Two-dimensional (2D) cell culture is a traditional cell culture system for cell growth in a monolayer [1]. 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 [8]. 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 [9]. 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 [1]. The 3D cell culture methods can be divided into scaffold-based 3D culture and scaffold-free 3D culture methods [10]. Scaffold-free methods include hanging drop, microwell-based maturation, rotational, and manetbased formation methods [11]. 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)] [12]. 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)] [12]. 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 [15]. The limitation of this type of bioprinter is a low resolution and cell damage owing to shear damage caused by pressure or mechanical force [16]. Third, stereolithography (SLA) bioprinting uses light to crosslink light-sensitive bioinks in a reservoir using a layer-by-layer process [Fig. 1(c)] [12]. 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)] [16]. 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 [13]. 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)] [20]. DLP technology is similar to SLA technology [20]. 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 [23]. 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.

Figure 1. Schematic diagrams showing the 3D bioprinting modalities. (a) Inkjetbased bioprinting, (b) extrusion-based bioprinting, (c) SLA bioprinting, (d) laserassisted bioprinting, and (e) DLP-based bioprinting.

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 [24]. 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 [12]. 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 [12]. 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 [28]. Second, biomaterials with fugitive roles can be removed to form internal channels and voids [28]. Third, support biomaterials provide mechanical support to form complex structures [28]. Finally, biomaterials with functional roles provide biochemical, electrical, and mechanical signals to influence cellular behavior [28]. These roles can be facilitated through the combination of some polymers.

2.1. Natural polymers

The natural polymers for 3D cell culture have similar properties to those of human ECM to mimic bioactivity [12]. 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)] [30]. 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 [35]. 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)] [31]. To overcome this limitation, chitosan and other bioinks are used in combination to achieve a higher cell viability [12]. 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].

Figure 2. Chemical structures of natural polymers such as (a) alginate, (b) gelatin, (c) collagen, (d) chitosan, (e) agarose, and (f) HA.

2.2. Synthetic polymers

Synthetic polymers are excellent sources for bioink manufacturing because of their specific physical properties and superior deposition [12]. However, the challenging issues of synthetic polymers are uncontrollable degradation and poor biocompatibility [12]. 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)] [39]. However, it has the disadvantage of poor cell adhesion due to its hydrophobicity [40]. PVA is a water-soluble polymer that has good biocompatibility, biodegradability, thermal stability, chemical stability, nontoxicity, and swelling stability [Fig. 3(b)] [39]. 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)] [31]. Despite having these advantages, PCL has limitations such as low bioactivity, low encapsulation efficiency, and burst release because of its hydrophobic properties [31]. PEG is a hydrophilic polymer with biocompatible, biodegradable, low toxicity, nonimmunogenic, and flexible properties [Fig. 3(d)] [12]. 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)] [43]. However, the disadvantages of TPU are its low mechanical strength and weak shape fixity [44].

Figure 3. Chemical structures of synthetic polymers such as (a) PLA, (b) PVA, (c) PCL, (d) PEG, and (e) TPU.

3.1. Tissue regeneration

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 [45]. 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)] [46]. The hair follicle scaffold showed high cell viability and good activity in vitro, and hair growth was observed in mice implanted with the scaffold [46]. A skin model using DLP-based 3D printing to produce a GelMA/HA-NB/LAP hydrogel was printable and biocompatible [Fig. 4(b)] [45]. 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 [45]. Jamalpour et al. [47] developed a canine model for interfacial tissue regeneration using a hybrid gelatin/PCL membrane with degradation, mechanical stability, and biological activity in vitro and in vivo [Fig. 4(c)]. The fabrication of vascularized bone tissue based on a dual 3D printing technique with SLA and fused deposition modeling (FDM) systems was reported by Hann et al. [Fig. 4(d)] [48]. A bone scaffold was created with an FDM-printed PVA sacrificial template [48]. In the sacrificial PVA template, bone tissue and vessel channels with human bone marrow mesenchymal stem cells (hMSCs) and HUVECs were constructed by SLA using GelMA/poly(ethylene glycol) diacrylate [48]. The hMSC-HUVEC co-culture system promoted osteogenic maturation and vascular network formation [48]. Antich et al. [49] developed articular cartilage engineering constructs using a hydrogel composed of HA and alginate polymer coprinted with PLA [Fig. 4(e)]. Articular cartilage fabricated with HA-based bioink had superior mechanical properties, including gelling abilities, printability, and degradability, and enhanced chondrogenesis [49].

Figure 4. Tissue regeneration using 3D bioprinting technology. (a) 3D-bioprinted multilayer composite scaffolds and hair follicle-like structures in vitro and in vivo. Reproduced with the permission from [46], Copyright 2022, Elsevier. (b) Bioprinted skin implanted into a skin defect area. Reproduced with the permission from [45], Copyright 2020, Elsevier. (c) Images of 3D-printed PCL membranes with different porosities and aminolyzed 3D-printed membranes after gelatin attachment. Reproduced with the permission from [47], Copyright 2022, Elsevier. (d) Image of a fabricated 3D construct and vessel layer. Confocal microscopic fluorescence images of HUVECs in a blood vessel and hMSCs in bone. Reproduced with the permission from [48], Copyright 2021, Elsevier. (e) Scheme of a 3D bioprinting process in articular cartilage engineering. Quantitative analysis of type II collagen, hyaline-specific chondrogenic marker genes, fibrotic markers, and hypertrophic markers in a bioprinted 3D hybrid construct. Reproduced with the permission from [49], Copyright 2020, Elsevier.

