3D printing of polymeric tissue engineering scaffolds using open-source fused deposition modeling

  • Ayse Selcen Alagoz
  • Vasif HasirciEmail author


Open-source printing is a field where the cost of printing additive manufacturing products is cheaper due to more economical software and parts to construct a product including those of tissue engineering scaffolds. In this manuscript, fused deposition modeling (FDM) is used as the main avenue of open-source use in 3D printing of tissue engineering scaffolds. Additive manufacturing enables the researchers to build 3D products with interior and exterior architectures precisely defined and produced using open-access software which dictates the printer to print the models or the data obtained by various imaging techniques. In this way, implants to suit the dimensions and the mechanical and physicochemical properties needed for an artificial extracellular matrix can be produced. The main limitations are the limited number of printing materials and their unknown compositions which make their biocompatibility an issue. With the recent developments of in-house filament production, this limitation is also being overcome.


Open-source 3D printing Fused deposition modeling (FDM) Tissue engineering Polymeric filament Scaffold 

1 Introduction

The aim of tissue engineering is to prepare substitutes for lost organs or repair tissue damage caused by diseases, injuries, aging, and trauma [1]. It uses three main constituents which are cells, scaffolds, and bioactive agents. The function of the scaffolds is to provide cells with a surface for attachment and growth (proliferation) and also to guide the newly forming tissue to mimic the form of the defect to accelerate the rate of healing; in other words, the scaffold essentially serves as a temporary extracellular matrix (ECM) [2]. One of the main challenges in scaffold fabrication is meeting the specific physical requirements of each native tissue such as morphology, as well as chemistry and mechanical properties. Scaffolds with appropriate porosity and interconnectivity play an important role for tissue engineering in order to achieve cell migration, nutrient and metabolite transfer, and vascularization [3]. Although various conventional processing techniques including solvent casting, particulate leaching, electrospinning, lyophilization (freeze-drying), wet spinning, and gas foaming are used to produce scaffolds, morphological properties (pore size, pore interconnectivity, pore shape, porosity) and geometry of the scaffolds cannot be precisely controlled [4]. Rapid prototyping, a new method for biomaterials and tissue engineering field, overcomes these limitations by using three-dimensional computed tomography data of the defect site to create a desired shape and fabricate a case-specific product with controlled properties including form [5].

Rapid prototyping (RP), also known as solid freeform fabrication (SFF) or additive manufacturing (AM), has been extensively used until now in various fields including automotive, aerospace, and biomedical application to fabricate predefined, complex-shaped 3D (three dimensional) structures, customized products, end-use parts, prototype models, and tissue engineering [6]. Currently, various types of RP technologies are accessible in the market such as fused deposition modeling (FDM), stereolithography (SLA), selective laser sintering (SLS), laminated object manufacturing (LOM), and inkjet printing. Each of these techniques is capable of fabricating products that have complex external and internal shapes and geometries, although their working principles differ significantly from each other [7]. They also have differences regarding cost, materials used, size of workspace required, and the way the layers are built [6].

FDM technology is one of the most well-known, simple, and low-cost rapid prototyping techniques, and it is easily integrated with different CAD software that is being used with commercial 3D printers. In this method, a molten thermoplastic polymer is extruded in filament form guided by a computer program to fabricate layer-by-layer the desired architecture [8]. The high cost of the professional FDM system is the barrier for the producer to manufacture their own design. The development of the open-source 3D printer systems has made FDM an inexpensive approach in contrast to the professional alternatives and lowered the barrier against entry to the field for people who want to explore and develop new ideas regarding 3D printing. In this system, a model can be created by any 3D graphics software capable of exporting STL (stereolithography) files, while professional printers have their own closed software programs [9].

Fully interconnected porous structure is very important for tissue engineering (TE) applications because of tissue ingrowth into the scaffolds to form new tissue, and also for the diffusion of nutrient in and removal of waste product away from scaffolds. Besides, the sizes of the pores should be larger than 150 μm to promote vascularization [10]. FDM technology has the ability to print TE scaffolds with high porosity, controlled pore size, and pore interconnectivity. In addition, it can print very complex architectures including patient specificity [11]. The FDM devices are either more affordable open-source devices with a limited range of printing materials or professional devices which are more expensive, sophisticated scientific laboratory equipment. The aim of the current review is to present the advantages and disadvantages of open-source FDMs and their use in tissue engineering field. Thus, the price advantage made open-source FDMs more affordable and scientists started to use these systems in their tissue engineering applications [9].


Fused deposition modeling (FDM), also known as fused filament fabrication (FFF) or filament freeform fabrication (FFF), was developed by Scott Crump in 1988 and commercialized by Stratasys in 1991 [12]. Although FDM is one of the most widely utilized 3D printing technologies in different areas, the main limitation of industrial FDM is the high cost of hardware and software used to create the machine [13, 14]. For this reason, open-source 3D printers have been developed and marketed worldwide to decrease the cost of building FDM devices through widespread internet access and low-cost computer system integration [15]. These systems have open software providing free access to the community in contrast to industrial FDMs which have closed software and utilize open spool thermoplastic filaments rather than cartridges found in industrial FDMs.

