Applied Science and Convergence Technology 2023; 32(1): 1-6
Published online January 30, 2023
Copyright © The Korean Vacuum Society.
aInfectious Disease Research Center, Korea Research Institute of Bioscience and Biotechnology (KRIBB), Daejeon 34141, Republic of Korea
bSKKU Advanced Institute of Nanotechnology (SAINT), Sungkyunkwan University, Suwon 16419, Republic of Korea
cDepartment of Nano Engineering, Sungkyunkwan University, Suwon 16419, Republic of Korea
†These authors contributed equally to this work.
This is an Open Access article distributed under the terms of the Creative Commons Attribution Non-Commercial License(http://creativecommons.org/licenses/by-nc/3.0) which permits unrestricted non-commercial use, distribution, and reproduction in any medium, provided the original work is properly cited.
To overcome the limitation of two-dimensional cell culture not being able to mimic the in vivo microenvironment, three-dimensional (3D) bioprinting technology for 3D cell culture has emerged as an innovative culture platform. 3D bioprinting technologies can be divided into five types: inkjet-based bioprinting, extrusion-based bioprinting, stereolithography bioprinting, laser-assisted bioprinting and digital laser processing-based bioprinting technology. The 3D printing strategies achieved through a combination of these technologies can be applied to develop tissue regeneration, drug evaluation and drug delivery systems. In addition, the choice of cells and biomaterials is an important factor in fabricating tissue/organ models. Biomaterials for 3D bioprinting can be divided into natural polymers (alginate, gelatin, collagen, chitosan, agarose, and hyaluronic acid) and synthetic polymers (polylactic acid, polyvinyl alcohol, polycaprolactone, polyethylene oxide and thermoplastic polyurethane). Depending on the goals of 3D bioprinting experiments, biomaterials can be used alone or in combination with various polymers. 3D bioprinting technology has the potential to be applied for personalized medicine, precision medicine and the fabrication of artificial tissue/organs.
Keywords: 3D bioprinting, 3D cell culture, Tissue regeneration, Drug evaluation, Drug delivery
Two-dimensional (2D) cell culture is a traditional cell culture system for cell growth in a monolayer . The limitation of 2D cell culture with monolayer cell growth is that it cannot mimic the in vivo microenvironment since it lacks the structure, biological signals, physiology, and extracellular matrix (ECM) of living tissue [1–5]. Thus, a 2D culture system without ECM may cause abnormal protein expression and abnormal cell metabolism [6,7]. On the other hand, threedimensional (3D) cell culture systems have the potential to mimic in vivo conditions because these systems can facilitate the conditions of complex cell−cell and cell–ECM interactions . In addition, 3D cell culture has unique properties, such as specific function and cell growth via the 3D formation of cell aggregation, spheroids, and organoids . Furthermore, 3D cell culture systems can narrow the gap between cellbased methods and animal models for studying the development of novel drugs and the repair and replacement of organs . The 3D cell culture methods can be divided into scaffold-based 3D culture and scaffold-free 3D culture methods . Scaffold-free methods include hanging drop, microwell-based maturation, rotational, and manetbased formation methods . Scaffold-based 3D cell culture is commonly used in 3D bioprinting technology.
The main 3D bioprinting technologies can be classified into five types (Fig. 1). First, inkjet-based bioprinting uses printers known as droplet bioprinters, and it uses heating reservoirs or piezoelectric actuators to eject bioink drops [Fig. 1(a)] . The advantages of inkjetbased bioprinting are a high printing speed and low cost, but the disadvantage is the narrow range of printable biomaterials [13,14]. Second, extrusion-based bioprinting uses pneumatic pressure or mechanical tools, such as pistons or screws [Fig. 1(b)] . Extrusion-based bioprinting has many advantages, such as a high cell density, large-scale biomimetic structure and the use of a wide range of biomaterials, including natural polymers and synthetic polymers . The limitation of this type of bioprinter is a low resolution and cell damage owing to shear damage caused by pressure or mechanical force . Third, stereolithography (SLA) bioprinting uses light to crosslink light-sensitive bioinks in a reservoir using a layer-by-layer process [Fig. 1(c)] . SLA bioprinting technology can be used to fabricate 3D patterned scaffolds with micro- and nanosizes, but it requires high-cost equipment and materials [17,18]. Fourth, laser-assisted bioprinting involves a system comprising an energy absorbing layer, donor ribbon, and bioink layer [Fig. 1(d)] . This type of bioprinter can use bioink with a high viscosity and resolution but causes cell damage due to a high laser energy [12,19]. Moreover, it has the disadvantages of a high cost and difficulty in use . Finally, digital light processing (DLP)-based bioprinting is method in which a photopolymer is cured by light with a plane-by-plane pattern [Fig. 1(e)] . DLP technology is similar to SLA technology . However, the DLP printer can be used to fabricate tissue constructs faster than the SLA printer, and it can produce cell patterns with complex morphologies at a high resolution, such as those of capillaries and perusable blood vessels [21,22]. The limitation of the DLP printer is that it requires expensive equipment and materials, can use only photopolymers and produce materials with cell cytotoxicity due to uncured photoinitiators . These 3D bioprinters can be used alone or in combination to fabricate 3D cell culture systems for tissue regeneration, drug evaluation and drug delivery systems.
