TL;DR

Key Facts

What Are Dental Metals and Ceramics?

Dental metals and ceramics are the primary structural materials used in indirect restorations — crowns, bridges, inlays, onlays, veneers, partial denture frameworks, and implant components. Metals are crystalline materials with free electrons in their structure, giving them characteristic properties: high electrical and thermal conductivity, metallic lustre, ductility (ability to deform before fracturing), and high strength in both tension and compression. Ceramics are inorganic, non-metallic materials (oxides, silicates, and mixed oxide systems) with ionic or covalent bonding, characterized by high hardness, brittleness, chemical inertness, low thermal and electrical conductivity, and aesthetically tooth-like optical properties.

The choice between metal, ceramic, or their combination (PFM) is determined by the mechanical demands of the restoration (occlusal load, span length, clinical crown height), the aesthetic requirement, the preparation reduction achievable without compromising pulp viability, and patient factors (bruxism, allergies, cost considerations). Understanding the fundamental properties of each material class allows the clinician to select rationally rather than by habit or assumption.

Why It Matters

Dental materials science is extensively tested on board examinations. High-yield topics include: gold alloy types (I–IV) and their clinical uses; ADA noble metal classification; the PFM bonding mechanism and CTE matching; ceramic types and their strength ranking; transformation toughening in zirconia; HF acid etching of ceramics for resin bonding; amalgam alloy composition and setting; and the consequences of CTE mismatch in PFM restorations. Understanding materials also underpins material selection in clinical practice — using the wrong material for a clinical situation is a source of preventable failures.

Key Mechanical Properties

Dental Metals and Alloys

Gold Alloys — ADA Types I–IV

Gold alloys are classified by the ADA based on hardness and composition into four types. The principal alloying elements in dental gold alloys are: gold (Au — corrosion resistance, biocompatibility); silver (Ag — increases hardness, reduces cost); copper (Cu — main hardening element — age-hardening by Cu₃Au precipitation); palladium (Pd — increases melting range, improves corrosion resistance, whitens colour); platinum (Pt — increases hardness and melting point, whitens colour); zinc (Zn — deoxidiser during casting).

Base Metal Alloys

Base metal alloys contain <25% noble metals and primarily use nickel (Ni), chromium (Cr), cobalt (Co), or molybdenum (Mo) as principal components. Key properties: higher strength than noble alloys (UTS often >700 MPa); lighter weight (important for RPD frameworks); stiffer (higher modulus); lower cost. Used for: RPD frameworks (Co-Cr alloy — “vitallium”; stiff, strong, thin clasp arms possible), full crown frameworks (Ni-Cr — now declining in favour of zirconia), implant components.

Nickel allergy: Ni-Cr alloys carry a significant risk in nickel-sensitive patients (~10–15% of women, 1–3% of men). Nickel ions released by corrosion can cause intra-oral mucosal reactions, and systemic sensitisation. Cobalt-chromium alloys have a better allergy profile than Ni-Cr. For patients with known or suspected nickel allergy: use Co-Cr, high-noble gold, palladium-silver, or titanium alloys; or all-ceramic alternatives.

Titanium and Titanium Alloys

Commercially pure titanium (cp-Ti, grades 1–4) and titanium-6 aluminium-4 vanadium alloy (Ti-6Al-4V, grade 5/ASTM F136) are the dominant implant materials in dentistry — all Brånemark-type implants and the vast majority of modern implants use cp-Ti or Ti-6Al-4V. Titanium’s unique combination of properties for implants: (a) osseointegration — titanium dioxide (TiO₂) layer forms spontaneously on the surface; this oxide layer is chemically compatible with bone and promotes direct bone-to-implant contact; (b) corrosion resistance — the TiO₂ passivation layer prevents further corrosion; (c) biocompatibility — minimal ion release; no allergic reactions; (d) mechanical properties — sufficient strength for implant function; lighter than gold or Co-Cr alloys. Disadvantages: high casting temperature and reactivity with oxygen (special vacuum casting equipment required); galvanic corrosion risk if in contact with dissimilar metals.

