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host’s angiogenesis potential and strives to stimulate it through optimizing the scaffold material, application of growth factors, or cell seeding in the scaffold. However,
the ingrowth of blood vessels into the scaffold is often slower than 5μm/h [77],
causing a prolonged healing period associated with major hypoxic conditions, especially in the central portion of the scaffold [8]. The vascular outgrowth, directed
from the implanted scaffold to natural bone, has been proposed as an alternative to
enhance vascularity. Other strategies of establishing vascularization of scaffold
include incorporating the scaffold with preformed microvasculature that can be
directly perfused by blood as soon as it establishes interconnection (inosculation) to
the vessels of host [78–80] or with microvascular fragments that promptly develop
into microvasculature network upon implantation [81–83].
Buser etal. demonstrated that creating small perforations in the cortical plate
permits enhanced migration of cells with angiogenic and osteogenic potential [84].
Decortication provides communication between marrow space and the scaffold.
This phenomenon could augment vascularization, promote growth factors, and subsequently enhance bone formation [85, 86].
Some conditions can enormously affect vascularity of bone; for instance, prior
treatment with radiation as an adjunct therapy for malignant head and neck lesions
could heavily enfeeble vascularity, decrease cellularity and osteogenic cell count, hinder bone remodeling, alter cytokine capacity, and create a hypoxic environment [87–
89]. Mineral component of bone is relatively resistant to radiation; on the contrary,
bone marrow containing BMSCs and endothelial cells of vasculature are strongly
radiosensitive. Delivery of cell-signaling factors such as VEGF, TGF-β1, and BMP or
gene therapy has been proposed to improve bone regeneration in radiated sites, yet
regeneration of irradiated critical-sized defects could not be accomplished [90–92].
P. Hazrati and A. Khojasteh
6 Treatment Planning
6.1 Scaffold Manufacturing Considerations
6.1.1 Material
Material selection is of signicant importance in designing scaffolds made by additive manufacturing (AM) methods. The size and load-bearing requirements of the
defective area inuence the biomaterial selection. To operate effectively invivo,
suitable biomaterials must exhibit process compatibility with the particular AM
technology used, as well as proper biochemical and physical features [93]. Many
varied materials have been used in fabricating biodegradable scaffolds. For example, because of their improved corrosion resistance and biomimicry, biodegradable
metal alloys, including magnesium (Mg) or zinc (Zn), have recently gained popularity [94–96]. On the other hand, these materials provide particular problems, such as
high melting temperatures, ammability, and the formation of metallic vapors, all of
which jeopardize process stability [97].
Due to their strong biocompatibility, bioresorption, and processability, various
polymers are viable materials for AM production of patient-specic implants [98].
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While polymers, either natural or synthetic, are great materials for accommodating
AM processing parameters, single material groups are restricted in their ability to
meet both AM processing and clinical usefulness criteria in biomedical applications. Polylactides, for example, have high tensile strength and low degradation
rate; therefore, they may last longer invivo than desired. While polyglycolic acid
(PGA) and poly(lactic-co-glycolic acid) (PLGA) have better mechanical qualities,
they degrade fast when employed as a bioresorbable scaffold, and their tensile
strength is decreased by 50% within 2weeks [99]. Furthermore, polymers’ pure use
is limited due to improper mechanical properties and inappropriate degradation
[100, 101]. Ceramics are often mixed with polymers to overcome this stumbling
block to create composite materials with superior mechanical and physiological
properties [102, 103]. Also, incorporation of this mineral phase provides osteoinductivity for the scaffold [104]. Ceramics’ brittleness and poor mechanical resistance barricade pure application of them in load-bearing sites [105]. Naturally, bone
is a composite tissue composed of a mineral component dominated by hydroxyapatite (HA) nanocrystals and an organic phase made up of extracellular matrix (ECM)
proteins, with collagen type 1 accounting for around 90% of the total [106]. Multimaterial constructs own a better capacity to imitate hybrid tissue architectures as
scaffold material because of the complicated arrangement and composition necessary for function [107, 108]. Accordingly, when designing a scaffold for bone
regeneration, a mixture of mineral-like and ECM-like components must be considered as scaffold material to achieve the desired outcomes.