In this system, the desired 3D model is initially created by open-source computer-aided design (CAD) software and saved in a stereolithography file format (STL) which describes the properties of the model but does not contain information regarding the printing parameters. For this reason, STL file is converted to G-code (geometric code) which contains a command to run the printer using Slicer software, an open-source software platform for medical image informatics, image processing, and 3D visualization. Then, G-code is transferred into SD card to be loaded into the 3D printer [16].

Open-source printers use thermoplastic polymers as continuous filaments with diameters of 1.75, 2.85, or 3.0 mm as the material to fabricate the 3D structure. In principle, molten filament heated inside the thermal nozzle is extruded onto the platform in a layer-by-layer fashion building up cross sections in z-direction until the whole structure is printed [17, 18]. The process is shown in Fig. 1. In this system, printing time is based on the size and shape of the structure. A small and thin object is printed in a matter of minutes while larger and more complex structures take more time [19].
Fig. 1

Scheme showing the 3D scaffold printing process

3 Polymeric materials used as filaments in open-source FDM 3D printers

Different types of thermoplastic polymers in filament form are used in some open-source FDM printers due to ease of processability, low equipment and material cost, and free software. Devices using polymeric filaments generally have a certain diameter (e.g., 1.75 mm) to fit the heated extruder head. A continuous flow of material without clogging the nozzle is needed during extrusion, and therefore, optimization of operation temperature, movement in x-y direction, filament extrusion rate, and base temperature is performed. A variety of commercial filaments such as poly(lactic acid) (PLA), acrylonitrile butadiene styrene (ABS), polycaprolactone (PCL), polyether ether ketone (PEEK), and polybutylene terephthalate (PBT) summarized in Table 1 are available in the market for 3D printing [22].
Table 1

Summary of commonly used thermoplastic filaments for 3D printing and their advantages and disadvantages






Poly(lactic acid)

Eco-friendly polymer

Low Tg

High elasticity

Easy processability


Lactic acid as a by-product


Low compressive strength

[20, 21]


Acrylonitrile butadiene styrene

Low Tg

Ease of processing

High mechanical strength

Small shrinkage ratio


Low cell integration

[22, 23, 24]



Low melting temperature

Thermal stability

Superior rheologic and viscoelastic features


Prolonged degradation

High stiffness




Polyether ether ketone


High mechanical strength

Low heat conductivity

Extremely high

Tm (350 °C)

[22, 26, 27]


Polybutylene terephthalate


High Tm (225 °C)


Tm, melting point; Tg, glass transition temperature

Poly(lactic acid) (PLA) is a biodegradable, thermoplastic, and eco-friendly polymer which can be derived from renewable resources such as sugar and starch or made from compounds of natural origin. These properties make it a great alternative to petroleum-based polymers [20]. PLA has been extensively used in 3D printing systems due to having a relatively low glass transition temperature (Tg) (where molecular chains gain some mobility but do not melt) in the range 55–65 °C. It has a melting point around 175 °C, high elasticity, and biocompatibility enabling its use in biomedical and biomaterials applications [29]. PLA can be easily processed into filaments to be utilized in 3D printing devices, and its printing temperature is higher than its melting temperature, generally in the range 200–230 °C. These temperatures vary to some extent based on the average molecular weight and crystallinity. It is immunologically inert and degrades within 6 months to 2 years depending on the molecular properties and also product dimensions, and the degradation by-product is lactic acid, a biological metabolite. The long degradation duration leads to a local accumulation of lactic acid, resulting in a localized decrease in pH (to about pH 5) that sometimes leads to inflammation and necrosis of the surrounding tissue. Other limitations of PLA are that it is prone to be brittle and has low mechanical properties (e.g., compressive strength) compared with the bone and this is a handicap for its use in hard tissue engineering applications. In order to overcome or decrease these limitations, PLA can be blended with other polymers or with ceramics which results in enhanced compressive strength and mineralization after implantation [21].

Acrylonitrile butadiene styrene (ABS) is another thermoplastic copolymer and is composed of acrylonitrile, 1,3-butadiene, and styrene units. This triblock copolymer of petrochemical origin is widely used as a filament material in open-source 3D printers due to its low glass transition temperature and ease of processing. The melting point of ABS is 105 °C, thus it is compatible with the heating capacity of FDM-type 3D printers. ABS derives its high mechanical strength from the acrylonitrile and butadiene units, and toughness from styrene units [23]. ABS is an amorphous polymer; therefore, the shrinkage ratio of ABS is very small during solidification and this leads to high accuracy and geometrical stability of the fabricated structure. All these make ABS an excellent material for 3D printing of tissue engineering scaffolds [24]. Nevertheless, ABS is not preferred for the production of tissue engineering scaffolds because it is non-biodegradable and has low cell integration [22].

In open-source FDM system, polycaprolactone (PCL) filaments are preferred synthetic polymer to produce scaffolds due to its low melting temperature (58–60 °C), thermal stability, superior rheological and viscoelastic features, prolonged degradation, and biocompatibility. Printed PCL is a good product for bone tissue engineering due to its long-term degradation (around 2 years) and high stiffness. However, changes in crystallinity of PCL in accordance with temperature can directly influence the properties of the scaffold such as degradation rate. In addition, percent crystallinity and hydrophobicity of the polymer can influence cell-scaffold interactions and cell spreading [25].

