3D Printing Applications for Craniomaxillofacial Reconstruction: A Sweeping Review

Abstract

The field of craniomaxillofacial (CMF) surgery is rich in pathological diversity and broad in the ages that it treats. Moreover, the CMF skeleton is a complex confluence of sensory organs and hard and soft tissue with load-bearing demands that can change within millimeters. Computer-aided design (CAD) and additive manufacturing (AM) create extraordinary opportunities to repair the infinite array of craniomaxillofacial defects that exist because of the aforementioned circumstances. 3D printed scaffolds have the potential to serve as a comparable if not superior alternative to the “gold standard” autologous graft. In vitro and in vivo studies continue to investigate the optimal 3D printed scaffold design and composition to foster bone regeneration that is suited to the unique biological and mechanical environment of each CMF defect. Furthermore, 3D printed fixation devices serve as a patient-specific alternative to those that are available off-the-shelf with an opportunity to reduce operative time and optimize fit. Similar benefits have been found to apply to 3D printed anatomical models and surgical guides for preoperative or intraoperative use. Creation and implementation of these devices requires extensive preclinical and clinical research, novel manufacturing capabilities, and strict regulatory oversight. Researchers, manufacturers, CMF surgeons, and the United States Food and Drug Administration (FDA) are working in tandem to further the development of such technology within their respective domains, all with a mutual goal to deliver safe, effective, cost-efficient, and patient-specific CMF care. This manuscript reviews FDA regulatory status, 3D printing techniques, biomaterials, and sterilization procedures suitable for 3D printed devices of the craniomaxillofacial skeleton. It also seeks to discuss recent clinical applications, economic feasibility, and future directions of this novel technology. By reviewing the current state of 3D printing in CMF surgery, we hope to gain a better understanding of its impact and in turn identify opportunities to further the development of patient-specific surgical care.

Graphical Abstract

1. INTRODUCTION

A large portion of our self-identity is housed within facial features. Patients with facial differences as a result of craniomaxillofacial (CMF) trauma, such as sports injuries and road accidents, congenital anomalies, infectious disease, tumor resection, or edentulism have reported associated stressors of self-acceptance, negative responses of others, and difficulty coping with associated impairments.^1–4 These associated impairments can include compromised sensory organ function, speech, and mastication.⁵ The field of CMF surgery treats a diverse range of ages and pathologies, both of which must be considered when CMF surgeons form and execute their surgical plan. Moreover, the CMF surgeon is challenged by varying load-bearing demands, proximity of sensory organs, as well as the microbiota of the oronasal cavity, all of which may complicate bone healing.⁶ A multitude of strategies to fit and fill these defects have been deployed and continue to evolve in the hope of restoring the form and function of any given defect site. The combined effort of computer-aided design/computer-aided manufacturing (CAD/CAM) and 3D printing (3DP) to create patient-specific devices based on individual computed tomography (CT) or magnetic resonance imaging (MRI) data is one such strategy that has drastically changed the approach to CMF defect repair.

Traditionally, CMF reconstruction relies upon harvesting bone from a healthy donor site such as the fibula. While autografting remains the gold standard today, it is limited by donor site morbidity, finite stock, risk of resorption, prolonged operative time and hospital stay, as well as patient-reported increases in postoperative pain.^7–9 Allografts and xenografts solve many of the autograft’s shortcomings, yet they are still associated with risk of disease transmission and antigenic reactions.^7,10 For these reasons, synthetic alloplasts, including metals, polymers, and ceramics, have been popularized within the field of CMF. The characteristics of any implantable material, resorbable or nonresorbable, should meet the needs of the biological environment it is placed within. For non-resorbable materials traditionally used for rigid fixation (e.g., titanium alloys), these needs include biocompatibility, mechanical strength, and osteogenic properties that facilitate bone healing, at a minimum. While this is equally as important for resorbable materials used in other forms of bone defect repair (e.g., ceramics and many synthetic polymers), they are also expected to degrade in a predictable and timely manner such that the defect site is mechanically supported as the material gradually becomes replaced with regenerated bone that is similar in quality to native bone.^11,12,10 This is achievable only if the material is pro-osteogenic, pro-angiogenic, and inflammation from its degradation is kept to a minimum so as to not interfere with any of these processes. A material that meets these requirements does not currently exist without the combined effort of bone tissue engineering strategies and three-dimensional (3D) printing.

With respect to conventionally used fixation devices, commercially available plates are produced in generalizable configurations, requiring the surgeon to manually bend them to a contour specific to the patient’s anatomy. They typically need to be adjusted again intraoperatively as changes to surgical margins may occur between the preoperative and intraoperative period.¹³ This can be physically challenging and time-consuming to the point that it significantly extends operative time, particularly for surgeons early on in their training.^13–15 Even still, complete bone-plate congruence without visible gaps is rarely achieved.^13,16 For more complicated cases, a plate may need to be bent repeatedly, resulting in residual stresses that can compromise its performance in vivo.^{13,14,17–19} Residual stresses generated from titanium plate contouring have been theorized to initiate crack formation that then grow larger under the loading stresses associated with mastication, leading to plate fracture and screw loosening.^13,17,19 Previous literature has found that in comparison to commercial plates made of the same material, 3D printed plates have superior mechanical properties.¹⁴ Furthermore, the need to bend 3D printed plates preoperatively may be minimized or eliminated, as they are printed to fit the exact shape of patient anatomy. This shortens operative time and reduces any unnecessary wear that may compromise performance.^14,20

In addition to fixation devices, 3D printing can be utilized to create other implantable and nonimplantable devices to improve CMF surgical care. Implantable scaffolds for bone defect repair are customizable to the nanoscale allowing them, unlike autografts, to precisely fit geometrically complex defects.^16,21–23 Alterations to scaffold architecture, including porosity and surface topography, lends to both augmented osteoconductivity and vascularization within defects.^10,24 Moreover, scaffolds may be further seeded and/or coated with mesenchymal stem cells, exogenous growth factors, antimicrobials, and regenerative pharmaceuticals to minimize risk of failure and further accelerate bone growth and neoangiogenesis.^10,25–27 Other uses for 3D printing within craniofacial surgery are patient-specific anatomic models or surgical guides for preoperative and intraoperative use. Clinical studies have found overwhelming evidence of improved surgical accuracy and aesthetic outcomes.^28–31

However, patient-specific 3D printed medical devices (3DMD) are still in their nascency, with many barriers to overcome before they become commonplace in the operating room. The aim of our review is to comprehensively assess the role of 3D printing in craniomaxillofacial reconstruction. The FDA regulatory status of 3DMD, as well as 3D printing techniques, biomaterials, and sterilization procedures suitable for 3DMD of the CMF skeleton are discussed. Finally, recent clinical applications, economic feasibility, and future directions of this novel technology are examined.

2. FDA REGULATORY OVERSIGHT

Attaining FDA approval of patient-specific 3D printed medical devices, whether for anatomical model, surgical guide, or implantable scaffold, is a complex feat. 3DMD devices are regulated by the FDA’s Center for Medical Devices and Radiological Health (CDRH). 3DMD with biologic components (i.e., stem cells and exogenous growth factors) are termed “combination products” and require the additional involvement of the Center for Biologics Evaluation and Research (CBER), adding further scrutiny/complexity to the approval process. This has led to the creation of the Breakthrough Device Program which aims to expedite the development, assessment, and review of combination products if they meet certain criteria.³² It is paramount that researchers determine what regulatory agency their device falls under, as this will dictate which regulations they must follow. Researchers should also anticipate any resistance that they may meet regarding their development process in the early stages of experimentation to accelerate future clinical applicability.³³ Furthermore, gaining a deeper understanding of this rigorous process will allow craniomaxillofacial surgeons to effectively deliver safe, innovative, and informed care to their patients.

The U.S. FDA seeks to align its quality system (QS) requirements for medical devices with those of the International Organization for Standardization (ISO). The ISO is a nongovernmental global network of national standards bodies; of which, the American National Standards Institute (ANSI) serves as the U.S. member body to the ISO. QS requirements for FDA-regulated medical devices are called current good manufacturing practices (CGMP) and can be found in part 820 under section 520(f) of the Federal Food, Drug and Cosmetic Act.³⁴ They are intended to provide a framework of basic requirements that manufacturers must abide by in their production process of medical devices for human use.³⁴ Manufacturers only need to follow those that are applicable to their specific device.³⁴ For QS regulation of a device constructed with additive manufacturing, QS requirements must be enforced at each phase of development. The set of ISO standards that are related to medical device biocompatibility is ISO 10993. Standards that fall within the scope of ISO 10993 are often referenced in FDA guidance documents for medical device manufacturers. Furthermore, the ISO technical committee (TC) responsible for publishing standards related specifically to additive manufacturing is called ISO/TC 261.