Dental Amalgam

Dental amalgam (a mercury-silver alloy) remains an important material despite controversy. Composition: Mercury (Hg, ~50% by weight) + alloy powder (lathe-cut particles or spherical particles, or admixed both). Alloy composition: silver (Ag, ~65%) — primary component, reacts with Hg to form γ₁ (Ag₂Hg₃) and γ₂ (Sn₈Hg) phases; tin (Sn, ~29%) — reacts with Hg; copper (Cu) — in high-copper alloys, reacts with Sn to form Cu₆Sn₅, preventing formation of the weak, corrosion-prone γ₂ phase; zinc (Zn, minor) — deoxidiser during manufacturing.

High-copper vs. low-copper amalgam: High-copper alloys (>6% Cu) have largely replaced low-copper alloys. The key improvement: high copper eliminates the γ₂ phase (tin-mercury) which was the main cause of amalgam corrosion, creep, and marginal breakdown. High-copper amalgams have superior strength, lower creep, and better corrosion resistance. Setting reactions: Ag₃Sn (γ phase, silver-tin particles) + Hg → γ₁ (Ag₂Hg₃) + γ₂ (Sn₈Hg) + unreacted Ag₃Sn. In high-copper: Cu reacts with Sn to form Cu₆Sn₅ (η phase) instead of γ₂.

Dental Ceramics

Feldspathic Porcelain

Feldspathic porcelain is a glassy ceramic based on potassium feldspar (KAlSi₃O₈) with leucite crystals dispersed in a glass matrix for limited reinforcement. It is the traditional dental porcelain and the basis from which all other dental ceramics evolved. Properties: excellent aesthetics (high translucency, natural light transmission, shade-matching capability); flexural strength ~60–100 MPa (weak — fractures under posterior occlusal loads without a supporting metal framework); thermally fired in layers (opaque, body, incisal, glaze). Primarily used: as the veneering material over PFM frameworks; for anterior veneers; and for characterisation of restorations. Not appropriate as a standalone crown material in the posterior region (insufficient strength).

Leucite-Reinforced Ceramic (IPS Empress)

Leucite-reinforced ceramic (IPS Empress, Ivoclar) introduced heat-pressing technology — a wax pattern of the restoration is invested and a leucite-glass ceramic is heat-pressed into the mould under vacuum and pressure, producing a denser, more uniform ceramic with fewer pores than powder-slurry built-up feldspathic porcelain. Leucite crystals (KAlSi₂O₆) dispersed in a glass matrix deflect crack propagation, improving strength to ~120–180 MPa. Etched with 9.5% HF acid for resin bonding. Primarily used for: anterior crowns; inlays/onlays; and veneers where strength demands are moderate. Not appropriate for posterior long-span bridges or in bruxists.

Lithium Disilicate (IPS e.max)

Lithium disilicate glass-ceramic (IPS e.max, Ivoclar) is the result of controlled crystallisation of a lithium-containing glass to produce interlocking lithium disilicate crystals (Li₂Si₂O₅) with an aspect ratio of ~5:1 within a glass matrix. The interlocking crystal morphology deflects crack propagation around the crystals rather than through them — producing flexural strength of ~350–500 MPa, approximately 4–5× stronger than leucite ceramics. Available as: heat-pressed ingots (IPS e.max Press — in different opacity grades: LT, MT, HT, MO; pressed at ~920°C); and pre-crystallised blocks for CAD/CAM milling (IPS e.max CAD — milled in a partially crystallised blue-phase state for ease of milling, then crystallised in a furnace to full strength at ~840°C). Bonded with resin cement after 4.9% HF acid etching (60 seconds) + silanisation. The HF concentration is lower (4.9%) for lithium disilicate than for leucite ceramics (9.5%) because the disilicate crystals are more susceptible to HF etching than leucite crystals.

Zirconia (3Y-TZP and 5Y-TZP)

Yttria-stabilised tetragonal zirconia polycrystal (Y-TZP) is the strongest dental ceramic. Zirconia exists in three crystal phases: monoclinic (stable room temperature), tetragonal (stable 1170–2370°C), and cubic (stable >2370°C). The addition of yttria (Y₂O₃) stabilises the tetragonal phase at room temperature. For 3Y-TZP (3 mol% yttria): predominantly tetragonal → highest fracture toughness via transformation toughening → flexural strength 900–1200 MPa; but opacity limits aesthetic use. For 5Y-TZP (5 mol% yttria): more cubic phase stabilised → higher translucency → strength 700–800 MPa (transformation toughening reduced due to less tetragonal phase available for transformation).

