell migration into the scaffold [31].

3.4 Scaffold architecture

Scaffold architecture plays a critical role in tissue engineering applications. High porosity and an interconnected pore network are essential scaffold properties that enable cell infiltration and proper nutrient diffusion throughout the construct and to the newly formed ECM. In addition, waste products must be able to diffuse outward through this porous structure, and scaffold degradation byproducts should be safely eliminated from the body without adversely affecting surrounding tissues or organs. One of the major challenges in tissue engineering is core degradation, which occurs due to insufficient vascularization and inadequate waste removal from the center of tissue-engineered constructs [32, 33].

3.5 Manufacturing technology

A key step toward making a polymer-scaffold-based construct viable for both medical and market success is the shift from lab-scale, single-sample production to cost-effective, small-batch manufacturing. The successful clinical translation of tissue engineering technologies is contingent upon establishing scalable manufacturing processes that comply with good manufacturing practice (GMP) guidelines [33].

4.1 Inkjet bioprinting

The earliest attempts to print living cells were carried out using modified commercial inkjet printers [39]. One of the first challenges of inkjet bioprinting was that printed cells dried out rapidly on the substrate, leading to cell death. Researchers overcame this limitation by embedding cells in water-rich polymer environments, giving rise to cell-loaded hydrogels that protected cell viability and supported tissue development [40]. Over the past few decades, advancements in inkjet printing have led to the development of several droplet generation and delivery methods. Among these, drop-on-demand (DOD), electrohydrodynamic (EHD), and continuous inkjet (CIJ) printing are the most widely recognized. While DOD and EHD techniques have already been successfully adapted for 3D bioprinting, CIJ printing has not yet been employed in this context. The following section focuses on the principles, advantages, and applications of DOD and EHD printing in 3D bioprinting [41–43].

DOD bioprinting can be actuated using thermal, piezoelectric, or acoustic mechanisms, each producing picoliter droplets with precise spatial control. In thermal DOD, vapor bubbles are generated to expel droplets, although transient heat exposure may affect cell viability. Piezoelectric DOD employs crystal deformation to generate pressure pulses with minimal thermal stress, while acoustic DOD uses pressure waves for gentle droplet ejection, albeit with trade-offs in printing speed and hardware complexity [44–47]. For stable jetting, bioinks typically require low viscosities (≈1–10 mPa s for DOD) and appropriate surface tension, with printability often assessed using the Z number (1/Oh), which falls within a well-defined practical window [48].

EHD bioprinting, by contrast, uses high electric fields to draw ultrafine jets, achieving sub-nozzle or even sub-micron resolution. This enables the precise patterning of skin cell micro-architectures but necessitates careful optimization to mitigate jet instability and field-induced stress on cells [49, 50]. Across both modalities, high cell viability depends on factors such as waveform tuning, nozzle design, humidity regulation, anti-clogging strategies, and bioink rheology optimization [51, 52].

In wound healing applications, DOD and EHD printing have been applied to deposit keratinocytes and fibroblasts within GelMA, alginate, hyaluronic acid (HA), and collagen-based hydrogels. These constructs are often co-loaded with growth factors or antimicrobials to enhance re-epithelialization, angiogenesis, and matrix organization in excision wound models [53–55].

4.2 Laser base bioprinting

Several droplet generation strategies have emerged from decades of inkjet printing advancements. The most recognized are DOD, EHD, and CIJ. To date, only DOD and EHD have been successfully integrated into 3D bioprinting. The working principles and applications of these two methods are highlighted below [41–43].

In DOD bioprinting, droplet ejection is achieved through thermal, piezoelectric, or acoustic actuation, each producing picoliter-scale droplets with precise spatial control. Thermal DOD relies on the rapid formation of vapor bubbles to expel droplets, piezoelectric DOD employs crystal deformation to generate pressure pulses, and acoustic DOD utilizes pressure waves for gentle droplet release. Each method offers distinct advantages and trade-offs in terms of cell viability, printing speed, and hardware complexity.

4.3 Laser-assisted bioprinting

Originally used for metal deposition, laser-assisted printing was later adapted into what is now known as laser-assisted bioprinting (LAB) [48]. This technique involves a laser pulse, a donor slide (or ribbon), and a receiver slide. The donor ribbon is specifically constructed with a transparent glass base, a thin metallic coating, and a layer of bioink on top. When struck by the laser pulse, the metal layer instantly vaporizes, creating the energy needed to propel the bioink onto the receiving slide.

This scaffold-free approach provides exceptionally high cell viability [49] and achieves spatial resolutions between 10–50 µm [39]. LAB also allows precise positioning of different cell types, with some studies reporting the ability to deposit even single cells per droplet. However, despite its remarkable accuracy, the technique is still costly and faces limitations in terms of long-term stability and large-scale application. Even so, LAB continues to show great promise, especially when integrated with other biofabrication methods [50–53].