Polyether ether ketone (PEEK), a semicrystalline thermoplastic polymer developed by Imperial Chemical Industries (ICI) in 1977, is another 3D printing material [26]. PEEK has excellent properties including cell compatibility, inertness, high mechanical strength, low heat conductivity, and radiolucency allowing radiographic assessment when implanted. Also, the modulus of elasticity of PEEK is close to that of the cortical bone. These features resulted in PEEK becoming a widely used polymer in orthopedic applications [27]. The main limitation of PEEK, however, is its extremely high melting point (350 °C) compared with other polymers. For this reason, it is generally processed with selective laser sintering (SLS) method. Recent advances in open-source 3D printers overcame this problem with the development of devices that can process materials using high temperatures leading to the use of PEEK filaments [22].

Polybutylene terephthalate (PBT) is another thermoplastic polyester used in 3D printing systems. However, its high melting point (225 °C) is a limitation during processing compared with PCL and PLA in spite of its biocompatibility. Up to now, PBT filament was used in only one study that produced a 3D-printed scaffold for bone tissue engineering [28].

4 Results

4.1 Comparison of open-source FDM 3D printers

Selection of the 3D printing method and device is an important aspect of every 3D printing application as the features they provide should meet the requirements of the said application. In this regard, open-source FDM 3D printers present some advantages compared with their industrial counterparts while falling short in some other aspects.

Lately, the development of commercial 3D printers has made FDM method inexpensive (average cost of a printer is in the range $500–2000) due to the open-source software and structural components (gears, pulleys, and any other part) which can be produced by the printers themselves because of their self-replicating properties. In this technology, open-source software provides easy accessibility advantage for users (many hobbyists, inventors, and researchers) by creating a community that shares models and designs with each other worldwide [30, 31]. Besides, these systems are compact in size and portable compared with the professional printers while also being easy to operate and safe to use when unsupervised.

The disadvantage of open-source FDM printers is that they generally have lower resolution than industrial printers due to still in the early stage of development, which allows fabrication of only coarse structures. Besides, printing of complex structures can take more time due to low printing speed and needs supervision to oversee the printing process to prevent or resolve any printing errors [32]. Another disadvantage is the availability of thermoplastic materials with optimal melt viscosity, the need for optimum viscosity to print a 3D form by extrusion. In addition, these systems are not able to incorporate living cells or biological agents because of high extrusion temperature they would destroy their delicate organization and 3D forms [33]. The absence of knowledge about the chemical composition and physicochemical properties of the ingredients prevent their use as biomaterials (e.g., scaffolds) because of biocompatibility issues. Cytotoxicity and extensive biocompatibility tests are necessary for these materials prior to use in any in vivo tissue engineering application. Besides, although composite materials with osteoinduction, osteoconduction, and osteointegration capabilities and high mechanical strength are required for load-bearing hard tissue applications (e.g., bone, cartilage), only polymeric materials are available on the market for 3D printers. These days, researchers are producing their own filaments using blends of polymer and ceramic powders (e.g., PCL and HAP) using extruders to overcome these limitations [34].

In summary, the advantages of the printers are their low cost, supervision-free operation, compact size, and low working temperatures while disadvantages are lower resolution, the narrow range of available materials, and unknown material compositions.

4.2 Open-source FDM 3D printers in tissue engineering applications

Although professional FDM 3D printers are frequently used in biomedical industries, their price is a prohibitive factor. Open-source FDM printers overcome this limitation and thus becoming affordable to wider medical researchers which have recently adopted 3D printing methods for the fabrication of tissue engineering products where extreme control over the shape and architecture is utmost importance for cell growth and tissue regeneration.

FDM is increasingly used in printing tissue engineering scaffolds. The products are mainly biodegradable and stiff materials like those for hard tissue implants such as those of bone substitutes. The material form is what distinguishes between the professional FDM and open-source FDM. In the open-source case, the main material feed is in the form of filaments as was in the original stages of its invention, and in the professional FDM, the material can be both in the form of filaments and as a powdered polymer which is melted and then extruded in the form of fibers onto the platform. Thus, the main difference is in the flexibility of the feed material choice. In both cases, the product is extruded in fiber form after an initial melting step. Open-source FDM technology has great potential in producing scaffolds for tissue engineering applications. Suitably designed internal (pore size, porosity, and pore interconnectivity) and external architecture (complex-shaped geometry) of the scaffolds are essential for cell growth, and therefore, tissue regeneration. In neither case, cells can be included in the extruded fiber due to its high temperature. However, industrial systems allow the choice of polymers from a limitless variety of blends and thus choice biocompatible polymers with engineered properties are possible whereas in the open-source the filaments are those provided by the companies unless if a tissue engineer produces the filaments himself. This limits the properties of the finished product. The mechanical properties of the scaffold should match that at the defect site so that the product does not collapse during the healing process. FDM technology allows controlling scaffold properties by extruding molten polymer they produce in-house or commercially available with layer-by-layer deposition guided by the computer program containing the data of the defective site. Thus, cell seeding is possible only after cooling of the printed scaffold and thus there is no difference in terms of the printed product serving as a cell carrier. The professional/industrial FDMs are either commercially available at a much higher cost due to proprietary hardware and firmware or are custom-made, which limit the economic feasibility of replication of three-dimensional printing (3DP) methods across the tissue engineering applications.