Polymers and bioceramics are frequently utilized as ECM-like and mineral components, respectively. The combination of polymers with bioceramics could help
achieve the desired ideal characteristics. However, fabricating scaffolds with a high
mineral concentration similar to the actual bone may not be the best way to promote
bone growth [1]. For instance, increasing the weight ratio of HA to high percentages
in composite scaffolds not only complicates the production procedure but also may
decrease osteogenic differentiation, decrease compressive strength, and proliferation of cells [109–111]. Increasing the proportion of β-TCP and bioglass has similar
effects on scaffolds’ biological and functional characteristics [39, 112–114]. The
proportion of both ceramic and polymer elements in the composite material has a
substantial impact on the composite scaffold’s invivo and invitro behavior [115].
Figure4 demonstrates materials utilized in scaffold fabrication.
Polymers
Polycaprolactone (PCL)
PCL is a biogenic polyphosphate (bio-polyP). It is the most widely utilized polymer
biomaterial in AM because of its high mechanical qualities, biocompatibility, signicant printing accuracy, relatively low expenses, FDA approval, osteoconductivity, exibility, lower pH changes, simplicity of processability, low melting point,
ease of mixing with other materials, and established efcacy in bone regeneration
[105, 116–122]. On the other hand, PCL cannot be solely employed for hard tissue
regeneration due to its ductility and low elastic modulus [100, 101]. Therefore,
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P. Hazrati and A. Khojasteh
Fig. 4 Materials utilized in the fabrication of polymer-ceramic scaffolds. Al alginate, BD bioden-
tine, BCP biphasic calcium phosphate, BO bio-oss, BG bioglass, CaP calcium phosphate, CaC
calcium carbonate, CH chitosan, CS calcium silicate, Col collagen, DBB deproteinized bovine
bone, DCPD dicalcium phosphate dihydrate, GL gelatin, Gr graphene, HA hydroxyapatite, MgO
magnesium oxide, MTA mineral trioxide aggregate, PCL polycaprolactone, PDLLA poly(D,Llactide acid), PEG poly(ethylene glycol), PEU poly(ester urea), PGF phosphate glass ber, PLLA
poly(l-lactic acid), PLA poly lactic acid, PLGA poly(lactic-co-glycolic) acid, PLG polylactide-coglycolide, POC poly(1,8-octanediol-co-citrate), PVA poly(vinyl alcohol), PHB poly(3hydroxybutyrate- co-3-hydroxyhexanoate), SF silk broin, TCP tricalcium phosphate. (Reprinted
with permission from [1])
many mineral components, like HA, tricalcium phosphate (TCP), calcium silicate,
and bioglass, are added to PCL in order to compensate mentioned shortcomings [1].
The combination of PCL and beta-tricalcium phosphate (β-TCP) is one of the most
frequently utilized mixtures of PCL in bone tissue engineering [123].
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Poly Lactic Acid (PLA)
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PLA might be considered a choice for hard tissue regeneration due to its superior
mechanical characteristics compared to other synthetic polymers. However, its
applicability is limited due to its lengthy biodegradation period and the diffusion of
acidic metabolites during this process [1, 124]. Poly(D,L-lactide acid) (PDLLA)
and poly(l-lactic acid) (PLLA) are other forms of this polymer with enhanced biocompatibility and mechanical stability [109, 125, 126]. HA, TCP, and bioglass are
the most frequent ceramics combined with PLA and its derivatives [1, 125].
Poly(Lactic-co-Glycolic Acid) (PLGA)
PLGA has a faster rate of degradation and lower osteoinductivity [127]. PLGA may
degrade to nontoxic acidic compounds, such as lactate, which might trigger an
inammatory reaction [128]. Hence, mineral components, especially HA and TCP,
are added to PLGA to cover its aws [1].
Alginate
Alginate is often used to repair cartilage due to its ability to induce chondrocyte
proliferation and compatibility with cartilaginous tissue [129, 130]. Although
because of its very low stiffness and having no interaction with cells, it is not appropriate for application in load-bearing areas [22, 131].