As characterized by the FDA, the additive manufacturing process of a medical device can be divided into five phases: design, software workflow, build, postprocessing, and final testing considerations (Figure 1A).³³ In the design phase, patient CT or MRI data are converted to a compatible format (e.g., DICOM file).³⁵ During the software workflow phase, the scan is segmented by an image segmentation system, such that the anatomical region of interest is isolated from the remainder of the scan. It can then be further optimized prior to conversion to a 3D printer compatible file (e.g., standard triangulation language (STL) file). Many patient-specific implants for cranial defects are designed using the mirror image reconstruction technique, where the healthy side contralateral to the defect is mirrored along the midsagittal plane.¹⁵ This technique becomes problematic if the defect crosses the midsagittal line. An alternate technique, called the baffle planner technique, does not require an intact contralateral side to model the implant.¹⁵ As a part of the build phase, selected materials are used to fabricate the device. Postprocessing may consist of, but is not limited to, cleaning unsintered raw materials, sterilization, packing, and labeling. The device can only then be tested and characterized (e.g., geometry, dimensions, mechanical properties) in its final form, after postprocessing, in what is known as the final testing considerations phase. Ultimately, any 3DMD device to be sold and marketed should be prepared using FDA-cleared segmentation software, printer, materials, and postprocesses.³⁵ An overview of considerations to address within each part of the 3DMD workflow is illustrated in Figure 1B.

Figure 1.

(A) The 3DP workflow is for manufacturing a patient-specific anatomical model, surgical guide, or implant. CT/MRI data are converted to a DICOM³⁶ and further processed in preparation for design, printing, and postprocessing. (B) This highlights the critical phases of the 3DMD workflow that are overseen by the FDA. Within each phase, the FDA emphasizes specifications that must be considered for approval of a premarket device. Created with BioRender.com.

The FDA classifies medical devices into the following regulatory classes, Class I, II, and III. Assigned rank of a device is determined by degree of risk posed to the patient and intended use and specialized indications for use. Consequently, class rank determines the type of premarket submission required for FDA clearance to market.³⁷ Class I and II devices, unless exempt, will require the submission of a premarket notification, also known as a 510(k). A class I or II device cannot be marketed until the FDA issues a 510(k) clearance stating that the device is substantially equivalent (SE) to an existent, FDA-approved predicate device.³⁸ A study published in 2018 performed a search of the FDA medical device database that determined 7% of 3DMD cleared through the 510(k) pathway thus far were for craniofacial application.³⁹ And as of 2023, only five 3DP software platforms (class II devices) currently on the market have been FDA-cleared to create anatomical models specifically for craniomaxillofacial reconstruction.³⁵ The companies that produce the 510(k)-cleared software include Materialise, 3D systems, Ricoh, Axial 3D, and Medviso.³⁵ For class III devices, a premarket approval application (PMA) is necessary. This is typically required if the device is new or if a device submitted through a 510(k) is deemed not substantially equivalent (NSE) to a predicate.⁴⁰ A PMA for a class III device necessitates submission of scientific evidence in support of clinically significant results with benefits that outweigh any associated risks of its use.⁴⁰ Moreover, whether the 3DMD is implantable, load-bearing, or patient-specific will also impact the type of information that should be included to establish device safety and efficacy in any given premarket submission. The CDRH and CBER can be contacted during the presubmission process for feedback regarding any component of the development process.³³

To better understand this complex topic, we present the following real-world example. In 2015, BioArchitects submitted a 510(k) premarket notification to the CDRH for FDA approval of a class II, 3DMD device.⁴¹ Their 3D printed titanium cranioplasty plate was deemed substantially equivalent to an existing, legally marketed predicate device used for the same purposes.⁴² Thus, the FDA granted them permission to market their device as seen on their Web site,⁴³ eliminating the need to complete a PMA. It is important to note that data pertaining to any FDA-cleared medical device are accessible in the FDA’s online “medical device databases” and can be searched by application route (e.g., 510(k), PMA) or device name.⁴⁴

A massive barrier to 3DMD currently is transporting them safely and in timely to the hospital care facility (HCF). The FDA has begun discussions with stakeholders to determine a regulatory approach that will allow these 3DMD to be manufactured closer to their end-user, the patient, at the point-of-care (POC). Three different POC scenarios have been proposed including (1) HCF use of a 3D printing medical device production system (MDPS), (2) traditional manufacturer on or near the HCF, and (3) 3D printing facility is created and run by the HCF.⁴⁵ In scenario 1, an FDA-cleared MDPS, which is a legally marketed bundle of materials, software, printer, and postprocessing equipment intended to create a 3DMD for an intended use (e.g., polymer-based cranioplasty plate), would be located in the HCF. QS regulatory compliance would ultimately fall on the manufacturer of the MDPS as long as the HCF uses the MDPS in accordance with its labeled use. In scenario 2, the HCF leases space to a traditional manufacturer of a finished device, who would be responsible for regulatory compliance of a given FDA-cleared 3DMD intended for a specific use. Scenario 3 would transfer the responsibilities of QS regulatory compliance entirely to the hands of the HCF. In this case, clinical and engineering training programs for 3DMD development may become a part of certification and licensing programs for surgeons that will frequently use this technology in their practice. Ultimately, the FDA will modify existent and/or formulate new regulations as the technology evolves and regulatory compliance is passed from traditional manufactures to the HCF. For this reason, it is of utmost importance that CMF surgeons become familiar with this process as point-of-care 3D printing facilities trend toward finding a commonplace within hospital systems.

3. BIOMATERIALS FOR 3D PRINTED BONY DEFECT REPAIR OF THE CMF SKELETON

3.1. Titanium and Titanium Alloys.

Titanium is the leading 3D printed biomaterial for craniomaxillofacial repair.⁴⁶ More specifically, it is the gold standard for rigid fixation of craniofacial fractures.⁴⁶ Direct Metal Laser Sintering (DMLS) or Electron Beam Melting (EBM) are utilized to generate customizable implants for orbital, mandibular, and cranial reconstruction.⁴⁷ Titanium and its alloys, classified as a-, (a + b)-, and b-type, are known to exhibit excellent tensile strength while remaining lightweight.^21,29 They are inherently osteointe-grative, facilitating bone growth into the implant; this is theorized to due to an early transition from proinflammatory (M1-phenotype) to reparative/anti-inflammatory (M2-phenotype) macrophage predominance shortly after placement.^29,48–50

Additionally, the surface layer of titanium-based implants forms an oxide film that is resistant to corrosion.²¹ A drawback of this biomaterial is the discrepancy in Young’s moduli (i.e., elastic modulus) between titanium alloys and cortical bone, leading to a phenomenon known as stress shielding.⁵¹ Stress shielding, due to differences in moduli between bone and titanium, is a cause of progressive implant loosening, often necessitating reoperation. Interestingly as a class, b-type titanium alloys exhibit the lowest elastic moduli, yet a- and (a + b)-type alloys are more widely used in practice.^51,52 In addition, micro- or nanoscale wear particles from titanium implants have been shown to dampen osteoblastic activity and augment osteoclast recruitment, further perpetuating the issue.^53,54 Investigations continue to (a) engineer b-type titanium alloys with both maximized dynamic strength and a Young’s modulus similar to native bone and (b) determine a surface treatment (e.g., abrasive smoothening and polishing, gritting, cold treatment, or passivation and anodization) that would generate a 3D printed surface topography to counteract potential loosening or migration.^24,47,51 There is also risk of susceptibility artifact with various imaging modalities when considering titanium implant placement.^55,56 When imaged with CT or CBCT, metal implants often result in beam hardening, streak artifact, and photon starvation, all of which obscure surrounding structures.^57,58 Furthermore, signal loss, geometric distortion and “pile-up” artifacts often impede MRI imaging if metal implants are present.⁵⁹ For these reasons, metal-free implants may be preferable in patients undergoing oncologic reconstruction, in which complete and accurate visualization of bone and soft tissue structures in the postoperative period is critical.⁵⁵ Shared complications of titanium-based implants in all CMF patients (i.e., adult and pediatric) that increase the risk of reoperation include postoperative pain, palpability, infection, and exposure. In this context, 3D printed synthetic polymers have become increasingly popularized as they provide self-limited rigid fixation with a lesser chance of requiring a second procedure.^60,61

3.2. Synthetic Polymers.

Biocompatible synthetic polymers that are under investigation or currently in use for craniofacial reconstruction include polycaprolactone (PCL), polylactic acid (PLA), polyglycolic acid (PGA), polylactide-co-glycolide (PLGA), poly(methyl methacrylate) (PMMA), polyethylene (PE), and polyetheretherketone (PEEK). Synthetic polymers are easily manufactured and cost-effective.⁶² Among them, PMMA, PE, and PEEK are nondegradable polymers.⁶³ Unlike titanium alloys, the appeal of the remaining degradable polymers is that they will not restrict growth in developing pediatric skeletal defects.^60,61,64 Moreover, synthetic polymers do not interfere with radiographic, CT, or MRI studies.⁶⁵ They are also lighter than titanium alloys, with fewer reports of pain and palpability.⁶⁶ However, synthetic polymers do not possess osteointegrative properties due to their biological inertness.⁶⁷ Additionally, elastic moduli of synthetic polymers are generally lower than those of native bone. While this eliminates the concern for stress shielding, it also implies a lessened load-bearing capacity.^5,65 Figure 2A–C represent the Von Mises stress distribution during mastication at single bite points (canine, first molar, second/3rd molar), as determined by finite element models of the human cranium performed by Prado et al.⁶⁸ Areas under high stress, as depicted in red, would require materials with high mechanical performance.⁶⁸ Scaffolds composed entirely of a synthetic polymer may inadequately support a bony defect at these points while bone regeneration transpires. For these reasons, they are often combined with other biomaterials or bioactive molecules to improve their mechanical strength and bone regenerative capacity, respectively. Some of which have been previously studied include bone-derived mesenchymal stem cells, adipose-derived mesenchymal stem cells (ASCs), stromal vascular fraction (SVF), decellularized cortical bone (DCB), polydopamine, tricalcium phosphate (TCP), hydroxyapatite (HA), gelatin, and collagen.^69–76 A higher rate of infection is also a frequently reported concern of synthetic polymer-based implants.⁷⁷ To minimize infection, bactericidal-coated scaffolds continue to be investigated and optimized (e.g., silver nanoparticle impregnation, zinc chelation, ε-Polylysine coating).^78–81 Despite many commonalities among synthetic polymers, each differs in its hydrophilicity, degradation rate, and biocompatibility, impacting their individual potential for clinical application.