The architectural design of the printed scaffolds, including parameters fiber diameter, pore size, porosity, and permeability, has an important role in tissue regeneration. These morphological properties, in addition to providing appropriate cell adhesion sites, affect the mechanical characteristics of the structures. Trachtenberg et al. have studied the effect of processing parameters like printing speed, fiber spacing, and extrusion rate on PCL scaffolds printed by an open-source FDM 3D printer (RepRap Mendel 3D printer) (Fig. 2). The properties studied were mainly morphological (fiber diameter, pore size, and porosity) and mechanical. While the porosity of the scaffolds increased with increasing printing speed, the fiber diameter decreased for the same fiber spacing. Also, pore size decreased with increased pressure, which caused both a decrease in porosity and an increase in fiber diameter. Printed scaffolds with lower porosity exhibited higher compressive modulus and yield strength [35].
Fig. 2

Programmed fiber spacing (s, mm), printing speed (F, mm/min), and operating pressure (P, psi) are specified in the Python code to print 0° and 90° PCL layers. One 0° and one 90° layer are considered to be a complete grid with square pores. Pore size (dp, mm) and fiber diameter (df, mm) can be measured using optical microscopy and later used to calculate the experimental fiber spacing to compare with its corresponding programmed value (s). (Reprinted with permission from Trachtenberg et al. (2014), Copyright 2014 John Wiley & Sons)

In another study, Temple et al. used an open-source custom-built 3D printer to produce PCL scaffolds with varying porosities for patient-specific vascularized craniomaxillofacial bone tissue engineering which were tested both in vitro and in vivo. Scaffolds with a lay-down pattern of 0/90/45/135° (angles between subsequent layers) and infill density (amount of print material within a product) ranging from 20% (highly porous, larger pores) to 80% (nearly impermeable, smaller pores) were fabricated and then filled with a solution of fibrinogen and thrombin mixed with aggregated human adipose-derived stem cells (hASCs). Upon polymerization of fibrinogen to fibrin, cell aggregates were entrapped in the pores of the scaffolds. Results showed that cell aggregates were uniformly distributed on scaffolds with 40% infill density, while they accumulated at the bottom of the structure for scaffolds with larger pore sizes (20–30% infill) and clogged the pores of scaffolds with smaller pore sizes (50–80% infill). Before implantation, scaffolds with hASCs were incubated in a vascularization medium for 18 days and then implanted subcutaneously in nude rats. After a week in in vivo, the center of the cell-seeded scaffolds exhibited greater host cell infiltration and vascularization compared with the unseeded scaffolds. Finally, large anatomically shaped mandibular and maxillary scaffolds with 40% infill density were printed using patient’s microCT data (Fig. 3) [36].
Fig. 3

Anatomically shaped scaffolds. Left: 3D models of the maxilla (top) and mandible (bottom). Right: 3D printed porous PCL scaffolds at 40% infill density. (Reprinted with permission from Temple et al. (2014), Copyright 2014 John Wiley & Sons)

Liu et al. printed screw-like macroporous orthogonal PLA scaffolds with a desktop 3D printer (D3DP) (Dot Go 3D Technology Corporation) and tested them in vitro and in vivo for anterior cruciate ligament (ACL) reconstruction (Fig. 4). In order to enhance osteoconductivity and cell adhesion, surfaces of the scaffolds were coated with hydroxyapatite (HA) by dip-coating technique after printing. PLA and PLA/HA scaffolds were then seeded with marrow mesenchymal stem cells (MSCs) derived from the bone marrow of New Zealand white rabbits. Another set of PLA/HA scaffold was seeded with MSCs suspended in Pluronic F-127. Results showed that the PLA/HA scaffolds seeded with MSCs suspended in Pluronic F-127 demonstrated the highest cell proliferation and osteogenesis in vitro on day 7. PLA, PLA/HA, and PLA/HA seeded with MSCs suspended in Pluronic F-12 were implanted into the femoral tunnel of rabbits for in vivo testing. After 4 and 12 weeks in in vivo, PLA/HA scaffolds treated with MSC carrying Pluronic F-127 exhibited higher bone infiltration and bone-graft interface formation within femoral tunnel compared with the PLA and PLA/HA scaffolds. These results showed that surgical implant fabrication in the clinic with D3DPs can be practical, efficient, and cost-effective [37].
Fig. 4

Schematic diagrams of the implant and tendon graft within the bone tunnel in ACL reconstruction. a The 3D perspective of the bone tunnel in ACL reconstruction. b The transverse section view of the bone tunnel (G: graft; S: screw-like scaffold; BT: bone tunnel; M: macropore). c 3D view of the theoretical design of PLA-based screw-like scaffold structure. d The PLA screw-like scaffold printed. e SEM image of the PLA scaffold surface with well-defined orthogonal structure. (Adapted from Liu et al. (2015), licensed under CC BY 4.0)