4.4 Extrusion-based bioprinting

Driven by pneumatic or mechanical pressure [46, 55], extrusion bioprinting is a method where bioink is dispensed through a nozzle to form predefined structures. Its main strength lies in fabricating constructs with very high cell densities [54]. To ensure a successful outcome, it’s critical to assess process parameters like viscosity, nozzle diameter, and the consequent shear stress before printing begins [56].

Highly viscous hydrogels, which often do not require chemical additives for curing, are frequently employed in extrusion-based bioprinting [57]. Because hydrogels behave as non-Newtonian fluids, their viscosity changes with shear rate. High viscosity increases shear stress, which can accelerate cell death. In contrast, shear-thinning behavior is particularly advantageous as it enhances flow control, improves printing accuracy, and reduces viscosity under applied force. However, excessively low viscosities compromise structural integrity and print fidelity, even though they may facilitate improved cellular infiltration [58].

5.1 Freeze-drying

The technique of freeze-drying operates on the principle of sublimation. It begins when polymers and ceramics are dissolved in an appropriate solvent (like water or an organic one) and then emulsified with an aqueous phase. This solution is then cooled below its freezing point. Afterward, heat is applied to trigger sublimation, where the frozen solvent transitions directly from a solid into a gas, leaving behind a solid scaffold that is then dried [36, 61]. This process produces porous structures by removing the solvent and concentrating solutes within the ice phase. Freeze-drying is particularly advantageous because it enables the rapid fabrication of highly porous scaffolds without exposing biological components to high temperatures that could damage their activity [62–65]. For instance, Wu et al. [66] developed a porous gelatin scaffold using a unidirectional freeze-drying technique, which was shown to be non-toxic to cartilage cells and to promote extensive cell growth and spreading.

5.2 Electrospinning

As a powerful and flexible technique, electrospinning is used to produce 3D networks of nanofibers [67]. The process operates on the principle of a high-voltage electric field ejecting continuous, ultra-thin fibers when electrostatic forces overpower the polymer solution’s surface tension [68]. The process requires a syringe pump with a spinneret, a metal collector, and a high-voltage power supply. As the charged jet of polymer solution is ejected, the solvent evaporates during its trajectory, leaving behind a solid nonwoven fibrous membrane. Electrospinning setups can be classified as horizontal or vertical, depending on the orientation of the syringe and electric field relative to the ground.

In a study, Tan et al. [69] developed biodegradable nanofiber scaffolds made from cellulose acetate butyrate and PEG using electrospinning. The resulting fibers demonstrated improved tensile strength, strong cell adhesion, and excellent biocompatibility with human dermal fibroblasts, while remaining non-toxic. These findings highlight the potential of such nanofibers as promising scaffolds for tissue engineering applications [69].

5.3 Solvent casting and particulate leaching

Particulate leaching combined with solvent casting is a widely used technique for producing highly porous scaffolds with controllable pore sizes. Its main advantages are simplicity and the absence of a need for specialized equipment [70–72]. The process begins by creating a polymer solution, which is then blended with salt particles of a predefined size. Next, the solvent is evaporated, trapping the salt particles within the polymer. The resulting composite is then placed in distilled water to dissolve the salt, leaving behind a highly porous structure. The final scaffold has a pore size that directly corresponds to the size of the salt particles used.

5.4 Sol–gel methods

The sol–gel method is widely used for the inorganic polymerization of metal alkoxides. In this approach, a surfactant is added to form a colloidal solution (sol), which then undergoes condensation to produce a gel. Using a hybrid sol–gel process, Chen et al. [73] developed bioactive glass (BG) ceramics containing sodium oxide. These materials exhibited improved mechanical stiffness while retaining biodegradability. Importantly, the properties of the resulting scaffold are highly dependent on the procedure, as well as on the specific additives and processing parameters applied [72].

6.1 Bone tissue engineering

Bone’s structure is a natural composite, divided into organic and inorganic phases [75]. The organic phase is dominated by type I collagen, while the inorganic phase consists largely of calcium phosphates (~70%), which exist predominantly in the crystalline structure of hydroxyapatite (Ca₅(PO₄)₃(OH)) [76]. To effectively design scaffolds for bone tissue engineering (BTE), it’s crucial to have a foundational knowledge of the bone regeneration process. This natural healing mechanism is initiated in response to events like injury, trauma, or the surgical implantation of a BTE scaffold [75]. Bone formation, or osteogenesis, proceeds through two distinct pathways: endochondral ossification and intramembranous ossification [77].