Figure 2.

Representation of Von Mises stress distribution across the facial skeleton based upon bite point at the (A) canine, (B) 1st molar, or (C) 2nd/3rd molars. Green and red depict low and high areas of stress, respectively. Created with BioRender.com.

3.2.1. Polycaprolactone.

Polycaprolactone, an FDA-ap-proved thermoplastic polyester, is the second most studied and used 3D printed material for craniofacial application.^{12,75,76,82–87} PCL is known to be biodegradable and highly biocompatible.²¹ PCL is commonly 3D printed using Fused Deposition Modeling (FDM), rendering it very cost-effective.²¹ Due to its hydrophobicity, the degradation rate in vivo is estimated to occur between 1 and 2 years.⁸⁸ This prolonged resorption time may interfere with proper bone regeneration within the defect site if a shorter healing period is expected. For this reason, PCL has been successfully copolymerized with hydrogels, like collagen, to augment its resorption kinetics while still providing the scaffold great mechanical strength.^89,90 The addition of collagen also improves the scaffold’s osteoconductive properties, which are lessened if PCL is used alone. Furthermore, several composite PCL-mineral scaffolds have been compared to evaluate potential osteointegrative enhancement. Mineral sources included HA, TCP, DCB, and Bio-Oss (BO) which is an inorganic bovine trabecular bone. Among all treatment groups, PCL-DCB and PCL-BO demonstrated the greatest osteoinductivity after 3 weeks in vitro which was theorized to be due to the presence of a collagen phase.⁷⁵

PCL has also been evaluated as a candidate for craniofacial defect repair in several in vivo studies.^86,91 Recently, the effect of oxygen loading on vascularization and bone regeneration was evaluated in critical-sized calvarial defects of a murine model.⁸⁶ Biodegradable synthetic microtanks housed within a porous ASC-seeded PCL-scaffold were hyperbarically loaded with pure oxygen. The microtanks were designed to release oxygen over hours after placement in vivo. Compared to non-O₂-loaded scaffolds, O₂-loaded scaffolds exhibited enhanced bone regeneration after 8 weeks.⁸⁶ O₂-loading may serve as a viable strategy in improving the hypoxic microenvironment known to exist within large nonvascularized scaffolds (>1 mm).⁸⁶ Additionally, a study performed by Singh et al. in a skeletally mature minipig model examined the bone regenerative capacity of PCL-DCB scaffolds that were intraoperatively infused with the autologous stromal vascular fraction (SVF) in critical-sized zygomatic arch defects. SVF is an easily accessible and readily available source of mesenchymal stem cells that can be extracted from autologous lipoaspirate tissue. Compared to acellular PCL-DCB scaffolds, the SVF group demonstrated superior osteointegration after 1-year in vivo, as deduced by requiring a significantly higher torque to fracture the bone-scaffold interface.⁹¹ This approach to bone regeneration is particularly appealing because, if clinically translated, the SVF could be harvested and extracted intraoperatively from patient’s adipose tissue. Moreover, as an autologous cell source with osteogenic potential, SVF-infused scaffolds may supply the benefits associated with an autograft without the comorbidities.

3.2.2. Polylactic Acid, Polyglycolic Acid, Polylactide-co-glycolide.

Polylactic acid, polyglycolic acid, and polylactide-co-glycolide are FDA-approved, biodegradable thermoplastic polyesters. PLA and PGA are both hydrolyzed to nontoxic, although acidic products. Local inflammatory responses (LIR) because of their metabolization have been previously reported.^11,92,93 This is of particular importance because inflammation often potentiates fibrosis which may compromise regenerated tissue function or even result in scaffold rejection.²³ Of note, implants placed in anatomical areas with little vascularization are at higher risk of LIR.⁹⁴ PLA is a stiff, hydrophobic polymer with an estimated degradation time of 6–24 months.^95,96 By contrast, PGA is more hydrophilic in nature and thus degrades rather rapidly within 1.5–3 months.^95,96 For this reason, PGA-fabricated scaffolds may not adequately provide mechanical support to a defect site while tissue regeneration occurs.⁹⁷ However, due to a similar elastic modulus to cortical bone, PGA remains an excellent candidate for bone regeneration if copolymerized with other biomaterials. PLA and PGA can be copolymerized, forming PLGA, to better suit the biological environment. The degradation rate of a PLGA-scaffold can be customized with molecular weight and PLA:PGA ratio adjustments.^92,96 For example, PLGA (LA/GA = 50/50) will degrade over 1 month, as opposed to PLGA (LA/GA = 85/15) which resorbs completely after 5–6 months due to a higher composition of hydrophobic PLA.⁹⁶ Thus far, preclinical studies utilizing 3D printed scaffolds fabricated with PLA, PGA, or PLGA have demonstrated promising results for CMF defect repair.^98–100

3.2.3. Poly(methyl methacrylate).

Poly(methyl methacrylate) is a nondegradable thermoplastic polyester. PMMA is hydrophobic with mechanical and elastic strength comparable to bone.¹⁰¹ PMMA commonly serves as an alternative to autografting for cranioplasty. Historically, liquid PMMA was poured directly into a defect to fit and fill it intraoperatively. Consequently, an exothermic hardening process occurred with an associated risk of local tissue necrosis.^102–104 Risk of burn injury is circumvented when PMMA is poured into a 3D printed mold of the defect intraoperatively or 3D printed as a porous PMMA-scaffold preoperatively. While autografting remains the gold standard in cranioplasty today, there is a wealth of evidence that autografts carry greater risk of resorption and failure.^102,105 In this context, alloplastics, such as titanium and PMMA, have been looked at more favorably. That being said, PMMA-fabricated scaffolds have yet to be FDA-cleared for use in cranioplasty due to very low quality of evidence about risk of infection.^101,106 More recently, 3D printed PMMA scaffolds have been successfully applied to midface and mandibular osseous defect repair with low infection rates.¹⁰⁷

3.2.4. Polyethylene.

Polyethylene is a nondegradable, inexpensive thermoplastic polymer that is often printed with Fused Deposition Modeling (FDM). In comparison to pure PE, high-density polyethylene (HDPE) is more frequently seen in clinical practice due to its superior durability and higher melting point (131 °C).¹⁰⁸ Since 1985, Stryker (Kalamazoo, MI) has produced porous HDPE (MEDPOR) and HDPE/titanium mesh (MEDPOR TITAN) implants for use in cranio-maxillofacial reconstructive or cosmetic cases.¹⁰⁹ Like other synthetic polymers, there is always a risk of infection and extrusion. The chances of these complications are lessened, however, if the elastic modulus of the biomaterial closely matches that of the native tissue.¹¹⁰ This is postulated to be due to more evenly distributed mechanical loading across the tissue-implant interface, preventing tension and micromovement that may lead to extrusion. For example, for a hypothetical auricular reconstruction, the modulus of a porous 3D printed HDPE-scaffold was engineered to more closely match that of auricular cartilage when compared to MEDPOR.⁸¹ Additionally, these HDPE-scaffolds have been experimentally loaded and/or coated with bioactive, antimicrobic, and angiogenic substances resulting in dampened inflammatory response and enhanced tissue ingrowth, outperforming their commercial counterparts.^81,111

3.2.5. Polyetheretherketone.

Polyetheretherketone is a well-described nondegradable, thermoplastic polyaromatic for use in craniofacial reconstruction. PEEK implants are commonly printed via FDM or Selective Laser Sintering.¹¹² This biomaterial is known for its mechanical strength, modulus similar to cortical bone, and excellent thermal resistance.⁷³ Unlike the previously mentioned thermoplastic polymers, structurally complex 3D printed PEEK implants can be sterilized by autoclaving without risk of deformation or changes to its biocompatibility.^73,112 It is important to note that a recently published PEEK Safety Profile by the FDA identified an association between PEEK used in cranioplasty and seizure or exposure of the implant with a moderate quality of evidence.¹¹³ Although it was emphasized that these are complications associated with other biomaterials used for cranioplasty as well.¹¹³ Other craniofacial applications have included correction of zygomatic, mandibular, and orbital rim or floor deformities.¹¹⁴ Composite scaffolds, such as PEEK/Hydroxyapatite, continue to be evaluated as a candidate for craniofacial defect repair.^74,115 Previous studies have demonstrated increased cellular adhesion, proliferation, and alkaline phosphatase activity with a corresponding decrease in tensile strength.^74,115 In this context, the load-bearing demands of each unique defect environment must be considered.