Ageing (low-temperature degradation, LTD): 3Y-TZP can undergo spontaneous tetragonal → monoclinic transformation in the presence of water at low temperatures (~150–400°C, or clinically at 37°C over years). This surface phase transformation causes microcracking and surface roughening, reducing strength and increasing wear of opposing dentition. Modern 3Y-TZP formulations with controlled grain size and density have significantly reduced ageing susceptibility. High-translucency 5Y-TZP is more resistant to ageing (more cubic phase; less tetragonal available to transform).

Bonding to zirconia: Standard HF acid etching does NOT work on zirconia (it has no glass phase to dissolve). Zirconia bonding requires: (a) alumina sandblasting (50 µm Al₂O₃ at 0.1–0.25 MPa) to create micromechanical retention; (b) MDP (10-methacryloyloxydecyl dihydrogen phosphate) primer — chemical adhesion through phosphate groups that bond to Zr-OH groups on the zirconia surface; products such as Panavia F, MDP-containing primers (Clearfil Ceramic Primer Plus, Z-Prime Plus) are used. The combination of sandblasting + MDP primer + resin cement is the evidence-based bonding protocol for zirconia.

Alumina and In-Ceram

Alumina (Al₂O₃)-based ceramics were the first high-strength dental ceramics developed as an alternative to PFM. In-Ceram Alumina (Vita) uses a slip-casting technique: alumina powder slurry is applied over a porous refractory die and sintered (partial sintering creates interconnected porosity); then glass (lanthanum-borosilicate) is infiltrated into the porous alumina framework under heat — the glass fills the pores, producing a dense, glass-infiltrated alumina with flexural strength ~400–600 MPa. In-Ceram Zirconia (alumina + 30% zirconia particles) achieves ~700 MPa. Now largely superseded by CAD/CAM zirconia (which achieves higher strength more consistently with simpler processing) for framework fabrication, though In-Ceram materials remain in use in some laboratories.

Porcelain-Fused-to-Metal (PFM) System

PFM Bonding Mechanisms

The bond between the metal coping and the veneering porcelain in PFM restorations depends on three mechanisms acting simultaneously:

1. Mechanical interlocking: The metal surface is sandblasted (50 µm Al₂O₃) and the framework surface is designed with beads, loops, or surface texture to create macromechanical retention. The liquid porcelain wets the metal surface and flows into micro-irregularities, creating interlocking after solidification.
2. Chemical bonding (metal oxide–glass bond): After sandblasting, the metal framework is oxidised in a degassing furnace at high temperature. Metal oxides form on the surface (for gold alloys: tin, indium oxides; for base metals: chromium oxide). The first firing of opaque porcelain at ~960°C allows these metal oxides to partially dissolve into the glass phase of the porcelain — the covalent/ionic Si-O-Metal bond creates a chemical union between the metal and the ceramic. This is the primary bonding mechanism.
3. Van der Waals forces: Minor contribution from close surface contact between the metal and ceramic at areas of intimate contact.

PFM Failure Modes

PFM restorations fail by one of three mechanisms: (a) porcelain fracture (chipping) — the most common failure mode (5-year chipping rate ~3–5%, increasing to ~10% at 10 years); typically occurs at areas of unsupported porcelain (inadequate metal framework support beneath the porcelain), at the porcelain-metal interface in areas where framework design created stress concentrations, or in bruxist patients with parafunctional heavy occlusal loads on porcelain; (b) porcelain delamination — separation of the porcelain from the metal as a layer, typically caused by CTE mismatch (metal CTE < porcelain CTE → porcelain in tension on cooling → delamination); inadequate oxidation of the metal surface (oxide layer too thin or too thick — optimal range is a thin adherent oxide); contamination of the metal surface before porcelain application (saliva, fingerprints, polishing compounds — all prevent oxide formation); (c) framework fracture — rare with adequate metal thickness; occurs in areas of insufficient metal cross-section at connector regions of bridges; related to inadequate design by the dentist (insufficient interproximal clearance for connector size) or laboratory (under-designed framework).