Ciurana et al. employed the open-source 3D printer called RepRap in fabricating of PLA scaffolds with three different lay-down patterns (0/90°, 45/135°, and 0/90/45/135°) (Fig. 5), fiber distance, and slenderness ratio (height to base ratio).
Fig. 5

Lay-down patterns. a 0/90°. b 45/135°. c 0/45°/90°/135°. (Adapted with permission from de Ciurana et al. (2013). Copyright 2019 Elsevier B.V)

Results showed that the lay-down pattern influenced porosity, and higher porosity was observed with 0/90° pattern compared with others. Also, fiber distance significantly increased the porosity of the structure which in turn adversely affected Young’s modulus. Finally, higher slenderness ratio (greater height, smaller base) was found to enhance Young’s modulus of the structures [38].

In another study, Kosorn et al. investigated PCL/PHBV blend scaffolds produced by an in-house built 3D printer for cartilage regeneration (Fig. 6). PCL/PHBV filaments with different weight ratios (100/0, 75/25, 50/50, and 25/75) were fabricated using an extruder, and then honeycomb-shaped scaffolds (0/90/45/135° lay-down pattern) were printed. The average porosity of the PCL/PHBV:100/0 and PCL/PHBV:50/50 were measured as 60 ± 2% and 51 ± 5% respectively. Compressive strength of the scaffolds was increased with increasing PHBV content of the filaments. In addition, a high rate of chondrocyte cell proliferation and glycosaminoglycan (GAG) secretion was observed on scaffolds containing higher PHBV fractions. All results showed that printed PCL/PHBV:50/50 scaffolds could be used for cartilage tissue engineering [39].
Fig. 6

Microstructures of PCL/PHBV:100/0 and 50/50 scaffolds viewed with SEM (a, b, e, f) and μ-CT (c, d, g, h). (Reprinted with permission from Kosorn et al. (2016) Copyright 2016 John Wiley & Sons)

Ritz et al. fabricated porous PLA cage-like scaffolds as well as solid discs filled or coated with collagen using an inexpensive 3D printer (Ultimaker 2+) in order to test the biocompatibility of the scaffolds with different cell types (osteoblasts, fibroblasts, and endothelial cells) and endotoxin contamination. All cell types proliferated on PLA solid discs with or without coating of collagen at 10 days which demonstrated printed PLA discs are biocompatible. Also, porous cage-like scaffolds were loaded with stromal-derived factor (SDF-1) immobilized in collagen in order to promote neovascularization. SDF-1 released steadily from scaffolds which supported endothelial cell growth and promoted neo-vessel formation. All results indicated that printed PLA scaffolds have potential in biomedical applications, especially bone tissue engineering [40] (Table 2).
Table 2

Application of open-source 3D printers in tissue engineering

Printing filament material

Target tissue

In vitro study

In vivo study

3D printer









Subcutaneous implantation in nude rat




Poly(lactic acid)


Poly(lactic acid)/


Anterior cruciate ligament


Femoral tunnel in New Zealand rabbit














Osteoblasts, fibroblasts, and endothelial cells

Ultimaker 2+








(Beta-tricalcium phosphate/alginate)


Human osteoblast (CRL-11372)



TCP/AA/Graphene oxide


Human osteoblast (hOB) cells







(Methacrylated gelatin)


Human osteosarcoma cell

MakerBot Replicator 2


Gremare et al. employed a custom-made 3D printer developed by Technoshop and printed PLA mesh scaffolds with different pore sizes (150, 200, and 250 μm) to determine the influence of design on biological and physiochemical properties of the tissue engineered bone. Pore sizes of printed scaffolds were slightly smaller than the designed. Although the pore sizes of the three scaffold types designed were different, all showed similar ultimate tensile strengths. They did not display any cytotoxic effect for human bone marrow stromal cells (hBMSC) homogenously distributed on the scaffolds regardless of pore sizes and were suitable for bone tissue engineering applications [41].

Diogo et al. produced beta-tricalcium phosphate (β-TCP)/alginate-blended scaffolds using an open-source 3D plotter, personal fabricator called Fab@home, in bone tissue engineering. Printing parameters were selected to fabricate scaffolds with β-TCP to alginate ratios of 50/50, 30/70, and 20/80% (w/w) with predefined mesh-like architecture. Results demonstrated that blend viscosity directly affected the precision and resolution of the scaffolds; higher viscosity (50/50) achieved higher precision in the printing of the scaffold. 50/50 scaffolds had higher Young’s modulus and similar compressive load than the native trabecular bone. Osteoblast attachment and proliferation were also higher on this scaffold [42]. In other studies of this tricalcium phosphate (TCP)/alginic acid (AA) 60/40 scaffolds (%), mimicking the ratio of inorganic/organic phases found in native bone, graphene oxide (GO) was added and extruded with the same printer in order to increase their mechanical and biological properties. Results showed that TCP/AA scaffolds functionalized with GO have higher swelling, porosity, and mechanical performance than TCP/AA scaffolds. Besides, existence of GO in the structures seeded with human osteoblast (hOB) cells, even at a low concentration, demonstrated a remarkable effect on the mineral absorption and deposition on the surface of the scaffolds [43].