3.3. Bioceramics.

Bioceramics generally utilized for bone regeneration have included calcium phosphates (CaP), calcium carbonates, calcium sulfates, and bioactive glasses.¹¹⁶ In contrast to synthetic polymers, bioceramics are inherently osteointegrative, osteoinductive, and osteoconductive.¹¹⁷ When fabricated as a porous 3D printed scaffold, bioceramics promote a strong tissue-scaffold interface, recruit mesenchymal stem cells (MSCs) to the defect site, and provide bioactive surfaces that facilitate osteogenesis, respectively. In their pure forms, however, they are limited by poor angiogenic properties and slow degradation rates, making these processes occur at an inadequate pace.^117,118 This issue is amplified at the center of large craniofacial defects, which are already under hypoxic conditions. Also, unlike synthetic polymers, bioceramics are mechanically weak and brittle, raising concern for their performative potential within load-bearing defects.^117,119 However, a recent study found that 3D printed β-TCP scaffolds were able to restore critically sized, full-thickness mandibulectomy defects in a rabbit model despite return to normal mastication after implantation.¹²⁰ This solidifies their potential to withstand areas of high mechanical stress if they are constructed properly on the microscale. Ongoing strategies to improve the osteogenic and mechanical potential of bioceramic scaffolds have included active component loading, doping with trace elements, bioactive surface coatings, as well as the addition of micro- and nanostructures or alterations to pore size.¹¹⁷ β-Tricalcium phosphate (β-TCP), a bioactive calcium phosphate, has garnered particular attention within the field of craniomaxillofacial reconstruction given its well-documented safety profile, 6–18-month resorption period, and similar composition to the mineral phase of human bone.^25,84,119 β-TCP resorption kinetics are much more favorable to hydroxyapatite, which exhibits a degradation rate of approximately 1%/year.¹²¹ Direct-Ink-Writing (DIW), or robocasting, with bioactive ceramic colloidal ink followed by sintering allows for customization of a defect-specific scaffold.²⁵ Sintering effectively improves the scaffold’s mechanical strength by densification.¹²¹ Previous studies have suggested optimal outcomes in treating mandibular defects of a rabbit model after scaffold sintering at 1100 °C for 4 h.^120,122 The in vivo performance of β-TCP scaffolds has been found to significantly improve with the addition of the FDA-approved pharmaceutical dipyridamole (DIPY). DIPY is an adenosine A_2A receptor indirect agonist with a bone regenerative capacity. DIPY-augmented β-TCP scaffolds have been investigated in criticalsized, craniofacial defects within skeletally mature and immature rabbit, ovine, or swine models.^122–130 Shen et al. successfully accelerated the degradation rate of β-TCP to 55%/year (calvarial defect) and 90%/year (alveolar defect) with the addition of a 1000 μM DIPY coating (Figure 3A,B).¹³⁰ This study, as well as other studies performed by this group, found that the β-TCP scaffold was replaced with regenerated, vascularized bone that is similar in its mechanical and histologic properties to native bone.^122–130 If evaluated in a skeletally immature animal model in proximity to a growth suture, the suture remained patent, even past the point of facial growth completion.^{126–128,131} This finding is critical as suture obliteration in pediatric patients is a feared complication of β-TCP scaffolds augmented with growth factor, recombinant human bone morphogenic protein (rhBMP-2).¹²⁸ Ultimately, replacement of exogenous growth factors, such as rhBMP-2, with regenerative pharmaceuticals, like DIPY, mitigates the deleterious effects associated with their supraphysiological dosing requirements.¹³² To summarize the differences between the aforementioned biomaterials, refer to Table 1 for their respective advantages and disadvantages.

Figure 3.

A drawback of 3D printed bioceramics for targeted bone regeneration is their resorption kinetics. With the addition of a 1000 μM DIPY coating, Shen et al. successfully altered the degradation rate of β-TCP scaffolds to 55%/year and 90%/year within (A) calvarial and (B) alveolar defects, respectively, in a rabbit model. Reproduced from ref 131. Copyright 2023, Wolters Kluwer Health Inc.

Table 1.

Overview of the Most Commonly Used Biomaterials in CMF Repair

Material	Advantages	Disadvantages	Ref
Titanium	Excellent for rigid fixation, strong, lightweight, osteointegrative, resistant to corrosion	Stress shielding and micro- or nano- wear particles may cause implant loosening, metal artifact with imaging that obscures surrounding anatomy	21, 29, 50, 51, 53, 54
Polycaprolactone (PCL)	Cost effective, easily copolymerized with other materials	1–2 years degradation time	21, 87–89
Polylactic acid (PLA)	Nontoxic degradation products, easily copolymerized	Acidity may cause local inflammatory response, degradation time of 6–24 months	11, 91–95
Polyglycolic acid (PGA)	Nontoxic degradation products, easily copolymerized	Acidity may cause local inflammatory response, degradation time of 1.5–3 months	94–99
Polylactide- co-glycolide (PLGA)	Degradation rate is customizable	More expensive	91, 95, 97–99
Poly(methyl methacrylate) (PMMA)	Nondegradable and modulus similar to bone	Scarce information about risk ofinfection	100–106
Polyethylene (HDPE)	Nondegradable, highly durable, inexpensive	Risk of infection and extrusion	80, 107, 109
Polyetheretherketone (PEEK)	Nondegradable, modulus similar to bone, autoclavable without risk of deformation	Possible risk ofseizure and extrusion when used in cranioplasty	72, 111–114
B-Tricalcium phosphate (B-TCP)	Osteointegrative, osteoinductive, and osteoconductive	Brittle, slow degradation rate, poor angiogenic properties	25, 83, 116–118

4. 3D PRINTING TECHNOLOGIES

According to the ISO and American Society for Testing and Materials (ASTM), there are seven overarching categories of additive manufacturing processes: vat polymerization, material jetting, binder jetting, material extrusion, powder bed fusion, sheet lamination, and directed energy deposition. These 3DP processes differ in their energy sources, suitable printing materials, and methods by which CAD data are used to fabricate individual layers and then solidify and/or fuse them to create the desired design.¹³³ Several printing techniques may belong within a single category, in which case small nuances distinguish them from one another. Many of these printing techniques have been patented or trademarked by their original developers, adding an additional layer of complexity to the nomenclature. The categories used to 3D print patient-specific anatomical models, surgical guides, or implantable devices in CMF reconstruction broadly include vat photopolymerization, material jetting, material extrusion, and powder bed fusion. Here we seek to discuss the intricacies of each 3D printing technique as they relate to craniofacial surgery.

4.1. Vat Photopolymerization.

The three key components of vat photopolymerization (VPPP) include a light source, a vat, and a light-curable resin.^133,134 For a 3DMD, the resin must be a nontoxic, biocompatible photoactive material.¹³⁴ A platform is submersed within a vat, or reservoir, of a liquid resin in its monomer or oligomer form (Figure 4A). A focused curing light irreversibly solidifies the liquid through photopolymerization of these monomers and oligomers.¹³³ The curing light path follows the segmented layers of the STL model. Once stabilized, the platform is lowered or raised by a constant height to build upon the previous layer until device completion. Layer thickness typically falls between 50 and 100 μm.¹³⁵ The printer will also build a removable latticework beneath any overhangs that may require additional support.^133,134 The completed model is withdrawn from the vat, support structures are removed, and the final model is cleansed and then cured within an UV chamber to complete the polymerization process. Techniques within this category are distinguished based upon light curing method, including laser (stereolithography (SLA)), projector (digital light processing (DLP)), light emitting diodes (LEDs), and oxygen (continuous digital light processing/continuous liquid light processing).¹³⁴

Figure 4.

(A) Schematic representation of stereolithography to create a 3DMD. A UV laser hits a vat filled with resin, resulting in a targeted photopolymerization reaction that allows for material solidification. (B) In DIW, a pressure controller regulates material flow while the nozzle moves in the x and y planes. Created with BioRender.com.