Chitosan
Chitosan, a natural polycationic linear polysaccharide generated by deacetylation of
chitin, has several benets, including biocompatibility, biodegradability, antifungal,
antibacterial, and anticancer activity, but it also has certain disadvantages, such as
insolubility in physiological pH [132–135]. Carboxymethyl chitosan (CMCS) has
improved water solubility, hence increasing the processability of the tissueengineered scaffolds. Grafting CMCS with waterborne polyurethane (WPU)
enhances mechanical properties, such as compression strength, and biological qualities, such as osteoblast adhesion and proliferation [136].
Bioceramics
Calcium Phosphate Compounds
Calcium phosphate compounds (mostly β-TCP) are structurally comparable to the
calcium phosphate substance that makes up 60–70% of normal bone, indicating
intrinsic biomimetic properties. It is biodegradable and has been shown to have
osteoconductive qualities [137]. β-TCP has good biodegradability, chemical interaction with bone, and wear resistance, which are all essential characteristics for
load-bearing craniofacial defect locations [138]. The proportion of β-TCP in composite materials has a direct and inverse relationship with the surface roughness and
contact angle, respectively.
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Hydroxyapatite (HA)
P. Hazrati and A. Khojasteh
Because of its stoichiometric closeness to the mineral phase of normal bone, HA
has been regarded as a scaffold material for bone healing with excellent biocompatibility [139]. However, because of its inherent poor interaction with the binder liquid during the 3D printing process, it is commonly mixed with other bioceramics or
biopolymers [140].
Bioglass
Bioglasses are also employed because of their high osteoconductivity and ability to
attach to hard tissue; nevertheless, they degrade slowly and induce cytotoxicity in
the surrounding tissue [141]. Their main benet is that they promote osteogenesis
and interact well with cells [13, 142], but they are too brittle to be used in loadbearing craniofacial locations [143]. A study has reported that bioglass affects
osteogenic differentiation better than HA [121].
6.1.2 Composition Technique
Melt blending, solvent-based blending, and powder blending are among the composition processes utilized to make printing materials [1]. Melt blending may be done
in two ways. To form a composite, either the ceramic powder is added to the molten
polymer or a combination of ceramic powder and polymer is melted. A solvent is
used to mix polymer and ceramic components in solvent-based blending. Extra
steps are usually required to eliminate the solvent before or after constructing the
structure. For sintering-based molding methods, such as stereolithography (SLA)
and digital light processing (DLP), powder blending is often employed. A combination of ceramic powder and polymer powder is created, and the ceramic and polymer components are fused using a laser or light. Solvent blending and melt blending
are the most common methods of preparing polymer-ceramic composite scaffolds [1].
6.1.3 Designing
Compilation of CAD/CAM technologies and tissue engineering-based approaches
of maxillofacial reconstructive procedures have yielded application of patientsspecic implants as a mean of bone tissue engineering. Unique scaffolds that are
highly compatible with the defect’s topography and meet each patient’s particular
needs are designed in a virtual environment, according to data provided by 3D
imaging modalities [13–15]. Metallic restorations of adjacent teeth could distort the
images and complicate the CAD process [57].
After obtaining 3D radiographs (i.e., CT) and conveying this dataset as standard
stereolithographic les to the CAD program, a 3D model of intended bones with the
defect is reconstructed (Fig. 5). These 3D models allow surgeons to witness the
exact shape and size of the defect rather than visualizing it. Afterward, the CAD
program determines the orientation and location of the osteotomy cuts on the virtual
model (Fig.6a). The surgeon’s judgment could modify these cuts; eventually, these
virtual osteotomy cuts could be replicated as surgical guides to instruct the surgeon
during the procedure.
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ab
c d
Fig. 5 Workow of virtually simulating surgical site before surgery. CT image of coronal (a),
axial (b), and sagittal (c) plane. (d) 3D reconstructed image of a defective left posterior mandible
ab
Fig. 6 (a) Dened position and angulation of osteotomy cuts on the virtual model; (b) the defec-
tive site lled by designed patient-specic scaffold
The empty space created by osteotomy is then lled with a patient-specic scaffold, which is designed according to the normal corresponding contralateral anatomic site (Fig.6b). The designed scaffold is then transformed into a CAM program
where, in addition to morphology, signicant properties such as porosity and interconnectivity are specied. As mentioned before, pore size greater than 300μm is a
prerequisite for proper vascular and bone ingrowth [31].