3.2. Drug evaluation

3D bioprinting technologies have been widely applied in the development of in vitro tissue/organ models for drug screening. Nie et al. [50] classified in vitro drug screening methods using 3D bioprinting technology into three types: mini-tissue, tissue/organ construct, and organ-on-a-chip. Hong et al. [51] developed 3D bioprinted drugresistant breast cancer spheroids for the quantitative evaluation of the drug resistance of cancer cells (MCF-7 cells). The bioinks used to build the spheroid cells were gelatin and alginate [Fig. 5(a)]. Breast cancer spheroids embedded in a gelatin-alginate gel matrix showed superior resistance to antitumor agents, including camptothecin and paclitaxel [51]. Gebeyehu et al. [52] reported chemotherapeutic drug screening using polysaccharide-based inks (Ink H4-RGD) in negative breast cancer (MDA-MB-231WT) and lung adenocarcinoma (HCC- 827) [Fig. 5(b)]. The 3D spheroid cells showed higher resistance to docetaxel, doxorubicin, and erlotinib compared to the 2D culture system [52]. Patrick Ulrich N’deh et al. [53] demonstrated a drug effect in human hepatoma cells (HepG2 cells) and human keratinocyte cells (HaCaT cells) cultured on an acrylonitrile butadiene styrene (ABS) 3D scaffold coated with gold nanoparticles. The 3D cultured cells on the ABS 3D scaffold promoted cell proliferation and drug resistance [Fig. 5(c)]. Kim et al. [54] monitored the drug evaluation based on a multichannel cell chip by extrusion-based bioprinting technology using polydimethylsiloxane (PDMS), PLA, and PCL polymers [Fig. 5(d)]. Additionally, the cultured cells (liver cells, lung cells, human glioblastoma cells, and human cervical carcinoma) in a multichannel cell chip grew in a 3D formation and had enhanced drug resistance [54]. Zhang et al. [55] developed an endothelialized myocardium-on-a-chip for cardiovascular drug testing through 3D bioprinting technology [Fig. 5(e)]. Endothelialized myocardium-on-a-chip has the potential for personalized drug screening to mitigate drug-induced cardiovascular toxicity or improve treatment efficacy [56].

Figure 5. Drug evaluation using 3D bioprinting technology. (a) Schematic diagram showing the formation of bulk and drug-resistant cancer spheroids using 3D bioprinting and imaging analysis of embedded spheroids. Reproduced with the permission from [51], Copyright 2022, Elsevier. (b) Photographic image of a ten-layer printed scaffold with Ink H4-RGD. Reproduced with the permission from [52], Copyright 2021, Nature. (c) Cancer model HepG2 and keratinocyte HaCaT cells were cultured in 3D and 2D culture systems. Reproduced with the permission from [53], Copyright 2020, MDPI. (d). Schematic diagram of a polydimethylsiloxane multichannel cell chip. Reproduced with the permission from [54], Copyright 2021, MDPI. (e) Schematics showing the procedure of fabricating endothelialized myocardium using a 3D bioprinting strategy. Reproduced with the permission from [55], Copyright 2016, Elsevier.

3.3. Drug delivery

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 [56]. 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 [57]. Economidou et al. [58] used SLA technology to design a microneedle array using a biocompatible resin for insulin delivery [Fig. 6(a)]. Additionally, Farias et al. [59] reported methacrylate-based custom hollow microneedle assembly based on SLA technology using a cell hydrogel to evaluate the potential of HepG2 cells [Fig. 6(b)]. Dual drug delivery systems that can release cefazolin (CFZ) and rifampicin (RFP) were developed by Lee et al. [60] [Fig. 6(c)]. The dual drug-based scaffold was manufactured with a CFZ-containing polycaprolacone 3D scaffold and RFP-loaded alginate hydrogel 3D scaffold [60]. It inhibited biofilm formation and enhanced antimicrobial activity [60]. Mei et al. [61] reported degradable antitumor scaffolds for controllable drug delivery to various tumor cells (human lung cancer cells, human osteosarcoma cells, mouse embryo osteoblast precursor cells, and mouse breast cancer cells) and in vivo [Fig. 6(d)]. The 3D printed porous scaffold was fabricated with PLA materials and methotrexate (MTX), an antitumor drug [61]. The PLA/MTX scaffolds showed a high tumor inhibition rate and reduced drug toxicity compared to intraperitoneal injection [61]. Herrada-Manchon et al. [62] studied 3D printed gummies with different shapes for oral and personalized drug dosage [Fig. 6(e)]. The 3D printed gummy was designed with ranitidine, gelatin, and xanthan gum, and pigment and sweeteners were added to improve the aroma and appearance [62].

Figure 6. Drug delivery system using 3D bioprinting technology. (a) 3D printed microneedle patch using SLA for intradermal insulin delivery. Reproduced with the permission from [58], Copyright 2019, Elsevier. (b) 3D printed microneedle device for microencapsulated cell extrusion. Reproduced with the permission from [59], Copyright 2018, MDPI. (c) Cefazolin-PCL scaffold for dual drug delivery systems. Reproduced with the permission from [60], Copyright 2022, MDPI. (d) Controllable drug delivery with 3D-printed degradable antitumor scaffolds. Reproduced with the permission from [61], Copyright 2021, Whioce. (e) Gummy dosage fabricated with 3D printing technology and drug dissolution profiles of gummy dosages. Reproduced with the permission from [62], Copyright 2020, Elsevier.

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.