To fabricate functional tissue/organ engineering, the use of bioprinting technology with a suitable choice of biomaterials and cell sources is essential [13,24,25]. Several studies using 3D printing technology for tissue/organ engineering demonstrated the capability to both encapsulate cells directly within scaffolds to build a tissue construct and print scaffolds for cell seeding . In this review, we describe the characteristics of biomaterials (natural polymers and synthetic polymers) for the fabrication of tissue/organ models. Additionally, this review will focus on 3D bioprinting applications, including tissue regeneration, drug evaluation, and drug delivery using various 3D bioprinting strategies.
Biomaterials called bioinks are key elements for bioprinting . Biomaterials should meet several requirements, such as bioprintability, biocompatibility, and biodegradability, for tissue engineering [26,27]. The polymers used in biomaterials can be classified as natural polymers, synthetic polymers or combinations of both . The roles of biomaterials can be divided into four classes according to their properties. First, biomaterials with structural roles promote cell proliferation, cell adhesion and ECM mimicking . Second, biomaterials with fugitive roles can be removed to form internal channels and voids . Third, support biomaterials provide mechanical support to form complex structures . Finally, biomaterials with functional roles provide biochemical, electrical, and mechanical signals to influence cellular behavior . These roles can be facilitated through the combination of some polymers.
The natural polymers for 3D cell culture have similar properties to those of human ECM to mimic bioactivity . Natural polymers used as bioink sources include alginate, gelatin, collagen, chitosan, agarose, and hyaluronic acid (HA). Alginate extracted from brown seaweed has been widely used for biomedical applications due to its good biocompatibility, low cost, low toxicity, and nonimmnunogenicity [Fig. 2(a)] [29,30]. However, alginate has disadvantages, including the difficulty of maintaining long-term stability and cell attachment [29,31]. The properties of gelatin include biocompatibility, high cell adhesion, cell remodeling, and nonimmunogenicity, but pure gelatin cannot be used as bioink owing to its low viscosity and weak mechanical strength [Fig. 2(b)] . Collagen derived from animal tendons is the main component of the ECM in actual tissues or organs [Fig. 2(c)] [32,33]. The advantages of collagen are good biocompatibility, biodegradation, high cell growth, high cell adhesion, and low antigenicity [12,30,34]. However, pure collagen is difficult to print because of its low viscosity . Chitosan has superior biocompatibility, biodegradability, bioactivity, antibacterial, nonallergenicity, and cost effectiveness because it is used in tissue engineering, including bone, skin, and liver engineering [30,35]. Chitosan has poor mechanical properties and low cell attachment [Fig. 2(d)] . To overcome this limitation, chitosan and other bioinks are used in combination to achieve a higher cell viability . Agarose is derived from certain red seaweed, and it has an excellent mechanical strength, low price, weak cell adhesion, and brittleness in the solid-state [Fig. 2(e)] [36,37]. HA with superior biocompatibility, hydrophilicity, and excellent resistance to compressive force is used to form hydrogels, but HA hydrogels have the disadvantage of rapid degradation and poor mechanical properties in the physiological microenvironment [Fig. 2(f)] [1,38].