6.2 Cartilage regeneration

The ECM of cartilage tissue relies on key components like glycosaminoglycans (GAGs) and proteoglycans (PGs) [78]. This specialized matrix serves two main functions: it provides a low-friction surface for joint articulation, and it absorbs mechanical shocks, protecting the underlying bone [79, 80]. The goal of cartilage tissue engineering (CTE) is to improve the body’s natural repair processes. This is often done by seeding scaffolds with cells like chondrocytes or mesenchymal stem cells (MSCs) and adding growth factors to promote cell adhesion, growth, and specialization [81–83]. A wide variety of biomaterials, from synthetic polymers like PEG, PLGA, and PCL to natural ones such as collagen, alginate, and hyaluronan, have been investigated. Furthermore, modern 3D bioprinting techniques now allow for the creation of patient-specific, anisotropic scaffolds, which can integrate more effectively with native cartilage tissue [78].

6.3 Skin and wound healing

Tissue engineering employs natural, synthetic, or semi-synthetic materials to mimic the ECM, thereby regulating collagen deposition, cytokine expression, signaling pathways, and tissue morphology. In doing so, it promotes angiogenesis, reduces inflammation, influences cell proliferation, and guides ECM remodeling [84–86]. Polymeric scaffolds, although non-pharmacological in nature, play a pivotal role in restoring, enhancing, or replacing the functions of damaged tissues and organs. In wound healing, they serve two primary functions. These biomaterials have a dual role: in one capacity, they provide a physical framework for endogenous cells to adhere, proliferate, and move, thereby enhancing tissue repair and wound closure; in another capacity, they act as a temporary protective barrier over the wound, creating an optimal environment to accelerate the body’s innate healing mechanisms [87].

6.4 Nerve regeneration

Bioengineered solutions are increasingly being explored for addressing clinical challenges such as the repair of peripheral nerve gaps. An effective nerve conduit must provide a microenvironment that supports cellular interactions and promotes axonal regeneration across the injured site [88]. For very small nerve injuries – where the gap is only a few millimeters – the recommended medical approach is tension-free end-to-end suturing, in which the severed nerve ends are stitched together without stretching or undue tension [89]. However, for larger gaps, scaffolds based on natural and synthetic polymers are being investigated to facilitate nerve tissue regeneration. To be effective, a scaffold should closely mimic the nanofibrous architecture of the native ECM, thereby stimulating the activity of surrounding cells and supporting functional nerve repair [90, 91]. Among the available fabrication methods, electrospinning has emerged as a particularly promising technique due to its ability to produce nanofibrous structures that closely replicate the ECM in both fiber size and spatial organization [92].

6.5 Cardiac and vascular application

Cardiovascular disease (CVD) is an umbrella term for a range of conditions affecting the heart and blood vessels. Of these disorders, coronary heart disease (CHD) is responsible for the majority of deaths [93, 94]. CHD occurs due to atherosclerosis, a process where fatty plaques build up inside the arteries, narrowing them and restricting blood supply to the heart, which can ultimately lead to ischemia. As an interventional strategy, stents are commonly deployed at lesion sites to restore perfusion in patients with 50%–70% stenosis [95]. However, earlier FDA-approved bioresorbable stents exhibited complications arising from their relatively high strut thickness, while permanent metallic stents are associated with long-term risks, such as the late catch-up phenomenon [93].

Polymer-based bioresorbable scaffolds (BRS) present a promising alternative to permanent implants, but their use is limited by lower mechanical strength compared to metallic scaffolds such as cobalt–chromium alloys or stainless steel. For example, poly(L-lactic acid) (PLLA)-based BRS have a tensile modulus nearly 100 times lower than that of metal stents. The mechanical behavior and crystallinity of PLA are largely influenced by the methyl side group present on its polymer backbone. PLA occurs in two main forms: poly(D,L-lactic acid) (PDLLA), which is amorphous, and PLLA, which is semi-crystalline with crystallinity levels reaching up to 70%. The higher crystallinity of PLLA contributes to its superior structural integrity and mechanical stability. In contrast, PGA, which lacks a methyl side group, allows tighter chain packing, resulting in a higher melting point, improved mechanical strength, and increased crystallinity. However, the absence of steric hindrance also makes PGA more susceptible to hydrolytic degradation [93, 96].

6.6 Drug delivery systems

A DDS refers to any platform designed to improve the safety and efficacy of therapeutics by enabling localized loading and controlled release of active compounds. Polymeric scaffolds are particularly attractive in this context, as they allow drugs to be encapsulated and released in a sustained manner at the target site. Controlled release minimizes adverse effects associated with fluctuating drug levels and repeated administrations of immediate-release formulations – an issue that becomes increasingly important as drug loading increases. Owing to their site-specific implantation, scaffolds provide a spatiotemporally regulated release profile that is superior to conventional delivery approaches [97, 98].

Beyond drug release, the physicochemical interactions between polymeric scaffolds and biological components promote tissue regeneration by facilitating cellular adhesion, proliferation, and integration with host tissue [99]. Importantly, scaffolds also act as reservoirs for growth factors, enabling the delivery of key signaling proteins from the transforming growth factor beta (TGF-β) superfamily, including differentiation, osteogenic, and angiogenic factors. Targeted delivery of these biomolecules accelerates regenerative processes and enhances functional tissue repair.