6.1.4 Fabrication Technique
Solvent casting, gas foaming, phase separation, particulate leaching, and freezedrying are some of the conventional techniques used to fabricate tissue scaffolds.
However, using these traditional fabrication procedures, it is not possible to produce
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patient-specic scaffolds with unique geometry and a highly porous structure with
linked pores [144, 145]. When used as an improved tissue scaffold production technique, 3D printing technology can construct complex geometries with specic benets, such as the ability to t into irregular defective areas and replicate tissue
complexity via precise placement and improved control over the porosity and pore
pattern [146–149]. When 3D printed patient-specic implants are employed for
functional bone regeneration in bony defects, it takes less time and needs fewer
screws for retention than the classic approaches [142]. Even though all 3D printers
follow a dened process to produce constructs (layer-by-layer placement of the
material through the prespecied spatial motion) [150], accuracy, expenses, postprinting procedures, and preferred primary material vary between different methods
[115, 151]. Because of its simplicity of use, the diversity of biocompatible polymers
and composites that may be employed and the accuracy with which porous structures can be created, FDM is particularly appealing for bone tissue engineering
applications [39]. FDM is the most frequent method of fabricating composite scaffolds [1]. In most situations, the melt blending process is employed to create the
composite used in the FDM procedure. DIW, sintering-based techniques, and LDM
are other methods of printing biodegradable scaffolds. However, these 3D printing
techniques have some drawbacks. For example, it has been shown that printing HA
scaffolds with a composite mix with a 40% ceramic ratio is not possible [110].
P. Hazrati and A. Khojasteh
6.2 Surgical Technique
General anesthesia with nasal intubation is required for the surgical procedure; however, local anesthesia with or without sedation may sufce when reconstructing
small areas. Placing the scaffold into the defective area could be accomplished during surgical removal of ablated tissue. However, a 1-year interval between resection
and regeneration procedures is recommended in the surgical protocols. Flap design
should provide the accessibility and visibility required for both resection and regeneration procedures. In vascular surgery, the location of neurovascular bundles should
be predesigned in the scaffold’s morphology. If the morphology of the defective area
allows, a series of tiny holes, 1–2mm deep could be constructed on the defect walls
with round burs under abundant irrigation to enhance the quantity of bleeding and
revascularization at the surgical site. Because utilizing the scaffold as a guide for
drilling the hole might cause the fragile scaffold to crack and fracture, a polytetrauoroethylene (PTFE) scaffold replica could be produced to serve as a guide for the
proper placement of the hole for the xation screw [15]. The patient- specic implant
is then placed at the defective site and secured to the remaining wall with the titanium
mini-screw(s) lengthen 6–8mm inserted through the predrilled hole. Plasma rich in
growth factor (PRGF) could be prepared before surgery and applied to the surgical
site in order to promote tissue healing and facilitate bone regeneration [15, 152]. An
addition to conventional coverage of scaffold with a membrane, a pedicled buccal fat
pad ap could be used to cover the scaffold. Figure7 shows application of B-TCP
scaffolds to treat a defect in the posterior mandible.
a
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133
b
c
f
de
Fig. 7 (a) Radiographic view of a large defect at the posterior mandible. (b) Implanting the
degradable scaffold and covering it with buccal fat pad pedicle. (c) Postsurgical radiographic view
of the defect. (d) Radiographic view after 12months showing signicant amounts of bone formation. (e) Radiographic view after 18months. (f) Clinical view of the defect site after 18months
showing bone formation and 90% defect lling
7 Prognosis
All the studies which conducted FBR in humans reported in the literature between
2009 and 2017 demonstrated high success rate with nite serious complications
[37]. Long-term stability has been reported in cases, where FBR is employed to
prepare the bone quantity for implant placement.
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