Ceramic Processing Methods

Coefficient of Thermal Expansion (CTE) in PFM Restorations

The CTE measures the fractional change in length per degree of temperature change (units: × 10⁻⁶/°C or ppm/°C). For PFM restorations, the relative CTE of the metal and porcelain is critical. The porcelain is fired at ~900–1100°C and cooled to room temperature. During cooling:

• If metal CTE = porcelain CTE: equal contraction during cooling → no residual thermal stresses → theoretically ideal but difficult to achieve in practice.
• If metal CTE > porcelain CTE (by ~0.5 × 10⁻⁶/°C): metal contracts more than porcelain → metal “squeezes” the porcelain → porcelain is placed in compression. Ceramics are ~10× stronger in compression than tension — this compressive prestress significantly strengthens the porcelain and compensates for the tensile stresses that arise during occlusal loading. This is the deliberately designed state for all PFM restorations.
• If metal CTE < porcelain CTE: metal contracts less than porcelain → porcelain contracts more than the metal restrains → porcelain is placed in tension. Ceramics are very weak in tension → porcelain cracks, crazes, or delamination occurs during cooling from the kiln. This is a catastrophic mismatch scenario. The usual cause: using the wrong porcelain (intended for a different alloy system) or a custom alloy without proper CTE characterisation.

Corrosion and Tarnish

Corrosion is the deterioration of metals by electrochemical reaction with the oral environment. In dentistry, the relevant corrosion reactions involve oxidation of base metals and formation of ionic products. Tarnish (surface discolouration without significant material loss — from sulphide formation, oxide layer formation) is distinct from true corrosion. Galvanic corrosion occurs when two dissimilar metals are in electrical contact in a conductive solution (saliva) — the more active metal in the galvanic series is corroded preferentially. Two amalgam restorations in opposing arches in contact produce galvanic corrosion; an amalgam restoration adjacent to a gold crown produces the same effect — the amalgam is the more active metal and corrodes at an accelerated rate. Noble metals (gold, platinum, palladium) are extremely corrosion-resistant. Base metals rely on a passive oxide layer (chromium oxide on Ni-Cr and Co-Cr alloys) for corrosion protection — disruption of this layer (mechanical damage, exposure to halides including fluoride in toothpaste) can cause pitting corrosion.

Clinical Considerations

• Ceramic surfaces must be re-polished or re-glazed after chairside adjustment: Grinding a ceramic surface with a rotary bur creates surface flaws (microcracks) that reduce the in-vivo strength of the ceramic by 30–50% compared with the as-manufactured polished surface. The standard protocol: adjust with a fine-grit diamond bur under water cooling (avoid dry grinding — heat causes surface damage); finish with progressively finer ceramic polishing discs and points; apply glaze (paste glaze fired in the porcelain furnace — adds a smooth glassy layer that blunts surface cracks) or use a high-shine ceramic polishing paste (acceptable substitute for single-unit restorations where re-firing is impractical). Never leave an adjusted ceramic surface unpolished.
• The zirconia-porcelain interface in veneered zirconia crowns is the site of the most common failure: Full-contour (monolithic) zirconia crowns have essentially eliminated the porcelain fracture problem in posterior teeth by eliminating the porcelain layer entirely — the zirconia is ground, polished, and stained to colour without a veneering porcelain layer. Veneered zirconia (zirconia coping + porcelain veneering layer) has a significantly higher chipping rate than monolithic zirconia or PFM (5-year chipping rate ~15–20% for veneered zirconia vs. ~3–5% for PFM) due to: (a) the zirconia–porcelain CTE mismatch (zirconia CTE ~10 × 10⁻⁶/°C; porcelain CTE ~8–9 × 10⁻⁶/°C — porcelain is placed in tension); (b) the high stiffness of the zirconia framework (which does not flex under load, concentrating stress at the zirconia-porcelain interface). Monolithic zirconia crowns are now preferred for posterior teeth; veneered zirconia is reserved for cases requiring maximum aesthetics (anterior teeth, specifically designed coping with adequate porcelain support).
• Fluoride-containing oral hygiene products may affect the surface of some ceramics: High-concentration fluoride gels (1.23% APF — acidulated phosphate fluoride) and high-fluoride toothpastes (5000 ppm fluoride — prescription-only) have been shown to roughen the surface of feldspathic and leucite-reinforced ceramics, increasing plaque retention and wear of opposing teeth. The effect is attributable to HF formed in the acidic (pH 3.5) APF gel selectively dissolving the glass phase of feldspathic ceramics. Clinical recommendation: patients with ceramic restorations should avoid prolonged high-concentration acidic fluoride gel use; neutral sodium fluoride (NaF) products are preferred. Zirconia and lithium disilicate are more resistant to HF dissolution (zirconia has no glass phase; lithium disilicate’s glass phase is less susceptible than feldspathic at dilute HF concentrations).
• The galvanic series determines which metal is sacrificed when two metals are in contact: The galvanic series in seawater (which approximates the ionic environment of saliva) ranks metals from most noble (least reactive — cathodic — corroded least) to least noble (most reactive — anodic — corroded most). In dental practice: gold > platinum > palladium > titanium (passive) > Co-Cr (passive) > Ni-Cr (passive) > amalgam > zinc. When amalgam contacts a gold restoration in the same occlusion, the amalgam is the anode and corrodes. The clinical consequence is usually not clinically significant for a brief contact — but if two restorations are in prolonged proximal contact or direct occlusal contact and both are metal, galvanic effects must be considered in material selection.
• Age-hardening (precipitation hardening) in gold alloys can increase their hardness and strength after fabrication: Type III and IV gold alloys containing copper (Cu) can be age-hardened by a two-step heat treatment: (1) high-temperature annealing at ~700°C followed by quenching → all Cu dissolved in Au-Ag matrix (disordered supersaturated solid solution, soft); (2) low-temperature heating at ~300–350°C for 15–30 minutes → ordered Cu₃Au or AuCu precipitates form within the matrix → increased hardness (Vickers hardness may increase by 30–50%). The laboratory applies this heat treatment after casting to optimise clasp spring properties in RPD frameworks. Clinically, extended casting torch exposure can partially re-anneal (soften) the alloy — excessive heating during castings or finishing reduces hardness below the designed value.