SLA, patented by Charles W. Hull in 1986, is a technique utilized regularly for 3D printing in complex CMF reconstruction. Anatomical models and surgical guides are fabricated using epoxy- or acrylic resins.¹³⁴ SLA-printed scaffolds for hard tissue repair can also be developed by using photocurable bioinks. These bioinks are hydrogel-based with or without additional cells or osteoconductive bioceramics (e.g., hydroxyapatite, β-TCP). For example, a scaffold with a perfusable vessel lumen was successfully constructed with a gelatin methacryloyl (GelMA)/nanocrystalline hydroxyapatite bioink for bony defect repair.¹³⁶ Recently, SLA-printed bioceramic scaffolds demonstrated robust osteointegration in a rat calvarial model.¹³⁷ In this study, HA and β-TCP were mixed with acrylic resin for SLA-printing, followed by thermal cycling of the final construct to remove the resin entirely ahead of implantation. Standard print speed ranges between 10 and 20 mm/h and accuracy is determined by diameter of the laser beam, with small spot sizes allowing for high resolution.¹³⁴

4.2. Material Jetting.

Like vat photopolymerization, material jetting requires a UV light source and a light-curable liquid resin. Unlike VPP, material jetting (MJT) uses thermal force to deliver microdroplets of low-viscosity feedstock, from cartridge to jet head, onto a build platform.¹³⁸ Photopolymer materials are extruded in a layered fashion and cured as successive layers are assembled. Layers can be as small as 16 μm in thickness, allowing for intricate geometrical structures with high accuracy.^139,140 Instead of printing a lattice support structure to later be removed as done with VPP processes, all MJT printers require a support material to deposit the build material(s). MultiJet, trademarked by 3DSystems, is a single build material printer supported by a wax.¹³⁹ By contrast PolyJet, a MJT system developed by Stratasys, can print multiple build materials supported by a gel.¹³⁹ This allows PolyJet printers to create multicolored anatomical models that highlight relevant vasculature and innervation or patient-specific hardware.¹⁴¹ During postprocessing, MultiJet and PolyJet printers remove support material by melting and waterjet/chemical bath, respectively.¹³⁹ PolyJet postprocessing steps are laborious in comparison to those of MultiJet. However, PolyJet is often still the preferred system due to its multimaterial capabilities. A recent study found that compared to surgical guides produced by SLA and digital metal printing (DMP), the PolyJet printer was the most accurate and demonstrated the highest reproducibility, even after 1-month storage of devices.¹⁴² Similarly, the smallest mean difference between 3D printed surgical guide and its CAD model was found when fabricated by a PolyJet printer compared to an SLA or MultiJet system.¹⁴³ Bone constructs were also printed with this AM technology. Similar to VPP, biocompatible photopolymers must be used but may be mixed and/or coated with osteoconductive materials, like polydopamine and hydroxyapatite.¹⁴⁴ Although generally, restriction of printable materials to photoreactive polymers is a significant limitation of vat polymerization and material jetting.

4.3. Material Extrusion.

Fused deposition modeling (FDM), trademarked by Stratasys, and fused filament fabrication (FFF) both use thermal material extrusion technology. Within a heated build chamber, FDM dispenses a spool of thermoplastic filament through a heated, pressurized nozzle onto a heated build plate.¹³³ Like the AM technologies discussed previously, a dissolvable support material is required for material extrusion. This is deposited simultaneously through a second nozzle and removed during postprocessing by hot water or solvent bath.¹⁴⁵ Instead of requiring light-curable resins to be photopolymerized, such as in VPP or MJT, molten layers printed with FDM or FFF simply fuse as they cool. Printable materials include thermoplastic polymers often used for CMF applications such as PEEK, PLA, PLGA, PCL, and bioceramics. It is important to note that FFF is not conducted within a heated build chamber, leading to uneven cooling¹⁴⁶ and inconsistencies in the device’s mechanical performance; thus, FDM is preferred if the 3DMD is intended for implantation. FDM has also been found to produce the most dimensionally accurate mandibular models among the material extrusion techniques.¹⁴⁷ That being said, FDM has been employed to fabricate thermoplastic polymer/ceramic composite scaffolds for bone regeneration, anatomical models for resident training, preoperative bending of off-the shelf fixation devices, or to facilitate accurate intraoperative reconstruction, as well as surgical guide fabrication for a range of CMF applications.^148–152

Direct-ink-writing (DIW) is also a material extrusion AM process but relies entirely on pressurized syringe pumps, without heat, for deposition (Figure 4B).¹⁵³ Because of this, emphasis is placed on fine-tuning the ink’s rheological properties, such as viscosity, yield stress, and elastic moduli for optimal printability.¹⁵³ As it relates to craniofacial surgery, DIW is capable of printing multimaterial structures with thermoplastic polymers (PLA, PGA, PLGA, PCL, PEEK), bioceramics, bioceramic/hydrogel composites, and less frequently titanium when incorporated into a water-based suspension for printing.^121,154,155 A support material may also be necessary for successful DIW and can be removed by dissolution or melting during postprocessing. Additional efforts to facilitate solidification of the 3DMD’s final form will vary by material but may include sintering, photocuring, or solvent evaporation.¹⁵³

4.4. Powder Bed Fusion.

Powder bed fusion (PBF) is an AM process that encompasses selective laser sintering (SLS), selective laser melting (SLM), direct metal laser sintering (DMLS), and electron beam melting (EBM). All materials printed by PBF exist as a powder feedstock supplemented with a feeder chamber. Powder is mechanically rolled evenly, and in successive 2D-cross sections, across a neighboring build chamber. Layer thicknesses are reported as 100–120 μm (SLS), 20–100 μm (SLM/DMLS), and 45–150 μm (EBM).¹⁵⁶ The powder is selectively consolidated by melting (SLM, EBM) or sintering (SLS, DMLS). In the case of SLS, SLM, and DMLS the heat source enabling solidification is a high-power laser and, in EBM, an electron beam. The surrounding, unconsolidated powder serves as a support to the 3DMD as it is built. Although, additional support may be warranted for structures with complex geometries. One advantage of PBF, unlike VPP, is unused powder can be recycled for future build cycles; however this should be done cautiously as studies have reported alterations to several mechanical properties of the final device.¹⁵⁷ In respect to CMF applications, EBM, SLM, and DMLS can print powdered titanium or TiAl6 V4 ELI, whereas SLS prints powdered thermoplastic polymers (e.g., PE, PEEK, PLA, PCL) or ceramics.¹⁵⁸ Postprocessing involves compressed air to remove any unconsolidated powder, as well as numerous grinding and/or polishing procedures to adjust surface roughness.¹⁵⁹ Overall, PBF AM processes are great options for printing patient-specific surgical plates, fixation devices, or scaffolds for CMF bone repair.^160–162

5. STERILIZATION TECHNIQUES

After printing, the devices must be adequately sterilized without compromising their final form. To be FDA-cleared for intraoperative use, all 3DMD require validation of a certain sterility assurance level (SAL).¹⁶³ The SAL is defined as the expected probability of a living microorganism following sterilization.¹⁶⁴ FDA requires an SAL < 10⁻³ for any device touching skin and SAL < 10⁻⁶ for an implantable device. Additionally, implantable devices necessitate pyrogenicity testing to rule out presence of Gram-negative endotoxins or material-mediated pyrogens that may cause a febrile reaction in the patient.¹⁶⁵

Sterilization techniques are distinguished by class on the basis of their known level of efficacy. Established Category A have been deemed safe and effective, with the most literature in support of their efficacy.¹⁶⁵ This class encompasses dry heat, ethyl oxide (EO) in a fixed chamber, steam (e.g., autoclave), and radiation (e.g., gamma or electron beam) sterilization.¹⁶⁵ Sterilizers within Established Category B do not have any FDA-recognized consensus standards but do have an established place in the literature.¹⁶⁵ Established Category B includes hydrogen peroxide, ozone, and EO in a flexible bag system.¹⁶⁵ The final category is known as Novel Sterilization Methods and includes vaporized peracetic acid, high intensity light, micro-wave radiation, sound waves, and ultraviolet light.¹⁶⁵ These methods have not been FDA-reviewed nor determined to adequately sterilize products, however, may become viable sterilization options in the future.¹⁶⁵

The mechanical and physical properties of the materials within a 3DMD determine the safest method for sterilization. Due to its high melting point and Young’s modulus, a 3DMD made from titanium is sterilizable by irradiation or autoclaving without risk of deformation.¹⁶⁶ Similarly, PEEK 3DMD can be sterilized by autoclave without significant changes to its dimensional accuracy.^73,107,167 By contrast, PMMA devices should not be sterilized by autoclave as conflicting data exists regarding structural compromise.^107,168 The remaining thermoplastic materials, PLA, PGA, PLGA, PE and PCL, are not survivable by autoclave, dry heat, or EO sterilization without altering their final structure or mechanical properties.^169,170 H₂O₂ plasma vapor sterilization, ethanol, and electron beam irradiation have been found to be the safest sterilization techniques for PGA/PLGA and polycaprolactone.^169,171 Steam sterilization by autoclave (121 °C for 40 min) has been found to alter the physiochemical properties and pH of porous bioceramics with a concern that this would affect their osteoinductivity/osteoconductivity in vivo.¹⁷²

6. ECONOMICS

Since the approval of 3D printing-related Current Procedural Terminology (CPT) codes for anatomic models and surgical guides in 2018, interest to incorporate 3DMD in routine surgical care has grown exponentially.¹⁷³ By 2028, the 3DMD market is projected to be a $6.9B industry with an annual growth rate 17.1%.¹⁷⁴ Establishment of a POC 3DP facility has enormous benefits to both patients and surgeons within the private and academic sectors. As it stands, extended manufacturing and transport times due to distant third-party manufacturers bottleneck widespread use of the technology. For example, 3DMD cannot be used in acute CMF trauma cases, although if available it would alleviate many of the challenges associated with reconstruction of extensive facial traumas.⁴ Additionally, the tendency to incorporate biologics (e.g., stem cells) into 3D printed scaffolds is not conducive to transport across great distances. Thus, HCFs can differentiate themselves by having such capabilities in-house. As of 2019, there were 113 hospitals in the United States with centralized 3D printing capabilities.¹⁷⁵ An HCF considering the establishment of their own POC facility must first consider the cost of all the required materials and equipment needed for the 3DP workflow.