In another study, Albrecht et al. printed 3D porous 0/90° lay-down–patterned scaffolds utilizing open-source MakerBot Replicator 2 printer to treat complex bone defects. Initially, filaments of PCL/HAP and PCL/PLA blends with diameters of 1.75 mm were fabricated by a compounding extruder, wound onto a spool to be fed to the 3D printer. Then the scaffolds were printed and incorporated with human osteosarcoma–loaded GelMA. Mechanical properties of unseeded scaffolds were determined, and results showed that the compressive moduli of PCL-PLA (159.2 MPa) were higher than PCL (44.3 MPa) and PCL-HAP (59.3 MPa). This work could be potentially useful in the treatment of patients with complex bone defects. Similar cell proliferation results were observed on the scaffolds after 5 days [34].

5 Conclusions

The 3D printing field is used in tissue engineering of mainly hard tissues like bone and cartilage simply because thermoplastic materials are extruded in a molten form to construct a scaffold. Its main advantage is the free nature of the software and the economical cost of the printed material which enabled even the laymen to get acquainted with the field. Open-source technology permits any user to build their own design and to easily access community documentation shared by users. However, the disadvantage of being restricted to use only commercial filaments with the composition of which being unknown, and therefore the questionable suitability for tissue engineering, was a limiting factor for biomedical devices to be implanted in the human body. Another disadvantage of the system is resolution which is influenced by certain parameters that lead to the low quality of the scaffolds. These parameters are polymeric materials that retain heat after printing which may result in uncontrolled deformations or undesirable fibers due to the movement of the nozzle from last depositing point to the next, which cause to discontinuous scaffolds. Besides, FDM compared with other rapid prototyping techniques such as stereolithography (SLA) only fabricates scaffolds with restricted shape and comparatively regular structure due to the resolution of the technique as the viscosity of the molten polymer causes a barrier in the achievable resolution of printing. Also, materials in the range of proper properties using FDM such as appropriate viscoelasticity, thermoplasticity, processing the capability of the polymer as a filament, and melting/solidification character are limited. Many efforts are required to overcome these barriers to fabricate proper scaffolds for clinical applications. On the other hand, FDM system allows producing customized implants for clinical applications. For example, standard implants designed for joint arthroplasty do not fit the defect site tightly, requiring specific implant dimensions and properties in order for perfect function. FDM system overcomes this problem and can help manufacture patient-specific implants by using CT data obtained from the patient’s defect pattern. Also, printed scaffolds which have predetermined internal and external architecture can perfectly match the defect site. However, many 3D printed scaffolds have lower mechanical properties than human cortical bone (120–450 MPa) and this restricts the use of the printed scaffolds at load-bearing sites in clinical applications.

Nowadays, with the in-house production of filaments from known materials for the targeted tissue, the use of open-source production is expanding and making tissue engineering without prohibitive costs possible. Also, the development of simple open-source gel printers using bioinks made according to the needs of the cells and tissues is about to make open-source-printed soft tissues similarly available to low-budget researchers. All these will benefit the biomaterials and the regenerative medicine fields and make them more productive.


Compliance with ethical standards

Conflict of interest

The authors declare that they have no conflict of interest.