Common Mistakes & Misconceptions

• Misconception: “High-noble means high-gold; predominantly base metal means no gold.”
  Correction: The ADA noble metal classification is based on total noble metal content. High-noble: ≥60% noble metals + ≥40% gold specifically. Noble: ≥25% noble metals — but this does not require gold to be the dominant noble metal; palladium-silver alloys (with little or no gold) can be “noble” if they contain ≥25% noble metals. Predominantly base metal: <25% noble metals — but this category can contain small amounts of gold (just insufficient to reach the 25% noble threshold). The classification tells you about corrosion resistance and biocompatibility broadly, not the specific composition.
• Misconception: “Zirconia can be etched with hydrofluoric acid for resin bonding, like other ceramics.”
  Correction: HF acid etching works by dissolving the glass phase of feldspathic-based ceramics (leucite, lithium disilicate) to create a micromechanically retentive surface. Zirconia has no glass phase — it is a polycrystalline ceramic without any glassy component. HF acid has essentially no effect on the zirconia surface (it does not create the required microroughness). Zirconia bonding requires: (1) alumina sandblasting for mechanical retention; (2) MDP-containing primer (or air-particle abrasion + Monobond Plus which contains MDP) for chemical adhesion to zirconia; (3) resin cement.
• Misconception: “The thicker the oxide layer on a PFM metal framework, the stronger the porcelain bond.”
  Correction: The optimal oxide layer for PFM bonding is thin (~50–100 nm) and adherent. An excessively thick oxide layer (from over-oxidation or repeated degassing cycles) is brittle and cohesively weak — the bond fails within the thick oxide layer (cohesive failure of the oxide) rather than at the metal-oxide or oxide-porcelain interface. The oxide layer must also be chemically compatible with the porcelain glass phase — which is why each porcelain must be used with its designated alloy family (porcelains are formulated to match specific metal oxide chemistry).
• Misconception: “Amalgam is now banned in dentistry.”
  Correction: While amalgam use has been restricted or phased down in several countries and patient groups (the Minamata Convention on Mercury led to restrictions in many jurisdictions; the European Union phase-down began in 2018 targeting most groups, with a general ban proposed for 2030), amalgam has not been universally banned and remains an approved material in many countries. It is particularly limited for primary teeth, pregnant and breastfeeding women, and children under 15 in EU regulations. In countries where it remains available, its high compressive strength, longevity, and cost-effectiveness make it still relevant — particularly in resource-limited settings. Students should know the current status in their jurisdiction rather than assuming a universal ban.
• Misconception: “Full-contour zirconia crowns are too hard and will excessively wear opposing teeth.”
  Correction: The clinical concern about zirconia wear of opposing teeth is material-state dependent. Rough, poorly polished, or unglazed zirconia surfaces (from chairside adjustment without subsequent polishing) cause significantly more opposing tooth wear than glazed or well-polished zirconia. Properly finished full-contour monolithic zirconia produces enamel wear comparable to or lower than some glass ceramics (lithium disilicate, feldspathic) when polished to an appropriate surface finish (≤0.2 µm Ra). The practical recommendation: full-contour zirconia must be polished to a mirror-like surface after any adjustments; rough adjustments should be finished with ceramic polishing systems before delivery.