The first requirement is a 3D printer and compatible segmentation software. The 3D printer itself has the largest upfront cost. Depending on the type, processing speed, manufacturer, and desired printable materials, cost ranges widely with a mean of approximately $100,000.^35,176 Several of the more inexpensive options with many potential CMF applications are material jetting, material extrusion, and stereolithography printers.^35,140,177 It is important to emphasize that only a select number of printers currently on the market are FDA-cleared for 3DMD fabrication.³⁵ Options that have been coined as low cost, such as an inexpensive material extrusion printer, have not yet been validated by the FDA, despite similar clinical performance.^178,179 The price of segmentation software typically requires a quote, as they depend on the size and output of the 3D printing system. Some software are accompanied by their own biomedical engineers available to collaborate directly with surgeons. Their labor fees in addition to any other individuals necessary to oversee and maintain the equipment must also be accounted for in the overall cost of a POC 3D printing facility.

An additional cost to consider is the required printing materials. Their prices fluctuate heavily based on supply and demand in a commodity-like market. For example, as of June 2023, polylactic acid was trading at $2.66/kg.¹⁸⁰ By contrast, PEEK is valued at hundreds of dollars per kilogram.¹⁸¹ A 1-year trend in cost per kilogram of common 3D printed materials can be seen in Figure 5A and B with the volatility and percent change per month of their respective costs over the past year depicted in Table 2.^{180,182–184} While these prices are seemingly inexpensive compared to the larger upfront costs discussed earlier, this would be a recurring expense for the HCF.

Figure 5.

(A) Cost per kilogram for various thermoplastic polymer raw materials over the past year (until July 2023) with projected values (P) for the next six months. (B) Cost per kilogram for raw titanium over the past year.

Table 2.

Trailing 12 Months of Reported Volatility (Standard Deviation) and % Change Per Month of Cost per Kilogram of Each Material

Material	Volatility	% Change per Month
PLA	$0.08	−0.36%
HDPE	$0.15	−2.22%
PMMA	$0.26	−1.21%
Titanium	$2.48	−5.57%

The average cost for an anatomical model produced by industry was $2467 with some reported to be as high as $6000.^150,185,186 Conversely, a recent study noted that the average cost for 3D printed anatomical models produced in-house was $2180.¹⁸⁵ Despite this initial cost, 3D printed anatomical model and surgical guides are estimated to save hospitals $5172 and $1918, respectively, due to reductions in operative time per case.¹⁷³ That being said, it is predicted that an HCF must print a volume of 63 models or guides annually to cover the costs of having the 3DP system itself.¹⁷³ Because of the extensive FDA-clearance process required for implantable 3DMD, there are few data reported currently on costs saved, if any, because of their use.

7. CLINICAL APPLICATIONS OF 3DMD IN CMF SURGERY

Considering both the barriers and benefits associated with 3DMD discussed thus far, it is understandable that anatomical models, surgical guides, and implants are slowly, yet excitedly, making their way into operating rooms for CMF application. Barriers have centered on the establishment of FDA-regulatory oversight and the expenses associated with this new and ever-developing technology. Cumulatively, the benefits of 3D printed anatomical models and surgical guides are theorized to include improved surgical accuracy, decreased operative time, and fewer complications. Additionally, in vitro and in vivo studies of patient-specific implants have established their potential to restore critical-sized defects in both load-bearing and nonload-bearing regions of the craniofacial region. However, surgeon familiarity, imaging quality, material, production equipment, and patient anatomy are all confounding variables that may hinder 3DMDs’ clinical performance. This section serves as a discussion of the most recent and representative clinical cases in craniomaxillofacial surgery that have utilized 3D printed models, guides, and implants produced by additive manufacturing processes.

7.1. Anatomical Models.

Anatomical models allow surgeons to tangibly visualize patient anatomy, developing a clearer, more efficient surgical plan for complex cranio-maxillofacial reconstructions. Anatomical model development is the product of a process known as virtual surgical planning (VSP). A digital simulation of the surgical intervention is generated from patient high resolution CT data collaboratively among the CMF reconstructive team and biomedical engineer. A study that 3D printed an auricular model for a microtia case highlights the dramatic difference that can exist between material cost alone and labor fees associated with model design (i.e., $1 vs $500, respectively).¹⁵⁰ This data is then used to print the patient-specific model in addition to surgical guides or implants when warranted. The broad scope of anatomical model applications in oncologic, traumatic, and congenital CMF reconstruction is described in Table 3. Within this subset of representative cases, SLA^187–192 and FDM^{150,190,193,194} were the most commonly employed AM techniques. Models were studied both preoperatively and intraoperatively for surgery simulation and resident training, patient education, situating hardware and/or implant placement, and bending off-the-shelf fixation devices to anatomical contours. Reported benefits broadly included decreased operative time, decreased need for surgical estimations and thus higher accuracy, and improved patient education and modulation of expectations.

Table 3.

Summary of Recent CMF Applications of Patient-Specific 3D Printed Anatomical Models

procedure name	sample size	patient age	material used	software used	printer used	sterilization	manufacturer	time to manufacture	cost per model	biomedical engineer used	result	complications	ref
Mandibular Oncologic Reconstruction	n = 5	18–63 yrs	a	Materialise	a (SLA)	a	a	a	a	Medical Modeling Inc.	Excellent bone-to-bone approximation, decreased intraoperative time and measurements, symmetric mandibular contour	None	188
Maxillary Oncologic Reconstruction	n = 16	26–85 yrs	Acrylic resin	SurgiCase CMF	a (SLA)	a	a	a	a	Medical Modeling Inc.	Prebent reconstructive locking plate, orthogonal jaw relationships	None	187
Mandibular Reconstruction (osteomyelitis)	n = 1	51 yrs	PLA, Formlabs white resin	3DSlicer, Autodesk MeshMixer	Form 2 Formlabs (SLA)	a	a	35 h	a	a	Prebent plates, decreased intraoperative time, improved accuracy of surgery	None	192
Midface Distraction (hemifacial microsomia syndrome)	a	a	a	Mimics, Materialise, 3matic	3D System Zprinter 650 (MultiJet)	a	a	7.4 h (complete skulls), 6.0 h (maxillofacial)	$61 (complete skulls), $40 (maxillofacial)	a	Subjective decreased intraoperative time, prebent plates, more thorough patient consultation	None	189
Face Vascularized Composite Allotransplantation	n = 1	21 yrs	Multimaterial	TeraRecon, Magics	Stratasys J750 (PolyJet)	a	a	28.2 h, 13.8 h	$762 (recipient), $695 (donor)	Cleveland Clinic Biomedical Engineering Department	Furthered understanding of spatial anatomy, improved education, but time-consuming	None	214
Craniofacial Reconstruction	n = 2	10–34 yrs	PLA	Rapid3D, De-Vide	MakerBot Replicator (FDM)	Hydrogen peroxide and gas plasma (Sterrad)	a	14 h	$25	a	Improved surgical accuracy	None	190
Cleft Lip and Palate Reconstruction	n = 30	a	ABS (bone) + TangoPlus (soft tissue)	Mimics	Fortus 3D Printer + Connex 3D Printer (PolyJet)	a	a	a	a	a	Improved patient education and satisfaction	None	215
Frontal Craniotomy and Bilateral Orbi-totomyb	n = 1	1.5 yrs	Multimaterial	Mimics	a (PolyJet)	a	a	38 h	a	a	Preoperative surgical practice and better impression of anatomic characteristics	None	216
Auricular Reconstruction (microtia)	n = 1	a	PLA	Amira + Blender,	Builder Premium 3D Printer (FDM)	Autoclave	a	4 h (Print only)	$1.10 (material), $501 (material and labor)	a	Greater repetitions to practice carving auricular framework, reduced estimation when shaping the ear	None	150
Craniosynostosis Repair	n = 2	0.5–2 yrs	Clear resin	Mimics	a (SLA)	Standard per institutional protocol	a	a	a	a	Merging of patient anatomy with normative anatomy and Intraoperative contouring of the calvaria on anatomic model	None	191
Mandibular Distraction Osteogenesis	n = 5	7–10 yrs	PLA	3DSlicer, Cura Software	a (FDM)	a	a	2 h	$25	Dremel 3D Idea Builder	Prebending of osteosynthesis material, improved learning for trainees	None	194
Rhinoplasty	n = 12	16–59 yrs	PLA	Blender	Builder Premium 3D Printer (FDM)	Autoclave	a	44 h	$5	a	Improved patient education and expectation setting, good intraoperative comparison to visualize desired outcome	None	193
Orbital Osteotomyc	n = 1	11 yrs	a	a	HP DesignJet 3D	a	a	a	a	a	Osteosynthesis material was precontoured on the 3D model	None	217

^aNot applicable.