  1. 1.
    R. Langer, J.P. Vacanti, Tissue engineering. Science. 260, 920–926 (1993)CrossRefGoogle Scholar
  2. 2.
    F. Asghari, M. Samiei, K. Adibkia, A. Akbarzadeh, S. Davaran, Biodegradable and biocompatible polymers for tissue engineering application: a review. Artif. Cells Nanomed. Biotechnol. 45(2), 185–192 (2017)CrossRefGoogle Scholar
  3. 3.
    C. Zhou, K. Yang, K. Wang, X. Pei, Z. Dong, Y. Hong, X. Zhang, Combination of fused deposition modeling and gas foaming technique to fabricated hierarchical macro/microporous polymer scaffolds. Mater. Des. 109, 415–424 (2016)CrossRefGoogle Scholar
  4. 4.
    N.G. Rim, C.S. Shin, H. Shin, Current approaches to electrospun nanofibers for tissue engineering. Biomed. Mater. 8(1), 014102 (2013)CrossRefGoogle Scholar
  5. 5.
    Y.F. Liu, X.T. Dong, F.D. Zhu, Adv. Mater. Res. 102–104, 550–554 (2010)CrossRefGoogle Scholar
  6. 6.
    N. Mohan, P. Senthil, S. Vinodh, N. Jayanth, A review on composite materials and process parameters optimisation for the fused deposition modelling process. Virtual Phys. Prototyp. 12(1), 47–59 (2017)CrossRefGoogle Scholar
  7. 7.
    K.S. Boparai, R. Singh, H. Singh, Development of rapid tooling using fused deposition modeling: a review. Rapid Prototyp. J. 22(2), 281–299 (2016)CrossRefGoogle Scholar
  8. 8.
    B. Zhang, B. Seong, V.D. Nguyen, D. Byun, 3D printing of high-resolution PLA-based structures by hybrid electrohydrodynamic and fused deposition modeling techniques. J. Micromech. Microeng. 26(2), 025015 (2016)CrossRefGoogle Scholar
  9. 9.
    J.M. Pearce, C. Morris Blair, K.J. Laciak, R. Andrews, A. Nosrat, I. Zelenika-Zovko, J. Sustain. Dev. 3(4), 17 (2012)Google Scholar
  10. 10.
    S. Bose, M. Roy, A. Bandyopadhyay, Recent advances in bone tissue engineering scaffolds. Trends Biotechnol. 30(10), 546–554 (2012)CrossRefGoogle Scholar
  11. 11.
    U. Jammalamadaka, K. Tappa, Recent Advances in Biomaterials for 3D Printing and Tissue Engineering. J. Funct. Biomater. 9(1), 22 (2018)CrossRefGoogle Scholar
  12. 12.
    C. Casavola, A. Cazzato, V. Moramarco, C. Pappalettere, Orthotropic mechanical properties of fused deposition modelling parts described by classical laminate theory. Mater. Des. 90, 453–458 (2016)CrossRefGoogle Scholar
  13. 13.
    L.M. Galantucci, F. Lavecchia, G. Percoco, Quantitative analysis of a chemical treatment to reduce roughness of parts fabricated using fused deposition modeling. CIRP Ann. Manuf. Technol. 59(1), 247–250 (2010)CrossRefGoogle Scholar
  14. 14.
    D. Pranzo, P. Larizza, D. Filippini, G. Percoco, Extrusion-based 3D printing of microfluidic devices for chemical and biomedical applications: a topical review. Micromachines 9(8), 374 (2018)CrossRefGoogle Scholar
  15. 15.
    L.M. Galantucci, I. Bodi, J. Kacani, F. Lavecchia, Analysis of dimensional performance for a 3D open-source printer based on fused deposition modeling technique. Procedia CIRP 28, 82–87 (2015)CrossRefGoogle Scholar
  16. 16.
    A. Brown, D. De Beer, P. Conradie, Development of a stereolithography (stl) input and computer numerical control (cnc) output algorithm for an entry-level 3-d printer. S. Afr. J. Ind. Eng 25, 39–47 (2014)Google Scholar
  17. 17.
    B.K. Gu, D.J. Choi, S.J. Park, M.S. Kim, C.M. Kang, C.H. Kim, 3-dimensional bioprinting for tissue engineering applications. Biomater. Res. 20(1), 12 (2016)CrossRefGoogle Scholar
  18. 18.
    I. Abudayyeh, B. Gordon, M.M. Ansari, K. Jutzy, L. Stoletniy, A. Hilliard, J. Interv, Cardiol. 31(3), 375–383 (2018)Google Scholar
  19. 19.
    R. Singh, J.P. Davim, Additive Manufacturing: Applications and Innovations, 1st edn. (CRC Press, NW, 2018)CrossRefGoogle Scholar
  20. 20.
    J. Torres, M. Cole, A. Owji, Z. DeMastry, A.P. Gordon, An approach for mechanical property optimization of fused deposition modeling with polylactic acid via design of experiments. Rapid Prototyp. J. 22(2), 387–404 (2016)CrossRefGoogle Scholar
  21. 21.
    J. Suganuma, H. Alexander, J. Appl. Biomater. 4(1), 13–27 (2005)CrossRefGoogle Scholar
  22. 22.
    M. Guvendiren, J. Molde, R.M.D. Soares, J. Kohn, Designing Biomaterials for 3D Printing. ACS Biomater. Sci. Eng. 2(10), 1679–1693 (2016)CrossRefGoogle Scholar
  23. 23.
    E.J. McCullough, V.K. Yadavalli, Surface modification of fused deposition modeling ABS to enable rapid prototyping of biomedical microdevices. J. Mater. Process. Technol. 213(6), 947–954 (2013)CrossRefGoogle Scholar
  24. 24.
    S. Chen, J. Lu, J. Feng, 3D-printable ABS blends with improved scratch resistance and balanced mechanical performance. Ind. Eng. Chem. Res. 57(11), 3923–3931 (2018)CrossRefGoogle Scholar