Synthetic polymers are excellent sources for bioink manufacturing because of their specific physical properties and superior deposition . However, the challenging issues of synthetic polymers are uncontrollable degradation and poor biocompatibility . Polylactic acid (PLA), polyvinyl alcohol (PVA), polycaprolactone (PCL), polyethylene oxide (PEG), and thermoplastic polyurethane (TPU) are widely used as synthetic polymers for tissue engineering. PLA is a biocompatible and biodegradable polyester-based polymer [Fig. 3(a)] . However, it has the disadvantage of poor cell adhesion due to its hydrophobicity . PVA is a water-soluble polymer that has good biocompatibility, biodegradability, thermal stability, chemical stability, nontoxicity, and swelling stability [Fig. 3(b)] . However, the disadvantages of PVA are poor affinity and low strength [40–42]. PCL has the advantages of good biocompatibility, biodegradability, flexibility, high permeability, nontoxicity, and low cost [Fig. 3(c)] . Despite having these advantages, PCL has limitations such as low bioactivity, low encapsulation efficiency, and burst release because of its hydrophobic properties . PEG is a hydrophilic polymer with biocompatible, biodegradable, low toxicity, nonimmunogenic, and flexible properties [Fig. 3(d)] . However, its disadvantage is low cell adhesion. The advantage of TPU is its superior mechanical properties, such as high elongation, excellent biocompatibility, high ductility, and good abrasion resistance for tissue engineering [Fig. 3(e)] . However, the disadvantages of TPU are its low mechanical strength and weak shape fixity .
Recently, 3D printing techniques have been widely applied in tissue engineering and regenerative medicine, such as in neural, hair follicle, bone, cartilage, and canine models . In this section, we present an overview of the printing in various tissue regeneration methods. The scaffolds of hybrid constructs using various polymers and combinations of various 3D printing technologies are essential for fabricating tissues or organs for tissue regeneration. Artificial hair follicle regeneration was facilitated with a multilayer composite scaffold based on human umbilical vein endothelial cells (HUVECs), dermal papilla cells (DPCs), epidermal cells (EPCs), and fibroblast cells (FBs) encapsulated into a gelatin/alginate hydrogel using rapid prototyping technology [Fig. 4(a)] . The hair follicle scaffold showed high cell viability and good activity in vitro, and hair growth was observed in mice implanted with the scaffold . A skin model using DLP-based 3D printing to produce a GelMA/HA-NB/LAP hydrogel was printable and biocompatible [Fig. 4(b)] . Human skin fibroblasts (HSFs) and HUVECs embedded into N-(2-aminoethyl)-4-(4-(hydroxymethyl)-2- methoxy-5-nitro-sophenoxy) butanamide (NB)-linked hyaluronic acid (HA-NB), gelatin methacrylate (GelMA), and photoinitiator lithium phenyl-2,4,6-trimethylbenzoylphosphinate (LAP) materials were implanted into skin defect rat and pig models and demonstrated skin regeneration . Jamalpour
3D bioprinting technologies have been widely applied in the development of in vitro tissue/organ models for drug screening. Nie
3D printed drug delivery systems have advantages, such as the ability to design customized drug products with high flexibility to select the shape, dose, and size of the dosage form to meet individual patient needs . Drug delivery systems designed by 3D printing technology can lead to personalized drug dosing, complex drug release profiles and personalized topical drug delivery as novel applications . Economidou
3D bioprinting technology has the ability to reconstruct complex structures that can be used to build tissue models using biomaterials and living cells. In addition, innovative 3D bioprinting technology can be applied in personalized medicine and precision medicine through the development of tissue regeneration, drug evaluation, and drug delivery systems. Despite the many recent achievements in 3D bioprinting, challenges remain regarding the cells, biomaterials, and 3D bioprinting technology used to construct tissue models. Additionally, it is difficult for current studies to totally recapitulate in vivo structural complexity and cellular organization. Furthermore, limitations of 3D bioprinting methods include technical challenges, such as biomaterial selection, cell deposition, and vascular network refinement. To overcome these problems, advances in 3D bioprinting need to be made to improve the printing resolution, printing speed, biocompatibility, development of novel biomaterials, in situ bioprinting, construction of functional structures, and scalability. Conclusively, the appropriate choice of combining bioprinting technology, biomaterials, and cell sources is essential for developing complex tissues, and establishing a physiologically relevant microenvironment with appropriate biological, chemical, and physical features remains the goal.
This research was supported by the National Research Council of Science & Technology (NST) grant by the Korea government (MSIT) (No. CAP22011-000).
The authors declare no conflicts of interest.