Related Topics

References & Sources

1. Anusavice KJ, Shen C, Rawls HR (eds) (2013). Phillips’ Science of Dental Materials, 12th ed. Saunders/Elsevier. [Standard dental materials science reference — mechanical properties, alloy classification, ceramic types, PFM bonding, CTE]
2. Craig RG, Powers JM (eds) (2012). Restorative Dental Materials, 13th ed. Mosby/Elsevier. [Comprehensive dental materials text — metals, ceramics, amalgam, composites]
3. Christel P, Meunier A, Dorlot JM, et al. (1989). Biomechanical compatibility and design of ceramic implants for orthopedic surgery. In: Bioceramics: Material Characteristics versus In Vivo Behavior. [Foundation reference for transformation toughening in zirconia — tetragonal to monoclinic transformation mechanics]
4. Pjetursson BE, Sailer I, Makarov NA, et al. (2015). All-ceramic or metal-ceramic tooth-supported fixed dental prostheses (FDPs)? A systematic review of the survival and complication rates. Dental Materials, 31(6):603–623. [PFM vs. all-ceramic survival comparison; chipping rates by material]
5. Sailer I, Fehér A, Filser F, et al. (2007). Five-year clinical results of zirconia frameworks for posterior fixed partial dentures. International Journal of Prosthodontics, 20(4):383–388. [Zirconia FPD clinical outcomes — framework survival, chipping rates]
6. Lohbauer U, Belli R, Cune MS, et al. (2011). Lifetime prediction of CAD/CAM dental ceramics. Journal of the Mechanical Behavior of Biomedical Materials, 4(7):1660–1667. [Ceramic strength, fatigue, and lifetime prediction — lithium disilicate and feldspathic porcelain]
7. Donovan TE, Cho GC (1999). Diagnostic provisional restorations in restorative dentistry: the blueprint for success. Journal of the California Dental Association, 27(2):178–179. [Provisional materials in prosthodontic treatment planning]
8. Wataha JC (2002). Alloys for prosthodontic restorations. Journal of Prosthetic Dentistry, 87(4):351–363. [Comprehensive review of dental alloys — noble metal classification, alloy properties, biocompatibility, corrosion]

Summary

Dental metals and ceramics span a wide range of compositions and properties matched to specific clinical applications. Gold alloys (Types I–IV, progressively harder) are the most biocompatible and corrosion-resistant dental metals with exceptional long-term clinical longevity — full gold remains the benchmark for posterior crown survival, at the cost of aesthetics. Base metal alloys (Ni-Cr, Co-Cr) provide high strength at lower cost but carry nickel allergy risk and lower corrosion resistance. Titanium is the dominant implant metal due to its TiO₂ passivation layer enabling osseointegration. Ceramics provide aesthetics and chemical inertness but are brittle — failure is always tensile fracture from surface cracks; all ceramic adjustments must be re-polished. Zirconia (3Y-TZP) achieves the highest ceramic strength (900–1200 MPa) through transformation toughening (tetragonal → monoclinic phase change under stress → compressive zone around crack tip); HT-zirconia (5Y-TZP) sacrifices some strength for higher translucency. Lithium disilicate (e.max) is the gold standard for anterior all-ceramic crowns (~350–500 MPa; etched with 4.9% HF + silane for resin bonding). PFM bond requires: sandblasting + metal surface oxidation + opaque porcelain oxide-glass dissolution; CTE of metal must be slightly higher than porcelain CTE to place porcelain in compressive prestress during cooling. Monolithic zirconia crowns are preferred for posterior teeth (eliminates porcelain-zirconia interface chipping risk); veneered zirconia has ~15–20% 5-year chipping rate.

About the Author

Dr. Andries Smith

Founder, Dental Panda

Dr. Andries Smith founded Dental Panda in 2020. As an immigrant to the United States, he had to take the INBDE exam, even though he was practicing dentistry for over 10 years. This revealed an opportunity. Andries noticed that INBDE prep course companies were putting profit over students. With his expertise and experience in dentistry, he created free dental wiki resources for students and the general public to have access to.