^bFrontoethmoidal meningoencephalocele.

^cOrbital hypertelorism.

7.2. Surgical Guides.

The primary function of 3D printed surgical guides is to serve as templates to facilitate safe and accurate bone cutting or to properly position the device placement. Table 4 is a representative summary of the various recent applications of surgical guides in CMF surgery. Although traditionally thought to be used strictly intraoperatively, surgical guides were also used preoperatively to perform mock surgeries on patient-specific anatomical models.¹⁹⁵ Very few identified studies disclosed the AM technique, but of those that did, SLA and FDM were the most common. Material of choice were various polymers, and in only one study was a titanium surgical guide made.³¹ In this case the titanium surgical guide was preferred, although much more expensive, to avoid risk of polymer deterioration and to guarantee guide stability and rigidity.³¹ Other parameters including cost, time to manufacture, and sterilization method were scarcely mentioned in the existing literature. Overarching benefits included decreased operative time, prevention of injury to critical nearby structures, accurate and symmetrical cutting and implant placement, and overall improved patient outcomes.¹⁹⁶ One caution that was raised in the case of oncologic reconstruction was the risk of tumor growth between the time of VSP and the day of surgery. In this case, it was encouraged that cutting guides be developed with a slight overestimation of defect size.¹⁹⁷

Table 4.

Summary of Recent CMF Applications of Patient-Specific 3D Printed Surgical Guide

procedure name	sample size	patient age	material used	software used	printer used	sterilization	manufacturer	time to manufacture	cost	biomedical engineer used	result	complications	ref
Craniosynostosis Repair	n = 4, n = 5	0.75–6yrs, 0.66–1.33yrs	a, Polyamide	Materialise, 3D Slicer + Free-form Plus + CranioNav	a (SLA, SLS)	a, Autoclave	a	a	a	Medical Modeling Inc,a	Decreased intraoperative time and reduced irregularities from preoperative osteotomy design and bone shaping, improved accuracy via repeated guide placement registration	None, None	218, 219
Facial Transplantation	n = 2	a	Appropriate medically approved plastic	Planmeca Romexis, 3D-Systems Geomagic Freeform	a	a	a	8 h	a	a	Decreased surgical time and simplified bony reconstruction	None	220
Nasal Prosthesis Insertion	n = 1	58yrs	ABS	NobelGuide, Amira 3.1.1, Rapidform XOS2, Rhino 3.0	Stratasys (FDM)	a	a	a	a	a	Accurate implant positioning and easier accessibility for hygienic maintenance	None	151
Neonatal Mandibular Distraction Osteogenesis	n = 1	0yrs	a	Medical Modeling	a	a	a	a	a	a	Proper fit of cutting guides led successful placement of device	None	221
Maxillofacial Reconstruction	n = 9	a	a	a	a	a	Stryker, Synthes	a	a	Materialise, KLS Martin	Higher proportion of bone segment contact and successful reconstruction of more subunits	None	222
Mandibular Reconstruction	n = 38	~51.0yrs	Polymer	SurgiCase SMF, Materialise	a (SLA)	a	a	a	a	Medical Modeling Inc.	Reduced operative time and improved reconstructive accuracy	None	197
Forehead Contouring	a	a	a	ProPlan CMF 3.0, 3-matic	a	a	a	a	a	Materialise	Mock surgery with guides and model allowed for optimal surgical planning	None	223
Soft Tissue Correction via Fat Grafting	n = 6	a	Biocompatible photopolymerizing resin	Mimics, Zbrush	Form 2 (SLA)	a	a	a	a	a	Quantifiable volume and location for facial filling leading to better patient outcome	None	196
Cranial Vault Reconstruction	n = 3	28–41yrs	a	Anatomage In-Vivo, Rhinoceros	a (SLA)	a	a	a	a	a	Reduced intraoperative time	None	28
Mandibular Step Osteotomy	n = 1	21yrs	Titanium	Mimics, Simplant Pro	a	a	a	a	a	a	Prevention of injury to adjacent structures, reduced operative time, and improved surgical outcomes	None	31

^aNot applicable.

7.3. Implants.

To date, 3D printed implants have been utilized largely in cranioplasties and nasal, orbital, maxillary, and mandibular reconstruction (Table 5). Titanium and thermoplastic polymers, predominantly PEEK, were the materials of choice for orbital, nasal, maxillary, and mandibular reconstruction.^{114,198–208} In addition to these materials, the use of 3D printed bioceramics has also been attempted in cranioplasties.^209–213 Of note, many of these studies were performed outside of the United States and largely in adults with few exceptions of pediatric patients. Implant cost, time to manufacture, and sterilization methods were rarely reported, despite being some of the most widely discussed limitations of 3DMD. Commonly reported benefits mentioned were physician and patient aesthetic satisfaction, improved implant fit to anatomical contours, and reduced operative time. However, 3 of the 17 identified studies reported development of an implant with an improper fit due to errors with the initial CT scan or conversion to CAD.^200–202 It must be emphasized that each transition step within the 3D printing process introduces an opportunity for inaccuracy. Imaging quality of CT- and MRI-data have been reported to modulate the mean absolute error between virtual 3D model and final printed product.³⁵ Furthermore, the use of CAD software to develop a 3D printer compatible file has also been shown to compromise the integrity of the original scan.³⁵ As a part of their quality system regulations, FDA-approved software platforms verify preservation of original scan data.³⁵ It is imperative that HCFs purchase FDA-cleared software and compatible printers when planning to establish their own POC 3D printing facility. Future clinical studies should aim to report as much information as possible regarding the various stages of their unique 3D printing workflow to facilitate reproducibility and build a foundational clinical database.

Table 5.