  25. 25.
    F. Liu, C. Vyas, G. Poologasundarampillai, I. Pape, S. Hinduja, W. Mirihanage, P. Bartolo, Structural evolution of PCL during melt extrusion 3D printing. Macromol. Mater. Eng. 303(2), 1700494 (2018)CrossRefGoogle Scholar
  26. 26.
    W. Wu, P. Geng, G. Li, D. Zhao, H. Zhang, J. Zhao, Influence of layer thickness and raster angle on the mechanical properties of 3D-printed PEEK and a comparative mechanical study between PEEK and ABS. Materials. 8(9), 5834–5846 (2015)CrossRefGoogle Scholar
  27. 27.
    M. Vaezi, S. Yang, Extrusion-based additive manufacturing of PEEK for biomedical applications. Virtual Phys. Prototyp. 10(3), 123–135 (2015)CrossRefGoogle Scholar
  28. 28.
    B.C. Tellis, J.A. Szivek, C.L. Bliss, D.S. Margolis, R.K. Vaidyanathan, P. Calvert, Trabecular scaffolds created using micro CT guided fused deposition modeling. Mater. Sci. Eng. C 28(1), 171–178 (2008)CrossRefGoogle Scholar
  29. 29.
    F.S. Senatov, K.V. Niaza, M.Y. Zadorozhnyy, A.V. Maksimkin, S.D. Kaloshkin, Y.Z. Estrin, Mechanical properties and shape memory effect of 3D-printed PLA-based porous scaffolds. J. Mech. Behav. Biomed. Mater. 57, 139–148 (2016)CrossRefGoogle Scholar
  30. 30.
    G.C. Anzalone, B. Wijnen, J.M. Pearce, Multi-material additive and subtractive prosumer digital fabrication with a free and open-source convertible delta RepRap 3-D printer. Rapid Prototyp. J. 21(5), 506–519 (2015)CrossRefGoogle Scholar
  31. 31.
    A. Goyanes, J. Wang, A. Buanz, R. Martínez-Pacheco, R. Telford, S. Gaisford, A.W. Basit, 3D Printing of Medicines: Engineering Novel Oral Devices with Unique Design and Drug Release Characteristics. Mol. Pharm. 12(11), 4077–4084 (2015)CrossRefGoogle Scholar
  32. 32.
    A.J.L. Morgan, L.H. San Jose, W.D. Jamieson, J.M. Wymant, B. Song, P. Stephens, D.A. Barrow, O.K. Castell, Simple and Versatile 3D Printed Microfluidics Using Fused Filament Fabrication. PLoS One 11(4), e0152023 (2016)CrossRefGoogle Scholar
  33. 33.
    H.N. Chia, B.M. Wu, Recent advances in 3D printing of biomaterials. J. Biol. Eng. 9(1), 4 (2015)CrossRefGoogle Scholar
  34. 34.
    L.D. Albrecht, S.W. Sawyer, P. Soman, 3D Print. Addit. Manuf. 3, 106–112 (2016)Google Scholar
  35. 35.
    J.E. Trachtenberg, P.M. Mountziaris, J.S. Miller, M. Wettergreen, F.K. Kasper, A.G. Mikos, Open-source three-dimensional printing of biodegradable polymer scaffolds for tissue engineering. J. Biomed. Mater. Res. A 102(12), 4326–4335 (2014)Google Scholar
  36. 36.
    J.P. Temple, D.L. Hutton, B.P. Hung, P.Y. Huri, C.A. Cook, R. Kondragunta, W.L. Grayson, Engineering anatomically shaped vascularized bone grafts with hASCs and 3D-printed PCL scaffolds. J. Biomed. Mater. Res. A 102(12), 4317–4325 (2014)Google Scholar
  37. 37.
    A. Liu, G.H. Xue, M. Sun, H.F. Shao, C.Y. Ma, Q. Gao, Z. Gou, S.G. Yan, Y.M. Liu, Y. He, Sci. Rep. 6, 1–13 (2016)CrossRefGoogle Scholar
  38. 38.
    J. De Ciurana, L. Serenó, È. Vallès, Selecting process parameters in RepRap additive manufacturing system for PLA scaffolds manufacture. Procedia CIRP. 5, 152–157 (2013)CrossRefGoogle Scholar
  39. 39.
    W. Kosorn, M. Sakulsumbat, P. Uppanan, P. Kaewkong, S. Chantaweroad, J. Jitsaard, K. Sitthiseripratip, W. Janvikul, J Biomed Mater Res B Appl Biomater 105, 1141–1150 (2016)CrossRefGoogle Scholar
  40. 40.
    U. Ritz, R. Gerke, H. Götz, S. Stein, P.M. Rommens, Int. J. Mol. Sci. 18(12), 2596 (2017)CrossRefGoogle Scholar
  41. 41.
    A. Grémare, V. Guduric, R. Bareille, V. Heroguez, S. Latour, N. L’heureux, J.C. Fricain, S. Catros, D. Le Nihouannen, Characterization of printed PLA scaffolds for bone tissue engineering. J. Biomed. Mater. Res. A 106(4), 887–894 (2018)CrossRefGoogle Scholar
  42. 42.
    G.S. Diogo, V.M. Gaspar, I.R. Serra, R. Fradique, I.J. Correia, Manufacture of β-TCP/alginate scaffolds through a Fab@home model for application in bone tissue engineering. Biofabrication 6, 25001 (2014)CrossRefGoogle Scholar
  43. 43.
    J.C. Boga, S.P. Miguel, D. de Melo-Diogo, A.G. Mendonça, R.O. Louro, I.J. Correia, In vitro characterization of 3D printed scaffolds aimed at bone tissue regeneration. Colloids Surf. B: Biointerfaces 165, 207–218 (2018)CrossRefGoogle Scholar

Copyright information

© Qatar University and Springer Nature Switzerland AG 2019

Authors and Affiliations

  1. 1.BIOMATEN, Center of Excellence in Biomaterials and Tissue EngineeringMiddle East Technical University (METU)AnkaraTurkey
  2. 2.School of Engineering, Department of Medical EngineeringAcibadem Mehmet Ali Aydinlar UniversityIstanbulTurkey

Personalised recommendations