Summary of Recent CMF Applications of Patient-Specific 3D Printed Implants

procedure name	sample size		patient age		material used	software used	printer used		sterilization	implant manufacturer	time to manufacture	cost	postoperative follow-up	result	complications	ref
Nasal Septal Deformity Reconstruction	n = 14		~38 ± 11 yrs		PCL (pore size: 500 μm)	a	a		Gamma Irradiation	a	a	a	T = 3 mo and ~4 yr	Long-term clinical efficacy and safety, maintenance of mechanical strength	None	198
Primary Reconstruction of Fronto-orbital Region	n = 1		13 yrs		PEEK	a	a		a	a	a	a	T = 2 mo	Some left-sided temporal hollowing, overall optimal aesthetic and safe results in fronto-orbital reconstruction	None	199
Orbital Wall or Maxilla-orbital Reconstruction	n = 26 3D: n = 12 Stock: n = 14		~42 yrs		3D: Ti6Al4V ELI Stock: Ti6Al4V ELI	Pro/ENGINEER (PTC)	Arcam (EBM)		a	Planmeca	a	a	T = 0 mo	Average decrease of operation time in 3D vs Stock Orbital: ~ 0.4h maxilla-orbital: ~ 3.82h	Implant incorrect shape due to error in CAD (n = 2)	200
Subperiosteal Implants for Atrophic Mandible	n = 10		~70 yrs		Ti6Al4V ELI	Mimics (Materialise), Studio 2012 (Geomagics), Meshmixer (Autodesk)	ProXDMP100, 3D System (DMLS)		a	IUXTA-3D (BTK)	2–3 weeks	a	T = 12 mo	100% survival rate, surgeon-determined good to excellent adaptation and fit	Implant incorrect shape due to scattering from nearby crowns in CT (n = 2)	201
Cranioplasty	n = 21		~28.6 yrs		Ti6Al4V ELI	3-Matic, Magics (Materialise)	Arcam A1 (EBM)		a	a	a	a	T = ~14.1 mo	Patients satisfied, good fixation of implants, and satisfactory symmetry	Infection + implant removal (n = 1)	209
Cranioplasty	n = 38		~43 ± 18 yrs		PEEK	Maxilim (Medicim NV), 3ds Max 2012 (Autodesk)	a		a	a	a	a	T = 6 mo	Comparable complication rates to literature on auto- or allograft cranioplasties. In 3 pts with infection, implant was successfully removed and reused after sterilization	Infection + removal (n = 5), hematoma (n = 4), CSF leak (n = 1). wound-re-lated (n = 1)	210
Cranioplasty	n = 6		~46 ± 21 yrs		PEEK	a	a		a	DePuy Synthes	~2 wks	$8493 ± $837.98	T = 4 yr	Good surgical outcomes with high patient satisfaction	None	211
Cranioplasty	n = 11		~46 yrs		PEEK	a	a		a	DePuy Synthes	a	a	T = ~6 mo	Satisfactory aesthetic results overall	None	212
Cranioplasty	n = 8		~44 yrs		Hydroxyapatite (pore size: 300 and 550 um)	a	a (SLA)		Gamma irradiation	3DCeram	2–3 weeks	$10,640 to 13,300* (adjusted for inflation)	T = 1, 6, and 12 mo	Satisfactory aesthetic results overall	None	213
Maxillofacial Reconstruction	n = 24 mandibular: n = 14 zygomatic: n = 3 orbital: n = 7		~30.8 yrs		PEEK	a	a		a	DePuy Synthes + Planmeca	a	a	T = ~16.2 mo	Fit to surrounding bone was perfect in 22 cases, yet outer contour of implant modified in 9 cases prior to fixation. Infection rate of 8.3% comparable with 3D printed PEEK studies	Wound dehiscence (n = 2)	114
Maxillary Reconstruction	n = 32		28.63 yrs		Ti6Al4V ELI	a	a		a	Planmeca	a	a	T = 0 mo	Precise fitting of all implants except one	Implant incorrect shape due to error in mandible position during CT (n = 1)	202
Maxillary or Mandibular Reconstruction		n = 33		3D: 56 ± 14 yrs; Stock: 55 ± 16 yrs	Grade 2 titanium	ProPlan CMF 2.0, 3-Matic 13.0 (Materialise)	a (SLM)	a		a	a	a	T = 0 mo	3D: superior accuracy of reconstructive outcomes	3D: None	203
		3D: n = 17 Stock: n = 16												No difference in total operative time between 3D and stock groups	Stock: Exposure + removal (n = 2)
Isolated Mandibular Advancement ± LeFort I Maxillary Osteotomy		n = 60; mandibular: n = 30; mandibular + maxillary: w = 30		~27 yrs	Grade 2 titanium	ProPlan CMF (Materialise)	a (SLS)	a		Materialise	a	a	T = 1 mo	No difference between planned and actual position of the mandible in either group when considering absolute values of the differences	a	204
Maxillary ± Mandibular Osteotomies		n = 43		~23 yrs	3D: Titanium	a	a	a		a	a	a	T = ~9.8 mo	Reduced operative time in 3D maxillary/stock mandibular plates subgroup vs 3D maxillary/3D mandibular and stock maxillary/stock mandibular plate groups without significant difference in complications	3D: Infection (n = 3) + removal (n = 2), palatal fistula + repair (n = 1)	205
		3D: n = 18 Stock: n = 15			Stock: a (Stryker)										Stock: Infection (n = 2) + removal (n = 1), exposure + removal (n = 1), occlusal relapse irevision (n = 2)
Mandibular Anterior Subapical Osteotomy (MASCO)		n = 11		~27 yrs	Ti6Al4V ELI	ProPlan CMF 3.0 (Materialise)	a	a		Materialise	a	a	T = 0 mo	High degree of accuracy between the virtual plan and immediate postoperative result	None	206
Mandibular Contour Reconstruction		n = 2		~37 yrs	Ti6Al4V ELI (pore size: 750–950 μm)	Meshmixer Ver. 3.5 (Autodesk), Magics (Materialise)	MetalSys250 (SLM)	a		a	a	a	T = 6 mo	Patients satisfied, surgical accuracy and postoperative stability without positional change of implant	None	207
Mandibular Contour Reconstruction		n = 21		a	3D: PEEK	3D-Surgery (a)	a	a		a	a	a	T = 0 mo	Reduced operative time and greater facial symmetry in 3D group	3D: Infection (n = 2)	208
		3D: n = 9 Stock: n = 12			Stock: porous PE (MED P OR, Stryker)										Stock: Infection (n = l)

^aNot applicable.

8. FUTURE DIRECTIONS

8.1. Automatization of 3D Printed Implant Design.

With the focus of HCFs and the U.S. FDA alike on the development of hospital-operated 3DP facilities, researchers have begun to investigate the automatization of the user-dependent aspects of the 3DP workflow. More specifically, the challenge of accurate 3DMD formulation lies within the conversion from original scan data to DICOM, and finally to a 3D printable file. This process is both timely and technically complex, posing a challenge to inexperienced users. To combat this issue, Li et al. developed a database of CT-based, healthy cranial scans injected with randomly generated artificial defects.²²⁴ The goal is to train deep learning algorithms, by way of data sets like these, to perform automatic defect reconstruction and implant generation.²²⁴ This would save time, reduce labor expenses, and improve implant design accuracy by eliminating risk of human error, which is a frequently reported cause of improper fit of 3D printed implants.^200–202 The deep learning algorithm can be refined to generate implant designs for real CMF defects more accurately if other researchers contribute their own scan data to the database.²²⁴

8.2. 3DMD + Augmented Reality.

With the latest advancements in augmented reality (AR) technology, commercial AR platforms have become available for surgical use. AR allows the user to virtually visualize an overlay of critical patient anatomy (e.g., arterial supply, innervation, musculature) that may not be visible in reality.²²⁵ The virtual, holographic projections are generated from patient CT/MRI data and registered in that specific surgical field. AR can be combined with 3D printing technology for resident training, preoperative planning, or intraoperatively to accurately position surgical guides and subsequent fixation devices. Thus, far, this technology has been successfully applied in craniosynostosis, face transplantation, mandibular and auricular reconstruction, as well as for patient-specific facial artery mapping.^{214,225–229} With adequate training, reported outcomes have included improved accuracy, minimization of operative injury, reduced costs, and shortened intraoperative time.²²⁵

8.3. In Situ Bioprinting.

In situ bioprinting is a rapidly advancing form of 3D bioprinting in which biomaterials/bioinks are used print acellular/cell-laden constructs, respectively, directly into the desired defect site.²³⁰ Subtypes include bedside mounted (e.g., traditional in situ bioprinter, robotic-arm assisted bioprinter) and hand-held (e.g., ionically activated, light activated) printers; both of which can be used in a sterile surgical suite. In situ bioprinting research for CMF repair remains in its preclinical stages, with critical-sized, murine calvaria defects.^231–233 For example, Kérourédan et al. printed endothelial cell-based microvascular networks in defects prefilled with collagen, MSCs, and VEGF.²³¹ This experimental group exhibited improved vascularization and bone regeneration compared with scaffolds randomly seeded with endothelial cells and negative controls. Relative to ex vivo bioprinting (i.e., premade constructs), clinical application of in situ bioprinting would eliminate many challenges associated with the production of 3DMD with biologic components. More specifically, printing directly onto the patient would minimize contamination risk and implantation delays associated with ex vivo manufacturing and transportation.²³¹

8.4. 4D Printing.

4D printing is defined by the ability of 3D printed structures to temporarily transform and reform to an original shape when cued by an external stimulus. Stimuli can include but are not limited to temperature, water or solvents, magnetism, and ultraviolet light.²³⁴ Shape transformations such as folding, curling, twisting, and linear expansion have all been achieved. This technology has the potential to optimize bone defect repair with 3D printed scaffolds. For example, Senatov et al. 3D printed a PLA/15 wt % HA porous scaffold by fused filament fabrication with heat-sensitive shape memory.²³⁵ Clinically, scaffolds can be surgically inserted in their compact, deformed shape and upon shape-memory activation, expand to precisely fit the bone defect.^235,236 As it stands, a limitation of 4D printing is identifying materials that are responsive to stimuli that are benign to humans. For example, heat-sensitive shape memory is only feasible if it is within the range of human body temperature.²³⁴ Soybean oil epoxidized acrylate fabricated scaffolds have recently been discovered to be capable of such transformations and have even demonstrated promising osteoconductive potential.²³⁴ This technology has also been harnessed to create 3DMD capable of shape change with tissue growth, with enormous implications for pediatric cranio-maxillofacial bone repair as the scaffold could expand with skeletal development.²³⁷

9. CONCLUSION

3D printing technology within the field of craniomaxillofacial surgery has grown exponentially in recent decades. Possible applications have proven to touch every aspect of the field with an overarching ability to improve functional and aesthetic outcomes as well as reduce the operative time and postoperative complications. Researchers continue to push the envelope with the capabilities of 3DMD, including shape memory devices. This has the potential to disrupt current standard-of-care for pediatric craniofacial patients. Craniofacial surgeons excitedly introduce anatomic models, surgical guides, and implantable devices into their operating rooms, proving or disproving their benefits on a case-by-case basis and furthering the research as a result. Meanwhile, the FDA is working diligently with stakeholders to discuss feasible ways to provide this care safely and accessibly to patients at the point-of-care. As discussed, 3DMD has the potential to become the new standard within CMF surgery. In order to prepare for 3DMD finding its eventual commonplace, it is imperative that this technology continue to be better understood, discussed, and refined for all of its applications within craniomaxillofacial surgery.